Principles of Regenerative Medicine

Principles of Regenerative Medicine

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Principles of Regenerative Medicine Anthony Atala,

MD

W.B. Boyce Professor and Director, Wake Forest Institute for Regenerative Medicine, Wake Forest School of Medicine, Professor, Virginia Tech-WakeForest University School of Biomedical Engineering and Sciences Winston-Salem, North Carolina, USA

Robert Lanza,

MD

Advanced Cell Technology, Worcester, Massachusettes, USA

James A. Thomson,

PhD

Wisconsin Regional Primate Research Center, Department of Anatomy, Madison, Wisconsin, USA

and

Robert M. Nerem,

PhD

Georgia Institute of Technology, Atlanta, Georgia, USA

Editorial Board Keith H.S. Campbell, Neal First, John D. Gearhart, William A. Haseltine, Peter Johnson, Robert Langer, Michael Lysaght, Antonios G. Mikos, David J. Mooney, Buddy D. Ratner, Alan J. Russell, Shay Soker, Joseph P. Vacanti, Catherine M. Verfaillie, Ian Wilmut, James J. Yoo, Leonard I. Zon

AMSTERDAM • BOSTON • HEIDELBERG • LONDON • NEW YORK • OXFORD PARIS • SAN DIEGO • SAN FRANCISCO • SINGAPORE • SYDNEY • TOKYO Academic Press is an imprint of Elsevier

Academic Press is an imprint of Elsevier 30 Corporate Drive, Suite 400, Burlington, MA 01803, USA. First edition 2008 Copyright © 2008 Elsevier, Inc. No part of this publication may be reproduced, stored in a retrieval system or transmitted in any form or by any means electronic, mechanical, photocopying, recording or otherwise without the prior written permission of the publisher. Permissions may be sought directly from Elsevier’s Science & Technology Rights Department in Oxford, UK: phone (44) (0) 1865 843830; fax (44) (0) 1865 853333; email: [email protected]. Alternatively you can submit your request online by visiting the Elsevier web site at http://elsevier.com/locate/permissions, and selecting. Obtaining permission to use Elsevier material. Notice No responsibility is assumed by the publisher for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions or ideas contained in the material herein. Because of rapid advances in the medical sciences, in particular, independent verification of diagnoses and drug dosages should be made. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress ISBN: 978-0-12-369410-2 For information on all Academic Press publications visit our web site at books.elsevier.com Typeset by Charon Tec Ltd (A Macmillan Company), Chennai, India www.charontec.com. Printed and bound in Canada 08 09 10 11 10 9 8 7 6 5 4 3 2 1

I would like to dedicate this textbook to the joys of my life – my wife, Katherine, and my children, Christopher and Zachary –Anthony Atala

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Contents

Preface List of Contributors Part I 1.

Introduction to Regenerative Medicine

Current and Future Perspectives of Regenerative Medicine

xiii xv 1 2

Mark E. Furth and Anthony Atala 2.

Fundamentals of Cell-Based Therapies

16

Ross Tubo 3.

Stem Cell Research

28

T. Ahsan, A.M. Doyle, and R.M. Nerem

Part II 4.

Biologic and Molecular Basis of Regenerative Medicine

Molecular Organization of Cells

49 50

Jon D. Ahlstrom and Carol A. Erickson 5.

Cell–ECM Interactions in Repair and Regeneration

66

M. Petreaca and M. Martins-Green 6.

Developmental Mechanisms of Regeneration

100

David L. Stocum 7.

The Molecular Basis of Pluripotency in Principles of Regenerative Medicine

126

Ariel J. Levine and Ali H. Brivanlou 8.

How Do Cells Change Their Phenotype

136

Peter W. Andrews and Paul J. Gokhale 9.

Somatic Cloning and Epigenetic Reprogramming in Mammals

148

Heiner Niemann, Christine Wrenzycki, Wilfried A. Kues, Andrea Lucas-Hahn, and Joseph W. Carnwath 10.

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins

168

William Gavin, LiHow Chen, David Melican, Carol Ziomek, Yann Echelard, and Harry Meade

Part III 11.

Cells and Tissue Development

Genetic Approaches in Human Embryonic Stem Cells and Their Derivatives

189 190

Junfeng Ji, Bonan Zhong, and Mickie Bhatia 12.

Embryonic Stem Cells: Derivation and Properties

210

Junying Yu and James A. Thomson

vii

viii

CONTENTS

13.

Stem Cells Derived from Amniotic Fluid and Placenta

226

Paolo De Coppi, Shay Soker, and Anthony Atala 14.

Stem Cells Derived from Cord Blood

238

Julie G. Allickson 15.

Multipotent Adult Progenitor Cells

258

Catherine M. Verfaillie, Aernout Luttun, Karen Pauwelyn, Jeff Ross, Lepeng Zeng, Marta Serafini, Yuehua Jiang, and Fernando Ulloa Montoya 16.

Bone Marrow Stem Cells: Properties and Pluripotency

268

Munira Xaymardan, Massimo Cimini, Richard D. Weisel, and Ren-Ke Li 17.

Hematopoietic Stem Cell Properties, Markers, and Therapeutics

284

S.M. Chambers, William J. Lindblad, and M.A. Goodell 18.

Neural Stem Cells

300

Yang D. Teng, Filipe N.C. Santos, Peter M. Black, Deniz Konya, Kook In Park, Richard L. Sidman, and Evan Y. Snyder 19.

Mesenchymal Stem Cells

318

Zulma Gazit, Hadi Aslan, Yossi Gafni, Nadav Kimelman, Gadi Pelled, and Dan Gazit 20.

Hepatic Stem Cells: Lineage Biology and Pluripotency

344

N. Cheng, Hsin-lei Yao, and Lola M. Reid 21.

Skeletal Muscle Stem Cells

386

Jason H. Pomerantz and Helen M. Blau 22.

Islet Cell Therapy and Pancreatic Stem Cells

398

Juan Domínguez-Bendala, Antonello Pileggi, and Camillo Ricordi 23.

Regenerative Medicine for Diseases of the Retina

418

Deepak Lamba and Thomas A. Reh 24.

Peripheral Blood Stem Cells

438

Shay Soker, Gunter Schuch, and J. Koudy Williams 25.

Prospects of Somatic Cell Nuclear Transfer-derived Embryonic Stem Cells in Regenerative Medicine

456

Z. Beyhan and J.B. Cibelli 26.

Somatic Cells: Growth and Expansion Potential of T Lymphocytes

468

Rita B. Effros 27.

Mechanical Determinants of Tissue Development

480

Jonathan A. Kluge, Gary G. Leisk, and David L. Kaplan 28.

Morphogenesis and Morphogenetic Proteins

498

A.H. Reddi 29.

Physical Stress as a Factor in Tissue Growth and Remodeling

512

Robert E. Guldberg, Christopher S. Gemmiti, Yash Kolambkar, and Blaise Porter 30.

Engineering Cellular Microenvironments Wendy F. Liu, Elliot E. Hui, Sangeeta N. Bhatia, and Christopher S. Chen

536

Contents

31.

Applications of Nanotechnology

554

Benjamin S. Harrison 32.

GeneChips in Regenerative Medicine

562

Jason Hipp and Anthony Atala

Part IV 33.

Biomaterials for Regenerative Medicine

Design Principles in Biomaterials and Scaffolds

579 580

Hyukjin Lee and Tae Gwan Park 34.

Naturally Occurring Scaffold Materials

594

Stephen F. Badylak 35.

Synthetic Polymers

604

M.C. Hacker and A.G. Mikos 36.

Hybrid, Composite, and Complex Biomaterials for Scaffolds

636

Gilson Khang, Soon Hee Kim, Moon Suk Kim, and Hai Bang Lee 37.

Surface Modification of Biomaterials

656

Andrés J. García 38.

Cell–Substrate Interactions

666

Aparna Nori, Evelyn K.F. Yim, Sulin Chen, and Kam W. Leong 39.

Histogenesis in Three-Dimensional Scaffolds

686

Nicole M. Bergmann and Jennifer L. West 40.

Biocompatibility and Bioresponse to Biomaterials

704

James M. Anderson 41.

Essential Elements of Wound Healing

724

William J. Lindblad 42.

Proteins Controlled with Precision at Organic, Polymeric, and Biopolymer Interfaces for Tissue Engineering and Regenerative Medicine

734

Buddy D. Ratner

Part V 43.

Therapeutic Applications: Cell Therapy

Biomineralization and Bone Regeneration

743 744

Jiang Hu, Xiaohua Liu, and Peter X. Ma 44.

Blood Substitutes: Reverse Evolution from Oxygen Carrying to Non-Oxygen Carrying Plasma Expanders

756

Amy Tsai, Marcos Intaglietta, and Mark Van Dyke 45.

Articular Cartilage

766

Francois Ng kee Kwong and Myron Spector 46.

Implantation of Myogenic Cells in Skeletal Muscles Daniel Skuk and Jacques P. Tremblay

782

ix

x

CONTENTS

47.

Islet Cell Transplantation

794

Juliet A. Emamaullee and A.M. James Shapiro 48.

Cell-Based Repair for Cardiovascular Regeneration and Neovascularization: What, Why, How, and Where Are We Going in the Next 5–10 Years?

812

Doris A. Taylor and Andrey G. Zenovich 49.

Retinal Pigment Epithelium Derived from Embryonic Stem Cells

852

Irina Klimanskaya 50.

Cell Therapies for Bone Regeneration

868

Rehan N. Khanzada, Chantal E. Holy, F. Jerry Volenec, and Scott P. Bruder 51.

Cell-Based Therapies for Musculoskeletal Repair

888

Wan-Ju Li, Kiran Gollapudi, David P. Patterson, George T.-J. Huang, and Rocky S. Tuan 52.

Hepatocyte Transplantation

912

Stephen C. Strom and Ewa C.S. Ellis 53.

Bioartificial Livers

928

Randall E. McClelland and Lola M. Reid 54.

Neuronal Transplantation for Stroke

946

Douglas Kondziolka and Lawrence Wechsler 55.

Cell-Based Drug Delivery

954

Grace J. Lim, Sang Jin Lee, and Anthony Atala

Part VI 56.

Therapeutic Applications: Tissue Therapy

Fetal Tissues

967 968

Seyung Chung and Chester J. Koh 57.

Engineering of Large Diameter Vessels

978

Saami K. Yazdani and George J. Christ 58.

Engineering of Small Diameter Vessels

1000

Chrysanthi Williams and Robert T. Tranquillo 59.

Vascular Assembly in Engineered and Natural Tissues

1020

Eric M. Brey and Larry V. McIntire 60.

Cardiac Tissue

1038

Milica Radisic and Michael V. Sefton 61.

Regenerative Medicine in the Cornea

1060

Heather Sheardown and May Griffith 62.

Alimentary Tract

1072

Mike K. Chen 63.

Liver Cell-Based Therapy – Bioreactors as Enabling Technology Jörg C. Gerlach, Mariah Hout, Keneth Gage, and Katrin Zeilinger

1086

Contents

64.

Intracorporeal Kidney Support

1106

James J. Yoo, Akira Joraku, and Anthony Atala 65.

The Kidney

1114

William H. Fissell and H. David Humes 66.

Genitourinary System

1126

Anthony Atala 67.

Tissue Engineering of the Reproductive System

1138

Stefano Giuliani, Laura Perin, Sargis Sedrakyan, and Roger De Filippo 68.

Therapeutic Opportunities for Bone Grafting

1164

Jeffrey O. Hollinger, John P. Schmitz, Gary E. Friedlaender, Chris R. Brown, Scott D. Boden, and Samuel Lynch 69.

Cartilage Tissue Engineering

1176

Paulesh Shah, Alexander Hillel, Ronald Silverman, and Jennifer Elisseeff 70.

Phalanges and Small Joints

1198

Makoto Komura, Daniel Eberli, James J. Yoo, and Anthony Atala 71.

Functional Tissue Engineering of Ligament and Tendon Injuries

1206

Savio L.-Y. Woo, Alejandro J. Almarza, Sinan Karaoglu, and Steven D. Abramowitch 72.

Tissue Therapy: Implications of Regenerative Medicine for Skeletal Muscle

1232

Shen Wei and Johnny Huard 73.

Tissue Therapy: Central Nervous System

1248

Jordan H. Wosnick, M. Douglas Baumann, and Molly S. Shoichet 74.

Peripheral Nerve Regeneration

1270

Mahesh C. Dodla and Ravi V. Bellamkonda 75.

Dental Tissue Engineering

1286

Yan Lin and Pamela C. Yelick 76.

Innovative Regenerative Medicine Approaches to Skin Cell-Based Therapy for Patients with Burn Injuries

1298

Jörg C. Gerlach, Steven E. Wolf, Christa Johnen, and Bernd Hartmann 77.

Military Needs and Solutions in Regenerative Medicine

1322

Sara Wargo, Alan J. Russell, and Colonel John B. Holcomb

Part VII 78.

Regulations and Ethics

Ethical Considerations

1333 1334

Louis M. Guenin 79.

To Make is to Know: The Ethical Issues in Human Tissue Engineering Laurie Zoloth

1346

xi

xii

CONTENTS

80.

US Stem Cell Research Policy

1354

Josephine Johnston 81.

Overview of FDA Regulatory Process

1366

Celia Witten, Ashok Batra, Charles N. Durfor, Stephen L. Hilbert, David S. Kaplan, Donald Fink, Deborah Lavoie, Ellen Maher, and Richard McFarland 82.

Current Issues in US Patent Law

1386

Patrea L. Pabst

Index

1402

Preface

The textbook Principles of Regenerative Medicine has been created to be the primary resource for scientists, clinicians, teachers, students, and the public at large in the area of regenerative medicine. I am honored to have had the opportunity to edit the first edition with our co-editors, Robert Lanza, Robert Nerem, and Jamie Thompson. The contributions of the editors and editorial board cannot be overestimated. We are indebted to their vision, and the strong foundation they have created, upon which the current text is built. The specialty of regenerative medicine continues to grow and change rapidly. There have been major areas of advances in just the last few years. The field now encompasses multiple areas of scientific inquiry, each complex, but together, a powerful combination of technologies such as stem cells, genetic reprogramming, nuclear transfer, cloning, genomics, proteomics, nanotechnology, and tissue engineering. We are on the verge of an era of translation of benchside discoveries to clinical therapies. We hope that this book will enlighten all of these areas, and supply guidance where it is needed most. The textbook was organized in a manner which builds upon the basic science of the field, and goes forward to possible clinical applications and clinical utility. The textbook is organized into seven major areas, starting with an Introduction to Regenerative Medicine that encompasses some of the fundamentals of the field. The Biologic and Molecular Basis of Regenerative Medicine covers the molecular, mechanistic and phenotypic aspects of cells and cloning. The third section, Cells and Tissue Development, deals with the various types of cells and determinants of tissue formation. A section is dedicated to the area of biomaterials, especially as it pertains to tissue engineering. The fifth and sixth sections cover the topics of therapeutic applications, and deal with cell and tissue therapy, respectively. The last section of the book is dedicated to the regulatory and ethical aspects of the field. This area is becoming increasingly more important as the nexus between science, safety, and ethics is constantly changing. The authors have been tasked with enlightening the reader with the scientific efforts that are likely to impact the future of the field. We are indebted to our authors who graciously accepted their assignments, and who have infused the text with their energetic contributions. We are especially indebted to our publisher, Academic Press, without whose trust and guidance this work would not have begun; and our developmental editor, Melissa Turner, without whose hard work it would not have been finished. Anthony Atala, M.D. For the Editors

xiii

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List of Contributors

Jon D. Ahlstrom Department of Molecular and Cellular Biology University of California at Davis

Stephen F. Badylak* McGowan Institute for Regenerative Medicine University of Pittsburgh

Taby Ahsan Parker H. Petit Institute for Bioengineering and Bioscience Georgia Institute of Technology

Ashok Batra Office of Cellular, Tissue and Gene Therapies Center for Biologics Evaluations and Research

Julie Allickson* Vice President Laboratory Operations, Research & Development Cryo-Cell International, Inc.

M. Douglas Baumann Chemical Engineering and Applied Chemistry University of Toronto

Alejandro J. Almarza Research Assistant Professor Musculoskeletal Research Center Department of Bioengineering University of Pittsburgh

Ravi V. Bellamkonda* Professor of Biomedical Engineering Neurological Biomaterials and Therapeutics Wallace H Coulter Department of Biomedical Engineering Georgia Institute of Technology/Emory University Atlanta

James M. Anderson* Department of Pathology University Hospitals of Cleveland

Nicole M. Bergman Department of Bioengineering Rice University

Peter Andrews* Department of Biomedical Science University of Sheffield Western Bank Sheffield, Great Britain

Z. Beyhan Cellular Reprogramming Laboratory Michigan State University

Hadi Aslan Hebrew University – Hadassah Medical Center Hebrew University Center for Converging Sciences & Technologies Skeletal Biotechnology Lab Anthony Atala* W.B. Boyce Professor and Director, Wake Forest Institute for Regenerative Medicine, Wake Forest School of Medicine, Winston-Salem, North Carolina, USA

Mickie Bhatia* Stem Cell and Cancer Research Institute (SCC-RI) Michael G. DeGroote School of Medicine, McMaster University Canada Sangeeta N. Bhatia Division of Medicine Brigham & Women’s Hospital Peter M. Black Harvard Medical School Department of Neurosurgery and Physical Medicine & Rehabilitation

xv

xvi

LIST OF CONTRIBUTORS

Helen Blau* Donald E. and Delia B. Baxter Professor Director, Baxter Laboratory in Genetic Pharmacology Stanford University School of Medicine Department of Microbiology and Immunology Clinical Sciences Research Center CA, USA Scott D. Boden Professor of Orthopaedic Surgery Director, Emory Orthopaedics & Spine Center

George J. Christ* Professor of Regenerative Medicine, Urology and Physiology & Member, Molecular Medicine Program and Virginia Tech-Wake Forest University School of Biomedical Engineering and Sciences Head of the Cell, Tissue & Organ Physiology Program Wake Forest Institute for Regenerative Medicine Wake Forest University Baptist Medical Center NC, USA

Eric M. Brey Department of Research Hines V.A. Hospital

Seyung Chung Childrens Hospital Los Angeles University of Southern California Keck School of Medicine

Ali H. Brivanlou* Professor and Head of the Laboratory of Molecular Vertebrate Embryology, Rockerfeller University

J.B. Cibelli* Cellular Reprogramming Laboratory Michigan State University

Chris R. Brown The Emory Spine Center

Massimo Cimini MaRS Center Toronto Medical Discover Tower

Scott P. Bruder DePuy Spine, Inc., a Johnson & Johnson Company S.M. Chambers Center for Cell & Gene Therapy Baylor College of Medicine Christopher S. Chen* University of Pennsylvania Translational Research Labs LiHow Chen GTC Biotherapeutics Mike Chen* Department of Surgery University of Florida Sulin Chen Department of Biomedical Engineering Johns Hopkins School of Medicine N. Cheng UNC School of Medicine

Paolo De Coppi* Department of General Paediatric Surgery Great Ormond Street Hospital and Institute of Child Health Mahesh C. Dodla Georgia Institute of Technology Juan Dominguez-Bendala University of Miami Leonard M. Miller School of Medicine Pancreatic Development & Stem Cell Laboratory Diabetes Research Institute AM Doyle Georgia Institute of Technology Parker H. Petit Institute for Bioengineering and Bioscience Charles N. Durfor Center for Devices and Radiological Health Yann Echelard GTC Biotherapeutics

List of Contributors

Rita B. Effros* Department of Pathology & Laboratory Medicine David Geffen School of Medicine at UCLA Jennifer Elisseeff* Department of Biomedical Engineering Johns Hopkins University Ewa C.S. Ellis Department of Pathology University of Pittsburgh Juliet A. Emamaullee University of Alberta Carol A. Erickson* Department of Molecular and Cellular Biology University of California at Davis Roger De Filippo* Childrens Hospital Los Angeles Division of Urology Donald Fink Office of Cellular, Tissue and Gene Therapies Center for Biologics Evaluations and Research William H. Fissell University of Michigan Medical School, Department of Internal Medicine Gary E. Friedlaender Department of Orthopaedic Surgery Yale University School of Medicine Mark E. Furth Department of Urology Wake Forest University Yossi Gafni Hebrew University – Hadassah Medical Center Hebrew University Center for Converging Sciences & Technologies Skeletal Biotechnology Lab Keneth Gage McGowan Institute for Regenerative Medicine, University of Pittsburgh

Andres Garcia* Associate Professor Woodruff Faculty Fellow Woodruff School of Mechanical Engineering Petit Institute for Bioengineering and Bioscience Georgia Institute of Technology William Gavin GTC Biotherapeutics Daniel Gazit* Hebrew University – Hadassah Medical Center Hebrew University Center for Converging Sciences & Technologies Skeletal Biotechnology Lab International Stem Cell Institute – Department of Surgery Cedars Sinai Medical Center Los Angeles, CA Zulma Gazit Hebrew University – Hadassah Medical Center Hebrew University Center for Converging Sciences & Technologies Skeletal Biotechnology Lab Christopher S. Gemmiti Georgia Institute of Technology Jörg C. Gerlach* McGowan Institute for Regenerative Medicine, Bridgeside Point Boulevard University of Pittsburgh Paul J. Gokhale University of Sheffield Institute for Animal Breeding (FAL) M.A. Goodell* Center for Cell & Gene Therapy Baylor College of Medicine May Griffith University of Ottawa Eye Institute Ottawa Hospital-General Campus

xvii

xviii LIST OF CONTRIBUTORS

Louis M. Guenin Department of Microbiology and Molecular Genetics Harvard Medical School Stefano Giuliani Division of Urology, Childrens Hospital Los Angeles, Saban Research Institute, Keck School of Medicine, University of Southern California Robert E. Guldberg* Professor, Associate Director, Institute for Bioengineering and Bioscience School of Mechanical Engineering Georgia Institute of Technology M.C. Hacker Rice University Laboratory of Biomedical Engineering Benjamin S. Harrison* Wake Forest University School of Medicine Bernd Hartmann Burn Center Unfalkrankenhaus Stephen H. Hilbert Center for Devices and Radiological Health Alexander Hillel Department of Otolaryngology – Head & Neck Surgery Johns Hopkins University School of Medicine MD, USA Jason Hipp* Department of Urology and Regenerative Medicine Wake Forest University Col. J. B. Holcomb U.S. Army Institute of Surgical Research Jeffrey O. Hollinger* Professor of Biomedical Engineering and Biological Sciences Director, Bone Tissue Engineering Center Carnegie Mellon University

Chantal E. Holy* Director of Scientific Affairs DePuy Spine 325 Paramount Drive Raynham, MA, USA Mariah Hout McGowan Institute for Regenerative Medicine, University of Pittsburgh Jiang Hu Department of Biologic and Materials Sciences University of Michigan George T.-J. Huang Division of Endodontics Baltimore College of Dental Surgery University of Marlyand Johnny Huard* The Growth and Development Lab, Childrens Hospital of Pittsburgh Elliot E. Hui Harvard – M.I.T. Division of Health Sciences and Biology Electrical Engineering and Computer Science H. David Humes* University of Michigan Medical School, Department of Internal Medicine Marcos Intaglietta Department of Bioengineering University of California San Diego Junfeng Ji McMaster Cancer and Stem Cell Biology Research Institute McMaster University Yueha Jiang Stem Cell Institute University of Minnesota Christa Johnen Charite-Campus Virchow Humboldt University

List of Contributors

Josephine Johnston* Associate for Law and Bioethics Director of Research Operations The Hastings Center, New York Akira Joraku Department of Regenerative Medicine, Wake Forest University Health Sciences, Winston Salem David L. Kaplan* Department of Biomedical Engineering Tufts University David S. Kaplan Center for Devices and Radiological Health Gilson Khang Department of Polymer NanoScience and Technology Chonbuk National University Rehan N. Khanzada Sr. Process Development Engineer Johnson & Johnson Regenerative Therapeutics, LLC Soon Hee Kim Department of Polymer NanoScience and Technology Chonbuk National University Moon Suk Kim Nanobiomaterials Laboratory Korea Research Institutes of Chemical Technology Nadav Kimelman Hebrew University – Hadassah Medical Center Hebrew University Center for Converging Sciences & Technologies Skeletal Biotechnology Lab

Yash Kolambkar Graduate Research Assistant Department of Biomedical Engineering Georgia Tech/Emory Center for the Engineering of Living Tissues Chester J. Koh Childrens Hospital Los Angeles University of Southern California Keck School of Medicine Makoto Komura The Department of Pediatric Surgery Tokyo University Hospital Douglas Kondziolka University of Pittsburgh Neurological Surgery Deniz Konya Department of Neurosurgery and Physical Medicine & Rehabilitation Harvard Medical School Wilfried A. Kues Department of Biotechnology Institute for Animal Breeding (FAL) Francois Ng kee Kwong Tissue Engineering, VA Boston Healthcare System Deepak Lamba Department of Biological Structure, School of Medicine University of Washington Hai Bang Lee* Nanobiomaterials Laboratory Korea Research Institutes of Chemical Technology

Irina Klimanskaya* Advanced Cell Technology Biotech Five

Hyukjin Lee Department of Biological Sciences Korea Advances Institute of Science and Technology

Jonathan A. Kluge Department of Biomedical Engineering Tufts University

Gary G. Leisk Department of Biomedical Engineering Tufts University

xix

xx

LIST OF CONTRIBUTORS

Kam W. Leong* James B. Duke Professor of Biomedical Engineering, Director of the Bioengineering Initiative, UK Ariel J. Levine Rockerfeller University Ren Ke Li* MaRS Center Toronto Medical Discover Tower Wan-Ju Li Cartilage Biology and Orthopaedics Branch National Institute of Arthritis Grace J. Lim* Medical Research Institute Department of Medical and Biological Engineering Kyungpook National University School of Medicine, South Korea Yan Lin The Forsyth Institute William J. Lindblad* Department of Pharmaceutical Sciences Massachusetts College of Pharmacy & Health Sciences Wendy F. Liu University of Pennsylvania Translational Research Labs Xiaohua Liu Department of Biologic and Materials Sciences University of Michigan Andrea Lucas-Hahn Department of Biotechnology Institute for Animal Breeding (FAL) Aernout Luttun Department of Medicine Stem Cell Institute University of Minnesota Samuel Lynch BioMimetic Therapeutics Inc.

Peter X. Ma* Professor, Fellow, American Institute for Medical and Biological Engineering Department of Biologic and Materials Sciences Department of Biomedical Engineering Macromolecular Science and Engineering Center The University of Michigan Ellen Maher Office of Cellular, Tissue and Gene Therapies Center for Biologics Evaluations and Research Manuela Martins-Green* Department of Cell Biology and Neuroscience University of California Randall E. McClelland* University of North Carolina School of Medicine Richard McFarland Office of Cellular, Tissue and Gene Therapies Center for Biologics Evaluations and Research Larry V. McIntire* The Wallace H. Coulter Chair and Professor Department of Biomedical Engineering Georgia Tech Harry Meade* Senior Vice President Research and Development GTC Biotherapeutics David L. Melican GTC Biotherapeutics A.G. Mikos* Rice University Laboratory of Biomedical Engineering Fernando Ulloa Montoya Department of Medicine Stem Cell Institute University of Minnesota

List of Contributors

Robert M. Nerem* Georgia Institute of Technology Parker H. Petit Institute for Bioengineering and Bioscience Heiner Niemann* Department of Biotechnology Institute for Animal Breeding (FAL) Aparna Nori Department of Biomedical Engineering Johns Hopkins School of Medicine Patrea L. Pabst* Pabst Patent Group LLP Kook In Park Department of Pediatrics and Brain Korea 21 Project for Medical Science Yonsie University College of Medicine Tae Gwan Park* Department of Biological Sciences Korea Advances Institute of Science and Technology David P. Patterson Cartilage Biology and Orthopaedics Branch National Institute of Arthritis Karen Pauwelyn Stem Cell Institute Katholike Universiteit Leuven Gadi Pelled Hebrew University – Hadassah Medical Center Hebrew University Center for Converging Sciences & Technologies Skeletal Biotechnology Lab Laura Perin Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem M. Petreaca Department of Cell Biology and Neuroscience University of California

Antonello Pileggi Research Assistant, Professor of Surgery Cell Transplant Center & Clinical Islet Transplant Program Diabetes Research Institute Division of Cellular Transplantation DeWitt Daughtry Department of Surgery University of Miami Miller School of Medicine Jason H. Pomerantz Baxter Laboratory in Genetic Pharmacology Stanford University School of Medicine Blaise Porter Georgia Institute of Technology Milica Radisic Assistant Professor Institute of Biomaterials and Biomedical Engineering Department of Chemical Engineering and Applied Chemistry Heart & Stroke/Richard Lewar Centre of Excellence University of Toronto Buddy D. Ratner* Director, University of Washington Engineered Biomaterials (UWEB) Michael L. and Myrna Darland Endowed Chair in Technology Commercialization Professor of Bioengineering and Chemical Engineering University of Washington A. Hari Reddi* Professor and Lawrence J. Ellison Chair University of California, Davis Sacramento, CA, USA Thomas A. Reh* Professor of Biological Structure Health Sciences Center University of Washington School of Medicine Lola M. Reid* University of North Carolina School of Medicine Cell and Molecular Physiology & Biomedical Engineering

xxi

xxii

LIST OF CONTRIBUTORS

Camillo Ricordi* Diabetes Research Institute (R-134) Miller School of Medicine University of Miami

Heather Sheardown* Associate Professor Department of Chemical Engineering McMaster University

Jeff Ross Stem Cell Institute University of Minnesota Medical School

Molly S. Shoichet* Professor, Chemical Engineering and Applied Chemistry Director, Undergraduate Collaborative Bioengineering, Canada Research Chair in Tissue Engineering University of Toronto Terrence Donnelly Centre for Cellular and Biomolecular Research

Alan J. Russell* McGowan Institute for Regenerative Medicine University of Pittsburgh Filipe N.C. Santos Depts. Of Neurosurgery and Physical Medicine & Rehabilitation Harvard Medical School John P. Schmitz San Pedro Facial Surgery Gunter Schuch Wake Forest University School of Medicine Institute for Regenerative Medicine Sargis Sedrakyan Department of Urology Children’s Hospital Los Angeles Keck School of Medicine University of Southern California Michael V. Sefton* University Professor Institute of Biomaterials and Biomedical Engineering, Michael E. Charles Professor, Department of Chemical Engineering and Applied Chemistry University of Toronto Marta Serafini Stem Cell Institute, Department of Medicine University of Minnesota Medical School Paulesh Shah Department of Biomedical Engineering Johns Hopkins University A.M. James Shapiro University of Alberta

M. Minhaj Siddiqui Massachusetts General Hospital Richard L. Sidman Department of Neurology Beth Israel-eaconess Medical Center Harvard Medical School Ronald Silverman Department of Biomedical Engineering Johns Hopkins University Daniel Skuk Human Genetic Unit Centre de Recherche du CHUL Evan Snyder* The Burnham Institute Shay Soker* Associate Professor of Regenerative Medicine and Surgical Sciences Head, Molecular and Cell Biology Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Myron Spector* Tissue Engineering, VA Boston Healthcare System David L. Stocum* Center for Regenerative Biolog and Medicine Indiana University-Purdue University Indianapolis

List of Contributors xxiii

Stephen C. Strom* Department of Pathology University of Pittsburgh

Deborah Vavoie Office of Cellular, Tissue and Gene Therapies Center for Biologics Evaluations and Research

Doris A. Taylor* Bakken Professor Director, Center for Cardiovascular Repair University of Minnesota

Catherine Verfaillie* Professor of Medicine Director, Stamcelinstituut, K U Leuven Onderwijs & Navorsing 1Herestraat 49, bus 80343000 Leuven

Yang D. Teng* Associate Professor, Harvard Medical School Director, Lab of Spinal Cord Injury & Neural Stem Cell Biology Neosurgery & PM&R, HMS/BWH/SRH James Thomson Department of Anatomy Wisconsin Regional Primate Research Center

F. Jerry Volenec DuPuy Spine Sara Wargo McGowan Institute for Regenerative Medicine Joseph W. Warnwath Department of Biotechnology

Robert T. Tranquillo* Department of Biomedical Engineering University of Minnesota

Lawrence Wechsler University of Pittsburgh

Jacques P. Tremblay* Human Genetic Unit Centre de Recherche du CHUL

Shen Wei The Growth and Development Lab, Childrens Hospital of Pittsburgh

Amy Tsai Department of Bioengineering University of California San Diego

Richard D. Weisel MaRS Center Toronto Medical Discover Tower

Rocky S. Tuan* Chief, Cartilage Biology and Orthopaedics Branch National Institute of Arthritis, and Musculoskeletal & Skin Diseases National Institute of Health MD, USA

Jennifer L. West* Rice University Department of Bioengineering

Ross S. Tubo* Senior Director, Stem Cell Biology, Genzyme Corp. Mark Van Dyke* The Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine

Chrysanthi Williams Bose Corporation ElectroForce Systems Group J. Koudy Williams Wake Forest University School of Medicine Institute for Regenerative Medicine Celia Witten* Office of Cellular, Tissue and Gene Therapies Center for Biologics Evaluations and Research

xxiv LIST OF CONTRIBUTORS

Steven E. Wolf Musculoskeletal Research Center (MSRC) University of Pittsburgh Steven E. Wolf Burn Center United State Army Institute of Surgical Research Savio L.-Y. Woo* Musculoskeletal Research Center (MSRC) University of Pittsburgh Jordan H. Wosnick Chemical Engineering and Applied Chemistry University of Toronto Christine Wrenzycki Department of Biotechnology Institute for Animal Breeding (FAL) Munira Xaymardan MaRS Center Toronto Medical Discover Tower Hsin-Lei Yao Cell and Molecular Physiology & Biomedical Engineering University of North Carolina School of Medicine Saami K. Yazdani Wake Forest University Baptist Medical Center Medical Center Boulevard Pamela C. Yelick* Associate Professor of Oral and Maxillofacial Pathology, School of Dental Medicine Genetics Cell, Molecular, and Developmental Biology

*Corresponding Authors

James J. Yoo* Wake Forest University School of Medicine Institute for Regenerative Medicine Medical Center Boulevard Junying Yu* The Genetics and Biotechnology Building University of Wisconsin-Madison Katrin Zeilinger Charite Campus Virchow Humbold University Lepeng Zeng Stem Cell Institute University of Minnesota Medical School Andrey G. Zenovich Center for Cardiovascular Repair Bonan Zhong McMaster Cancer and Stem Cell Biology Research Institute McMaster University Carol A. Ziomek Vice President of Development, GTC Biotherapeutics Laurie Zoloth* Feinberg School of Medicine Northwestern University

Part I Introduction to Regenerative Medicine

1 Current and Future Perspectives of Regenerative Medicine Mark E. Furth and Anthony Atala

REGENERATIVE MEDICINE: CURRENT AND FUTURE PERSPECTIVES Progress and Challenges for Cell-Based Regenerative Medicine Regenerative medicine seeks to devise new therapies for patients with severe injuries or chronic diseases in which the body’s own responses do not suffice to restore functional tissue. A recent publication from the US National Academy of Sciences, Stem Cells and the Future of Regenerative Medicine (Committee, 2002), identified a wide array of major unmet medical needs which might be addressed by regenerative technologies. These include congestive heart failure (approximately 5 million patients in the United States) (Murray-Thomas and Cowie, 2003), osteoporosis (10 million US patients), Alzheimer’s and Parkinson’s diseases (5.5 million patients each), severe burns (0.3 million), spinal cord injuries (0.25 million), and birth defects (0.15 million). Another area of critical need is diabetes mellitus (16 million US patients and more than 217 million worldwide) (Smyth and Heron, 2006). Patients with type 1 diabetes lack pancreatic beta-cells, essential for the production of insulin, because of autoimmune destruction and represent from 10% to 20% of the total. Many patients with type 2 diabetes also show insufficient pancreatic beta-cell mass. Thus, patients in both groups potentially might be treated if methods could be developed to promote endogenous regeneration of beta-cells or to provide enough surrogate beta-cells and pancreatic islets for transplantation (Weir, 2004). The therapeutic use of growth factors and cytokines to stimulate the production and/or function of endogenous cells represents the area of regenerative medicine that, arguably, has shown the greatest clinical impact to date (Ioannidou, 2006). Regenerative therapies comprising living cells also have entered into practice, initially through the widespread adoption of both allogeneic and autologous bone marrow transplantation (Thomas, 1999). The presence of hematopoietic progenitor and stem cells with great replicative capacity in vivo, and their ability to reenter the bone marrow niche from the circulation, enabled this major medical advance. Subsequently, the development of methods to expand ex vivo and deliver such cell types as keratinocytes and chondrocytes, through advances in cell culture and scaffold technologies, led to successful tissue engineering for wound repair (Johnson, 2000; Lavik and Langer, 2004). Despite significant challenges in development and manufacturing, several bioartificial skin graft and cartilage replacement products have achieved regulatory approval (Lysaght and

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Reyes, 2001; Naughton, 2002; Lysaght and Hazlehurst, 2004). These therapies validate the potential of cell-based regenerative approaches. The extension to new therapeutic areas, especially the development of neo-organs with complex threedimensional structure, will depend on complementary advances in biology, materials science, and engineering. A major limitation remains the ability to provide oxygen and nutrients to neo-tissues both in vitro and after implantation. Advances in scaffold composition and design, in bioreactor technology, and in the use of pro-angiogenic factors may all help to overcome this barrier and are discussed in depth in other chapters of this book. Here we will focus mainly on sources of cells for regenerative medicines. A primary issue remains the choice between using a patient’s own cells, or those of a closely matched relative, versus those from an unrelated allogeneic donor. More broadly, future developments depend heavily on increased understanding and effective utilization of multiple classes of progenitor and stem cells. When populations that include precursor cells (i.e. cells not yet fully differentiated and capable of significant proliferation) can be obtained from a small biopsy of a patient’s tissue, and these cells are able to expand and differentiate in culture and/or after implantation back into the patient, autologous therapies are feasible. These have the great advantage of avoiding the risk of immune rejection based on differences in histocompatibility antigens, so that the use of immunosuppressive drugs is not required. However, there is a substantial practical appeal to “off the shelf ” products that do not require the cost and time associated with customized manufacturer of an individual product for each individual recipient (Lysaght and Hazlehurst, 2004). Among the approved bioengineered skin products, Dermagraft (Smith & Nephew) and Apligraf (Organogenesis) utilize allogeneic cells expanded from donated human foreskins to treat many unrelated patients. Despite the genetic mismatch between donor and recipient, the skin cells in Dermagraft and Apligraf do not induce acute immune rejection, possibly because of the absence of antigen-presenting cells in the grafts (Briscoe et al., 1999; Horch et al., 2005). Thus, these products can be utilized without immunosuppressive drug therapy (Moller et al., 1999). Eventually, the donated skin cells may be rejected, but after sufficient time has passed for the patient’s endogenous skin cells to recover and take their place. Products based on autologous cells also have achieved regulatory approval and reached the market. In particular, Genzyme Biosurgery has developed Epicel, a permanent skin replacement product for patients with life-threatening burns, and Carticel, a chondrocyte-based treatment for large articular cartilage lesions. In each case seed cells are obtained from a small biopsy of the patient’s tissue. These cells are expanded in culture, processed, and returned to the patient. New Therapies Using Autologous Cells Recent clinical studies highlight ongoing efforts to develop new autologous cell-based therapies. The recognition that, in addition to hematopoietic stem cells, bone marrow also contains mesenchymal stem cells (MSC) and endothelial progenitor cells (EPC), has spurred ongoing efforts to use autologous marrow cells for blood vessel tissue engineering and for treatment of myocardial infarction. In the case of engineering of blood vessels, vascular grafts of autologous bone marrow cells seeded onto biodegradable synthetic conduits or patches have been implanted in children with congenital heart defects (Shin’oka et al., 2005). Safety data on 42 patients with a mean follow-up period of 490 days post-surgery appeared very encouraging, with no major adverse events reported. The grafted engineered vessels remained patent and functional. Moreover, there was evidence that the vessels increased in diameter as the patients grew, thus highlighting a critical potential advantage of regenerative therapies incorporating living cells. Further advances in blood vessel engineering will likely arise from multidisciplinary approaches demanding advances at the interface of biology and engineering. In recent preclinical studies scaffolds for neo-vessels

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blending collagen type I and elastin with polylactic-co-glycolic acid (PLGA) were fabricated by electrospinning and showed compliance, burst pressure, and mechanical properties comparable to native vessels (Stitzel et al., 2006). The electrospun vessels also displayed good biocompatibility both in vitro and after implantation in vivo. When seeded with endothelial and smooth muscle cells, or progenitor MSCs and/or EPCs, these constructs may provide a basis to produce functional vascular grafts suitable for clinical applications such as cardiac bypass procedures. The seeding process itself may demand future advances, since it will be difficult for cells to penetrate a nanofibrillar structure in which pore spaces are considerably smaller than the diameter of a cell (Lutolf and Hubbell, 2005). Electrospinning actually may be used to incorporate living cells into a fibrous matrix. A recent proof of concept study documented that smooth muscle cells could be concurrently electrospun with an elastomeric poly(ester urethane)urea, leading to “microintegration” of the cells in strong, flexible fibers with mechanical properties not greatly inferior to those of the synthetic polymer alone (Stankus et al., 2006). The cell population retained high viability and, when maintained in a perfusion bioreactor, the cellular density in the electrospun fibers doubled over 4 days in culture. One can imagine that in the future, progenitors of vessel cells may be harvested from a patient, incorporated into an electrospun matrix and incubated in a bioreactor, first to drive expansion and differentiation and then, via pulsed flow, to promote vessel maturation (Niklason et al., 1999). Similar strategies may be attempted to treat patients with congestive heart failure (Krupnick et al., 2004). Already, a number of clinical studies have been carried out on the injection of autologous bone marrow cells, sometimes unfractionated sometimes enriched for stem/progenitor cells, into the heart after myocardial infarction (Stamm et al., 2006). The initial rationale for this approach came from experiments in rodents interpreted as demonstrating the production of new cardiomyocytes through the transdifferentiation of hematopoietic stem cells. Evidence for myogenesis of grafted cells, whether from the hematopoietic lineage or, as seems much more plausible, from mesenchymal progenitors, remains sparse. However, some controlled studies do indicate potential clinical benefits from the autologous cell therapy. This may result from the production of angiogenic factors by the injected cells rather than from integration of donor cells into either muscle or new blood vessels. Nonetheless, although still a daunting challenge, the application of regenerative medicine principles to repair damaged cardiac muscle now seems within the possible realm (Dimmeler et al., 2005). The correct choice of cell source, the development and maturation of tissue engineered cardiac patches, and overcoming chronic fibrotic scarring remain hurdles to be overcome. In another example of regenerative therapy utilizing autologous cells, here following the general paradigm first established for skin, bladder urothelial and smooth muscle cells were expanded in culture from small biopsies and seeded on scaffolds to produce tissue engineered neo-bladders. Such constructs were implanted in seven pediatric patients with high-pressure or poorly compliant bladders, some of whom have now been followed for over 5 years (mean 46 months) (Atala et al., 2006). The results are strongly encouraging and should lead to larger scale studies of safety and efficacy, targeting product approval after regulatory review. Cell Sources The ability to produce enough cells of the necessary types from the skin, cartilage, or bladder for bioengineered products depended on the presence of stem and progenitor cells in the corresponding adult tissues. It also required the development of culture methods that both permit the expansion of the precursor cells and allow enough differentiation for generation of the desired neo-tissue. Implementation of this strategy for regenerative medicine, based on expansion of autologous cells, cannot yet be extended to all tissues and organs. In some cases it is not clear how to obtain biopsies containing progenitor or stem cells, or even whether such cells exist. In other cases, culture conditions for expansion of the precursor cell population are not yet available.

Current and Future Perspectives of Regenerative Medicine

The future development of cell-based regenerative medicine depends on further translation of basic discoveries regarding the identity and behavior of stem cells into practical clinical applications. Important targets include cells of organs for which orthotopic transplantation already has been established as an important mode of therapy, but for which the supply of donor organs does not meet the current need. Examples include cells of the heart, kidney, liver, and pancreas, specifically insulin-producing beta-cells. In addition production of neurons and other cells of the nervous system may permit therapy of degenerative diseases for which no effective treatment yet exists. Mammalian stem cells have been divided into two general categories: embryonic and adult. Embryonic stem (ES) cells and the comparable embryonic germ (EG) cells appear to give rise to all specialized cell types, with the exception of a limited set of extra-embryonic cells. Adult stem cells, which may actually derive from fetal, neonatal, or truly adult tissue, show varying degrees of restriction to particular lineages. ES Cells ES cells and EG cells appear very similar (we will use “ES” to refer to both) and will likely have comparable medical applications. In fact, a recent report indicates that ES cells, which are derived from the inner cell mass of early embryos, most closely resemble early germ cells (Zwaka and Thomson, 2005). The ES cells can self-renew apparently without limit in culture, although mechanisms underlying this capacity remain incompletely understood (Rao, 2004; Stewart et al., 2006) and established ES lines may display some genomic instability. Furthermore, ES cells are broadly pluripotent (Evans and Kaufman, 1981; Martin, 1981; Shamblott et al., 1998; Amit et al., 2000). This great degree of plasticity represents both the strongest attraction and a significant potential limitation to the use of ES cells for regenerative medicine. A major remaining challenge is to direct the efficient production of pure populations of specific desired cell types from human ES cells (Odorico et al., 2001). ES cells appear unique among normal stem cells in being tumorigenic, forming teratomas that contain cell types representing all three EG layers in a disorganized form (Martin, 1981; Thomson et al., 1998; Cowan et al., 2004). For clinical use it will be important to exclude undifferentiated stem cells from any products derived from ES cells (Lawrenz et al., 2004). Strategies have been envisaged to increase safety by introducing into ES cells a “suicide” gene, for example that encoding the thymidine kinase of Herpes simplex virus, which would render any escaping tumor cells sensitive to the drug ganciclovir (Odorico et al., 2001; Schuldiner et al., 2003). However, the genetic manipulation is itself not without risk, and the need to validate the engineered cell system would likely extend and complicate regulatory review of therapeutic products. A central issue that must be addressed for tissue engineered products derived from ES cells, and also from any non-autologous adult stem cells, is immune rejection based on mismatches at genetic histocompatibility loci. It generally has been assumed that, because human ES cells and their differentiated derivatives can be induced to express high levels of major histocompatibility complex (MHC) Class I antigens (e.g. HLA-A and HLA-B), any ES cell-based product will be subjected to graft rejection (Drukker et al., 2002). Therapeutic cloning offers a potential means to generate cells with the exact genetic constitution of each individual patient, so that immune rejection of grafts based on mismatched histocompatibility antigens should not occur. The approach entails transferring the nucleus of a somatic cell into an enucleated oocyte (SCNT), generating a blastocyst, and then culturing the inner cell mass to obtain an ES cell line (Colman and Kind, 2000). If required, genetic manipulation of the cells may be carried out to correct an inherited defect prior to production of the therapeutic graft (Rideout III et al., 2002). Despite a published claim (Hwang et al., 2005) later withdrawn, the generation of human ES cells by SCNT has not yet been achieved. However, the concept of therapeutic cloning to provide cells for tissue engineering applications has been clearly validated in a large animal model. Adult bovine fibroblasts were used as nuclear donors and bioengineered tissues were generated from cloned cardiac, skeletal muscle, and kidney cells (Lanza et al., 2002). The grafts, including functioning

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renal units capable of urine production, were successfully transplanted into the corresponding donor animals long term with no evidence for rejection. Although SCNT is the subject of political, ethical, and scientific debate (Hall et al., 2006), intense efforts in both the private sector (Lysaght and Hazlehurst, 2003) and academic institutions are likely to yield cloned human lines in the near future. The reprogramming of somatic cell nuclei to yield pluripotent cells after introduction into the cytoplasm of enucleated eggs raises the possibility that additional means may be found to create cells with expanded potential to yield desired differentiated cell types. Counter to intuition, it appears that nuclei taken from certain terminally differentiated cells, such as postmitotic neurons, are readily reprogrammed to yield pluripotent cells by SCNT (Eggan et al., 2004). Nuclei from more differentiated cells may actually be superior for this purpose than nuclei of adult stem cells (Inoue et al., 2006; Sung et al., 2006), although the opposite trend was noted in studies using nuclei from neuronal lineage cells (Blelloch et al., 2006). In addition, fusion of somatic cells to ES cells can reprogram the somatic nuclei to an embryonic state (Cowan et al., 2005). Most remarkably, the expression of a small set of genes usually associated with ES cells (e.g. Oct3/4, Sox2, c-Myc, and Klf4) can induce an “embryonic” state, including pluripotency and the capacity to form teratoma tumors, in at least some somatic cells (fibroblasts) (Takahashi and Yamanaka, 2006). The properties and differentiation potential of a number of human ES cell lines obtained by traditional means from early embryos currently used for research have been reviewed recently (Hoffman and Carpenter, 2005). The clinical application of ES cells for tissue engineering will depend on the development of robust methods to isolate and grow them under conditions consistent with Good Manufacturing Practice and regulatory review for safety. In particular, it is important to eliminate the requirement for murine feeder cells by using human feeders or, better, feeder-free conditions. In addition, development of culture conditions without the requirement for non-human serum would be advantageous. Progress has been made in the derivation and expansion of human ES cells with human feeder cells (Amit et al., 2003; Hovatta et al., 2003; Yoo et al., 2005; Stacey et al., 2006) or entirely without feeders (Amit et al., 2004; Carpenter et al., 2004; Beattie et al., 2005; Hovatta and Skottman, 2005; Klimanskaya et al., 2005; Sjogren-Jansson et al., 2005). Perhaps the greater challenge remains in directing the differentiation of human ES cells to a given desired lineage with high efficiency. The underlying difficulty is that ES cells are developmentally many steps removed from adult, differentiated cells, and to date we have no general way to deterministically control the key steps in lineage restriction. Presumably, the same problem would be encountered with ES cells generated by SCNT or other means of reprogramming somatic cell nuclei. To induce differentiation in vitro ES cells are allowed to attach to plastic in monolayer culture or, more frequently, to form aggregates called embryoid bodies (Itskovitz-Eldor et al., 2000). Over time within these aggregates cell types of many lineages are generated, including representatives of the three germ layers. The production of embryoid bodies can be enhanced and made more consistent by incubation in bioreactors (Gerecht-Nir et al., 2004). Further selection of specific lineages generally requires sequential exposure to a series of inducing conditions, either based on known signaling pathways or identified by trial and error. In most cases lineage-specific markers are expressed by the differentiated cells, but cells often do not progress to a full terminally differentiated phenotype. As summarized in recent reviews, the cell lineages which have been generated in vitro include, among others, several classes of neurons, astrocytes, oligodendrocytes, multipotent mesenchymal precursor cells, osteoblasts, cardiomyocytes, keratinocytes, pneumocytes, hematopoietic cells, hepatocytes, and pancreatic beta-cells (Nir et al., 2003; Tian and Kaufman, 2005; Raikwar et al., 2006; Trounson, 2006). In general, it appears easier to obtain adult cells derived from ectoderm, including neurons, and mesoderm, including cardiomyocytes, than cells derived from endoderm (Trounson, 2006). This may help determine the first areas in which ES-derived cells enter clinical translation, once the barriers discussed above are

Current and Future Perspectives of Regenerative Medicine

surmounted. Dopaminergic neurons generated from primate and human ES cells already have been tested with encouraging results in animal models of Parkinson’s disease (Perrier et al., 2004; Sanchez-Pernaute et al., 2005). Promising data also have been obtained with ES-derived oligodendrocytes in spinal cord injury models (Keirstead et al., 2005; Mueller et al., 2005). Cardiomyocytes derived from human ES cells, similarly, are candidates for future clinical use (He et al., 2003; Nir et al., 2003; Goh et al., 2005; Lev et al., 2005). However, the functional criteria that must be met to ensure physiological competence will be stringent because of the risk of inducing arrhythmias (Caspi and Gepstein, 2006; Passier et al., 2006). The robust generation of pancreatic beta-cells and bioengineered islets from human ES cells or other stem cells would represent a particularly important achievement, with potential to treat diabetes (Weir, 2004; Nir and Dor, 2005). Clusters of insulin-positive cells, resembling pancreatic islets and expressing various additional markers of the endocrine pancreatic lineage, have been produced from mouse ES cells (Lumelsky et al., 2001) and also from non-human primate and human ES cells (Assady et al., 2001; Lester et al., 2004; Brolen et al., 2005; Baharvand et al., 2006). The production of beta-like cells can be enhanced by the expression of pancreatic transcription factors (Miyazaki et al., 2004; Shiroi et al., 2005). However, the assessment of differentiation must take into account the uptake of insulin from the growth medium, in addition to de novo synthesis (Paek et al., 2005). It seems fair to conclude that the efficient production of functional beta-cells from ES cells remains a difficult objective to achieve. As in other bioengineering applications with ES-derived cells, efforts to reverse diabetes also will depend on the complete removal of non-differentiated cells to avoid the formation of teratoma tumors, which were observed after implantation of ES-derived beta-cells in an animal model (Fujikawa et al., 2005). Adult Stem Cells Despite the acknowledged promise of ES cells, the challenges of controlling lineage-specific differentiation and eliminating residual stem cells are likely to extend the timeline for a number of tissue engineering applications. In many cases adult stem cells may provide a more direct route to clinical translation. Lineage-restricted stem cells have been isolated from both fetal and postnatal tissues based on selective outgrowth in culture and/or immunoselection for surface markers. Examples with significant potential for new applications in regenerative medicine include neural (Baizabal et al., 2003; Goh et al., 2003), cardiac (Beltrami et al., 2003; Oh et al., 2003), muscle-derived (Cao et al., 2005), and hepatic stem cells (Kamiya et al., 2006; Schmelzer et al., 2006). A significant feature of each of these populations is a high capacity for self-renewal in culture. Their ability to expand may be less than that for ES cells, but in some cases the cells have been shown to express telomerase and may not be subjected to replicative senescence. These adult stem cells are multipotent. Neural stem cells can yield neurons, astrocytes, and oligodendrocytes. Cardiac stem cells are reported to yield cardiomyocytes, smooth muscle, and endothelial cells. Muscle-derived stem cells yield skeletal muscle and can be induced to produce chondrocytes. Hepatic stem cells yield hepatocytes and bile duct epithelial cells. The lineagerestricted adult stem cells all appear non-tumorigenic. Thus, unlike ES cells, it is likely that they could be used safely for bioengineered products with or without prior differentiation. It is possible that some lineage-specific adult stem cells are capable of greater plasticity than might be supposed based solely on their tissue of origin. For example, there is evidence that hepatic stem cells may be induced to generate cells of additional endodermal lineages such as the endocrine pancreas (Yang et al., 2002; Nakajima-Nagata et al., 2004; Yamada et al., 2005; Zalzman et al., 2005). This type of switching of fates among related cell lineages may prove easier than inducing a full developmental program from a primitive precursor such as an ES cell. Another class of adult cells with enormous potential value for regenerative medicine is the MSC, initially described in bone marrow (Bruder et al., 1994; Pittenger et al., 1999). These multipotent cells are able to give rise

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to differentiated cells of connective tissues including bone, cartilage, muscle, tendon, and fat. The MSC have, therefore, generated considerable interest for musculoskeletal and vascular tissue engineering (Gao and Caplan, 2003; Tuan et al., 2003; Barry and Murphy, 2004; Guilak et al., 2004; Risbud and Shapiro, 2005). Cells with similar differentiation potential and marker profiles have been isolated from a number of tissues in addition to the bone marrow. A notable source is the adipose tissue in which the cells are abundant and easily obtained by processing of suction-assisted lipectomy (liposuction) specimens (Zuk et al., 2001; Gimble and Guilak, 2003). In general it seems better to view MSC as mixed populations of progenitor cells with varying degrees of replicative potential, rather than homogeneous stem cells. However, some classes of MSC, including lines cloned from single cells in skin (Bartsch et al., 2005), have been maintained in culture for extended periods. A very small subset of mesenchymal cells from bone marrow, termed multipotent adult progenitor cell (MAPC), reportedly are capable of extensive self-renewal and of differentiation into cell lineages not observed with typical MSC, including examples from each EG layer (Jiang et al., 2002). Cells originating in a developing fetus and isolated from amniotic fluid or chorionic villi are a new source of stem cells of great potential interest for regenerative medicine (De Coppi et al., 2001; Siddiqui and Atala, 2004; Tsai et al., 2006). Fetal-derived cells with apparently similar properties also have been described in the amnion of term placenta (Miki et al., 2005). Amniotic fluid stem (AFS) cells and amniotic epithelial cells can give rise to differentiated cell types representing the three EG layers (Siddiqui and Atala, 2004; Miki et al., 2005). Formal proof that single AFS cells can yield this full range of progeny cells was obtained using clones marked by retroviral insertion (unpublished data from A. Atala laboratory). The cells can be expanded for well over 200 population doublings with no sign of telomere shortening or replicative senescence, and retain a normal diploid karyotype. They are readily cultured without need for feeder cells. The AFS cells express some markers in common with ES cells, such as the surface antigen SSEA4 and the transcription factor Oct3/4, while other markers are shared with mesenchymal and neural stem cells. A broadly multipotent cell population obtained from umbilical cord blood may have certain key properties in common with AFS cells, and was termed “unrestricted somatic stem cells” (USSCs) (Kogler et al., 2004). This population may overlap with or be identical to the so-called “umbilical cord matrix stem” (UCMS) cells isolated from Wharton’s jelly (Mitchell et al., 2003; Weiss et al., 2006). The full developmental potential of the various stem cell populations obtained from fetal and adult sources remains to be determined. It is possible that virtually all of the cell types that might be desired for tissue engineering could be obtained from AFS cells, equivalent stem cells from placenta, those from the non-hematopoietic subset in umbilical cord blood, or comparable populations. Similar approaches to those being taken with ES cells, such as genetic modification with expression vectors for lineage-specific transcription factors, may help in the generation of those differentiated cell types for which it proves difficult to develop a straightforward induction protocol using external signals. However, it will remain necessary to show, beyond induction of a set of characteristic markers, that fully functional mature cells can be generated for any given lineage. Immune Compatibility The growing number of choices of cell sources for bioengineered tissues opens up a range of strategies to obtain the desired differentiated cell populations. The issue of immune compatibility remains central. Although life-long immunosuppression can be successful, as in conjunction with orthotopic organ transplantation, it would be preferable to design bioengineering-based products that will be tolerated by recipients without the need for immunosuppressive drugs. The only cell-based therapies guaranteed to be histocompatible would contain autologous cells or those derived by therapeutic cloning (assuming mitochondrial differences are not critical) (Lanza et al., 2002). When a perfectly matched, personalized therapeutic product is not available, there still should be ways to limit the requirement for immunosuppression. First, there may be a strong intrinsic advantage to developing cell-based products from certain stem cells because there is evidence

Current and Future Perspectives of Regenerative Medicine

that they, and possibly differentiated cells derived from them, are immune privileged. Second, it may be possible to develop banks of cells that can be used to permit histocompatibility matching with recipient patients. Human ES cells express low levels of MHC Class I antigens (HLA-A, HLA-B) and are negative for MHC Class II (HLA-DR) (Drukker et al., 2002). Differentiated derivatives of the ES cells remain negative for MHC II but show some increase in MHC Class I that is further up-regulated by exposure to interferon. These observations gave rise to the natural assumption that ES cells and their differentiated progeny would be subjected to rejection based on MHC mismatches, and led to a search for strategies to induce immunological tolerance in recipients of transplanted cells derived from ES lines (Drukker, 2004). However, it was observed that ES cells in the mouse and comparable stem cells from the inner cell mass of the embryo in the rat could be transplanted successfully in immune competent animals despite mismatches at the MHC loci. Furthermore, rodent ES cells may be able to induce immune tolerance in the recipient animals (Fandrich et al., 2002). Even more remarkably, human ES cells and differentiated derivatives were not rejected by immune competent mice in vivo, nor did they stimulate an immune response in vitro by human T-lymphocytes specific for mismatched MHC. Rather, the human cells appeared to inhibit the T-cell response (Li et al., 2004). An independent study using mice with a “humanized” immune system confirmed a very low T-cell response to human ES cells and differentiated derivatives (Drukker et al., 2006). MSC from bone marrow and their differentiated derivatives also have been shown both to escape an allogeneic immune response and to possess immunomodulatory activity to block such a response (Bartholomew et al., 2002; Le Blanc, 2003; Potian et al., 2003; Aggarwal and Pittenger, 2005). The effect likewise is observed with MSC isolated from adipose tissue (Puissant et al., 2005). The successful therapeutic use of allogeneic MSC has been confirmed in animal models (Arinzeh et al., 2003; De Kok et al., 2003). Therefore, beyond the application of MSC as regenerative cells, it is possible that they could be employed to induce immune tolerance to grafts of other cell types. The mechanisms underlying the immunodulatory properties of MSC are under active investigation and understanding them may have profound impact on regenerative medicine (Plumas et al., 2005; Krampera et al., 2006; Sotiropoulou et al., 2006). Other stem cell populations should be examined for their ability to escape and/or modulate an allogeneic immune response. While it is important to exercise caution in interpreting the laboratory results and in designing clinical trials, there is some reason to hope that the use of allogeneic stem cell-based bioengineered products will not necessarily imply the need for life-long treatment with immunosuppressive drugs. In the first FDA-approved clinical trial of allogeneic human neural stem cells, in children with a Neural Ceroid Lipofuscinosis disorder known as Batten disease (Taupin, 2006), immunosuppressive therapy will be utilized for the initial year after cell implantation and then reevaluated. Banking of stem cells for future therapeutic use extends possibilities both for autologous and allogeneic therapy paradigms, even if it turns out that histocompatibility matching is important for stem cell-based therapies. Amniocentesis specimens, placenta, and cord blood represent sources from which highly multipotent adult stem cells can be obtained and typed with minimal invasiveness. Prospective parents could opt for collection and cryopreservation of such cells for future use by their children in the event of medical need. Furthermore, collection and typing of a sufficient number of samples (ca. 100,000 for the US population) to permit nearly perfect histocompatibility matching between unrelated donors and recipients would be readily achieved. Similarly, collection and banking of cells from adult adipose tissue appears straightforward. Although it would entail a greater level of effort and could be politically controversial, it also might be feasible to prepare and bank a relatively large set of human ES lines to facilitate histocompatibility matching. One recent study suggests that a surprisingly modest number of banked lines or specimens could provide substantial ability to match donor cells to recipients (Taylor et al., 2005). Taken together with the low immunogenicity of certain stem cells, these results support the concept that allogeneic bioengineered products may not

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inevitably demand intensive immunosuppressive treatment, even if it proves impossible to develop general methods to induce selective immunological tolerance.

CONCLUSIONS Regenerative medicine is a highly interdisciplinary field. Future progress will continue to depend on synergies between advances in biology, chemistry, and engineering. Yet the development of new therapies may be rate limited by the need to identify and obtain stem and progenitor cells capable of yielding desired specialized cell types safely and efficiently. Exciting new work indicates unexpected paths that may provide novel solutions to two critical problems: sourcing of progenitors for a potentially unlimited range of specialized cell types and overcoming the need for life-long immunotherapy associated with allogeneic therapies.

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2 Fundamentals of Cell-Based Therapies Ross Tubo

INTRODUCTION Cell-based therapies have been proposed as a solution for a multitude of clinical problems ranging from structural repair of localized tissue damage to physiological restoration of systemic defects (Green et al., 1979; Caplan et al., 1998; Li et al., 1998). The successful treatment of such varied unmet medical needs ultimately depends upon the ability of cells to respond to their environment and function in a clinically relevant manner. This represents one of the most simple, and yet most complex principles for cell-based therapies. Many factors contribute to deciding on the most appropriate cell-based therapy for any given patient. The clinical problem and type of the tissue repair desired are primary factors. Whether the repair tissue is to be permanent or temporary, structural or biological are important considerations. For instance, replacement of permanent structure may require an autologous cell therapy, while temporary restoration of biology may be better suited for allogeneic cells. Autologous cell-based therapies represent our best clinical success in terms of permanent structural repair, harnessing the intrinsic capabilities of patient-derived cells to repair their own damaged tissues (Peterson et al., 2000). Studies examining the potential for allogeneic somatic cells for restoration of biology have also been successfully completed, resulting in Food and Drug Administration (FDA) approval for use of three allogeneic tissue-engineered products (Lysaght and Hazlehurst, 2004). The potential for use of allogeneic stem cells for structural repair of biological correction remains the subject of vigorous debate and research (Rao and Civin, 2006). Our knowledge of cells and their interaction with extracellular matrices and biological factors have continued to grow during the past 20–25 years, with significant progress being made in the in vitro generation of threedimensional tissue-engineered constructs of skin, cartilage, and blood vessels. We have learned the importance of providing proper physical and biological context in order to elicit the desired cellular response. Understanding these interactions will continue to guide the future development of clinically useful engineered tissues or organs in the practice of regenerative medicine. RATIONALE FOR CELL-BASED THERAPIES The inability of most adult tissues to regenerate themselves following injury has led to the development of cellbased strategies for structural repair or restoration of tissue physiology. Moreover, our ability to culture just about any somatic cell type has made it possible to consider the development of cell-specific culture systems for rapid proliferative expansion of such cells to treat previously unmet medical needs.

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Cell-based therapies generally fall into two main categories: (1) autologous cells for permanent structural repair and (2) allogeneic cells for short-term structural repair or restoration of physiological function. Autologous cells are derived from the patient to be treated, while allogeneic cells are derived from a donor. Allogeneic cell therapies developed in the past include cultured dermal fibroblasts and keratinocytes as dermal/epidermal constructs for the repair of cutaneous wounds (Parenteau et al., 2000), cultured kidney epithelia for renal assist (Humes and Szczypka, 2004), hepatocytes for liver function (Chan et al., 2004), pancreatic islets for diabetes (Ryan et al., 2002), and hematopoietic stem cells for bone marrow transplantation and immune reconstitution in leukemia and other cancers. Structural repair using autologous cells seems to be the most straightforward type of cell-based therapy, where the role of the cells is to produce a permanent repair tissue having the structural characteristics of the tissue from which they were derived. Allogeneic cells are expected to elicit a physiological response from the host by the transient production of tissue stimulatory molecules, which alters host disease biology resulting in restoration of physiological function. Use of allogeneic cells for short-term physiological restoration or stimulation of host repair is slightly more complicated, given the potential for immunological rejection of donor cells. Lastly, long-term correction of physiology, as is necessary for replacing organ function, is clearly the most sophisticated and problematic therapy. Careful attention needs to be given to physical structure, biological function, and the immunological component for a successful cell therapy.

Autologous Cell-Based Therapies (Unmet Medical Need) The two earliest examples of successful cell-based therapies for structural repair are cultured autologous epidermal keratinocytes (Epicel) for permanent skin replacement in severe burns (Gallico et al., 1984) and cultured autologous articular chondrocytes (Carticel) for repair of a patient’s own damaged articular cartilage (Brittberg et al., 1994). These products represent the first, the second, and the only autologous cell-based therapies ever commercialized. Epicel, the First Autologous Cell Therapy Human epidermal keratinocytes (HEKs) or skin cells can be proliferatively expanded by culture on a mouse 3T3-fibroblast feeder layer under very specific conditions. Single cell suspensions of HEKs are prepared by enzymatic digestion of host skin tissue and placed in monolayer culture on the feeder layer. The feeder layer provides the appropriate physical niche and biological milieu for rapid expansion (Rollins et al., 1989). HEKs change their morphology and characteristic in vivo gene expression pattern when placed in vitro. This phenomenon, generically called dedifferentiation, is a process quite characteristic of any cell type subjected to proliferative expansion in vitro (Haudenschild et al., 2001). For HEKs, dedifferentiation is characterized by rapid change in cellular morphology, increased cell proliferation, and decreased expression of keratins normally found in epidermis with increased expression of keratins found in proliferating cells (Lersch et al., 1989). When propagated HEKs are subsequently applied to the host, they sense their environment and respond by “redifferentiating,” expressing genes and proteins characteristic of HEKs found in skin. When applied to the patient, Epicel grafts have the appearance of a patchwork quilt. The grafts are quite fragile, being only 2–3 HEK cell layers thick (Figure 2.1). As such, they are very sensitive to microbial infection and physical manipulation. The nascent epithelial tissue attaches to the wound bed and further redifferentiates, having three to four differentiated layers of epidermis within about 7–10 days. Over time the epithelium develops into a fully functional epidermis and modulates the development of a neo-dermis or new dermis having all the histological hallmarks of a fully functional dermal–epidermal junction with rete ridges within a year (Compton et al., 1993).

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Figure 2.1 Confluent cultured autologous human epidermal keratinocytes are affixed to petrolatum gauze (left) prior to shipment and subsequent application to patient. The nascent epithelial tissue is only 2–3 cell layers thick, as shown in the cross-section of graft on the right (hematoxylin/eosin stained).

Carticel, the Second Autologous Cell Therapy The ability of cells to dedifferentiate, proliferate, and then redifferentiate and express a mature phenotype is central to the success of autologous cell therapy. As with Epicel, the paradigm used for Carticel is similar to that for any other autologous cell therapy. Cells are isolated by enzymatic digestion of a sample of the patient’s own tissue, subjected to proliferative expansion in cell culture, and then returned to the patient for treatment (Figure 2.2). Cells are cultured under conditions designed to increase the number of cells in a timely fashion while maintaining their ultimate ability to re-express a differentiated phenotype. Maintenance of this functional ability is required for proper tissue function in vivo. Cultured human autologous chondrocytes (HACs), delivered as a cell suspension underneath a periosteal patch to the subchondral bone surface of a localized or focal defect in articular cartilage, will redifferentiate and express protein and proteoglycan consistent with hyaline-like articular cartilage tissue (Brittberg et al., 1994). Human articular chondrocytes in tissue normally produce hyaline articular cartilage comprised of type II collagen and aggrecan in articular cartilage tissue. When isolated from tissue and placed in monolayer culture, expression of hyaline cartilage-specific genes is down-regulated (Haudenschild et al., 2001) and characterized by decreased expression of type II collagen and aggrecan, increased expression of type I collagen and versican, and subsequent cell proliferation. Once the cultured cells are returned to the environment of the knee joint, for example, they read the biological cues from the host extracellular matrix and growth factor milieu and redifferentiate, expressing genes more consistent with hyaline tissue (Brittberg et al., 1994). Epicel and Carticel represent life-saving and life-changing autologous cell-based biological solutions for which there were previously no treatments available. Epicel is used as a life-saving treatment for catastrophic burns of greater than 75% of total body surface area. Carticel is a life-changing treatment for repair of damaged articular cartilage and restoration of joint function. Autologous Structure – ACG – The Challenge of In Vitro Structure Related to In Vivo Function An ongoing challenge in autologous cell therapy is the development of “ready to use”tissue-engineered constructs for tissue replacement or repair. This problem revolves around the production of enough tissue architecture

Fundamentals of Cell-Based Therapies

Periosteal flap taken from medial tibia

Periosteal flap sutured over lesion

Lesion

Biopsy of healthy cartilage

Injection of cultured chondrocytes under flap into lesion

Enzymatic digestion

Cultivation for 11–21 days (10-fold increase in number of cells)

Trypsin treatment

Suspension of 2.6  106  5  106 cells

Figure 2.2 Chondrocyte transplantation in the right femoral condyle. The distal part of the femur and proximal part of the tibia are shown. Cells were isolated following enzymatic digestion of normal tissue. Cells were cultured in cell-specific media to increase the number of cells for subsequent administration to the patient (reprinted from Brittberg et al., 1994, with permission).

in vitro to allow for immediate and appropriate function in vivo. Sometimes the structure of the nascent tissue can adversely affect in vivo function. For example, articular chondrocytes cultured under conditions of high density in the presence of TGF-beta will produce cartilaginous tissue having nearly all the histological hallmarks of hyaline cartilage (Peel et al., 1998). However, when this three-dimensional tissue-engineered construct, composed of cells, extracellular matrix, and factors, is placed in a cartilage defect it does not heal. The tissue developed in vitro does not permit integration of the repair cartilage within the damaged host tissue. This is in contrast to placing a single-cell suspension in the defect without extracellular matrix, as is done in the Carticel procedure, where the “undifferentiated” cells attached to the bone redifferentiate in such a way so as to provide a better opportunity for integration of the nascent cartilage to the host tissue (Shortkroff et al., 1996). More recently, it has been reported that mesenchymal stem cell (MSC) constructs comprising “sheets” of cells, similar to the sheets obtained in HEK culture, have been used to successfully treat damaged myocardium (Miyahara et al., 2006). These adipose-derived MSCs reportedly differentiate into cardiomyocytes and vascular endothelial cells. When transplanted to the myocardium as a cultured sheet of cells they reportedly reversed cardiac wall thinning in the scar area and improved cardiac function in rats with myocardial infarction (Miyahara et al., 2006). Thus, some nascent structure may permit cell delivery without interfering with beneficial cellular function. The physical organization of cells, whether cell suspension, sheet, or three-dimensional construct, remains an important consideration for developing a cell-based therapy. Cultured cells respond to varied extracellular matrix and growth factor signals by producing varied extracellular matrix proteins themselves (Wakitani et al., 1989; Ben-Yishay et al., 1992; Solchaga et al., 2005).

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Given that cells respond differently to different environments, whether grown on or in materials such as collagen, polylactide/polygalactide (PLA/PGA), hyaluronic acid, or other polymers, in the presence of varied growth factors and morphogens, understanding the cellular response to these agents at the molecular level is critical to the development of successful cell-based products (Lutolf and Hubbell, 2005; Lisignoli et al., 2006). Restoration of Physiology Using Tissue-Engineered Organ Equivalents While our understanding of the interactions of cells, extracellular matrices, and biological factors in some ways seems to be in its infancy, tremendous progress has been made in the development of ex vivo organ constructs for renal assist (Humes and Szczypka, 2004) and bladder function (Atala, 2004). The renal assist device makes use of the intrinsic ability of cells from freshly digested kidney tissue to assemble spontaneously in a threedimensional cartridge, through which the patient blood is then passed. This ex vivo dialysis system is currently in a clinical trial. Clinical studies have also been conducted using a more traditional tissue-engineering approach for bladder replacement. Small biopsies of bladder tissue are obtained from the patient requiring bladder replacement. The epithelial and smooth muscle cells are cultivated separately, loaded sequentially onto a three-dimensional biodegradable construct, cultured for a brief period of time and transplanted. The results are truly astonishing and represent the first successful functional replacement of an organ using an “organ” engineered in vitro (Atala, 2004). Allogeneic “Ready-to-Use” Cell Therapies Two primary reasons have led investigators away from autologous cells toward allogeneic. First, some clinical indications do not require permanent survival of the applied cells, but rather temporary production of a biological agent that will restore host tissue function. Second, providing a patient with their own cells is inherently expensive due to logistical and manufacturing issues. Autologous cells require several weeks for the isolation, propagation, and return of cells to the patient. Delivery of functional cells providing more immediate clinical benefit is the goal of allogeneic or donor-derived cell therapies. Since allogeneic cells are likely to be rejected immunologically, they are likely better suited for indications not requiring permanent survival of the applied cells, but rather a temporary production of a biological agent to restore host tissue function. Dermal ulcers are small non-healing cutaneous wounds (10–50 cm2) which can be induced to heal by covering them with allogeneic skin wound dressings (Parenteau et al., 2000; Metcalfe and Ferguson, 2005). Dermal ulcers have been treated with a temporary epithelial dressing (allogeneic HEK, Acticel) or living-skin equivalents (Dermagraft, Apligraft). Each of these wound-healing dressings was derived from cells isolated from the donor tissue. These allogeneic cells can be propagated under the same conditions as autologous cells, but the expectation for their clinical use is for the temporary covering of cutaneous wounds to facilitate their healing. Allogeneic cells are intended to be “ready to use” by definition. Somatic cells – The logistical difficulties of providing a patient with their own cells and the inherent expense of the procedure would be significantly reduced by using allogeneic donor cells. Since each autologous sample is treated as its own manufacturing lot, it must be subjected to individualized culture, quality control testing, and preparation for delivery to a patient. Allogeneic cell preparation would allow for bulk quality control and manufacturing of one batch of cells to treat multiple patients, thus reducing expense. Furthermore, one batch of cells may be used to treat more than one clinical indication. Stem cells, the new frontier – Perhaps the cell type which has captivated the most attention from both scientists and lay people are stem cells. Stem cells fall broadly into two categories: embryonic or adult tissue derived. Embryonic stem (ES) cells are derived from the inner cell mass of developing embryos, whereas adult stem cells have been derived from a variety of adult tissue sources including bone marrow, dermis, adipose

Fundamentals of Cell-Based Therapies

Figure 2.3 Adult bone-marrow-derived MSCs were cultured under conditions to promote differentiation to the muscle (left), neural (middle), and cartilage (right) lineages. MSCs cultured in the presence of low serum formed myotubes, while those cultured in the presence of forskolin or TGF-beta differentiated into neural (nestin positive) or cartilage cells (type II collagen positive), respectively. tissue, and others (Pittenger et al., 1999; Jiang et al., 2002; Gimble and Guilak, 2003; Verfaillie et al., 2003; Bartsch et al., 2005). The bone marrow provides an attractive source of easily accessible adult pluripotent stem cells. The specialized microenvironment within the connective tissue framework of adult bone marrow supports the existence of at least two distinct populations of stem cells: one hematopoietic and the other mesenchymal. Hematopoietic stem cells (HSCs) in the adult ultimately give rise to all components of the immune and blood systems, while MSCs have the potential to give rise to cells of varied lineages, including bone, cartilage, and adipose tissues. The MSC population can be isolated from the bone marrow and expanded in culture in the absence of differentiation for at least 30–40 population doublings (Lodie et al., 2002). Even after expansion, MSCs can still differentiate to cells of multiple lineages (Bruder et al., 1997). Because MSCs have been shown to give rise to adipocytes, osteoblasts, chondrocytes, myoblasts, neurons, and other cell types (Figure 2.3), they are an intriguing alternative source of cells for cellular replacement therapies. ES cells have also been shown to exhibit pluripotent differentiation potential in vitro and in vivo (Schuldiner et al., 2000; Stojkovic et al., 2004). ES cells can spontaneously differentiate in culture into a layer of beating myocardium (He et al., 2003). These kinds of studies demonstrate the tremendous potential for ES cells; however, the exact culture conditions required to reproducibly induce ES cell differentiation in a controllable fashion remains the subject of intense study. Similarly, undifferentiated ES cells spontaneously form teratomas when injected subcutaneously in immune compromised mice (Przyborski, 2005). Histological analysis of ES cell implants reveals that tissue of cardiac, neural, and other tissue lineages spontaneously originate from the same population transplanted ES cells, again illustrating the tremendous differentiation potential of ES cells, and highlighting the need for further study to determine precise control of differentiation. Some investigators are engineering their ES cells to express conditional suicide genes to reduce the risk of inappropriate ES cell differentiation. Several recent papers suggest that the utility of adult stem cells may not be limited to in vitro differentiation for direct cell replacement of damaged tissues. Recent evidence suggests that the adult murine bone marrow cells possess the intrinsic capability to differentiate into β-cells after total bone marrow transplantation in nondiabetic animals (Ianus et al., 2003). In addition, transplantation of bone marrow (Zorina et al., 2003) and bone-marrow-derived stem cells was shown to activate endogenous tissue regeneration, specifically β-cell regeneration in the pancreas (Hess et al., 2003). There is also an emerging evidence which points to the transplant of bone-marrow-derived stem cells as having additional benefits including recruitment of endogenous stem cells (Kocher et al., 2001), vascularization of damaged tissue (Rafii and Lyden, 2003), and, as discussed further below, immune transplant tolerance (Bartholomew et al., 2002).

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Intramyocardial injection of bone-marrow-derived MSCs overexpressing Akt (Akt-MSCs) inhibits ventricular remodeling and restores cardiac function measured 2 weeks after myocardial infarction (Gnecchi et al., 2006). When injected into infarcted hearts, the Akt-MSC-conditioned medium significantly limits infarct size and improves ventricular function relative to controls. Support to the paracrine hypothesis is provided by data showing that several genes, coding for factors (VEGF, FGF-2, HGF, IGF-I, and TB4) that are potential mediators of the effects exerted by the Akt-MSC-conditioned medium, are significantly up-regulated in the Akt-MSCs, particularly in response to hypoxia. Taken together, our data support Akt-MSC-mediated paracrine mechanisms of myocardial protection and functional improvement. Immunosuppressive Properties of MSCs MSCs have been reported to be immunomodulatory both in vitro and in vivo. They express very low levels of co-stimulatory molecules and HLA Class I on their cell surface and lack HLA Class II expression (Devine and Hoffamn, 2000; Di Nicola et al., 2002). Class II expression did increase upon MSC differentiation. The immunophenotype of MSCs suggested that they may play a role in modulating T-cell proliferation and the immune response. MSCs have been shown to inhibit a mixed lymphocyte reaction (MLR) using purified CD3 T-cells and third party dendritic cells as antigen presenting cells (Tse et al., 2003). Both autologous and allogeneic MSCs suppress lymphocyte proliferation (Di Nicola et al., 2002). It has been postulated that MSCs may suppress T-cell proliferation by several mechanisms: secretion of growth factors, such as TGF-beta or HGF, suppression of pro-inflammatory (TH-1) cytokines, stimulation of anti-inflammatory (TH-2-type) cytokines, and up-regulation of pro-apoptotic cell surface molecules (Bartholomew et al., 2002; Di Nicola et al., 2002). Studies are ongoing to further elucidate the underlying mechanisms for MSC-mediated suppression of T-lymphocyte proliferation. It has been reported that intravenously administered allogeneic MSCs are not rejected in a baboon model due to lack of immune recognition. Furthermore, bone marrow transplantation of baboon MSCs into MHCmismatched recipients prior to a third party skin graft led to prolonged graft survival (Bartholomew et al., 2002). Taken together, these data suggest that MSCs not only possess immunosuppressive properties that inhibit T-cell proliferation in vitro, but they also have immunomodulating properties which may enhance graft survival in vivo. MSCs also reduced the incidence and severity of graft-versus-host disease (GVHD) during allogeneic transplantation and although the mechanisms remain to be elucidated, the data offer insight into the potential use of MSCs for induction of tolerance for reduction of GVHD, rejection, and modulation of inflammation (Aggarwal and Pittenger, 2005). Adult bone-marrow-derived MSCs appear to offer several advantages over autologous cell therapies and even ES cells. First, MSCs exhibit multi-potential differentiation in a well-controlled, predictable fashion, in contrast to ES cells. Second, the fact that they appear to down-modulate the host (recipient) immune response (GVHD) may permit their persistence for the longer term, similar to autologous cells. Although this is not necessarily a functional advantage of MSCs over autologous cells, the production of multiple treatment doses from a single donor source is quite attractive from a manufacturing and quality control perspective, thereby reducing costs associated with personalized medicine. Commercialization of a Cell Therapy Commercialization of cellular therapies is not easy. Autologous cells, while clinically successful, may not be commercially successful, due in part to the fact that they are logistically difficult and inherently expensive to produce. Clinical evaluation of such therapies is complicated and time consuming. Moreover, the regulatory and reimbursement issues can be very challenging. That having been said, cell therapies can significantly enhance the quality of human healthcare for serious unmet medical needs.

Fundamentals of Cell-Based Therapies

The appeal of allogeneic stem cells is obvious: one cell source for multiple indications; potential for an off-the-shelf product; improved quality control; and reduced cost of goods. However, before we get too carried away with the “promise” of stem cells, we need to do a reality check. We need to apply the same fundamental principles to stem cells that were applied to autologous cell therapies. Many questions remain to be addressed before the potential of stem cell therapy can be realized, such as: Can the cells be routinely isolated and propagated? Can the cells terminally differentiate into the cell type of interest? In vitro and in vivo? What is their potency and purity? How long do the cells persist in vivo? Can the purity of the expanded cells be established prior to shipment? Having the answers to these questions is critical for the successful commercialization of stem cell therapy. Ensuring Production of the Best Quality Cell Therapy Products Measurement of identity and functionality of cells following proliferative expansion are the two key features of ensuring the best possible quality of cell-based products, autologous or allogeneic. Cell surface makers can be used to assess identity and purity of the expanded cell population. Differentiation assays can be used to assess functionality of the cells in vitro. In vivo studies are required to determine differentiation and persistence in vivo. The principle that cultured cells can dedifferentiate and undergo proliferative expansion in vitro, and then redifferentiate when placed in vivo is central to the success of cell-based therapies. Stem cells may propagate in a multi-potent state and then differentiate in vivo. Our ability to assay for this activity in vitro, as a matter of “quality control,” is critical to the ultimate success of a cell-based therapy for tissue replacement or repair in vivo. Given that cells in culture respond differently to varied culture conditions and environmental cues (Haudenschild et al., 2001; Lodie et al., 2002; Solchaga et al., 2005), it is important to confirm that the cell types being propagated are indeed the desired cell type and that they are capable of the intended function.

CONCLUSIONS Inadequate therapies to repair injured tissue, replace failing organs, and restore structural and metabolic functions remain a driving force behind the demand for cell-based therapies. Cells represent a “lowest common denominator” of sorts for cell-based therapies; their numbers are expandable, they are programmed by the environment within which they find themselves to respond and produce a biological response. The challenge is to harness the tremendous potential within this tiny unit, ultimately providing the proper structure and biological function necessary for successful treatment of the clinical problem at hand. Generally speaking cell-based therapies fall into two broad categories of use: (1) cells for permanent structural repair or replacement (e.g. cultured keratinocytes as skin replacement, chondrocytes for repair of cartilage or visco-uretal reflux) and (2) cells for correction of a physiological or metabolic problem. Understanding the nature of the problem you are trying to treat and role that the cell may play in solving the problem is critical to developing a successful cell therapy. Issues to be considered include whether the cells are for structural replacement or restoration of metabolism. If structural replacement, autologous cells are likely the cell of choice. If the goal is the correction of metabolism, the length of time required to see physiological benefit and subsequent immunological status may influence the source of cells to be used. Scientists in regenerative medicine have strived to understand the interaction of cells, extracellular matrices, and biological factors as they have endeavored to develop tissue-engineered constructs for repair and replacement of damaged tissue. Understanding how to produce a “simple” functional tissue in vitro, by harnessing our knowledge of these building blocks remains a very complex and yet exciting problem for us to solve. While our understanding of the mechanisms underlying the interactions of cells, extracellular matrices, and biological factors continues to grow, we continue to take advantage of the “intrinsic knowledge” that the cell retains to accomplish the goal of tissue repair.

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ACKNOWLEDGMENTS I would like to thank Drs. Tracey Lodie, Ajeeta Dash, and Michelle Youd for their insightful review of this document, and Ms. Maureen Swartz for her administrative contributions.

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Humes, H.D. and Szczypka, M.S. (2004). Advances in cell therapy for renal failure. Transpl. Immunol. 12: 219–227. Ianus, A., Holz, G.G., Theise, N.D. and Hussain, M.A. (2003). In vivo derivation of glucose-competent pancreatic endocrine cells from bone marrow without evidence of cell fusion. J. Clin. Invest. 111: 843–850. Jiang, Y., Vaessen, B., Lenvik, T., Blackstad, M., Reyes, M. and Verfaillie, C.M. (2002). Multipotent progenitor cells can be isolated from postnatal murine bone marrow, muscle, and brain. Exp. Hematol. 30: 896–904. Kocher, A.A., Schuster, M.D., Szabolcs, M.J., Takuma, S., Burkhoff, D., Wang, J., Homma, S., Edwards, N.M. and Itescu, S. (2001). Neovascularization of ischemic myocardium by human bone-marrow-derived angioblasts prevents cardiomyocyte apoptosis, reduces remodeling and improves cardiac function. Nat. Med. 7: 430–436. Lersch, R., Stellmach, V., Stocks, C., Giudice, G. and Fuchs, E. (1989). Isolation, sequence, and expression of a human keratin K5 gene: transcriptional regulation of keratins and insights into pairwise control. Mol. Cell. Biol. 9: 3685–3697. Li, R.K., Yau, T.M., Sakai, T., Mickle, D.A. and Weisel, R.D. (1998). Cell therapy to repair broken hearts. Can. J. Card. 14: 735–744. Lisignoli, G., Cristino, S., Piacentini, A., Cavallo, C., Caplan, A.I. and Facchini, A. (2006). Hyaluronan-based polymer scaffold modulates the expression of inflammatory and degradative factors in mesenchymal stem cells: involvement of Cd44 and Cd54. J. Cell. Physiol. 207: 364–373. Lodie, T.A., Blickarz, C.E., Devarakonda, T.J., He, C., Dash, A.B., Clarke, J., Gleneck, K., Shihabuddin, L. and Tubo, R. (2002). Systematic analysis of reportedly distinct populations of multipotent bone marrow-derived stem cells reveals a lack of distinction. Tissue Eng. 8: 739–751. Lutolf, M.P. and Hubbell, J.A. (2005). Synthetic biomaterials as instructive extracellular microenvironments for morphogenesis in tissue engineering. Nat. Biotechnol. 23: 47–55. Lysaght, M.J. and Hazlehurst, A.L. (2004). Tissue engineering: the end of the beginning. Tissue Eng. 10: 309–320. Metcalfe, A.D. and Ferguson, M.W. (2005). Harnessing wound healing and regeneration for tissue engineering. Biochem. Soc. Trans. 33: 413–417. Miyahara, Y., Nagaya, N., Kataoka, M., Yanagawa, B., Tanaka, K., Hao, H., Ishino, K., Ishida, H., Shimizu, T., Kangawa, K., Sano, S., Okano, T., Kitamura, S. and Mori, H. (2006). Monolayered mesenchymal stem cells repair scarred myocardium after myocardial infarction. Nat. Med. 12: 459–465. Parenteau, N.L., Hardin-Young, J. and Ross, R.N. (2000). Skin. In: Lanza, R., Langer, R. and Vacanti, J. (eds.), Principles of Tissue Engineering. San Diego: Academic Press, pp. 879–890. Peel, S.A.F., Chen, H., Renlund, R., Badylak, S. F. and Kandel, R.A. (1998). Formation of a SIS-cartilage composite graft in vitro and its use in the repair of articular cartilage defects. Tissue Eng. 143–155. Peterson, L., Minas, T., Brittberg, M., Nilsson, A., Sjogren-Jansson, E. and Lindahl, A. (2000). Two- to 9-year outcome after autologous chondrocyte transplantation of the knee. Clin. Orthop. Relat. Res. 374: 212–234. Pittenger, M.F., Mackay, A.M., Beck, S.C., Jaiswal, R.K., Douglas, R., Mosca, J.D., Moorman, M.A., Simonetti, D.W., Craig, S. and Marshak, D.R. (1999). Multilineage potential of adult human mesenchymal stem cells. Science 284: 143–147. Przyborski, S.A. (2005). Differentiation of human embryonic stem cells after transplantation in immune-deficient mice. Stem Cells 23: 1242–1250. Rafii, S. and Lyden, D. (2003). Therapeutic stem and progenitor cell transplantation for organ vascularization and regeneration. Nat. Med. 9: 702–712. Rao, M.S. and Civin, C.I. (2006). How many human embryonic stem cell lines are sufficient? A US perspective. Stem Cells March 16 (Epub ahead of print). Rollins, B.J., O’Connell, T.M., Bennett, G., Burton, L.E., Stiles, C.D. and Rheinwald, J.G. (1989). Environment-dependent growth inhibition of human epidermal keratinocytes by recombinant human transforming growth factor-beta. J. Cell. Physiol. 139: 455–462. Ryan, E.A., Lakey, J.R., Paty, B.W., Imes, S., Korbutt, G.S., Kneteman, N.M., Bigam, D., Rajotte, R.V. and Shapiro, A.M. (2002). Successful islet transplantation: continued insulin reserve provides long-term glycemic control. Diabetes 51: 2148–2157. Schuldiner, M., Yanuka, O., Itskovitz-Eldor, J., Melton, D.A. and Benvenisty, N. (2000). Effects of eight growth factors on the differentiation of cells derived from human embryonic stem cells. Proc. Natl Acad. Sci. 97: 11307–11312.

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Shortkroff, S., Barone, L., Hsu, H.-P., Wrenn, C., Gagne, T., Chi, T., Breinan, H., Minas, T., Sledge, C.B., Tubo, R. and Spector, M. (1996). Healing of chondral and osteochondral defects in a canine model: the role of cultured chondrocytes in regeneration of articular cartilage. Biomaterials 17: 147–154. Solchaga, L.A., Temenoff, J.S., Gao, J., Mikos, A.G., Caplan, A.I. and Goldberg, V.M. (2005). Repair of osteochondral defects with hyaluronan- and polyester-based scaffolds. Osteoarthritis Cartilage 13: 297–309. Stojkovic, M., Lako, M., Strachan, T. and Murdoch, A. (2004). Derivation, growth and applications of human embryonic stem cells. Reproduction 128: 259–267. Tse, W.T., Pendleton, J.D., Beyer, W.M., Egalka, M.C. and Guinan, E.C. (2003). Suppression of allogeneic T-cell proliferation by human marrow stromal cells: implications in transplantation Transplantation 75: 389–397. Verfaillie, C.M., Schwartz, R., Reyes, M. and Jiang, Y. (2003) Unexpected potential of adult stem cells. Ann. NY Acad. Sci. 996: 231–234. Wakitani, S., Kimura, T., Hirooka, A., Ochi, T., Yoneda, M., Owaki, H., Ono, K. and Yasui, N. (1989). Repair of rabbit articular surfaces with allografts of chondrocytes embedded in collagen gels. J. Jpn Ortho. Assoc. 63: 529–538. Zorina, T.D., Subbotin, V.M., Bertera, S., Alexander, A.M., Haluszczak, C., Gambrell, B., Bottino, R., Styche, A.J. and Trucco, M. (2003). Recovery of the endogenous beta cell function in the NOD model of autoimmune diabetes. Stem Cells 21: 377–388.

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3 Stem Cell Research T. Ahsan, A.M. Doyle, and R.M. Nerem

INTRODUCTION Regenerative medicine is an emerging branch of medicine whose goal is to restore organ and/or tissue function using a biological approach. A growing crisis in organ transplantation and an aging population have driven a search for new and alternative therapies. There currently are approximately 90,000 patients on the US transplant waiting list. Despite growing numbers of donors, the availability of suitable organs is still insufficient. This discrepancy is only likely to increase during the next 25 years, given that the population of those 65 years and older is projected by the US Census Bureau to more than double. Recent advances in stem cell technology have shown great promise and propelled regenerative medicine to the forefront of both scientific research and public consciousness. While some believe the therapeutic potential of stem cells has been overstated in the media, an analysis of the potential benefits of stem cell-based therapies indicates that 128 million people in the United States alone may benefit, with the largest impact on patients with cardiovascular disorders, autoimmune diseases, and diabetes (Figure 3.1) (Perry, 2000). The enthusiasm surrounding stem cells is related in part to their potential to treat a broad range of clinical pathologies. Some identified stem cell targets, such as neurological diseases, spinal cord injuries, diabetes, and cardiovascular diseases, currently have few accepted treatments or no cures. In other conditions, such as bone fracture healing or cartilage repair, stem cells may improve upon therapies currently in use. Stem cells may change the very nature of medicine: they have the potential to address the cell sourcing issue of tissue

Cardiovascular Autoimmune Diabetes Osteoporosis Cancer Alzheimer’s Parkinson’s

Total: 128 million

Other 0

20 40 Millions of People

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Figure 3.1 Persons in the United States affected by diseases or injuries that may be helped by stem cell research (Perry, 2000).

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engineering and to customize therapies for individual patients. As a self-renewing source of allogeneic cells, stem cells make off-the-shelf products a more probable and closer reality. Through either autologous adult stem cells or somatic nuclear transfer, cell-based therapies can be genetically matched to each patient, addressing immunocompatibility and disease transfer concerns. Stem cells in regenerative medicine can serve not only as cellular machinery, but also as gene delivery vehicles and systems to promote further understanding of development. Due to their self-renewing, proliferating, differentiating, and distribution potential in vivo, stem cells are a natural choice as gene delivery vehicles (Lemoine, 2002). Stem cells grown in vitro or implanted in animals can serve as model systems to study many basic science questions. Our understanding of the mechanisms that govern development is expanded using the spontaneous differentiation of embryonic stem cells (ESCs) in embryoid bodies, the directed differentiation of adult stem cells in response to chemical and/or physical cues, and the homing and engraftment of stem cells in animals. Additionally, in vitro models of development provide a unique opportunity to study mutations that would otherwise be lethal in vivo. As a result of increased understanding from the use of these various stem cell model systems, strategies may be developed that focus on preventative medicine. The potential of stem cells and regenerative medicine is too vast to cover in its entirety in a single chapter. As a result, we have largely focused this discussion on an overview of genetically unmodified human stem cells and their current status in clinical applications of regenerative medicine. It is important to note, however, that the extensive amount of work done with animal stem cells and in animal models is not only a basis for human applications, but also indicates the long range potential of stem cells in regenerative medicine. Ultimately, the intent of this introductory chapter is to address the range of stem cell technology and leave to subsequent chapters the more exhaustive and in-depth analyses of specific stem cells and their applications. This chapter gives an overview of the different types of stem cells, the modes of stem cell modulation in vitro, the general strategies of regenerative medicine, and the role of stem cells in various clinical applications.

STEM CELLS A stem cell is an unspecialized cell that can both self-renew (reproduce itself) and differentiate into functional phenotypes. Stem cells can originate from embryonic, fetal, or adult tissue and are broadly categorized accordingly. ESCs are commonly derived from the inner cell mass of a blastocyst, an early (4–5 days) stage of the embryo. Embryonic germ cells (EGCs) are isolated from the gonadal ridge of a 5–10 week fetus. In particular, EGCs are derived from the primordial germ cells, which ultimately give rise to eggs or sperms in the adult. Adult stem cells differ from ESCs and EGCs in that they are found in tissues after birth, and to date, have been found to differentiate into a narrower range of cell types, primarily those phenotypes found in the originating tissue. A major value of stem cells in regenerative medicine is their potential to become different cell types. Our current understanding of differentiation is based on a hierarchical tree structure in which a few unspecialized stem cells branch to ultimately yield a larger number of mature cellular phenotypes (Figure 3.2). Stem cells divide to generate at least one daughter cell that retains the stem cell identity, resulting in a perpetuating population (Ho, 2005). They can also give rise to progenitors, or precursor cells, which typically differentiate into tissue-specific cell types and are only capable of symmetric division. Yet these progenitors play a major role in vivo that may be beneficial for cell-based therapies: symmetric division of rapidly proliferating progenitors allows exponential yield of terminally differentiated cells. Thus, as a system, this hierarchical structure allows for a small perpetual population of stem cells to give rise as needed to large numbers of differentiated cells. The hierarchical tree structure of differentiation is based on observations from developmental biology. Differentiation during embryogenesis begins with gastrulation, when cells separate into three structural

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Differentiation potential

Stem cells

Functional capacity

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Progenitor cells

Terminally differentiated cells

Figure 3.2 The hierarchical structure of differentiation. Stem cells become progenitors that yield terminally differentiated cells. Stem cells have the most differentiation potential, while differentiated cells have the greatest functional capacity.

layers: endoderm, mesoderm, and ectoderm. While together the three layers (or lineages) give rise to all the cells in the body, each layer proceeds through stages of differentiation to generate an independent subset of phenotypes. It has been found that the ectoderm includes skin and neural cells; the mesoderm includes cardiovascular, blood, and skeletal cells; and the endoderm includes cells of the gastrointestinal tract. Analogous to the developmental scheme, the hierarchical model is used to refer differentiation potential both in vivo and in vitro. A totipotent cell, such as the fertilized egg, is capable of differentiating not only into all three lineages, but also yields the extra-embryonic cells that support fetal development. Pluripotent cells, such as ESCs and EGCs, can differentiate into all three lineages. Committed to a specific lineage, adult stem cells are considered multipotent; they are able to form more than one cell type, but usually within the same lineage. Differentiation of cells from one lineage to another is referred to as stem cell plasticity or transdifferentiation. While some studies claim stem cell plasticity, possible alternative explanations make this topic controversial (Eisenberg and Eisenberg, 2003; Quesenberry et al., 2004; Wagers and Weissman, 2004; Lakshmipathy and Verfaillie, 2005). Therefore, it is possible that in future the hierarchical model of differentiation will be challenged and a new paradigm proposed. ESCs In 1981, Martin and Evans and Kaufman isolated and cultured pluripotent cells from the inner cell masses of mouse embryos. These key events in the mouse model were pivotal for the subsequent derivations in 1998 of the first human ESCs (Thomson et al., 1998; Reubinoff et al., 2000). Human ESCs have been defined to (a) be isolated from the inner cell mass of the blastocyst, (b) proliferate extensively in vitro (concomitantly expected to maintain high levels of Oct-4 expression, telomerase activity, and a normal karyotype), and (c) retain the potential to differentiate into cell types of all three lineages (Hoffman and Carpenter, 2005). Established human ESC lines were typically derived from embryos destined for destruction at in vitro fertilization clinics. To generate a single ESC line, the 30–34 cells of the inner cell mass of a pre-implantation

Stem Cell Research

blastocyst are removed and expanded in vitro. The number of human ESC lines is rapidly increasing worldwide, helping to advance the knowledge base related to these pluripotent cells. It is now known that the genomic expression of individual lines varies (Rao et al., 2004). Resultant characteristics of the cells, as well as differences in the overall efficiency of cell line isolation, likely depend on the quality of the embryo, its precise stage of development, and the means of cell isolation. Culture conditions for human ESCs have previously relied on xenogeneic components. The original human ESC lines were grown in medium supplemented with animal sera and/or maintained on mitotically inactivated mouse feeder layers. The use of xenogeneic components raises the concern of introducing nonhuman pathogens in clinical therapies. The currently available lines of human ESCs that have been exposed to animal contaminants are consequently unlikely to ever be used in future clinical applications. To address this concern, recent efforts have attempted to maintain ESCs on human feeder layers (Richards et al., 2002; Amit et al., 2003; Hovatta et al., 2003; Lee et al., 2005) and avoid animal sera-based medium supplements (Amit et al., 2004; Li et al., 2005b). Other efforts have focused on using growth factors together with protein substrates (Levenstein et al., 2005) or even synthetic polymers (Li et al., 2005a). While these adjusted conditions of culture have been shown to be somewhat effective, it is still not clear which specific mechanisms are critical to maintain ESCs undifferentiated. In any case, well-defined non-xenogeneic culture conditions will be critical in advancing human ESC-based therapies. The differentiation potential of human ESCs can be determined either in vivo or in vitro. In spontaneous differentiation models, undifferentiated cells are allowed to form three-dimensional (3D) cell clusters, which are assessed for the presence of expressed phenotypes. The in vivo model involves injecting cells into immunocompromised mice and analyzing the formed teratoma. An easier, yet still informative, in vitro model of differentiation consists of removing the human ESCs from the feeder layer and culturing them in suspension to form embryoid bodies. Spontaneous differentiation in in vivo and in vitro models may underestimate the number of phenotypes generated by pluripotent cells. Directed differentiation by controlling the chemical and/or mechanical environment may reveal a greater extent of the differentiation potential. In all of these models, cells are usually only qualitatively assessed for their potential to spontaneously differentiate into cells of ectoderm, mesoderm, and endoderm lineages. More quantitative techniques to assess lineage commitment, however, are needed to fully assess pluripotency. While much is known about the differentiation capabilities of mouse ESCs, the full potential of human ESCs is still being determined. The phenotypes derived from human ESCs are listed in Table 3.1. In general,

Table 3.1 Differentiated cell types derived from human embryonic stem cells Differentiation General Ectoderm Neuroprogenitors

References Itskovitz-Eldor et al. (2000); Schuldiner et al. (2000); Dvash et al. (2004) Carpenter et al. (2001); Reubinoff et al. (2001); Schuldiner et al. (2001); Park et al. (2004); Perrier et al. (2004); Schulz et al. (2004); Li et al. (2005a); Nistor et al. (2005)

Mesoderm Cardiomyocytes Hematopoietic progenitors Leukocytes Endothelial cells

Xu et al. (2002); Kehat et al. (2003) Kaufman et al. (2001); Chadwick et al. (2003); Vodyanik et al. (2005) Zhan et al. (2004) Levenberg et al. (2002)

Endoderm Insulin positive cells Hepatocyte-like cells

Assady et al. (2001); Segev et al. (2004) Rambhatla et al. (2003); Lavon et al. (2004)

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however, the categorization of phenotype is indistinct. Differentiated cells are often assessed based on specific markers, such as protein or gene expression. Unfortunately, the natural variability of cells and the span of certain markers across phenotypes necessitates that a panel of markers, rather than a single one, be assessed to converge on phenotype assignment. Characterization is even more complicated in stem cell studies, as unspecialized cells often simultaneously express markers of multiple phenotypes. EGCs EGCs are isolated from the gonadal ridge of a 5–10 week fetus. In particular, EGCs are derived from the primordial germ cells that in vivo give rise to eggs or sperms in the adult (Shamblott et al., 1998; Turnpenny et al., 2003). These cells have been shown to be pluripotent in vitro and were initially derived at a similar time as the ESCs, but there has been markedly less attention given to these EGCs. Although EGCs and ESCs seem to share certain characteristics, there are also intrinsic differences. EGCs are isolated from post-implantation fetal tissue recovered after termination of pregnancy. There are fewer opportunities to obtain tissue to isolate EGCs when contrasted with ESCs, but the derivation is considered easier and results in a propagated cell line more frequently (80% versus 15% of attempts) (Aflatoonian and Moore, 2005). There are still only a few EGC lines in existence, most of which are not readily available for the general scientific community to study. In part due to the limited number of investigators working with these cells, there are currently no standard procedures for derivation and propagation of these cells in vitro. Along with the fact that prolonged culture of EGCs is difficult even on mouse feeder layers (Shamblott et al., 1998), the EGC lines have yet to be well characterized (Aflatoonian and Moore, 2005). EGCs do differentiate using the embryoid body model, similar to ESCs. In that model, EGC pluripotency has been shown, as subpopulations express markers of various phenotypes, including neural, endothelial, muscle, and endodermal. These differentiating cells have then been isolated and expanded further in vitro (Shamblott et al., 2001) to generate more uniform populations of cells. As of yet, however, there have been no attempts to use directed differentiation to generate homogenous populations of differentiated cells. More extensive study of these cells, in terms of derivation, propagation, and differentiation, is needed before they can be considered a favored cell source for regenerative medicine applications. Adult Stem Cells Adult stem cells are those cells found in tissues after birth that are able to self-renew and yield differentiated cell types. Initially it was thought that adult stem cells were only located in a limited selection of organs and could differentiate into just those phenotypes found in the originating tissue. The field is still developing, however, and recent studies have identified stem cells in more tissues and indicate a greater range of potential than that originally believed. Already stem cells have been derived from human bone marrow (Edwards, 2004), blood (Ogawa, 1993; Asahara et al., 1997), brain (Steindler and Pincus, 2002), fat (Zuk et al., 2002), liver (Tosh and Strain, 2005), muscle (Alessandri et al., 2004), pancreas (Zulewski et al., 2001), and umbilical cord blood (Erices et al., 2000; Benito et al., 2004). As with many rapidly expanding fields, the use of non-standardized methods makes interpreting results from different investigators difficult, and this thus has led to controversy. Since adult stem cells are often a very small percentage of the total cells isolated from a given tissue, generating a pure population is difficult. In many cases different investigators use different means of isolating the stem cells from a given tissue. The question then arises whether the stem cells generated from the various techniques are identical or distinct stem cell populations. This difficulty is further exacerbated as these cells are commonly identified using a range of criteria, such as isolation procedure, morphology, protein expression, etc., leaving some question as to the defining characteristics of these stem cell populations.

Stem Cell Research

The potential to yield mature phenotypes is typically shown through either differentiation in vitro using biochemical cues or implantation in vivo in immunosuppressed mice. The lack of lineage tracing and clonal expansion in some studies has called into question whether observed phenotypes are due to the differentiation potential of a stem cell or to a heterogeneous initial population. As standardized protocols develop for adult stem cells, more rigorous criteria will develop for determining stem cell populations and their differentiation potential. There is a growing argument that all adult stem cells may have a signature expression profile. It is possible that self-renewing capabilities combined with multipotency, regardless of the cell origin, are associated with a set of characteristic properties. While such properties have not yet been determined, one candidate may be dye exclusion. When stained with Hoechst, some adult stem cells have been found to actively exclude the dye using transmembrane pumps. These cells have been coined “side population cells,” as they appear in a peripheral area when analyzed by flow cytometry using a UV laser. Originally identified in murine bone marrow (Goodell et al., 1996), the commonality of this functional property across adult stem cells has best been shown in the mouse model, where side population cells have been found in muscle, liver, lung, brain, kidney, heart, intestine, mammary tissue, and spleen (Asakura and Rudnicki, 2002). Expression of the ABCG2 protein, which plays a role in the transmembrane pump (Scharenberg et al., 2002), may be a convenient expression marker of this functional property. It is still unclear, however, which signature expressions, if any, are inherently associated with all adult stem cells. While adult stem cells may ultimately be derived from practically every tissue in the body, there is a subset, based on ease of isolation, availability, or potency, that is most likely to contribute to regenerative medicine. These stem cells, and the phenotypic lineages they have been shown to generate, are indicated in Table 3.2. Bone marrow- and blood-derived stem cells are fairly easy to isolate and have been the most thoroughly investigated. Both contain hematopoietic stem cells (HSCs) (Ogawa, 1993; Tao and Ma, 2003), which give rise to blood cells, and endothelial progenitor cells (EPCs) (Asahara et al., 1997; Kocher et al., 2001). Bone marrow additionally contains mesenchymal stem cells (MSCs) (Pittenger et al., 1999; Jiang et al., 2002), which have been shown to differentiate into mesodermal phenotypes, including orthopedic and vascular. The low yield of stem cells from marrow and blood motivates efforts to find alternative adult stem cell sources. HSCs and MSCs can also be derived from

Table 3.2 Differentiated cells derived from human adult stem cells Tissue source

Cell type

Derived cells

References

Blood

HSC EPC

Blood cells Endothelial cell

Ogawa (1993) Asahara et al. (1997)

Bone Marrow

HSC EPC MSC

Hepatocyte, blood cells Endothelial cell Adipocyte, cardiomyocyte, chondrocyte, endothelial cell, neuron, osteocyte, thymic cell

Alison et al. (2000); Tao and Ma (2003) Kocher et al. (2001) Pittenger et al. (1999); Liechty et al. (2000); Sanchez-Ramos et al. (2000); Woodbury et al. (2000); Jiang et al. (2002); Oswald et al. (2004)

Fat

PLA

Adipocyte, chondrocyte, myocyte, neural progenitor, osteocyte

Zuk et al. (2002); Ashjian et al. (2003); Huang, J.I., et al. (2004)

Umbilical Cord Blood

HSC MPC

Blood cells Adipocyte, endothelial cell, blood cells, osteoblast

Broxmeyer et al. (1989) Erices et al. (2000); Chiu et al. (2005)

PLA: processed lipoaspirate

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umbilical cord blood (Broxmeyer et al., 1989; Erices et al., 2000). As a widely available source of stem cells with extensive expansion capabilities in vitro, stored umbilical cord blood is considered an exciting resource for regenerative medicine applications (Chiu et al., 2005). One plentiful autologous adult stem cell source is fat. Lipoaspirate-derived stem cells have yet to be thoroughly investigated, but have already been shown to differentiate into multiple phenotypes (Zuk et al., 2002; Ashjian et al., 2003; Huang, J.I., et al., 2004). Overall, the proven differentiation potential of human adult stem cells is limited. Research in stem cell plasticity and animal adult stem cells, however, implies that the full potential of human adult stem cells is likely to be more extensive than has been currently shown. Issues in Stem Cell-Based Therapies Stem cells are attractive for use in cell-based therapies due to the very attributes that define them. Because they are self-renewing and can differentiate into mature cell types, in theory stem, cells can serve as a limitless supply of cells and a source for a wide range of phenotypes. In practice, however, each type of stem cell has its own advantages and disadvantages. ESCs and EGCs are similar in that they are highly proliferative and pluripotent, which serve as both advantages and disadvantages in cell-based therapies. For culture in vitro, their ability to generate the large number of cells often required for therapies, as well as their potential to yield whichever phenotype may be of interest, is considered beneficial. For implantation in vivo, however, the concern arises that these same attributes will either allow ESCs to proliferate limitlessly and form tumors or differentiate uncontrollably into undesirable cell phenotypes. Other current concerns relate to immunological issues. ESCs are commonly cultured with xenogeneic elements, which may induce an immune response or transfer cross-species pathogens. Additionally, ESCs by nature will be an allogeneic cell source, whose transplantation into a human patient would require lifelong immunosuppression. Some research addresses the immunorejection concerns. Chimeric studies indicate that immunoacceptance may be achieved by transplanting donor stem cells not only to the site of repair, but also to the bone marrow (Adams et al., 2003). The donor cells would then contribute to the hematopoietic and lymphatic systems and promote immunoacceptance. Conversely, a nuclear transfer technique may avoid immunorejection by genetically matching the implanted cells to the recipient. In this technique, the nuclear material from a somatic cell is inserted into an enucleated oocyte. This oocyte is induced to form a blastocyst, from which an ESC line is derived. It is then possible to generate an ESC line, with typical proliferative and pluripotent characteristics, that is genetically identical to the individual recipient of the cellular implant. Arguably the greatest hurdle for the use of ESCs in cell-based therapies is the ethical debate and the subsequent political, legal, and social consequences. ESC isolation from the inner cell mass of a blastocyst results in the destruction of the pre-implantation embryo. The crux of the ethical debate surrounds the destruction of an entity that would otherwise form a living human being. Recently published in the same issue of Nature were two proof-of-principle studies in mice for approaches that may circumvent this ethical concern. Lanza and colleagues showed that a single cell embryo biopsy could be used to generate an ESC line, leaving intact the developmental capacity of the embryo (Chung et al., 2005). In a separate approach, Meissner and Jaenisch (2005) modified the nuclear transfer technique to include a step that turns off the cdx2 gene, without which the blastocyst cannot implant on the uteral wall. The derived ESC line, later modified to restore the cdx2 gene, would then have been derived from an entity that never had the potential to form a human being. While scientists alone cannot resolve the ethical debate, it is clear from these scientific efforts that there are ongoing attempts to facilitate the translation of ESC research to medical advances. Adult stem cells are already used in some cell-based therapies, but are expected ultimately to be used in many more applications. Unlike ESCs and EGCs, adult stem cells are not mired in major ethical issues and

Stem Cell Research

allow for the use of autologous cells for individually customized therapeutic applications, avoiding some immunological concerns. The various types of adult stem cells share similar obstacles toward their use in therapies. Stem cells derived from adult tissues are usually very limited in number. Moreover, available adult stem cell numbers in most tissues decrease with age, over the same period when the need for those cells usually increases. The large numbers of cells usually required for therapies likely will drive the need to expand adult stem cells in vitro, where they have been found to be very slow growing in culture. The potential impact of adult stem cells in clinical applications is immense, so efforts to address the technical hurdles are ongoing. Both embryonic and adult stem cell sources are likely to have an impact on cell-based therapies in the future. The limitations discussed above range from ethical concerns to scientific challenges. Additional regulatory issues must also be addressed, though precautions will be fewer for autologous adult stem cells minimally manipulated ex vivo compared to allogeneic, potentially teratoma-forming, ESCs. Extensive ongoing research, however, indicates the confidence of both researchers and clinicians in our ability to overcome these obstacles and in the potential of stem cells to have a positive impact on clinical applications.

STEM CELL MODULATION IN VITRO Stem cells, like all cells, are influenced by their microenvironment, including chemical and physical cues. In vitro, these cues can serve to influence stem cell fate (e.g. maintain stem cells undifferentiated or promote differentiation along a pathway) and/or to facilitate regenerative medicine applications (e.g. expand stem cells to large numbers or promote uniformly differentiated populations). Until now, chemical cues have been the primary means by which stem cell self-renewal and differentiation have been influenced. Soluble factors and substrate coatings (Figure 3.3) have been used in maintaining stem cells undifferentiated, as well as in promoting a particular differentiation pathway. The literature in this area is vast and best reviewed elsewhere within a more specific context. Recent efforts have begun focusing on controlling the cellular microenvironment by engineering 3D biomaterials and/or applying physical forces (Figures 3.3 and 3.4). As the number of

Soluble factors

Biomaterials

Applied forces

Ectoderm

(Neuron)

Mesoderm Differentiation

Embryonic stem cell

(Endothelial cells) Self-renewal

Endoderm

(β-cell)

Figure 3.3 Cues in the microenvironment that affect stem cell fate. This schematic indicates the effect of chemical and physical cues on embryonic stem cell fate, including self-renewal processes and differentiation toward all three germ lineages.

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2D configuration Applied force

Soluble factors Cells Substrate

3D configuration

Figure 3.4 The different configurations and environmental cues during cell culture. In two dimensions, cells (gray circles) may be (a) adhered to a surface via a protein substrate (black strands), (b) exposed to soluble factors (red circles) in the medium, and (c) subjected to applied forces (green arrows) via surface distention or fluid motion. In three dimensions, cells may be (a) seeded onto or embedded within a scaffold with matrix molecules, (b) exposed to soluble factors in the medium, and (c) subjected to applied forces via scaffold deformation, fluid motion, or fluid pressurization.

studies on human stem cells is limited and these are based on prior work in animal cells, this discussion will review in vitro mechanical modulation of cells from multiple species. The field of biomaterials has developed a wide range of 3D scaffolds that customize the physical microenvironment within which cells reside. Some studies have found that a 3D culture environment can provide cues that are otherwise missing from 2D stem cell cultures (Tun et al., 2002). In particular, scaffolds can enhance ESC self-renewal (Nur-E-Kamal et al., 2005) and allow the propagation of hematopoietic progenitor cells without the normally requisite growth factors or stromal cells (Bagley et al., 1999; Ehring et al., 2003). Other studies indicate the capacity of stem cells to spontaneously differentiate to cells of all three lineages in 3D (Levenberg et al., 2003). Subsequent work along that line has focused on using 3D biomaterials to direct differentiation of stem cells toward a variety of phenotypes, including hematopoietic (Liu et al., 2003), neural (Ma et al., 2004), and orthopedic (Chaudhry et al., 2004; Hwang et al., 2005). The use of 3D environments has utility beyond simply promoting stem cell self-renewal and differentiation. The very nature of a 3D environment allows an organization of matrix (Grayson et al., 2004) and the formation of structures (Levenberg et al., 2003) that are not otherwise possible on flat surfaces. This may be particularly useful in tissue engineering, in which initial studies have used a 3D scaffold to serve both as a physical cue for stem cell differentiation, as well as the basis of a tissue construct (Awad et al., 2004; Betre et al., 2006). Additionally, a 3D environment may physically entrap or be bound to chemical cues to provide controlled spatial and temporal gradients (Batorsky et al., 2005). Biomaterials themselves can provide both chemical and physical cues that influence stem cell fate. It is thought that biological matrix components, such as collagen, fibrin, and laminin, provide bioactive cues typically seen in vivo that are difficult to replicate using synthesized polymers. Biological components have thus become the basis of gels (Chen et al., 2003) or have served as coatings (Levenberg et al., 2003) in 3D scaffolds. The influence of biological factors on stem cell differentiation was elucidated in one study in which ESCs differentiated toward the tissue-specific lineages when seeded onto extracts from cartilage versus basement

Stem Cell Research

membrane (Philp et al., 2005). The effect of these biologically based scaffolds is more than compositional, as chemically similar collagen in macroscopically different 3D configurations (sponge versus gel) results in different differentiation patterns (Chen et al., 2003). The importance of macroscopic 3D architecture was corroborated in one study where a porous polymeric scaffold without any matrix molecules (Gerecht-Nir et al., 2004b) promoted differentiation of seeded ESCs. It is not only scaffold architecture, however, but also mechanical properties, that play a role in differentiation. Scaffolds that are too stiff have been shown to inhibit embryoid body growth, cavitation, and differentiation (Battista et al., 2005). This suggests the need for more research using engineered scaffolds in which protein presentation, macroscopic architecture, mechanical stiffness, and degradation rates can all be tailored. In general, the engineering of scaffolds has already become quite sophisticated, at times even using biologically derived and synthetic components together. In relation to stem cell research, one example of an innovative-engineered scaffold uses silk, a natural polymer, that can be customized in terms of mechanical and degradation characteristics in 3D configurations (Wang et al., 2005b). Physical forces, such as compression, tension, and shear, have long been applied to cells via bioreactors, a term commonly used for systems with controlled culture conditions. Some bioreactor systems are used to study the modulation of cells and tissues by well-defined cues. Once the appropriate cues for a given application are determined, bioreactors can be appropriately designed to scale up modulation to large numbers of cells and tissue samples. With the recent commercial availability of a few systems, studies that utilize bioreactors and are designed to understand the importance of environmental cues have become more numerous, with some now focusing on stem cells. Initial use of bioreactors with stem cells revolved around non-adherent cells, namely hematopoietic and neural progenitors, in suspension cultures to accelerate and augment expansion kinetics and capabilities, respectively. Stir-based and perfusion bioreactors have been used with hematopoietic progenitor cells, in which the increase in cellular yield is attributed to frequent medium changes, as well as controlled oxygen and cytokine concentration gradients (reviewed by others: Nielsen, 1999; Cabrita et al., 2003). Similar stir-based bioreactors have also been used with neural progenitors, where the main objective is to provide fluid motion to regulate neurosphere diameter, a characteristic correlated to proliferation rates and differentiation potential (Kallos and Behie, 1999; Kallos et al., 1999; Sen et al., 2002; Alam et al., 2004). The approach of allowing limited cell aggregation (cells come together to form a cluster), without sphere agglomeration (clusters come together to form larger bodies), is now being applied to ESC studies. The embryoid body model of differentiation is being studied in some fluid shear stress-based bioreactors that control sphere morphology (Dang et al., 2004; Gerecht-Nir et al., 2004a; Bauwens et al., 2005) and promote differentiation toward a particular phenotype (Schroeder et al., 2005). Although these bioreactors often generate a poorly controlled microenvironment (Konstantinov et al., 2004), they are easily operated and can be scaled up for clinical or manufacturing purposes. The fundamental mechanisms that regulate stem cell responses to applied forces are commonly investigated using well-characterized bioreactor systems for 2D cultures. In systems where cyclic tensile strain (10%, 0.5 Hz) has been applied to silastic membranes seeded with cells, it has been found that both ESCs (Saha et al., 2006) and MSCs (Lee et al., 2005) proliferate and retain their original differentiation potential. Based on the assumption that cells functionally adapt to their microenvironment, many investigators have chosen to mimic certain aspects of in vivo mechanical environments in their studies on differentiation. Endothelial cells that line vascular vessels in situ experience varying levels of fluid shear stress as blood flows past. Similar fluid shear stresses applied in vitro to ESCs (Ahsan and Nerem, 2005; Yamamoto et al., 2005), circulating EPCs (Yamamoto et al., 2003), and mesenchymal progenitor cells (Wang et al., 2005a) have indeed resulted in an increase in protein expression typical of the endothelial phenotype. While bioreactors designed for 2D cell cultures cannot truly mimic in vivo conditions or create 3D tissues, they provide simplified mechanical environments that allow for careful study of stem cell responses.

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Bioreactors have previously been used for tissue engineering with differentiated cells seeded onto scaffolds to create 3D constructs. As stem cell research progresses, this approach is now incorporating the use of undifferentiated cells. Most studies in this area have until now focused on MSCs and orthopedic applications. MSCs seeded onto a fibrous 3D construct and then subjected to fluid perfusion resulted in an increased cell density and a more uniform cellular distribution (Zhao and Ma, 2005), whereas MSCs embedded in agarose and subjected to compression resulted in cells with a chondrogenic phenotype (Huang, C.Y., et al., 2004). These studies applied a single homogenous input to an entire population of cells. A recent study used a more complicated system that differentially applied complex inputs to subsets of the original cell population. In a layered microporous tubular scaffold, Flk-1 cells were seeded onto the lumen and submitted to fluid shear stress and circumferential strain. Due to the complex geometry and multiple inputs, the stimuli sensed by each cell were dependent on the spatial location of the cell within the scaffold. Analogous to the organization within a blood vessel, cells lining the lumen assumed an endothelial morphology, while cells in deeper layers took on smooth muscle cell-like characteristics (Huang et al., 2005). Although limited in scope, this study and others support the concept that the in vitro microenvironment can be used to differentially regulate stem/progenitor cell self-renewal and differentiation processes, potentially within a single tissue-engineered construct.

REGENERATIVE MEDICINE Regenerative medicine focuses on strategies to repair, regenerate, and/or replace tissues and organs. The goal in each of these cases is to restore tissue and organ function through the delivery of cells, signaling molecules, and/or support structures. Disease can be thought of as a failure of the normal biological repair mechanisms that are present in the body. If one can detect disease at an early stage, even at a pre-clinical stage, and intervene by enhancing/inducing biological repair, then it may be possible to restore normal biological function without creating new tissue. In this case repair is at a local, cellular level. Once disease progresses to a more advanced stage (i.e. a clinical stage), then it may be necessary to regenerate or create new tissue in order to restore full function. Finally, when such an approach is not sufficient, then the strategy may actually require replacement of the tissue so as to restore the structure and full biological function. This includes the mechanical/electrical/chemical aspects of function. It should be noted, however, that although one can attempt to define each of these three mechanisms so as to distinguish between them, the fact of the matter is that many regenerative medicine therapies incorporate multiple elements. Thus, as an example, inducing repair may in the long term lead to the creation (i.e. regeneration) of new tissue. Another example is the introduction of a replacement that acts as a trigger for a repair and/or regenerative response that ultimately restores function. Cells are the machinery that promote tissue regeneration and, specifically, stem cells are a useful source for transplantation or tissue engineering. The cells, however, may originate from a variety of locations and be at varying levels of commitment. Certain regenerative medicine approaches may rely on autologous adult stem cells being recruited from the host, such as an osteoinductive graft for critical bone defects, into which stem cells and osteoblasts from the recipient’s own tissues migrate to the site of repair. On the other hand, with the capacity to self-renew and differentiate in vitro, stem cells could be a means by which to generate large homogenous populations of normal cells, either undifferentiated or committed, for tissue engineering or transplantation. In tissue engineering, cells are used to grow 3D constructs in vitro for implantation. Transplant examples include bone marrow (which contains marrow-derived stem cells) to treat various blood disorders and chondrocytes for articular cartilage repair (Brittberg et al., 1994). Overall, stem cells in regenerative medicine may be allogeneic or autologous, added exogenously or recruited from the host, and potentially expanded or differentiated in vitro. Complex strategies may eventually be developed to combine approaches, perhaps exploiting the effects of co-culture by implanting donor cells of a particular phenotype, that together with host cells, result in a desired regenerative response.

Stem Cell Research

As regenerative medicine covers a wide spectrum of clinical applications and approaches, the field and the research that supports it include a range of disciplines and professions (i.e. basic scientists, engineers, and clinicians). The intersection of stem cell technology and regenerative medicine can be categorized by various criteria, such as stem cell type, technology, or approach. In the end, however, regenerative medicine is a subset of medical treatments, and so here the discussion is organized based on clinical application. Neural Applications Neural applications in regenerative medicine include trauma and diseases, such as spinal cord injuries and Parkinson’s, respectively. Spinal cord injury therapies may require multiple cell types, including neurons and oligodendrocytes, to help regenerate transected tissues. Parkinson’s is a degenerative condition in which dopaminergic cells are lost, resulting in motor dysfunction such as bradykinesia, rigidity, and tremors. Clinical studies related to neural applications have focused on Parkinson’s and aim to restore the presence of dopaminergic neurons. Transplantation of cadaveric and adrenal dopaminergic neurons has been shown to have little or short-lived effects (Quinn, 1990). Beyond survival, integration of transplanted neurons with the host tissue is thought to be pivotal for long-term success. It is thought that stem cells for transplantation may restore normal neural function by either integrating and forming working neurons or acting as a trigger to promote neurogenesis by host cells. Implantation of undifferentiated cells are feared to form teratomas with undesirable cell types, so a favored strategy for Parkinson’s therapy is to use stem or progenitor cells committed to the neurogenic pathway prior to transplantation. Fetal mesencephalic tissue containing dopamine-producing neural progenitors has been transplanted in multiple clinical studies. An “open label” clinical study during the mid-1980s transplanted fetal tissue and found mixed, but promising, results. This was followed by two independent National Institutes of Health (NIH)-funded double-blinded clinical trials: one led by Freed et al. (2001) and the other by Olanow et al. (2003), with each study using slightly different sample preparation and surgical procedures for fetal tissue transplantation. Neither study showed a significant improvement when comparing entire patient populations and in a few cases, unfortunately, side effects actually included periods of increased Parkinson’s symptoms (Hagell et al., 2002). The study results indicated, however, that increasing the number of transplanted cells may be beneficial for patients with milder cases of Parkinson’s. Due to the logistical and technical issues related to fetal tissue harvest, including the lack of tissue standardization and low cellular yield, alternate cell sources are required. Xenotransplantations, using cells from fetal pigs, were found to be safe, but failed to promote significant improvement in patients (Schumacher et al., 2000). The potential to generate large numbers of stem cell-derived dopaminergic neurons in vitro could have a meaningful effect on therapies for Parkinson’s. While clinical studies have mostly focused on Parkinson’s, there are many opportunities for stem cells to impact treatment of both neurodegenerative diseases and neural injuries. Cardiovascular Applications Cardiovascular applications of regenerative medicine include myocardial repair, blood vessel substitutes, and valvular replacements. Each application has unique challenges. In myocardial repair, the ideal repair response includes revascularization of ischemic tissue and electrical synchrony with host cardiomyocytes; blood vessel substitutes need to remain patent, and preferably are vasoactive as well; and valvular replacements must persist in a mechanically severe environment. In all three applications, the use of either allogeneic or autologous stem cells may be beneficial. A myocardial infarct starts a cascade of events that can lead to congestive heart failure. Initial events include ischemia-induced myocardial necrosis and dysfunction. Necrotic cells are removed through an immunological response and eventually a scar tissue is formed. As a result of this process, heart muscle contractility and

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remodeling are adversely affected, resulting in the loss of cardiac function. Clinical trials have investigated the use of stem cells as a post-infarction therapy using a myriad of approaches. Some studies used freshly isolated autologous bone marrow-derived mononuclear cells, delivered to the myocardium anytime from less than 3-day post-infarct to late stages of congestive heart failure, and showed improvements in common indicators of cardiac function (Assmus et al., 2002; Strauer et al., 2002; Perin et al., 2003; Tse et al., 2003). Other studies have used ex vivo expanded autologous blood-derived EPCs (Assmus et al., 2002) or selected (CD133) autologous marrowderived cells and also found promising results. Future possible cell sources may include ESC-derived cardiomyocytes (Caspi and Capstein, 2004) and endothelial cells (Levenberg et al., 2002), as well as muscle- (Qu-Petersen et al., 2002), adipose- (Planat-Benard et al., 2004a, b), and umbilical cord- (Murohara et al., 2000) derived stem cells. Nonetheless, these initial clinical studies using bone marrow- and blood-derived stem cells are among the most advanced applications of stem cells in cell-based therapies to date. Another application of stem cells in cardiovascular regenerative medicine is as a cell source for engineered tissues. Cardiovascular tissues synthesized in vitro include substitute blood vessels, myocardial patches, and valvular replacements. In substitute blood vessels, large (6 mm) synthetic vessels have been found to be somewhat successful and remain patent. Yet small diameter synthetic vessels, as would be used for coronary bypass, quickly become occluded. It is thought that an endothelial cell lining, as found in native vessels, would provide an anti-thrombogenic inner layer that would prevent clot formation. Future generations of blood vessel substitutes are likely to be capable of vasoactivity and long-term remodeling, to which smooth muscle cells in the medial layer are critical. Similarly, valvular replacements or myocardial patches expected to remodel over years will likely need endothelial and interstitial cells or cardiomyocytes, respectively. Stem cells, either allogeneic or autologous, may provide means to generate these diverse vascular phenotypes in vitro. Orthopedic Applications One current stem cell-based orthopedic therapy includes bone marrow-derived MSC transplantation for osteogenesis imperfecta, a genetic disorder in which osteoblasts synthesize defective collagen type I, which leads to a variety of skeletal pathologies. In limited clinical studies in children, it has been found that allogeneic bone marrow-derived mesenchymal cells engraft in multiple skeletal sites and improve bone growth velocity (Horwitz et al., 2002). In other applications, stem cells are recruited from the host to help regenerate tissues. Cartilage repair techniques, such as microfracture (Steadman et al., 2001), expose vascularized bone that then forms a conduit by which marrow cells, including MSCs, can access the defect site. One well-established cellular therapy in orthopedics, autologous chondrocyte transplantation for articular cartilage, may be improved through the use of stem cells. In this procedure, originally published by Brittberg et al. (1994) and subsequently commercialized by Genzyme Biosurgery under the name Carticel®, chondrocytes are harvested from a non-load bearing region of the knee, expanded in vitro, injected into an articular cartilage defect, and covered with a periosteal flap. Donor site morbidity is an undesirable consequence of this procedure. A stem cell-derived chondrocyte may provide a marked improvement on this already wellestablished orthopedic therapy. Future applications of stem cells in orthopedic regenerative medicine include tissue engineering. The scope of in vitro engineered tissues currently being studied in orthopedics includes bone, articular cartilage, temporal mandibular cartilage, meniscus, muscle, tendon, and ligament. Similar to many tissue engineering applications, cell sourcing of terminally differentiated or appropriate progenitor cells is problematic and stem cells are an option. One preliminary study already used stem cells for orthopedic tissue engineering. In just a few patients, marrow-derived osteoprogenitor cells were grown on porous hydroxyapatite scaffolds that were then implanted into critical length defects in long bones. In three patients, radiographs indicated callus

Stem Cell Research

formation along the implants and good integration with the adjacent host bone (Quarto et al., 2001). This study, albeit very limited, shows the promise of stem cells in orthopedic tissue engineering. Metabolic and Secretory Applications The cell types in metabolic and secretory organs are among those in the body that have the most complex functional properties. The Edmonton protocol has shown the value of islet transplantation in addressing insulin regulation in patients with type I diabetes (Shapiro et al., 2000). Islets are a collection of endocrine cells in the pancreas responsible for insulin secretion used in metabolizing glucose. Beta cells, which constitute 80–85% of the islets, sense blood sugar levels and secrete appropriate amounts of insulin in response. These cells are destroyed by an abnormal immunological response in individuals with type I diabetes. Among the key aspects that led to successful insulin-independence in the Edmonton protocol was the large number of islets transplanted into the patients. Collecting those large numbers is problematic due to the paucity of available donor organs and the difficulty in islet isolation (Kobayashi et al., 2004). As a result, there is great interest in generating insulin-secreting cells from stem cells. The literature in this area, however, is conflicting and controversial. While some studies reported the derivation of beta cells from pancreatic adult stem cells, lineage tracing and evidence of clonal expansion to support those claims were lacking. Embryoid bodies, used to differentiate human ESCs, include a small number of insulin producing cells. Better characterization is needed to determine whether those cells are beta cell precursors, neural cells, or extra-embryonic endodermal cells (Otonkoski et al., 2005). While developments in this area are significant, further basic science studies are required before stem cells provide an alternative therapy for diabetes. Hematopoietic and Autoimmune Applications Bone marrow transplantation, which originated in the 1950s, is now known to include the transfer of multiple stem cell types, including hematopoietic and MSCs. It is the capacity of HSCs to yield blood components or even to restore the entire immune system that is the basis of marrow-derived cell therapies. Hematological malignancies, such as leukemia, sickle cell, and aplastic anemia, arise as a result of abnormalities in marrowderived cells. Transplantation of allogeneic bone marrow- (Arcese et al., 1999) and umbilical cord- (Benito et al., 2004) derived stem cells treats pathologies in hematopoiesis and the immune system by providing a new source of blood and immune cells. Autoimmune diseases are another application of stem cells in regenerative medicine. Such diseases, which can affect either specific organs or the entire system, include multiple sclerosis, rheumatoid arthritis, and systemic lupus erythematosus. Conventional treatments for these conditions include immunosuppression, which can be effective but not curative. Recently, refractory cases of autoimmune diseases are being treated with severe immunosuppression, to the extent of immunoablation, followed by allogeneic or autologous stem cell transplantation (Jantunen and Luosujarvi, 2005). Subsequent treatment with stem cell mobilizers is meant to allow the transplanted cells to rebuild the entire immune system. In most cases, transplanted cells originate from the bone marrow, but now cells obtained from peripheral blood are also being used. In one recent study, a trial of 85 patients with progressive multiple sclerosis found that greater than 60% of the patients benefited from this procedure (Fassas et al., 2002). Another ongoing study focuses on using this approach to arrest the progression of the disease in patients that are less effected (Havrdova, 2005). Treatment regimens that are better tolerated will need to be developed before this approach becomes a widely accepted therapy. Once the risks associated with this therapy are sufficiently low, extensions of this approach may be applied to other uses, such as boosting immune systems after aggressive chemotherapy during cancer treatment.

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CONCLUSION Stem cell technology shows potential in contributing to regenerative medicine. Self-renewing with the potential to differentiate into specialized phenotypes, stem cells may be derived from embryonic, fetal, or adult cells or tissues. These cells are allogeneic or autologous, added exogenously or recruited from the host, and potentially expanded and/or differentiated in vitro. In regenerative medicine, stem cells can serve as the machinery to repair, regenerate, and/or replace tissues and organs. The ethical, regulatory, and scientific hurdles will need to be overcome for each stem cell type before clinical use. Applications of stem cells in regenerative medicine will help to confront the organ transplantation crisis and allow customization of therapies for each patient.

ACKNOWLEDGMENTS The authors thank the Georgia Tech/Emory Center for the Engineering of Living Tissues (National Science Foundation Engineering Research Center: NSF EEC-9731643), an NIH Biotechnology Training Program (T32GM08433), and the Ruth L. Kirschstein National Research Service Award (1F32HL076978-01A1) for financial support.

INTERNET RESOURCES National Institute for Health (2005). Stem Cell Information. stemcells.nih.gov National Marrow Donor Program (2005). www.marrow.org United Network for Organ Sharing (2005). Transplant Living. www.transplantliving.org United States Census Bureau (2005). www.census.gov

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Tosh, D. and Strain, A. (2005). Liver stem cells – prospects for clinical use. J. Hepatol. 42(Suppl1): S75–S84. Tse, H.F., Kwong, Y.L., Chan, J.K., Lo, G., Ho, C.L. and Lau, C.P. (2003). Angiogenesis in ischaemic myocardium by intramyocardial autologous bone marrow mononuclear cell implantation. Lancet 361(9351): 47–49. Tun, T., Miyoshi, H., Aung, T., Takahashi, S., Shimizu, R., Kuroha, T., Yamamoto, M. and Ohshima, N. (2002). Effect of growth factors on ex vivo bone marrow cell expansion using three-dimensional matrix support. Artif. Organs 26(4): 333–339. Turnpenny, L., Brickwood, S., Spalluto, C.M., Piper, K., Cameron, I.T., Wilson, D.I. and Hanley, N.A. (2003). Derivation of human embryonic germ cells: an alternative source of pluripotent stem cells. Stem Cells 21(5): 598–609. Wagers, A.J. and Weissman, I.L. (2004). Plasticity of adult stem cells. Cell 116(5): 639–648. Wang, H., Riha, G.M., Yan, S., Li, M., Chai, H., Yang, H., Yao, Q. and Chen, C. (2005a). Shear stress induces endothelial differentiation from a murine embryonic mesenchymal progenitor cell line. Arterioscler. Thromb. Vasc. Biol. 25(9): 1817–1823. Wang, Y., Kim, U.J., Blasioli, D.J., Kim, H.J. and Kaplan, D.L. (2005b). In vitro cartilage tissue engineering with 3d porous aqueous-derived silk scaffolds and mesenchymal stem cells. Biomaterials 26(34): 7082–7094. Yamamoto, K., Sokabe, T., Watabe, T., Miyazono, K., Yamashita, J.K., Obi, S., Ohura, N., Matsushita, A., Kamiya, A. and Ando, J. (2005). Fluid shear stress induces differentiation of flk-1-positive embryonic stem cells into vascular endothelial cells in vitro. Am. J. Physiol. Heart Circ. Physiol. 288(4): H1915–H1924. Yamamoto, K., Takahashi, T., Asahara, T., Ohura, N., Sokabe, T., Kamiya, A. and Ando, J. (2003). Proliferation, differentiation, and tube formation by endothelial progenitor cells in response to shear stress. J. Appl. Physiol. 95(5): 2081–2088. Zhao, F. and Ma, T. (2005). Perfusion bioreactor system for human mesenchymal stem cell tissue engineering: dynamic cell seeding and construct development. Biotechnol. Bioeng. 91(4): 482–493. Zuk, P.A., Zhu, M., Ashjian, P., De Ugarte, D.A., Huang, J.I., Mizuno, H. Alfonso, Z.C., Fraser, J.K., Benhaim, P. and Hedrick, M.H. (2002). Human adipose tissue is a source of multipotent stem cells. Mol. Biol. Cell 13(12): 4279–4295. Zulewski, H., Abraham, E.J., Gerlach, M.J., Daniel, P.B., Moritz, W., Muller, B., Vallejo, M., Thomas M.K. and Habener, J.F. (2001). Multipotential nestin-positive stem cells isolated from adult pancreatic islets differentiate ex vivo into pancreatic endocrine, exocrine, and hepatic phenotypes. Diabetes 50(3): 521–533.

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Part II Biologic and Molecular Basis of Regenerative Medicine

4 Molecular Organization of Cells Jon D. Ahlstrom and Carol A. Erickson

INTRODUCTION Multicellular tissues exist in one of two types of cellular arrangements – epithelial or mesenchymal. Epithelial cells adhere tightly to each other and to an organized extracellular matrix (ECM) called the basal lamina, thereby producing a sheet of cells with an apical or adhesion-free surface, and a basal side that adheres to the ECM. Mesenchymal cells, in contrast, are individual cells with a bipolar morphology that are held together as a tissue within a loose ECM and are frequently motile. The first tissue to arise in multicellular organisms is the epithelium, which then gives rise to mesenchymal cells through a process called the “epithelial-to-mesenchymal transition” (EMT). Numerous important EMTs occur during development. During gastrulation in amniotes (reptiles, birds, and mammals), the first major EMT occurs when the epithelial epiblast gives rise to mesoderm (reviewed in Leptin, 2005). EMTs also occur later in development, such as the delamination of neural crest cells from the neural tube, the invasion of endothelial cells into the cardiac jelly to form the cardiac cushions, formation of the sclerotome (connective tissue precursors) from epithelial somites, and the creation of palate mesenchymal cells at the seam where the palate shelves fuse (Shook and Keller, 2003; Hay, 2005). The reverse process of mesenchymal-to-epithelial transition (MET) is likewise crucial to development, and examples include the condensation of mesenchymal cells to form somites and the notochord, kidney tubule formation from nephrogenic mesenchyme (Barasch, 2001), and the creation of heart valves from cardiac mesenchyme (Eisenberg and Markwald, 1995). In the adult organism, EMTs and METs occur during wound healing and tissue remodeling (Kalluri and Neilson, 2003). The conversion of transformed epithelium into metastatic cancers is also an EMT process (Thiery, 2002), as is the disintegration of epithelial kidney tissue into fibroblastic cells during end-stage renal disease (Iwano et al., 2002). The focus of this chapter is on the regulation of molecules that control the organization of cells into epithelium or mesenchyme. First, we will discuss the cellular changes that occur during EMTs, including changes in cell–cell and cell–ECM adhesions, stimulation of cell motility, and the increased protease activity that accompanies invasion of the basal lamina. Then we will review the molecules and mechanisms that control EMTs or METs, from the signal transduction pathways to the transcription factors that orchestrate this intricate process. Many molecular mechanisms that regulate EMTs or METs are known; however, the picture is not yet complete and many more players and pathways remain to be discovered.

CELLULAR MECHANISMS OF THE EMT The conversion of an epithelial sheet into individual migratory cells requires the coordinated changes of many distinct families of molecules. As an example of an EMT, we give a brief overview of sea urchin gastrulation, where the individual cells undergoing an EMT can be observed directly (for a recent review, see Shook and Keller, 2003). Upon fertilization of the sea urchin oocyte, the embryo develops into a hollow sphere of epithelial cells (blastula) consisting of a basal domain with a supporting basal lamina on the inner surface of the

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sphere and an apical domain with cell–cell adhesions on the outer surface of the sphere. As the primary mesenchyme cells (PMCs) detach from the epithelium to enter the blastocoel, the apical adherens junctions that tether them to the epithelium are endocytosed (Miller and McClay, 1997), the PMCs lose cell–cell adhesion, and they gain adhesion to the basal lamina (Fink and McClay, 1985). The basal lamina is degraded at sites where PMCs enter the blastocoel (Katow and Solursh, 1980). Thus, the basic steps of an EMT are: (1) the loss of cell–cell adhesion and gain of cell–ECM adhesion, (2) the stimulation of cell motility, and (3) invasion of the basal lamina. Next we examine the components of the EMT in more detail. Changes in Cell–Cell Adhesion Epithelial cells are held together by specialized cell–cell junctions including adherens junctions (Perez-Moreno et al., 2003), desmosomes (Getsios et al., 2004), and tight junctions (Matter and Balda, 2003). These are localized near the apical surface and establish the apical and basal polarity of the epithelium (Ebnet et al., 2004). In order for an epithelial sheet to produce individual migrating cells, cell–cell adhesions must be disrupted. The principal component of the adherens junctions and desmosomes that mediates cell–cell adhesions are the transmembrane proteins of the cadherin superfamily (Wheelock and Johnson, 2003). Cadherins are essential for establishing adherens junctions and desmosomes and maintaining the epithelial phenotype (reviewed in Gumbiner, 2005). E-cadherin and N-cadherin (“E” for epithelial and “N” for neuronal) are classic cadherins that interact homotypically through their extracellular IgG domains with like-cadherins on adjacent cells. Function-blocking antibody against E-cadherin causes the epithelial Madin–Darby canine kidney (MDCK) cell line to dissociate into single migratory cells (Imhof et al., 1983), and E-cadherin-mediated adhesion is necessary to maintain the epithelial integrity of embryonic epidermis (Levine et al., 1994). E-cadherin is also sufficient to promote cell–cell adhesion and assembly of adherens junctions. Overexpression of E-cadherin in fibroblasts will cause them to aggregate tightly together (Nagafuchi et al., 1987). Partial or complete loss of E-cadherin in carcinomas (epithelial cancers) is associated with increased metastasis (Wheelock et al., 2001), and conversely, overexpressing E-cadherin in cultured cancer cells reduces their invasiveness in vitro (Frixen et al., 1991) and in vivo (Navarro et al., 1991). In a mouse model for β-cell pancreatic cancer, the loss of E-cadherin is the rate-limiting step for transformed epithelial cells to become invasive (Perl et al., 1998). Changes in cadherin expression, also known as cadherin switching, are characteristic of an EMT or an MET. For example, epithelia that express E-cadherin will downregulate this cadherin at the time of the EMT and express a different cadherin such as N-cadherin (for review, see Gumbiner, 2005). When mesenchymal tissue becomes epithelial again (MET), such as during kidney formation, N-cadherin is lost and E-cadherin is re-expressed (Kuure et al., 2000). Cadherin switching also occurs during the EMT that generates the neural crest. Just before neural crest cells detach from the neural tube, N-cadherin is downregulated and replaced by cadherin-11 and cadherin-7 expression (Nakagawa and Takeichi, 1995). When neural crest cells cease migration and coalesce into ganglia, they express N-cadherin again (Pla et al., 2001). The injection of functionblocking antibodies against N-cadherin into the neural tube promotes premature migration of neural crest cells (Bronner-Fraser et al., 1992), and forced expression of N-cadherin prevents neural crest delamination (Nakagawa and Takeichi, 1998). However, the loss of cadherins is not always sufficient for an EMT. In the N-cadherin knockout mouse, the neural tube is ill-formed (cell adhesion defect); however, an EMT is not induced by the loss of N-cadherin (Radice et al., 1997). In culture, cadherin switching is not sufficient for an EMT to occur in TGF-β-induced mammary epithelial cells, although cadherin switching is necessary for cell motility (Maeda et al., 2005). Hence, cadherins are essential for maintaining epithelial integrity, but cadherin switching is only one of several steps to complete an EMT. There are several ways through which cadherin expression and function can be regulated. The transcription factors that directly regulate an EMT such as Snail/Slug, Sip1, δEF-1, Twist, or E2A repress transcription of

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E-cadherin (reviewed in De Craene, 2005). E-cadherin protein activity can also be regulated by trafficking and turnover (Bryant and Stow, 2004). The precise endocytic pathways for E-cadherin are still unclear, and there is evidence for both caveolae-dependent endocytosis (Lu et al., 2003) and clathrin-dependent endocytosis of E-cadherin (Ivanov et al., 2004). E-cadherin can also be ubiquitinated in cultured cells by the E3-ligase, Hakai, which targets E-cadherin to the proteasome (Fujita et al., 2002). Another mechanism to disrupt E-cadherin function is through extracellular proteases such as matrix metalloproteases (MMPs), which degrade the extracellular domain of E-cadherin and consequently reduce cadherin-mediated cell adhesion (Egeblad and Werb, 2002). Some or all of these mechanisms may occur simultaneously during an EMT to disrupt cell–cell adhesion. In addition to their role in cell–cell adhesion, cadherins also function as cell-signaling molecules. Intracellularly, classical cadherins interact with α- and β-catenin, which in turn link with the actin cytoskeleton (Tepass et al., 2000). Hence, β-catenin is an important structural component of the cytoskeleton. β-Catenin can also function in cell signaling when it translocates to the nucleus and acts as a co-activator of the lymphoid enhancer-binding factor/T-cell factor (LEF/TCF) transcription factor family (Sharpe et al., 2001). β-Catenin is pivotal for regulating most EMTs. In vertebrates, β-catenin is required for gastrulation, and misexpression of β-catenin results in ectopic gastrulation events (Moon and Kimelman, 1998). β-Catenin is also necessary for the EMT during cardiac cushion development (Liebner et al., 2004). In breast cancer, β-catenin expression is highly correlated with metastasis and poor survival (Cowin et al., 2005), and blocking β-catenin function in tumor cells inhibits their invasion in vitro (Wong and Gumbiner, 2003). It is unclear whether β-catenin overexpression alone is sufficient for all EMTs. If β-catenin is misexpressed in cultured cells, it causes apoptosis (Kim et al., 2000); however, misexpressing a stabilized form of β-catenin in mouse epithelial cells in vivo causes metastatic skin tumors (Gat et al., 1998). Therefore, the central role of cadherins in an EMT may not be solely due to their cell–cell adhesive function, but also to cadherin regulation of the β-catenin signaling pathway. In support of this view, ectopic cadherin expression in Xenopus embryos sequesters β-catenin to adhesion junctions and consequently inhibits β-catenin migration to the nucleus (Fagotto et al., 1996). In E-cadherin misexpression studies in metastatic cancer cells, the suppression of cancer cell invasion does not require cell–cell adhesion, as only the cytoplasmic β-catenin-binding domain of E-cadherin and not the extracellular adhesion domain is required (Wong and Gumbiner, 2003). In summary, cell–cell adhesions depend on cadherins, and cadherins can regulate additional EMT events through β-catenin signaling. Cell–ECM Adhesion Changes Changes in the way that cells interact with the ECM are also important for EMTs and METs. During sea urchin gastrulation, PMCs lose cell–cell adhesions but simultaneously acquire adhesion to the basal lamina through which they invade (Fink and McClay, 1985). Cell–ECM adhesion is mediated principally by integrins (reviewed in Hynes, 2002). Integrins are transmembrane proteins composed of two non-covalently linked subunits, α and β, and require Ca2 or Mg2 for binding to ECM components such as fibronectin, laminin, and collagen. The cytoplasmic domain of integrins links to the cytoskeleton and interacts with other signaling molecules. Changes in integrin function are required for many EMTs. For example, in neural crest delamination, β1 integrin is necessary for neural crest adhesion to fibronectin and becomes functional just a few hours before the EMT (Delannet and Duband, 1992). Likewise, while epiblast cells undergo an EMT to form mesoderm during mouse gastrulation, the cells exhibit increased adhesion to ECM molecules (Burdsal et al., 1993). In both of these cases, blocking integrin function with function-blocking antibodies prevents cell migration. Integrin changes are also associated with increased metastasis in certain cancers (reviewed in Hood and Cheresh, 2002). One molecule that coordinates the loss of cell–cell adhesion with the gain of cell–ECM adhesion during EMT is the GTPase Rap1. In several cultured cell lines, the endocytosis of E-cadherin activates the Ras family member Rap1. Activated Rap1 is required to form integrin-mediated adhesions, as overexpression of the

Molecular Organization of Cells

Rap1-inactivating enzyme, Rap1GAPv, blocks integrin-ECM adhesion formation (Balzac et al., 2005). The molecules with which Rap1 interacts to activate integrin function are not yet known. Hence, cell–ECM adhesions are maintained by integrins, and changes in cell–ECM interactions are also important for EMTs. Stimulation of Cell Motility In order for epithelial cells to undergo an EMT they must become migratory. The gain of cell motility is distinct from simply losing cell–cell adhesions. For example, in EpH4 cells that undergo an EMT after activation of the transcription factor Jun, there is a complete loss of epithelial polarity, but cell migration is not activated (Fialka et al., 1996). Similarly there are two steps during the EMT that generates the cardiac cushion cells: first, the cardiac endothelium is “activated,” whereby the cells lose their adhesions to each other, become hypertrophic, and polarize the Golgi toward one end of the cell. Second, these activated cells become motile and invasive. Curiously it is estimated that only 7% of activated endothelial cells ever invade a collagen gel in in vitro invasion assays (Boyer et al., 1999). The activation and dispersion steps in the EMT are separable and are regulated by different signaling pathways (Markwald et al., 1977; Krug et al., 1985; Runyan et al., 1990). The cellular changes that are responsible for activating cell motility are not understood. However, in many EMTs, there is an upregulation of integrins (e.g. in the cardiac cushion precursors, integrin α6 is upregulated; Boyer et al., 1999). Potentially the ability to adhere to the ECM is sufficient to stimulate motility. Additionally, activation of members of the Rho family of GTPases is required for organizing actin to generate filopodia, lamellipodia, and focal contacts (reviewed in Burridge and Wennerberg, 2004). In many EMTs the loss of Rho family members inhibits the EMT (e.g. RhoB (Liu and Jessell, 1998) and rac (our unpublished data) are required for the neural crest EMT). The extent to which activation of cell motility is needed for the EMT and how it is regulated will be the subject of future research. Invasion of the Basal Lamina In most EMTs epithelial cells penetrate the underlying basal lamina. The basal lamina stabilizes epithelial integrity and generally acts as a barrier to migratory cells (Erickson, 1987). One mechanism that cells use to breach the basal lamina is to produce enzymes that degrade it, including plasminogen activator and MMPs. Plasminogen activator is associated with a number of EMTs, including neural crest delamination and the formation of cardiac cushion cells during heart morphogenesis. Experimentally, blocking plasminogen activity will reduce the number of migratory neural crest cells (Erickson and Isseroff, 1989) or migratory cardiac cells (McGuire and Alexander, 1993). MMPs are also important to a number of EMTs. MMP-2 is necessary for the EMT that generates neural crest cells, because when inhibitors of MMP-2 are added to chicken embryos in vivo, or if MMP-2 translation is blocked with MMP-2 antisense oligonucleotides, neural crest delamination – but not neural crest migration – is inhibited (Duong and Erickson, 2004). In mouse mammary cells, MMP-3 is sufficient for an EMT in vitro and in vivo (Sternlicht et al., 1999). MMP-3 induces an alternatively spliced form of Rac1 (Rac1b), which then causes an increase in reactive oxygen species (ROS) intracellularly. Either Rac1b activity or ROS is necessary and sufficient for an MMP-3-induced EMT. Rac1b or ROS can also induce the expression of the transcription factor Snail (Radisky et al., 2005). The role of Rac1b or ROS in controlling other EMT events during development or disease is not yet known.

MOLECULAR CONTROL OF THE EMT The initiation of an EMT or an MET is a tightly regulated event during development and tissue repair, because deregulation of epithelial organization is disastrous to the organism. A variety of external and internal signaling mechanisms coordinate the complex events of the EMT, and can be disrupted or reactivated during disease

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processes. Many of the molecules that trigger EMTs or METs have been identified, and in some cases the downstream effectors are known. Yet, in general, complete signaling cascades have not been elucidated. EMT can be induced by either ECM components or diffusible signaling molecules and these inductive signals act either directly on cell adhesion molecules or by regulating the EMT transcriptional cascades. Next we will discuss the role of signaling molecules and ECM in triggering the EMT, and then describe the transcriptional programs that are activated.

Signaling Molecules During development, five main cellular signaling pathways are employed – the TGF-β, Wnt, receptor tyrosine kinase (RTK), Notch, and Hedgehog pathways (Gerhart, 1999). These pathways and the soluble ligands that activate them have a role in triggering EMTs. Although the activation of a single signaling pathway can be sufficient for an EMT, in most cases the EMT is coordinated by a combination of signaling molecules. TGF-β Pathway The TGF-β superfamily includes TGF-β, activin, and bone morphogenetic protein (BMP) families. These ligands signal through receptor serine/threonine kinases to activate a variety of signaling molecules including Smads, mitogen-activated protein kinase (MAPK), and PI3K (Derynck and Zhang, 2003). Most EMTs studied are induced in part, or solely, by TGF-β superfamily members (Zavadil and Bottinger, 2005). TGF-β2 and TGF-β3 have sequential and necessary roles in signaling the formation of heart valves from cardiac endothelium (Camenisch et al., 2002a), and TGF-β3 triggers an EMT in the fusing palate (Nawshad et al., 2004). In experimental models, TGF-β has context-dependent effects, acting as a growth suppressor on normal tissue, but as an EMT inducer in later stages of cancer progression. For example, transgenic mice expressing TGF-β1 in keratinocytes are more resistant to the development of chemically induced skin tumors than controls, suggesting a growth-inhibiting effect of TGF-β1. However, a greater portion of the tumors that do form in the keratinocyte-TGF-β1 transgenic mice are highly invasive spindle-cell carcinomas, indicating that TGF-β1 also promotes an EMT (Cui et al., 1996). Similar effects of TGF-β are observed in breast cancer progression, where the TGF-β pathway inhibits initial tumor growth, but promotes metastasis to the lung (Siegel et al., 2003). Expression of dominant-negative TGF-βR II in cancer cells transplanted into nude mice blocks TGFβ-induced metastasis (Portella et al., 1998). In cultured breast cancer cells, TGF-β in combination with activated Ras induces an irreversible EMT (Janda et al., 2002), and in cultured pig cells TGF-β and epidermal growth factor (EGF) synergistically stimulate the EMT (Grande et al., 2002). Some of the downstream effectors of TGF-β signaling in EMTs have been determined. One mode of TGF-β action is to cause the dissociation of cell–cell adhesions. For example, in TGF-β-induced EMTs of mammary epithelial cells, TGF-βR II directly phosphorylates the polarity protein, Par6, and phosphorylated Par6 causes the E3 ubiquitin ligase, Smurf1, to target the GTPase, RhoA, for degradation. RhoA is required for the stability of tight junctions and loss of RhoA leads to their dissolution (Ozdamar et al., 2005). The loss of tight junctions causes changes in cell polarity. Exactly how the ubiquitination of RhoA leads to the loss of tight junctions is not yet known. Besides the action of TGF-β signaling on cell–cell adhesion, the TGF-β pathway also regulates EMT genes. TGF-β signaling through serine/threonine kinases results in the phosphorylation and activation of several Smads that regulate gene expression (reviewed in Shi and Massague, 2003). Smad3 may be the molecule that signals the TGF-β-induced EMT. The deletion of Smad3 in a mouse model leads to the inhibition of injuryinduced lens and kidney tissue EMT (Roberts et al., 2005). The precise role and mechanism of Smads in the EMT remain to be elucidated.

Molecular Organization of Cells

Wnt Pathway The Wnt family of ligands also has a central role in many EMTs. Wnt ligands signal through seven-pass transmembrane proteins of the Frizzled family, and activate G-proteins, PI3K, and β-catenin (Huelsken and Behrens, 2002). Wnt6 is sufficient for the induction of Slug transcription in the neural crest and perturbation of the Wnt pathway reduces neural crest formation (Garcia-Castro et al., 2002). Wnts can also signal an MET; for example, Wnt4 is necessary to induce the coalescence of nephrogenic mesenchyme into epithelial tubules during murine kidney formation (Stark et al., 1994), and Wnt6 is necessary and sufficient for the MET that forms somites (Schmidt et al., 2004). As with the TGF-β superfamily, Wnt signals both adhesion molecules and transcription factors. One mode of Wnt11 activity, which regulates zebrafish gastrulation, is to stimulate the GTPase Rab5c, which results in the endocytosis of E-cadherin and consequently the loss of cell–cell adhesion (Ulrich et al., 2005). Wnt signaling also activates transcription of genes that coordinate the EMT, often through the stabilization of β-catenin and the subsequent nuclear β-catenin co-activation of LEF/TCF transcription factors. Signaling by RTK Ligands The RTK family of receptors and the growth factors that activate them also regulate EMTs or METs. RTKs are activated by their respective ligands, which causes receptor dimerization and results in the autophosphorylation of tyrosine residues intracellularly. These cytoplasmic phosphotyrosines act as docking sites for intracellular signaling molecules or adapter proteins, which in turn activate signaling components such as Ras/MAPK, Rac, PI3K, and JAK/STAT (reviewed in Schlessinger, 2000). Next we cite a few examples. Hepatocyte growth factor (HGF), also known as scatter factor, acts through the RTK c-met. HGF is important for the MET in the developing kidney, since HGF/SF function-blocking antibodies inhibit the assembly of metanephric mesenchymal cells into kidney epithelium in organ culture (Woolf et al., 1995). HGF signaling is required for the EMT that produces myoblasts (limb muscle precursors) from somite tissue in the mouse, because in knockout mice for c-met, myoblasts fail to migrate into the limb bud (Bladt et al., 1995). Fibroblast growth factor (FGF) signaling regulates the EMT during mouse gastrulation. In FGFR1 mouse mutants, E-cadherin is not downregulated, β-catenin does not move into the nucleus, snail is not expressed in gastrulating cells, and gastrulation does not occur. Interestingly, if E-cadherin function is also inhibited in FGFR1 mutants by the addition of function-blocking E-cadherin antibodies, the EMT proceeds normally. The suggested mechanism is that failure to remove E-cadherin allows E-cadherin to sequester free β-catenin and therefore attenuate later Wnt signaling required to complete gastrulation events (Ciruna and Rossant, 2001). FGF signaling also stimulates cell motility and MMP activation. In studies using cultured cancer cells, sustained FGF2 signaling results in cell motility, MMP-9 activation, and the ability to invade ECM (Suyama et al., 2002). Insulin growth factor (IGF) signaling can also induce an EMT. In cultured epithelial cells, IGFR1 complexes with E-cadherin and β-catenin, and the ligand IGF-II causes nuclear translocation of β-catenin, activation of the transcription factor TCF-3, degradation of E-cadherin, and subsequent EMT (Morali et al., 2001). Another RTK receptor known for its role in EMTs is the ErbB2/HER-2/Neu receptor, whose ligand is heregulin/neuregulin. Overexpression of HER-2 occurs in 25% of human breast cancers, and misexpression of HER-2 in mouse mammary tissue in vivo is sufficient to cause metastatic breast cancer (Muller et al., 1988). Herceptin® (antibody against the anti-HER-2 receptor) treatment is effective in reducing the recurrence of HER-2-positive metastatic breast cancers (Goldenberg, 1999). HER-2 signaling activates snail expression in breast cancer (Moody et al., 2005). Another example of the importance of RTKs in the EMT is the mechanism used by the bacterium Helicobacter pylori to promote the breakdown of gastric epithelium that causes peptic ulcers and gastric adenocarcinoma.

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This disease process requires that the bacterial protein CagA be transferred into gastric epithelial cells. Once in epithelial cells, CagA is phosphorylated at tyrosine residues located at its C terminus, and results in the activation of RTK signaling pathways. When CagA is expressed in MDCK cells, the cells lose cell–cell adhesions and epithelial polarity, exhibit cell migration, and gain the ability to invade ECM (Bagnoli et al., 2005). Whether or not this CagA-mediated EMT also occurs in vivo is not yet known. Notch Pathway The Notch signaling family is well known for its role in cell specification, and is now emerging as an important regulator of EMTs (Huber et al., 2005). When the Notch receptor is activated by its ligand Delta, the intracellular portion of the Notch receptor ligand is cleaved and transported to the nucleus where it regulates gene expression (Mumm and Kopan, 2000). In zebrafish Notch1 mutants, cardiac endothelium expresses very little snail and does not undergo the EMT required to make the cardiac cushions (Timmerman et al., 2004). In this study, similar results are obtained by treating embryonic heart explants with inhibitors of the Notch pathway. Conversely, misexpression of activated Notch1 is sufficient to activate snail expression and promote an EMT in cultured endothelial cells. Notch signaling is also important for TGF-β-induced EMT. Upon TGF-β treatment of cultured kidney, mammary, and epidermal epithelial cells, the transcription of the transcriptional repressor Hey1 and the Notch-ligand Jagged1 is stimulated in a Smad3-dependent process. The use of antisense oligonucleotides against hey1 mRNA, siRNA against jagged1 mRNA, or γ-secretase inhibitor (GSI) treatment (to block Notch receptor activation) all inhibit TGF-β-induced EMT in these cell lines (Zavadil et al., 2004). ECM Signaling In addition to diffusible signaling molecules, the extracellular environment also regulates EMTs or METs. When lens or thyroid epithelium is embedded in collagen, these tissues undergo an EMT (reviewed in Hay, 2005). Integrin signaling appears to be important in this transition, because if function-blocking antibodies against integrins are added in the collagen gels, the EMT is inhibited (Zuk and Hay, 1994). Hyaluronan is another ECM component that regulates EMTs. In the hyaluronan synthase-2 knockout mouse (Has2 –/–, results in defects in hyaluronan synthesis and secretion), the cardiac endothelium fails to undergo an EMT and produces the migratory mesenchymal cells critical for heart valve formation (Camenisch et al., 2000). The role of hyaluronan in this EMT may be to activate the RTK ErbB2/HER-2/Neu, because treating cultured Has2 –/– heart explants with heregulin (ligand for ErbB2) rescues the EMT. Consistent with this hypothesis, treating cardiac explants with hyaluronan activates ErbB2, and blocking ErbB2 signaling with the drug herstatin reproduces the Has2 knockout phenotype (Camenisch et al., 2002b). A third ECM component that is important for EMTs is the gamma-2 chain of laminin 5, which is cleaved from laminin 5 by MMP-2. The gamma-2 chain causes the scattering and migration of epithelial cancer cells (Koshikawa et al., 2000), and may be a marker of epithelial tumor cell invasion (Katayama et al., 2003). Integrins are the major mediators of cell interactions with the ECM, but integrins are also involved in cell signaling. Integrins play important roles in regulating cell survival, proliferation, cytoskeletal rearrangements, cell polarity, and cell motility (reviewed in Hood and Cheresh, 2002). One of the intracellular mediators of integrin signaling is integrin-linked kinase (ILK). ILK interacts with the cytoplasmic domains of the β1 and β3 integrin subunits, and ILK can be activated by integrin, TGF-β, Wnt, or RTK signaling (for a review, see Oloumi et al., 2004). Overexpression of ILK in cultured breast or colon cancer cells leads to translocation of β-catenin to the nucleus, activation of Lef-1/β-catenin as transcription factors, and downregulation of E-cadherin (Novak et al., 1998). Inhibition of ILK in cultured colon cancer cells leads to the stabilization of GSK-3β activity, decreased nuclear β-catenin localization, and results in the suppression of lef-1 and snail transcription (Tan et al., 2001).

Molecular Organization of Cells

The EMT Transcriptional Program All of the molecules that regulate cell–cell adhesion, cell–ECM interactions, cell motility, and basal lamina invasion are encoded by DNA. Therefore, at the heart of an EMT are the transcription factors that control the expression of genes that are required for an EMT. Although many of the transcription factors that regulate an EMT have been identified, these complex transcriptional networks are still being defined. Here we review the transcription factors that control EMTs, and then review how the transcriptional activity and protein function of these transcription factors are regulated. Transcription Factors that Regulate EMTs The Snail family of zinc-finger transcription factors, including Snail and Slug, is emerging as the central regulator of adhesion and cell movement during EMTs (for recent reviews, see Barrallo-Gimeno and Nieto, 2005; De Craene et al., 2005). Snail and Slug are transcriptional repressors that are evolutionarily conserved in vertebrates and invertebrates, and are expressed singly or in combination during every EMT yet examined. Snail was first described in Drosophila, and snail mutants fail to express mesodermal markers or undergo the epithelial invagination that produces mesoderm (Alberga et al., 1991). In the Snail knockout mouse, migratory cells with mesodermal markers form a type of mesoderm; however, these presumptive mesenchymal cells still retain apical/basal polarity, adherens junctions, and express E-cadherin mRNA (Carver et al., 2001). Hence, Snail is only necessary for a part of the process that generates mesoderm. One of the known roles of Snail and Slug in an EMT is to repress the transcription of E-cadherin and thus promote the loss of cell–cell adhesion (reviewed in De Craene et al., 2005). Snail represses the e-cadherin promoter by recruiting the mSin3A co-repressor complex and histone deacetylases (Peinado et al., 2004a). Snail is also a transcriptional repressor of the tight junction proteins, Claudin and Occludin (Ikenouchi et al., 2003). The misexpression of Snail and Slug also leads to the transcription of genes important for cell motility. In MDCK, the misexpression of Snail indirectly leads to the expression of fibronectin and vimentin, which are important for mesenchymal cell motility (Cano et al., 2000), and Slug induces RhoB expression, a GTPase involved in motility, in avian neural crest cells (Del Barrio and Nieto, 2002). In MDCK cells, the misexpression of Snail also promotes mmp-9 transcription and basal lamina invasion through a yet unknown pathway (Jorda et al., 2005). Although Snail and Slug are transcriptional repressors, they somehow activate other EMT genes, and the process has not yet been elucidated. Two other zinc-finger transcription factors regulate EMTs. Delta-crystallin enhancer-binding factor 1 (δEF1), also known as ZEB1, is necessary and sufficient for an EMT in mammary cells transformed by the transcription factor c-Fos (Eger et al., 2005). Smad-interacting protein-1 (Sip1), also known as ZEB2, is structurally similar to δEF1, and Sip1 overexpression is sufficient to downregulate E-cadherin, dissociate adherens junctions, and increase motility in MDCK cells (Comijn et al., 2001). Both δEF1 and Sip1 can bind to the E-cadherin promoter and repress transcription (reviewed in De Craene et al., 2005). The basic helix-loop-helix (bHLH) transcription factors Twist and E2A also play roles in EMTs. Twist is expressed during Drosophila gastrulation, and the double twist and snail mutant has a more severe gastrulation phenotype than either mutant alone, suggesting that snail and twist have distinct functions. Twist1 is not necessary for mouse gastrulation, yet Twist1 mouse mutants do have neural tube fusion, limb, and somite defects (Chen and Behringer, 1995). Twist is also necessary for the EMT that generates the mouse neural crest (Soo et al., 2002). E2A is not necessary for many EMTs, since mouse mutants for E2A survive and are only defective in B cell production (Zhuang et al., 1994). However, overexpression of E2A in MDCK cells promotes tumor invasion (Perez-Moreno et al., 2001). In MDCK cells, Snail is more efficient at promoting the initial invasion of ECM, whereas E2A is better at inducing later angiogenesis (Peinado et al., 2004b). Twist and E2A can also both repress E-cadherin transcription (De Craene et al., 2005).

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Another important family of EMT transcription factors is the LEF/TCF transcription factor family. The limiting factor for LEF/TCF activation is the availability of β-catenin. β-Catenin levels are negatively regulated by GSK-3β or antigen-presenting cell (APC), and a surplus of β-catenin becomes available after being freed from disassembled adherens junctions (Stockinger et al., 2001). Forced expression of Lef-1 in the presence of stabilized β-catenin causes the downregulation of E-cadherin and promotes EMT in cultured colon cancer cells. Inhibition of Lef-1 misexpression (by removing Lef-1 retrovirus from the culture medium) causes cultured cells to revert back to an epithelium (Kim et al., 2002). LEF/TCF transcription factors directly activate genes that regulate cell motility. The LEF/TCF pathway activates the promoter of the L1 adhesion molecule, and L1 is associated with increased motility and invasive behavior of colon cancer cells (Gavert et al., 2005). β-Catenin and LEF/TCF also activate the fibronectin gene (Gradl et al., 1999). Finally, LEF/TCF transcription factors activate genes that stimulate basal lamina invasion, including mmp-3 and mmp-7 (Crawford et al., 1999; Gustavson et al., 2004). The Regulation of Transcription Factors that Control EMT To fully understand the transcriptional network that regulates EMTs, we should also know how EMT-inducing transcription factors are regulated. Transcription factor activity can be controlled both at the level of transcription as well as at the protein level by nuclear import/export or protein degradation. The activation of snail transcription in Drosophila requires the transcription factors Dorsal (NF-κB) and Twist, and the Snail promoter includes both Dorsal and Twist binding sites (Ip et al., 1992). The human Snail promoter also has functional NF-κB sites (Barbera et al., 2004). In cultured human cells transformed by Ras and induced by TGF-β, NF-κB is essential for EMT initiation and maintenance (Huber et al., 2004). A Snail transcriptional repressor has also been identified. In breast cancer cell lines, metastasis-associated protein 3 (MTA3) binds directly to and represses the transcription of Snail in combination with the Mi-2/NuRD complex (Fujita et al., 2003). MTA3 is induced by the estrogen receptor (ER, nuclear hormone) pathway, and the absence of ER signaling or MTA3 leads to the activation of Snail. This suggests a mechanism whereby loss of the ER in breast cancer contributes to metastasis. The role of MTA3 in other EMTs is not known. Slug transcriptional regulators have also been identified. In Xenopus, the Slug promoter has functional LEF/TCF binding sites (Vallin et al., 2001), and in the mouse, MyoD (transcription factor central to muscle cell development) binds to the Slug promoter and activates Slug transcription (Zhao et al., 2002). In humans, the oncogene E2A-HLF (Inukai et al., 1999), and the pigment cell regulator, microthalamia-associated transcription factor (MITF) (Sanchez-Martin et al., 2002), also bind to the Slug promoter and activate transcription. Lef-1 transcription is directly activated by Smad 2/4 (TGF-β signaling), and the phosphorylated complex of Smad 2/4 in the nucleus can promote Lef-1 transcription in the absence of nuclear β-catenin during fusion of the mouse palate (Nawshad and Hay, 2003). The misexpression of Snail also activates the transcription of δEF-1 and Lef-1 through a yet unknown mechanism (Guaita et al., 2002). The complete transcriptional networks that orchestrate an EMT remain to be elucidated. In addition to controlling gene expression, another way to regulate the activity of transcription factors is at the protein level, including protein stability (targeting to the proteasome) and nuclear localization. GSK-3β, the same protein kinase that phosphorylates β-catenin and targets it for destruction, also phosphorylates Snail. The human Snail protein contains two GSK-3β phosphorylation consensus sites between amino acids 97 and 123. Blocking GSK-3β stabilizes Snail expression and results in the loss of E-cadherin in cultured epithelial cells (Zhou et al., 2004; Yook et al., 2005). Hence, Wnt signaling stabilizes (and therefore activates) both β-catenin and Snail by inhibiting GSK-3β. Lysyl-oxidase-like proteins, LOXL2 and LOXL3, are two molecules that prevent GSK-3β-mediated phosphorylation of Snail, and thus stabilize Snail activity. LOXL2 and LOXL3 form a complex with Snail near the GSK-3β phosphorylation sites, thus preventing GSK-3β from

Molecular Organization of Cells

phosphorylating Snail. Expression of LOXL2 or LOXL3 prevents Snail protein destruction and induces an EMT in culture (Peinado et al., 2005). In addition to targeting Snail to the proteasome, the activity of Snail as a transcriptional repressor also depends on nuclear localization. Snail contains a nuclear export sequence (NES) at amino acids 132–143 that is sufficient and necessary for the export of Snail from the nucleus to the cytoplasm, and depends on the calreticulin nuclear export pathway (Dominguez et al., 2003). This NES sequence is activated by phosphorylation of the same lysine residues that GSK-3β acts upon, suggesting a mechanism whereby phosphorylation of Snail by GSK-3β leads to the export of Snail from the nucleus, although this has not yet been shown directly. While GSK-3β can cause the export of Snail from the nucleus, the phosphorylation of human Snail by p21-activated kinase 1 (Pak1) at Ser246 promotes the nuclear localization of Snail (and therefore Snail activation) in breast cancer cells. Knocking down Pak1 by siRNA blocks Pak1-mediated Snail phosphorylation, increases the cytoplasmic accumulation of Snail, and reduces the invasive behavior of breast cancer cells (Yang et al., 2005). The protein that imports Snail into the nucleus in human cells is not yet known, although a Snail importer has already been described in zebrafish. The zinc-finger transporting protein LIV1 is required for Snail to localize to the nucleus during zebrafish gastrulation, and LIV1 is activated by STAT3 signaling (Yamashita et al., 2004). In zebrafish, the protein kinase that phosphorylates Snail to activate the translocation of Snail to the nucleus has not yet been identified. Therefore, both the stability and the subcellular localization of snail are important for snail function in the EMT.

CONCLUSION Over the past 20 years since the term “EMT” was coined (Greenburg and Hay, 1982), great strides have been made in this rapidly expanding field of research. EMT and MET events occur during development and disease, and many of the molecules that regulate the EMT or MET have been characterized, thanks in large part to the advent of cell culture models. Despite this progress, our picture of the EMT is still not complete and there are major gaps in our knowledge of the EMT regulatory networks. Mounting evidence suggests that disease processes such as the metastasis of epithelial-derived cancers and kidney fibrosis are regulated by the same molecules that create migratory and invasive cells from an epithelium during development. A clearer understanding of EMT and MET pathways in the future will no doubt lead to more effective strategies for tissue engineering and novel therapeutic targets.

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Molecular Organization of Cells

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Gumbiner, B.M. (2005). Regulation of cadherin-mediated adhesion in morphogenesis. Nat. Rev. Mol. Cell Biol. 6: 622–634. Gustavson, M.D., Crawford, H.C., Fingleton, B. and Matrisian, L.M. (2004). Tcf binding sequence and position determines beta-catenin and Lef-1 responsiveness of MMP-7 promoters. Mol. Carcinog. 41: 125–139. Hay, E.D. (2005). The mesenchymal cell, its role in the embryo, and the remarkable signaling mechanisms that create it. Dev. Dyn. 233: 706–720. Hood, J.D. and Cheresh, D.A. (2002). Role of integrins in cell invasion and migration. Nat. Rev. Cancer 2: 91–100. Huber, M.A., Azoitei, N., Baumann, B., Grunert, S., Sommer, A., Pehamberger, H., Kraut, N., Beug, H. and Wirth, T. (2004). NF-KB is essential for epithelial–mesenchymal transition and metastasis in a model of breast cancer progression. J. Clin. Invest. 114: 569–581. Huber, M.A., Kraut, N. and Beug, H. (2005). Molecular requirements for epithelial–mesenchymal transition during tumor progression. Curr. Opin. Cell Biol. 17: 548–558. Huelsken, J. and Behrens, J. (2002). The Wnt signalling pathway. J. Cell Sci. 115: 3977–3978. Hynes, R.O. (2002). Integrins: bidirectional, allosteric signaling machines. Cell 110: 673–687. Ikenouchi, J., Matsuda, M., Furuse, M. and Tsukita, S. (2003). Regulation of tight junctions during the epithelium– mesenchyme transition: direct repression of the gene expression of claudins/occludin by snail. J. Cell Sci. 116: 1959–1967. Imhof, B.A., Vollmers, H.P., Goodman, S.L. and Birchmeier, W. (1983). Cell–cell interaction and polarity of epithelial cells: specific perturbation using a monoclonal antibody. Cell 35: 667–675. Inukai, T., Inoue, A., Kurosawa, H., Goi, K., Shinjyo, T., Ozawa, K., Mao, M., Inaba, T. and Look, A.T. (1999). SLUG, a ces-1-related zinc finger transcription factor gene with antiapoptotic activity, is a downstream target of the E2A-HLF oncoprotein. Mol. Cell 4: 343–352. Ip, Y.T., Park, R.E., Kosman, D., Yazdanbakhsh, K. and Levine, M. (1992). Dorsal–twist interactions establish snail expression in the presumptive mesoderm of the Drosophila embryo. Gene Dev. 6: 1518–1530. Ivanov, A.I., Nusrat, A. and Parkos, C.A. (2004). Endocytosis of epithelial apical junctional proteins by a clathrinmediated pathway into a unique storage compartment. Mol. Biol. Cell 15: 176–188. Iwano, M., Plieth, D., Danoff, T.M., Xue, C., Okada, H. and Neilson, E.G. (2002). Evidence that fibroblasts derive from epithelium during tissue fibrosis. J. Clin. Invest. 110: 341–350. Janda, E., Lehmann, K., Killisch, I., Jechlinger, M., Herzig, M., Downward, J., Beug, H. and Grunert, S. (2002). Ras and TGFβ cooperatively regulate epithelial cell plasticity and metastasis: dissection of Ras signaling pathways. J. Cell Biol. 156: 299–314. Jorda, M., Olmeda, D., Vinyals, A., Valero, E., Cubillo, E., Llorens, A., Cano, A. and Fabra, A. (2005). Upregulation of MMP-9 in MDCK epithelial cell line in response to expression of the Snail transcription factor. J. Cell Sci. 118: 3371–3385. Kalluri, R. and Neilson, E.G. (2003). Epithelial–mesenchymal transition and its implications for fibrosis. J. Clin. Invest. 112: 1776–1784. Katayama, M., Sanzen, N., Funakoshi, A. and Sekiguchi, K. (2003). Laminin gamma 2-chain fragment in the circulation: a prognostic indicator of epithelial tumor invasion. Cancer Res. 63: 222–229. Katow, H. and Solursh, M. (1980). Ultrastructure of primary mesenchyme cell ingression in the sea urchin Lytechinus pictus. J. Exp. Zool. 213: 231–246. Kim, K., Lu, Z. and Hay, E.D. (2002). Direct evidence for a role of β-catenin/LEF-1 signalling pathway in induction of EMT. Cell Biol. Int. 26: 463–476. Kim, K., Pang, K.M., Evans, M. and Hay, E.D. (2000). Overexpression of β-catenin induces apoptosis independent of its transactivation function with LEF-1 or the involvement of major G1 cell cycle regulators. Mol. Biol. Cell 11: 3509–3523. Koshikawa, N., Giannelli, G., Cirulli, V., Miyazaki, K. and Quaranta, V. (2000). Role of cell surface metalloprotease MT1MMP in epithelial cell migration over laminin-5. J. Cell Biol. 148: 615–624. Krug, E.L., Runyan, R.B. and Markwald, R.R. (1985). Protein extracts from early embryonic hearts initiate cardiac endothelial cytodifferentiation. Dev. Biol. 112: 414–426. Kuure, S., Vuolteenaho, R. and Vainio, S. (2000). Kidney morphogenesis: cellular and molecular regulation. Mech. Dev. 92: 31–45.

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5 Cell–ECM Interactions in Repair and Regeneration M. Petreaca and M. Martins-Green

INTRODUCTION For many years, the extracellular matrix (ECM) was thought to serve only as a structural support for tissues. However, as early as 1966, Hauschka and Konigsberg showed that interstitial collagen promoted the conversion of myoblasts to myotubes, and, shortly thereafter, it was shown that both collagen (Wessells and Cohen, 1968) and glycosaminoglycans (Bernfield et al., 1972) play a crucial role in salivary gland morphogenesis. Based upon these findings as well as other pieces of indirect evidence, Hay (1977) put forth the idea that the ECM is an important component in embryonic inductions, a concept which implicated the presence of binding sites (receptors) for specific matrix molecules on the surface of cells. The stage was then set to begin to investigate in detail the mechanisms by which ECM molecules influence cell behavior. Bissell et al. proposed the model of “dynamic reciprocity.” In this model, ECM molecules interact with receptors on the surface of cells which then transmit signals across the cell membrane to molecules in the cytoplasm; these signals initiate a cascade of events through the cytoskeleton into the nucleus, resulting in the expression of specific genes, whose products, in turn, affect the ECM in various ways (Bissell et al., 1982). It has become clear that this concept is essentially correct (Ingber, 1991; Boudreau et al., 1995); cell–ECM interactions participate directly in promoting cell adhesion, migration, growth, differentiation, and programmed cell death (also called apoptosis), as well as in modulation of the activities of cytokines and growth factors, and in directly activating intracellular signaling. Most of what we know about the molecular basis of cell–ECM interactions in these events comes from studies that have used induced mutations, experimental perturbations in vivo, and cell/organ cultures. Below, we will first briefly discuss the composition and diversity of some of the better known ECM molecules and their receptors, then discuss selected examples that illustrate the dynamics of cell–ECM interactions during wound healing and regeneration, as well as the potential mechanisms involved in the signal transduction pathways initiated by these interactions. Finally, we will discuss the implications of cell–ECM interactions in regenerative medicine. COMPOSITION AND DIVERSITY OF THE ECM The ECM is a molecular complex that consists of collagens and other glycoproteins, hyaluronic acid, proteoglycans, glycosaminoglycans and elastins, and that harbors molecules such as growth factors, cytokines, and matrix-degrading enzymes and their inhibitors. The distribution and organization of these molecules are not static, but rather vary from tissue to tissue and during development from stage to stage (Ffrench-Constant and Hynes, 1989; Laurie et al., 1989; Sanes et al., 1990; Martins-Green and Bissell, 1995; Tsuda et al., 1998; Werb and Chin, 1998; Zhu et al., 2001), which has significant implications for tissue function (Sechler et al., 1998; Xu et al.,

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1998; So et al., 2001). Mesenchymal cells are immersed in an interstitial matrix that confers specific biomechanical and functional properties to connective tissue (Culav et al., 1999; Suki et al., 2005). In contrast, epithelial and endothelial cells contact a specialized matrix, the basement membrane, via their basal surfaces only, conferring mechanical strength and specific physiological properties to the epithelia (Edwards and Streuli, 1995; Fuchs et al., 1997; Dockery et al., 1998). This diversity of composition, organization, and distribution of ECM results not only from differential gene expression of the various molecules in specific tissues, but also from the existence of differential splicing and post-translational modifications of those molecules. For example, alternative splicing may change the binding potential of proteins to other matrix molecules (Ffrench-Constant and Hynes, 1989; Chiquet-Ehrismann et al., 1991; Wallner et al., 1998; Ghert et al., 2001; Mostafavi-Pour et al., 2001) or to their receptors (Aota et al., 1994; Mould et al., 1994; Akiyama et al., 1995; Cox and Huttenlocher, 1998), and variations in glycosylation can lead to changes in cell adhesion (Dean et al., 1990; Anderson et al., 1994; Vlodavsky et al., 1996; Schamhart and Kurth, 1997; Cotman et al., 1999). In addition, the presence of divalent cations such as Ca2+ (Paulsson, 1988; Ekblom et al., 1994; Wess et al., 1998) can affect matrix organization and influence molecular interactions that are important in the way ECM molecules interact with cells (Sjaastad and Nelson, 1997; Kielty et al., 2002). Growth factors and cytokines interact with the ECM in a variety of ways which allows them to mutually affect each other (Nathan and Sporn, 1991; Adams and Watt, 1993); they can stimulate cells to alter the production of ECM molecules, their inhibitors and/or their receptors (Streuli et al., 1993; Schuppan et al., 1998; Verrecchia and Mauviel, 2002; Gratchev et al., 2005). TGFβ for example, upregulates the expression of matrix molecules and of inhibitors of enzymes that degrade ECM molecules, the combination of which increases ECM levels (Wikner et al., 1990; Bonewald, 1999; Kutz et al., 2001). The ECM can also influence the local concentration and biological activity of growth factors and cytokines by serving as a reservoir that binds them and protects them from being degraded, by presenting them more efficiently to their receptors, or by affecting their synthesis (Roberts et al., 1988; Chiquet-Ehrismann et al., 1991; Flaumenhaft and Rifkin, 1992; Lamszus et al., 1996; Miao et al., 1996; Kagami et al., 1998; Banwell et al., 2000; Schonherr and Hausser, 2000; Miralem et al., 2001; Rahman et al., 2005). Examples of this include the increased production of TNFα by neutrophils after binding to fibronectin (Nathan and Sporn, 1991), the dependence of HGF (hepatocyte growth factor)-mediated hepatocyte proliferation on heparan sulfate proteoglycans (Sakakura et al., 1999), and the increased ability of VEGF (vascular endothelial growth actor) to induce breast cancer cell proliferation and migration in the presence of fibronectin or heparin (Miralem et al., 2001). Growth factor binding to ECM molecules may also exert an inhibitory effect; SPARC (secreted protein acidic and rich in cysteine)/osteonectin binds multiple growth factors, preventing receptor binding and/or downstream signaling events (Lane and Sage, 1994; Kupprion et al., 1998; Francki et al., 2003). In some cases, only particular forms of these growth factors and cytokines bind to specific ECM molecules, for example, PDGF (platelet derived growth factor) (LaRochelle et al., 1991; Pollock and Richardson, 1992), VEGF (Poltorak et al., 1997), and the chemokine cIL-8 (previously called cCAF (chicken chemotactic and angiogenic factor)). cIL-8 is a small cytokine that is overexpressed during wound repair and in the stroma of tumors (Martins-Green and Bissell, 1990; Martins-Green et al., 1992), and is secreted as a 9 kDa protein, although it can be processed by plasmin to yield a 7 kDa protein. Both forms of the protein are found in association with interstitial collagen, but only the smaller form binds to laminin or tenascin, while neither form binds to fibronectin, collagen IV, or heparin (Martins-Green and Bissell, 1995; Martins-Green et al., 1996). Importantly, binding of specific forms of these factors to specific ECM molecules can lead to their localization to particular areas of tissues and affect their biological activities. Another feature of ECM/growth factor interactions that has been more recently characterized involves the ability of specific domains of various ECM molecules, including laminin-5, tenascin-C, and decorin, to bind and activate growth factor receptors (Tran et al., 2004). The epidermal growth factor (EGF)-like repeats of

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laminin (Panayotou et al., 1989; Schenk et al., 2003; Koshikawa et al., 2005) and tenascin-C bind and activate the EGF receptor (EGFR) (Swindle et al., 2001). In the case of laminin, the EGF-like repeats interact with EGFR following their release by matrix metalloproteinase (MMP)-mediated proteolysis (Schenk et al., 2003; Koshikawa et al., 2005), whereas tenascin-C repeats are thought to bind EGFR in the context of the full-length protein (Swindle et al., 2001). Decorin also binds and activates EGFR, although this binding occurs via leucine-rich repeats rather than EGF-like repeats (Iozzo et al., 1999; Santra et al., 2002). The ability of ECM molecules to serve as ligands for growth factor receptors may facilitate a stable signaling environment for the associated cells due to the inability for the ligand to either diffuse or be internalized, thus serving as a long-term pro-migratory and/or pro-proliferative signal (Tran et al., 2004, 2005).

RECEPTORS FOR ECM MOLECULES In order to establish that ECM molecules themselves directly affect cellular behavior, it was important to identify transmembrane receptors for the specific sequences present on these molecules. As early as 1973, it was observed that during salivary gland morphogenesis near the sites of glycosaminoglycan deposition, the intracellular microfilaments contracted (Bernfield et al., 1973). These investigators proposed that the ECM could “be involved in regulating microfilament function,” suggesting that these molecules can specifically interact with cell-surface receptors. It was subsequently shown that various ECM molecules contain specific amino acid motifs that allow them to bind directly to cell-surface receptors (Humphries et al., 1991; Hynes, 1992; Gullberg and Ekblom, 1995). The best characterized motif is the tripeptide RGD, first found in fibronectin (Pierschbacher and Ruoslahti, 1984; Yamada and Kennedy, 1984). Peptides containing this amino acid sequence promote adhesion of cells and inhibit the adhesive properties of fibronectin. This and other amino acid adhesive motifs have been found in laminin, entactin, thrombin, tenascin, fibrinogen, vitronectin, collagens I and VI, bone sialoprotein, and osteopondin (Humphries et al., 1991). Integrins, a family of heterodimeric transmembrane proteins composed of α and β subunits were the first ECM receptors to be identified (Hynes, 1987). At least 18α and 8β subunits have been identified so far; they pair with each other in a variety of combinations, giving rise to a large family which recognizes specific sequences on the ECM molecules (Figure 5.1). Some integrin receptors are very specific, whereas others bind several different epitopes, which may be on the same or different ECM molecules (Figure 5.1), thus facilitating plasticity and redundancy in specific systems (Hynes, 1992; Cotman et al., 1998; Dedhar, 1999; Hynes, 1999). Although the α and β subunits of integrins are unrelated, there is 40–50% homology within each subunit with the highest divergence in the intracellular domain of the α subunit. All but one of these subunits (β4) have large extracellular domains and very small intracellular domains (Briesewitz et al., 1995; Fornaro and Languino, 1997). The extracellular domain of the α subunits contains four regions that serve as binding sites for divalent cations, which appear to augment ligand binding and increase the strength of the ligand–integrin interactions (Gailit and Ruoslahti, 1988; Loftus et al., 1990; Dickeson et al., 1997; Pujades et al., 1997; Leitinger et al., 2000). Although not as extensively studied as the integrins, it has been found that transmembrane proteoglycans can also serve as receptors for ECM molecules (Rapraeger et al., 1987; Jalkehen et al., 1991; Couchman and Woods, 1996; McFall and Rapraeger, 1998). Several proteoglycan receptors that bind to ECM molecules have been isolated and characterized: syndecan, CD44, RHAMM (receptor for hyaluronate-mediated motility), and phosphacan (Grumet et al., 1994; Couchman and Woods, 1996; Entwistle et al., 1996; Liu et al., 1998). Syndecan binds cells to matrix via chondroitin- and heparan-sulfate glycosaminoglycans, whose composition varies based upon the type of tissue in which syndecan is expressed; the differential glycosaminoglycan modifications alter the binding capacity of particular ligands (Kim et al., 1994; Salmivirta and Jalkanen, 1995). Syndecan also associates with the cytoskeleton, promoting intracellular signaling events and cytoskeletal reorganization through

Cell–ECM Interactions in Repair and Regeneration

II b

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Figure 5.1 Representative members of the integrin family of ECM receptors and their respective ligands. These heterodimeric receptors are composed of one α and one β subunit, and are capable of binding a variety of ligands, including Ig superfamily cell adhesion molecules, complement factors, and clotting factors in addition to ECM molecules. Cell–cell adhesion is largely mediated through integrin heterodimers containing the β2 subunits, while cell–matrix adhesion is mediated primarily via integrin heterodimers containing the β1 and β3 subunits. In general, the β1 integrins interact with ligands found in the connective tissue matrix, including laminin, fibronectin, and collagen, whereas the β3 integrins interact with vascular ligands, including thrombospondin, vitronectin, fibrinogen, and von Willebrand factor. Abbreviations: CO, collagens; C3bi, complement component; FG, fibrinogen; FN, fibronectin; FX, Factor X; ICAM-1, intercellular adhesion molecule-1; ICAM-2, intercellular adhesion molecule-2; ICAM-3, intercellular adhesion molecule-3; LN, laminin; OSP, osteopontin; TN, tenascin; TSP, thrombospondin; VCAM-1, vascular cell adhesion molecule-1; VN, vitronectin; vWF, von Willebrand factor.

activation of Rho GTPases (Carey, 1997; Granes et al., 1999; Saoncella et al., 1999; Bass and Humphries, 2002; Yoneda and Couchman, 2003). The CD44 receptor also carries chondroitin sulfate and heparan sulfate chains on its extracellular domain (Milstone et al., 1994), and undergoes tissue-specific splicing and glycosylation to yield multiple isoforms; these may play roles in cell adhesion as well as in ligand binding (Brown et al., 1991; Ehnis et al., 1996; Tuhkanen et al., 1997). One of the extracellular domains of CD44 is structurally similar to the hyaluronan-binding domain of the cartilage link protein and aggrecan, which suggested that CD44 could serve as a hyaluronan receptor. Using a variety of techniques involving antibody binding and mutagenesis, it has been shown that this domain of CD44 as well as an additional domain outside this region can interact directly with hyaluronan (Miyake et al., 1990; Peach et al., 1993; Bajorath et al., 1998); these regions can also mediate CD44 binding to other proteoglycans, although hyaluronic acid is its primary ligand (Marhaba and Zoller, 2004). In addition, studies have shown that CD44 can also interact with collagen, laminin, and fibronectin (Jalkanen and

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Jalkanen, 1992; Ishii et al., 1993, 1994; Ehnis et al., 1996), although the exact binding sites of these molecules to CD44, as well as the functional significance of such interactions in vivo are not well understood (Ponta et al., 2003). RHAMM has been identified as an additional hyaluronic acid receptor (Hardwick et al., 1992), which is responsible for hyaluronic-acid-mediated cell motility in a number of cell types, and also appears to be important in trafficking of hematopoietic cells (Hall et al., 1994; Masellis-Smith et al., 1996; Pilarski et al., 1999; Savani et al., 2001). Cell-surface receptors other than integrins or proteoglycans have also been identified as receptors for ECM molecules. A non-integrin 67 kDa protein known as the elastin-laminin receptor (ELR) recognizes the YIGSR sequence of laminin and the VGVAPG sequence of elastin, sequences not recognized by integrins; the ELR co-localizes with cytoskeleton-associated and signaling proteins upon laminin ligation, suggesting a role in laminin-mediated signaling (Grant et al., 1989; Massia et al., 1993; Bushkin-Harav and Littauer, 1998), and has more recently been implicated in the signaling downstream of elastin and laminin during mechanotransduction (Spofford and Chilian, 2003). A second receptor, CD36, functions as a scavenger receptor for long chain fatty acids and oxidized LDL, but also binds collagens I and IV, thrombospondin, and malaria-infected erythrocytes to endothelial cells and some types of epithelial cells (Febbraio et al., 2001). Each of these ligands has a separate binding site, but all are located in the same external loop of CD36 (Asch et al., 1993), and the intracellular signals occurring after ligand binding lead to activation of a variety of signal transduction molecules (Huang et al., 1991; Lipsky et al., 1997). Indeed, the anti-angiogenic effects of thrombospondin are dependent upon signaling downstream of CD36 (Jimenez et al., 2000, 2001; Isenberg et al., 2005). Furthermore, alternative splice variants of tenascin-C interact with cell-surface annexin II, which may mediate the cellular responses to this particular form (Chung and Erickson, 1994). In addition, ECM molecules have been shown to bind and activate tyrosine kinase receptors, including the EGFR via EGF-like domains (see above) as well as the discoidin domain receptors DDR1 and DDR2. DDR1 and DDR2 function as receptors for various collagens and mediate cell adhesion and signaling events (Vogel et al., 1997). The DDR receptors have also been implicated in ECM remodeling, as their overexpression decreases the expression of multiple matrix molecules and their receptors, including collagen, syndecan-1, and integrin α3, while simultaneously increasing MMP activity (Faraci et al., 2003; Ferri et al., 2004).

SIGNAL TRANSDUCTION EVENTS DURING CELL–ECM INTERACTIONS The interactions between ECM molecules and their receptors as described above can transmit signals directly or indirectly to signaling molecules within the cell, leading to a cascade of events and the coordinated expression of a variety of genes involved in cell adhesion, migration, proliferation, differentiation, and death (Figure 5.2). There is increasing evidence that cell–ECM interactions, especially through integrins, activate a variety of signaling pathways that can be linked to those specific functions. Some of the signaling events important in these cellular processes are discussed below. Adhesion and Migration It is now well established that, upon ligand binding, integrins can directly induce biochemical signals inside cells (Kumar, 1998; Dedhar, 1999). The cytoplasmic domain of integrins interacts with the cytoskeleton, suggesting that ECM signaling through integrins is transduced via the cytoskeletal elements and can induce cell shape changes which, in turn, may lead to growth, migration, and/or differentiation (van der Flier and Sonnenberg, 2001; Hynes, 2002). For example, cell migration is promoted when fibronectin binds simultaneously to integrins through its cell-binding domain and to proteoglycan receptors through its heparin-binding domain (Bernfield et al., 1992; Hardingham and Fosang, 1992; Hynes, 1992; Giancotti, 1997; Schlaepfer and Hunter,

Cell–ECM Interactions in Repair and Regeneration

Figure 5.2 Schematic diagram of cell–ECM interactions present during the healing and regenerative responses. Such interactions between the ECM receptors and their respective ligands initiate signal transduction cascades culminating in a variety of cellular events important in repair and regeneration, including changes in cellular adhesion and migration and altered rates of proliferation and apoptosis. The presence and/or extent of such changes may influence the balance of repair and regenerative responses to favor one outcome over another; thus, interventions that alter ECM signaling events may shift this balance to favor tissue regeneration and thus decrease scarring.

1998; Dedhar, 1999; Mercurius and Morla, 2001). These receptors interact and colocalize in areas of adhesion where microfilaments associate with the β1 subunit of the integrin receptor via structural proteins such as talin and α-actinin present in the actin cytoskeleton of the focal adhesions. The cytoplasmic domain of the β1 subunit also interacts directly with the focal adhesion tyrosine kinase pp125FAK which, when activated, undergoes autophosphorylation on tyrosine 397 (Hildebrand et al., 1995); this phosphotyrosine residue subsequently serves as the binding site for the SH2 domain of the non-receptor tyrosine kinase c-Src. In turn, c-Src phosphorylates many components of the focal adhesion plaques, including paxillin, tensin, vinculin, and the protein p130cas. Paxillin has been implicated in the regulation of integrin-mediated signaling events and motility; paxillin-deficient fibroblasts exhibit reduced phosphorylation of signaling molecules downstream of integrin ligation, with a concomitant reduction in cell motility (Hagel et al., 2002). The specific role of tensin in the process of adhesion/de-adhesion during migration is not known; however, it interacts with both the cytoskeleton and with other phosphorylated signaling molecules via its SH2 domain, and may thus mediate signals between the plasma membrane and the cytoskeleton and/or facilitate signaling events (Lo, 2004). p130cas activation promotes its interaction with the adaptor molecules Crk and Nck, which appear to form a scaffold for localized activation of Rac-GTPase and the MAP/JNK kinase pathways, thus facilitating migration (Dolfi et al., 1998; Kiyokawa et al., 1998; Klemke et al., 1998; Cho and Klemke, 2002). In addition, it has also

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been shown that c-Src phosphorylates focal adhesion kinase (FAK) on tyrosine 925 which serves as a site for binding of Grb2/Sos complex with subsequent activation of Ras and the MAP kinase cascade (Schlaepfer and Hunter, 1996, 1997, 1998; Schlaepfer et al., 1997), which may also be involved in adhesion/de-adhesion and migration (Giancotti, 1997; Schlaepfer and Hunter, 1998; Dedhar, 1999; Ly and Corbett, 2005). Proliferation and Survival ECM interaction with its receptors can promote cell proliferation and survival, often in conjunction with growth factors or cytokine receptors. Such cooperative effects may occur in a direct manner, as in situations in which the EGF-like repeats of ECM molecules bind and activate growth factor receptors, leading to cell proliferation (Swindle et al., 2001; Tran et al., 2004). However, more is known regarding the importance of indirect cooperative effects, particularly those involved in the anchorage dependence of cell growth. Anchorage is required for cells to enter S phase; even in the presence of growth factors, cells will not enter the DNA synthesis phase without being anchored to a substrate (Zhu and Assoian, 1995). Thus, adhesion of cells to ECM molecules plays a very important role in regulating cell survival and proliferation (Giancotti, 1997; Mainiero et al., 1997; Murgia et al., 1998). Integrin–ECM binding leads to the activation of Fyn and its subsequent interaction with the adaptor protein Shc, which recruits Grb2 and thus activates the Ras/ERK cascade, leading to the phosphorylation of the Elk-1 transcription factor and the expression of early response genes involved in cell cycle progression (Wary et al., 1998; Aplin et al., 2001); integrin ligation is also important for the efficient and prolonged activation of MAPK by growth factors, which may explain, in part, the anchorage dependence of growth factor-mediated proliferation (Aplin and Juliano, 1999; Roovers et al., 1999). It has also been shown that cooperation between integrins and growth factors involves the activation of phosphatidylinositol phosphate kinases, thus increasing the levels of phosphatidylinositol bis-phosphate (PIP2). PIP2 then serves as substrate for phospholipase Cγ (PLCγ), which is activated by growth factors as well as by integrin ligation, ultimately leading to the activation of protein kinase C (PKC) and the promotion of cell proliferation (Housey et al., 1988; Schwartz, 1992; Cybulsky et al., 1993). Furthermore, PI-3 kinase activated downstream of Ras can rescue cells in suspension from undergoing apoptosis via the activation of the Akt serine/threonine kinase (Khwaja et al., 1997). Signaling downstream of cell–ECM binding may also promote degradation of cell cycle inhibitors, thus facilitating cell proliferation; indeed, fibronectin-mediated adhesion leads to the degradation of p21 in a Rac1and Cdc42-dependent manner (Bao et al., 2002). The importance of the Rac/JNK pathway in integrin-mediated proliferation is underscored by studies involving a β1 integrin cytoplasmic domain mutant, which decreased the activation of the Rac/JNK pathway and also negatively affected fibroblast proliferation and survival; these effects were rescued by the expression of constitutively active Rac1 (Hirsch et al., 2002). Likewise, other studies involving integrin inhibition or knockout yield similar negative effects on cell proliferation due to changes in signaling. For example, studies of mice lacking the α1β1 integrin, which is a primary collagen receptor, showed that the fibroblasts of these mice have reduced proliferation even though they attach normally (Pozzi et al., 1998). In addition, mammary epithelial cells over-expressing a dominant negative β1 integrin subunit exhibit reduced proliferation due to a combination of decreased MAPK and Akt activation (Faraldo et al., 2001); Akt activation is also diminished in cells over-expressing the β1 integrin mutant mentioned above (Hirsch et al., 2002). Differentiation Interaction of cells with ECM molecules, hormones, and growth factors is required to activate genes that are specific for differentiation. Interestingly, the latter studies have shown that the cell–ECM interactions that result in the differentiated phenotype are those that fail to activate Shc and the MAP kinase cascade, at least in some cases. This has been shown for endothelial cells in which the interaction of α2β1 with laminin, which does not activate the Shc pathway, leads to formation of capillary-type structures (Kubota et al., 1988), whereas

Cell–ECM Interactions in Repair and Regeneration

the interaction of α5β1 in the same cells with fibronectin results in proliferation (Wary et al., 1998). Similar observations have been made with primary bronchial epithelial cells when they are cultured on collagen matrices (Moghal and Neel, 1998). The formation of endothelial capillary-like tubes also relies upon additional signaling pathways, such as occur upon activation of integrin-linked kinase (ILK); over-expression of this kinase can rescue tube formation in the absence of ECM molecules (Cho et al., 2005), while expression of dominant negative ILK prevents tube formation in the presence of ECM and VEGF (Watanabe et al., 2005). Other differentiated phenotypes likewise require integrin-mediated signaling events. Indeed, TGF-β1-mediated myofibroblast differentiation, an event important in both wound healing and liver regeneration, requires the ligation of specific integrins as well as the activation of FAK and its associated signaling pathways (Thannickal et al., 2003; Lygoe et al., 2004). Apoptosis Signal transduction pathways that lead to apoptosis have been delineated for endothelial cells and leukocytes and appear to involve primarily tyrosine kinase activity (Fukai et al., 1998; Ilan et al., 1998; Kettritz et al., 1999; Avdi et al., 2001). For example, the neutrophil apoptosis stimulated by TNF-α is dependent upon β2 integrinmediated signaling events involving the activation of the Pyk2 and Syk tyrosine kinases as well as JNK1 (Avdi et al., 2001). In other cell types, alterations in the ligand presentation by ECM can also regulate apoptosis. Studies have suggested that integrin ligation by soluble, rather than intact, ligands can function as integrin antagonists and promote apoptosis rather than survival or proliferation (Brooks et al., 1994; Vogel et al., 2001; Stupack and Cheresh, 2002); such soluble ligands may be created by matrix degradation during tissue remodeling, and thus promote apoptosis. The apoptosis stimulated by soluble ligands or other antagonists appears to occur via the recruitment and activation of caspase 8 by the clustered integrins, without any requirement for death receptors (Stupack et al., 2001). However, the recruitment process itself is not well understood.

CELL–ECM INTERACTIONS DURING HEALING OF SKIN WOUNDS Interactions of cells with ECM molecules play a crucial role during wound healing and regeneration. It is the continuous crosstalk between cells and the surrounding matrix environment that contribute to the processes of clot formation, inflammation, granulation tissue development, and remodeling, and during regeneration, the matrix interactions are important in restoration of the damaged tissue. As we will see, many different lines of experimental evidence have shown that the basic cellular mechanisms that result in these events involve cell adhesion/de-adhesion, migration, proliferation, differentiation, and apoptosis (Figure 5.2). Adhesion and Migration Shortly after tissue damage and during the early stages of wound healing, there is a release of blood contents and tissue factors into the area of the wound, leading to platelet activation and adhesion, and the formation of a vascular plug containing primarily platelets, plasma fibronectin, and fibrin (crosslinked by factor XIII), but also including small amounts of tenascin, thrombospondin, and SPARC. During this process, activated mast cells degranulate, releasing vasodilating and chemotactic factors that will bring polymorphonucleocytes to the wound site. These events constitute the early stages of the inflammatory response. The fibrin–fibronectin meshwork provides a provisional matrix which serves as substrate for the subsequent migration of leukocytes and keratinocytes during the very early stages of healing when inflammation and wound closure are occurring. Leukocyte interactions with ECM molecules via integrin receptors affect many of the functions of these cells, in particular those that lead to cell adhesion and migration or to production of inflammatory mediators (Rosales and Juliano, 1995; Romanic et al., 1997; Wei et al., 1997; Vaday and Lider, 2000). An example of the latter

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involves the stimulation of pro-inflammatory cytokine release by tissue macrophages upon binding to low molecular weight hyaluronic acid via CD44 (Hodge-Dufour et al., 1997). Because some inflammatory molecules can be damaging to tissues when produced in excess, the course of inflammation can be affected significantly by the types of ECM encountered by these leukocytes (Wei et al., 1997; Vaday and Lider, 2000). ECM molecules can also facilitate leukocyte chemotaxis into the inflamed area by binding chemokines, thus creating a stable chemotactic gradient to promote a specific directional migration (Patel et al., 2001); mutant chemokines unable to bind glycosaminoglycans were unable to promote chemotaxis in vivo, underscoring the importance of ECM binding in leukocyte recruitment (Proudfoot et al., 2003). During re-epithelialization of cutaneous wounds, keratinocytes migrate over the provisional matrix primarily composed of fibrin/fibrinogen, fibronectin, vitronectin, tenascin, and collagen type III. These cells express α2β1, α3β1, α5β1, α6β1, α5β4, and αv integrin receptors for these ECM molecules, which, in conjunction with MMPs, facilitate their migration to close the wound (Cavani et al., 1993; Juhasz et al., 1993; Gailit et al., 1994; O’Toole, 2001; Li et al., 2004). The importance of individual matrix components in re-epithelialization is underscored by studies done in mice lacking these molecules; for example, fibrinogen-deficient mice experienced disordered re-epithelialization (Drew et al., 2001). This keratinocyte migration may also require new laminin deposition, as an antibody against laminin inhibited keratinocyte migration on fibronectin or collagen (Decline and Rousselle, 2001). Cell–ECM interactions are equally important in the closure of other epithelial wounds. Studies examining the sequential deposition of ECM molecules after wounding of retinal pigment epithelial cells showed “de novo” fibronectin deposition 24 h after wounding, which is followed by deposition of collagen IV and laminin. This sequence of matrix deposition is tightly linked to adhesion and migration of cells to close the wound (Kamei et al., 1998), and inhibition of integrin-matrix binding using antibodies or cyclic peptides can prevent both cell adhesion and migration, implicating cell–ECM interactions in the observed epithelial closure (Hergott et al., 1993; Hoffmann et al., 2005). A similar sequence of events is observed during the repair of airway epithelial cells after mechanical injury (Pilewski et al., 1997; White et al., 1999; Sacco et al., 2004); functional inhibition of fibronectin or various expressed integrins likewise diminished cell migration and healing of this epithelium (Herard et al., 1996; White et al., 1999). As healing progresses, embryonic-type cellular fibronectin produced by macrophages and fibroblasts in the wound bed contributes to formation of the granulation tissue, a provisional connective tissue containing nascent blood vessels and multiple types of ECM molecules (Li et al., 2003). This fibronectin serves as substrate for the migration of the endothelial cells that form the vasculature of the wound bed, myofibroblasts, and lymphocytes that are chemoattracted to the wound site by a variety of small cytokines (chemokines) secreted by both macrophages and fibroblasts (Greiling and Clark, 1997; Feugate et al., 2002b). These chemokines belong to a large superfamily, and have been characterized in humans, other mammals, and in avians (Rossi and Zlotnik, 2000; Gillitzer and Goebeler, 2001). Chemokine-mediated chemoattraction of cells involved in granulation tissue formation, in conjunction with the interaction of these cells with ECM via cell-surface receptors, results in processes that lead to cell adhesion and migration into the area of the wound to form the granulation tissue (Lukacs and Kunkel, 1998; Martins-Green and Feugate, 1998; Feugate et al., 2002b). One of the most extensively studied chemokines with functions important in wound healing is IL-8 (Martins-Green and Bissell, 1990; Martins-Green et al., 1992; Martins-Green and Hanafusa, 1997; MartinsGreen and Feugate, 1998; Martins-Green 2001; Feugate et al., 2002a, 2002b). This has been well illustrated in studies performed using cIL-8/cCAF and chicks as model system. cIL-8 is stimulated to high levels shortly after wounding in the fibroblasts of the wounded tissue (Martins-Green and Bissell, 1990; Martins-Green et al., 1992), and thrombin, an enzyme involved in coagulation that is activated upon wounding, stimulates these cells to overexpress cIL-8 (Vaingankar and Martins-Green, 1998; Li et al., 2000). This chemokine then chemoattracts monocyte/macrophages and lymphocytes (Martins-Green and Feugate, 1998). We have shown that thrombin

Cell–ECM Interactions in Repair and Regeneration

can promote further increases in hIL-8 levels by stimulation of hIL-8 expression in THP-1 differentiated macrophages (Zheng et al., 2007). Expression of cIL-8 remains elevated during granulation tissue formation due to its secretion by fibroblasts, the endothelial cells of the microvasculature of the wound, and macrophages, as well as from its binding to the interstitial collagens, tenascin, and laminin present in the granulation tissue (Martins-Green and Bissell, 1990; Martins-Green et al., 1992; Martins-Green et al., 1996). Furthermore, both hIL-8 and cIL-8 are angiogenic in vivo, and, in the case of cIL-8, the angiogenic portion of the molecule is localized in the C-terminus of the molecule (Martins-Green and Feugate, 1998; Martins-Green and Kelly, 1998). Based on the pattern of expression and functions of IL-8, it appears that this chemokine participates both in inflammation; via chemotaxis for specific leukocytes, and in the formation of the granulation tissue via stimulation of angiogenesis and ECM deposition (Martins-Green and Hanafusa, 1997; Martins-Green 2001; Feugate et al., 2002b). ECM interactions with endothelial cells are crucial in the cell migration and in the development of blood vessels during granulation tissue formation (Cockerill et al., 1995; Baldwin, 1996; Hanahan, 1997; Kumar et al., 1998; Li et al., 2003). Human umbilical vein endothelial cells migrate and arrange themselves in tubular structures when cultured for 12 h on a matrix isolated from Engelbreth-Holm-Swarm (EHS) tumors (a basement membrane-like matrix consisting primarily of laminin but also containing collagen IV, proteoglycans, and entactin/nidogen) (Kubota et al., 1988; Grant et al., 1989; Lawley and Kubota, 1989). When these cells are cultured on collagen I, however, tubular structures do not form in this period of time (Kubota et al., 1988), but if they are grown for a week inside collagen gels, giving the endothelial cells time to deposit their own basement membrane, tubes do develop (Montesano et al., 1983; Madri et al., 1988; Bell et al., 2001). The much more rapid tubulogenesis that occurs on EHS suggests that one or more components of the basement membrane plays an important role in the development of the capillary-like structures, a speculation confirmed both in culture and in vivo (Sakamoto et al., 1991; Grant et al., 1992). Indeed, preincubation of these endothelial cells with antibodies to laminin, the major component of basement membrane, prevents the formation of tubules in vitro (Kubota et al., 1988). Furthermore, synthetic peptides containing the sequence SIKVAV derived from the A chain of laminin induce endothelial cell adhesion and elongation and promote angiogenesis (Grant et al., 1992), while peptides containing the sequence YIGSR derived from the laminin B1 chain promote endothelial tube formation (Grant et al., 1989), although YIGSR peptides block angiogenesis in vivo (Sakamoto et al., 1991; Grant et al., 1992) and inhibit endothelial cell migration in vitro (Sakamoto et al., 1991). The mechanisms behind the ability of the YIGSR synthetic peptide to yield such different results in vivo may result from competition of this peptide with laminin for receptor binding, as this YIGSR peptide is known to block laminin binding to cells and block migration. If such competition does occur, the binding of the soluble YIGSR peptide to this receptor rather than YIGSR in the normal context of the complete laminin protein may alter downstream signaling events due to changes in the mechanical resistance and ligand presentation afforded by soluble, rather than intact, ligand, as has been suggested for integrin signaling (Vogel et al., 2001; Stupack and Cheresh, 2002). Regardless of the actual mechanism of action, the fact that soluble receptor-binding regions of ECM molecules may yield results different from those of the intact molecule may be of particular importance during matrix degradation, which releases ECM fragments. For example, matrix-degrading enzymes are activated during angiogenesis to facilitate the migration and invasion of endothelial cells into adjacent tissues and matrix; this matrix degradation may provide angiogenic or anti-angiogenic factors via release from the matrix or by appropriate cleavage of ECM molecules such as laminin (Werb et al., 1999; Rundhaug, 2005). In vivo, angiogenic sequences or factors could be provided locally, and when they have served their purpose, inhibition of further action could similarly be initiated by suitable cleavage to create CDPGYIGSR-NH2 or some other comparable factor present in the ECM (Sakamoto et al., 1991). Therefore, the way matrix molecules are locally cleaved and/or factors are locally released could have important consequences for the formation of the granulation tissue.

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Proliferation Immediately after wounding, the epithelium undergoes changes that lead to wound closure. During this re-epithelialization period, the keratinocytes trailing behind those at the front edge of migration replicate to provide a source of cells to cover the wound. Basement membrane-type ECM still present on the basal surface of these keratinocytes may be important in maintaining this proliferative state. In support of this possibility is the finding that during normal skin remodeling, fibronectin associated with the basal lamina of epithelia is crucial for maintaining the basal keratinocyte layer in a proliferative state for constant replenishment of the suprabasal layers (Nicholson and Watt, 1991). It has also been shown using a dermal wound model, that basement membrane matrices are able to sustain the proliferation of keratinocytes for several days (Dawson et al., 1996). The component of the basement membrane involved in this proliferation may be laminin, as laminin 10/11 can promote keratinocyte proliferation in vitro (Pouliot et al., 2002). In contrast, the fibrin-containing provisional matrix may prevent excessive keratinocyte proliferation, as the keratinocytes of fibrinogen-deficient mice do proliferate abnormally during re-epithelialization (Drew et al., 2001). As re-epithelialization is occurring, the granulation tissue begins to form. This latter tissue is composed of fibroblasts, myofibroblasts, monocytes/macrophages, lymphocytes, endothelial cells of the microvasculature, and ECM molecules, including embryonic fibronectin, hyaluronic acid, type III collagen, and small amounts of type I collagen (Clark, 1996). These ECM molecules, in conjunction with growth factors released by the platelets and secreted by the cells present in the granulation tissue, provide signals to the cells which lead to their proliferation (Tuan et al., 1996; Bissell, 1998). However, ECM molecules themselves such as fibronectin, as well as specific fragments of fibronectin, laminin, collagen VI, and SPARC/osteonectin, have been shown to stimulate fibroblast and endothelial cell proliferation (Bitterman et al., 1983; Panayotou et al., 1989; Atkinson et al., 1996; Grant et al., 1998; Kapila et al., 1998; Ruhl et al., 1999; Sage et al., 2003). In the case of laminin, this proliferative activity appears to be mediated by its EGF-like domains (Panayotou et al., 1989), suggesting a potential dependence upon the activation of EGFR (Schenk et al., 2003; Koshikawa et al., 2005). In contrast, ECM molecules and/or peptides derived from their proteolysis can have inhibitory effects on cell proliferation; intact decorin (Sulochana et al., 2005) and SPARC (Funk and Sage, 1991; Chlenski et al., 2005), as well as peptides derived from decorin (Sulochana et al., 2005), SPARC (Sage et al., 2003), collagens XVIII and XV (endostatin) (O’Reilly et al., 1997; Sasaki et al., 2000), and collagen IV (tumstatin) (Hamano et al., 2003) have anti-angiogenic effects due to their inhibition of endothelial cell proliferation. ECM molecules may also cooperate with growth factors in the proliferation of fibroblasts and the development of new blood vessels in the granulation tissue. During this angiogenic process, growth factors such as VEGFs and fibroblast growth factors (FGFs) associate with ECM molecules and stimulate proliferation of endothelial cells which then migrate to form the new microvessels (Miao et al., 1996; Ikuta et al., 2000, 2001; Sottile, 2004); indeed, recent studies suggest that some anti-angiogenic molecules, including thrombospondin and endostatin, may inhibit angiogenesis by competition with these growth factors for ECM binding (Gupta et al., 1999; Reis et al., 2005). Conversely, ECM–growth factor interactions can be inhibitory, for example, VEGF binding of SPARC can inhibit VEGFinduced proliferation (Kupprion et al., 1998). In addition, the proliferation stimulated by growth factors may be dependent upon the presence of specific ECM molecules; for example, TGF-β1 stimulation of fibroblast proliferation is dependent upon fibronectin (Clark et al., 1997). Differentiation As healing progresses during the formation of granulation tissue, some of the fibroblasts differentiate into myofibroblasts; they acquire the morphological and biochemical characteristics of smooth muscle cells by expressing α-smooth muscle actin (Desmouliere and Gabbiani, 1994; Desmouliere et al., 2005). Matrix molecules are important in this differentiation process. For example, heparin decreases the proliferation of

Cell–ECM Interactions in Repair and Regeneration

fibroblasts in culture and induces the expression of α-smooth muscle actin in these cells. In vivo, the local application of tumor necrosis factor α leads to the development of granulation tissue, but the presence of cells expressing α-smooth muscle actin was only observed when heparin was also applied (Desmouliere et al., 1992). These results suggest that some of the properties of heparin not related to its anticoagulant effects are important in the induction of α-smooth muscle actin. This function may be related to the ability of heparin and heparin sulfate proteoglycans to bind cytokines and/or growth factors, such as TGFβ that regulate myofibroblast differentiation (Kim. and Mooney, 1998; Kirkland et al., 1998; Menart et al., 2002; Li, J. et al., 2004). Specific interactions with the ECM are also important for myofibroblast differentiation; inhibition of the ED-A-containing form of fibronectin or αv or β1 integrins can block TGF-β1-mediated myofibroblast differentiation (Serini et al., 1998; Kato et al., 2001; Lygoe et al., 2004). In addition, cardiac fibroblasts undergo myofibroblast differentiation when plated on collagen VI (Naugle et al., 2005). Interstitial collagens have also been shown to play a role in the acquisition of the myofibroblastic phenotype. When fibroblasts are cultured on relaxed collagen gels or collagen-coated plates, they do not differentiate (Tomasek et al., 1992; Naugle et al., 2005); however, if they are grown on anchored collagen matrices where the collagen fibers are aligned (much like in the granulation tissue) they show myofibroblast characteristics (Bell et al., 1979; Arora et al., 1999). These observations led to the hypothesis that myofibroblast differentiation is regulated by mechanical tension; more recent studies in vivo, during wound healing, and in vitro have suggested that this hypothesis is, in fact, correct (Hinz et al., 2001; Wang et al., 2003). Apoptosis Apoptosis also plays a role during normal wound healing as the granulation tissue evolves into scar tissue. As the wound heals, the number of fibroblasts, myofibroblasts, endothelial cells, and pericytes decreases dramatically, matrix molecules, especially interstitial collagen, accumulate, and a scar forms (Clark, 1996). In this remodeling phase of healing, cell death by apoptosis leads to elimination of many cells of various types at once without causing tissue damage (Clark, 1996). For example, studies using transmission electron microscopy and in situ end-labeling of DNA fragments have shown that many myofibroblasts and endothelial cells undergo apoptosis during the remodeling process. Morphometric analysis of the granulation tissue showed that the number of cells undergoing apoptosis increases around days 20–25 after injury and this results in a dramatic reduction in cellularity after day 25 (Desmouliere et al., 1995); similar results were noted in cardiac granulation tissue following infarction (Takemura et al., 1998). Moreover, using model systems that mimic regression of granulation tissue, it has been shown that release of mechanical tension triggers apoptosis of human fibroblasts and myofibroblasts (Fluck et al., 1998; Grinnell et al., 1999; Bride et al., 2004). In these models, apoptotic cell death was regulated by interstitial-type collagens in combination with growth factors and mechanical tension and did not require differentiation of the fibroblasts into myofibroblasts, strongly suggesting that contractile collagens determine the susceptibility of fibroblasts of the wound tissue to undergo apoptotic cell death (Fluck et al., 1998; Grinnell et al., 1999). Further studies have also implicated the interactions between thrombospondin-1 and the αvβ3 integrin-CD47 complex in the mechanical tension-mediated stimulation of fibroblast apoptosis (Graf et al., 2002). Such apoptosis may be required for resolution of wound healing and the prevention of scarring. Indeed, fibroblast/myofibroblast apoptosis is reduced in keloid and hypertrophic scars, resulting in the excessive matrix accumulation and scarring (Ladin et al., 1998; Saed et al., 1998; Ishihara et al., 2000). In keloid scars, this decreased apoptosis may be due to p53 mutations and/or growth factor receptor over-expression (Ladin et al., 1998; Saed et al., 1998; Messadi et al., 1999; Ishihara et al., 2000; Moulin et al., 2004); in contrast, it is thought that apoptotic failure in hypertrophic scars results from an over-expression of tissue transglutaminase, leading to increased matrix breakdown and decreased collagen contraction (Linge et al., 2005). In addition to cell death by apoptosis, it has also been shown that bronchoalveolar lavage fluid collected during lung

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remodeling after injury can promote fibroblast cell death by a process that is distinct from that of necrosis or apoptosis (Polunovsky et al., 1993). Although this process of cell death has not been extensively studied, it suggests that there are other processes of programmed cell death which are distinct from apoptosis and occur preferentially in association with wound repair.

CELL–ECM INTERACTIONS DURING REGENERATION True tissue regeneration following injury rarely occurs in vertebrate species, but it does occur in specific instances, such as during in fetal cutaneous wound healing, liver regeneration, and urodele amphibian limb regeneration. Unlike wound healing in normal adult animals, which is characterized by scarring, fetal cutaneous wounds heal without fibrosis and scar formation, leading to regeneration of the injured area. Similarly, after injury, injured liver very effectively restores both normal function and normal organ size by proliferation and differentiation of pre-existing cell types. The contribution of cell–ECM interactions to regeneration in fetal healing and liver regeneration are discussed below (Figure 5.3). Fetal Wound Healing Adhesion and Migration Scarless fetal wounds have significant differences in cell–ECM interactions in the injured area when compared with scarring adult wounds; these changes occur due to alterations in the composition of the ECM molecules, the rate of their appearance after wounding, and their duration in the wound area. One crucial ECM molecule in fetal wound healing is hyaluronic acid, which appears to be necessary for the regenerative response; its removal from fetal wounds promotes a healing response more similar to that of adults (Mast et al., 1992), and treatment of normally scarring wounds or wound organ cultures with hyaluronic acid decreases scarring (Iocono et al., 1998a, b; Hu et al., 2003). Hyaluronic acid is present at higher levels (Krummel et al., 1987; Sawai et al., 1997) and for a longer duration in fetal skin wounds compared with adult wounds; the latter may result, in part, from the reduced activity of hyaluronidase in fetal wounds (West et al., 1997). Fetal fibroblasts also express higher levels of the hyaluronic acid receptor CD44 (Adolph et al., 1993), thus increasing receptor–ligand interactions which promote Healing with scar formation (adult healing)

Healing with regeneration (fetal healing)

↓ Hyaluronic acid, ↑ decorin, presence of ED-A fibronectin

↑ Hyaluronic acid, ↓ Decorin,

↑ TGF-1, disorganized collagen deposition

↓ TGF-1, ↑ collagen organization

↑ Myofibroblast differentiation ↑ contraction

↓ Myofibroblast differentiation ↓ contraction

↑ Scar formation ↓ Regeneration

↓ Scar formation ↑ Regeneration

Figure 5.3 A comparison of particular cell–ECM interactions occurring in scar-forming adult healing versus those occurring during regenerative fetal healing. As shown in this diagram, unique subsets of ECM molecules are associated with scarring versus regenerative healing. As such, therapeutic alteration of ECM composition may allow physicians to modulate healing to promote tissue regeneration. Additional therapeutic approaches may be generated upon further investigation into the importance of additional cell–ECM interactions in scarring and regenerative responses.

Cell–ECM Interactions in Repair and Regeneration

fibroblast migration (Huang-Lee et al., 1994). Increased fetal hyaluronic acid may also facilitate fibroblast migration by decreasing or preventing expression of TGF-β1, a factor that increases collagen I deposition (Ignotz and Massague, 1986) and inhibits fibroblast migration (Ellis et al., 1992; Hu et al., 2003). Tenascin C is induced more rapidly and to a greater extent in fetal wounds, thus modulating cell adhesion to fibronectin (Whitby and Ferguson, 1991; Whitby et al., 1991). Fibronectin levels also increase more quickly in fetal wounds than adult wounds (Longaker et al., 1989). This increased expression of tenascin and fibronectin is associated with concomitant increases in the expression of integrins that serve as their receptors. In particular, the α5 subunit, αvβ3, and αvβ6 integrins, which bind fibronectin and/or tenascin, are upregulated in the wounded fetal epithelium (Cass et al., 1998). The combined rapid increases in fibronectin and tenascin, coupled with increased expression of their respective integrin receptors in epithelial cells, are likely important in facilitating cell migration and re-epithelialization in fetal wounds. In addition, fetal fibroblasts produce more collagen (Adzick et al., 1985; Longaker et al., 1990; Lovvorn et al., 1999; Gosiewska et al., 2001), particularly collagen type III (Hallock et al., 1988), than adult cells, and the organization of the fibrils in the fetal wound appears normal, while that of the adult wound exhibits an organization indicative of scarring (Whitby and Ferguson, 1991). The changes in the collagen levels and organization in fetal wounds may result from the increased expression in fetal fibroblasts of the collagen receptor DDR1, which is important in collagen expression and organization (Chin et al., 2001). Furthermore, hyaluronic acid increases collagen synthesis in vitro, and may thus contribute to increased collagen deposition in fetal wounds (Mast et al., 1993). In spite of the increased collagen production by fetal fibroblasts, the fetal wounds do not exhibit excessive collagen deposition and fibrosis; this may be due to rapid turnover of these ECM components by proteasemediated degradation. For example, levels of urokinase plasminogen activator (uPA) and MMPs are increased while the levels of their endogenous inhibitors, PAI-1 and tissue inhibitor of metalloproteinases (TIMPs), are decreased in the fetal wounds, ultimately promoting matrix degradation and turnover (Huang et al., 2002; Peled et al., 2002; Dang et al., 2003). Not only does this prevent fibrosis, it also likely facilitates cell migration by reducing matrix density and increases the generation of proteolytic matrix fragments that modulate various stages of wound repair, as mentioned above for laminin and collagen fragments that can alter angiogenesis during granulation tissue formation. Proliferation As mentioned above, during fetal wound healing, increased levels of hyaluronic acid are present and in vitro studies indicate that hyaluronic acid decreases fetal fibroblast proliferation (Mast et al., 1993). However, early studies comparing fetal wounds with those of newborns and adults showed an increase in fibroblast number in the wounded area in the fetal wounds, and fetal fibroblasts proliferate more rapidly than adult cells (Adzick et al., 1985; Khorramizadeh et al., 1999). It is unclear how these findings may be reconciled; however, it is possible that hyaluronic acid prevents excessive fibroblast proliferation in fetal wounds. Another critical event in wound healing is re-epithelialization, which requires both keratinocyte migration and proliferation. Keratinocyte proliferation is decreased in mice lacking CD44 expression in keratinocytes (Kaya et al., 1997), suggesting that interactions between hyaluronic acid and CD44 may be important for keratinocyte proliferation during healing, and thus more effective re-epithelialization. This finding may explain, in part, the enhanced rate of healing seen in wounds treated with hyaluronic acid. Differentiation Fetal wounds have a decreased number of myofibroblasts, which appear in the wounded site earlier and remain a shorter time than in adult wounds; in fact, one study showed a lack of α-smooth muscle actin-expressing myofibroblasts in the wounds of early-stage fetuses (Estes et al., 1994). This is associated with a general lack of

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contraction in the fetal wounds themselves (Krummel et al., 1987). Increased levels of hyaluronic acid present during fetal wound healing may alter the differentiation and/or contractility of myofibroblasts in the wound site; studies in vitro have shown that addition of hyaluronic acid decreases fibroblast contraction of collagen matrices (Huang-Lee et al., 1994). This may be due, in part, to reduced expression of TGF-β1, a major inducer of myofibroblast differentiation and fibrosis. Indeed, incisional adult wounds treated with hyaluronic acid healed more rapidly with a significant decrease in TGF-β1 levels (Hu et al., 2003). The large amounts of hyaluronic acid in fetal wounds may thus explain the greatly reduced levels of TGF-β1 in fetal wounds (Nath et al., 1994; Chen et al., 2005). Downregulation of TGF-β1 in adult wounds produces a decrease in scarring similar to that observed with hyaluronic acid treatment (Choi et al., 1996). Conversely, studies have shown that the addition of TGF-β1 to normally scarless fetal wounds induces a more scarring phenotype, with myofibroblast differentiation, wound contraction, and fibrosis (Lin et al., 1995; Lanning et al., 1999). Thus, hyaluronic acid-mediated inhibition of TGF-β1 expression may be critical in scarless fetal healing. If any TGF-β1 is present during fetal wound healing, it may be regulated by inhibitory ECM molecules present in the injured area. One such inhibitor is fibromodulin, which is capable of binding TGF-β1 and preventing receptor binding and is expressed to a greater extent in fetal wounds relative to adult wounds (Hildebrand et al., 1994; Soo et al., 2000). Another molecule that may alter TGF-β1 activity is decorin, although the function of decorin in modulating TGF-β1 activity is somewhat controversial; some studies indicate that decorin binding decreases TGF-β1 activity (Noble et al., 1992), while others suggest that this interaction either has no effect on TGF-β1 or even actually increases activity (Hausser et al., 1994; Takeuchi et al., 1994). The outcome of decorin– TGF-β1 binding may depend upon the microenvironment, and this has not been extensively studied in fetal wounds. Regardless, decorin levels are decreased in scarless wounds, resulting in decreased decorin–TGF-β1 interactions and altered TGF-β1 activity (Beanes et al., 2001). Decreased activity of this growth factor, combined with extremely low levels of expression in fetal wounds, results in decreased fibrosis, myofibroblast differentiation, and wound contraction, leading to regeneration rather than scarring. Apoptosis Little is known regarding the apoptotic process in fetal wounds, and whether this differs from that of adult wounds. However, as in adult healing, multiple cell types present within the fetal granulation tissue likely disappear via apoptosis. It is also apparent that any myofibroblasts that do differentiate during granulation tissue formation disappear rapidly (Estes et al., 1994), perhaps due to an altered rate of apoptosis in these wounds. If changes in apoptotic efficiency do indeed occur, they may result from the decreased contraction, and thus decreased mechanical tension, in fetal wounds (Krummel et al., 1987), as well as altered collagen levels within the collagen matrix (Adzick et al., 1985; Longaker et al., 1990; Lovvorn et al., 1999; Gosiewska et al., 2001). It is also possible that apoptosis is not as critical in the healing of fetal wounds as in adult wounds; leukocyte influx and myofibroblast differentiation appear to be minimal in fetal wounds, and thus may not require large numbers of cells to undergo apoptosis for regeneration to occur (Estes et al., 1994; Harty et al., 2003). Liver Regeneration Adhesion and Migration ECM–cell interactions are also altered during mammalian liver regeneration, leading to changes in adhesion and migration. One major molecule upregulated after liver injury is laminin (Martinez-Hernandez et al., 1991; Kato et al., 1992). Hepatocytes isolated soon after liver injury and plated on laminin attach more efficiently than non-injured hepatocytes suggesting a concomitant increase in laminin-binding integrins (Carlsson et al., 1981; Kato et al., 1992). Collagens I, III, IV, and V increase in regenerating liver several days after injury. Hepatocytes isolated from this stage of regenerating liver show increased adhesion to collagen, which may

Cell–ECM Interactions in Repair and Regeneration

indicate increased expression of collagen adhesion receptors (Kato et al., 1992). The increased levels of laminin and collagen IV during regeneration may also promote hepatocyte migration, as both the basal and stimulated migration of hepatocytes is enhanced on laminin and collagen IV relative to other types of ECM (Ma et al., 1999). Proliferation In response to liver injury, hepatocytes proliferate to restore normal liver function and size. In vitro studies show that laminin enhances hepatocyte proliferation in general and in response to EGF; thus, the increased laminin present in regenerating tissue may facilitate proliferation (Hirata et al., 1983; Kato et al., 1992). Both the mRNA and the protein levels of plasma fibronectin and its receptor α5β1 integrin increase in regenerating liver following injury (Gluck et al., 1992; Kato et al., 1992; Pujades et al., 1992), which may also increase proliferation. Indeed, intraperitoneal injection of plasma fibronectin further stimulates proliferation in the regenerating liver (Kwon et al., 1990b). The primary growth factor responsible for hepatocyte proliferation is HGF; thus, processes that stimulate HGF production and/or release from matrix components will also increase hepatocyte numbers in regenerating liver. Heparan sulfate proteoglycans that are upregulated after injury bind HGF and promote its mitogenic activity (Matsumoto et al., 1993; Kato et al., 1994; Lai et al., 2004). Various proteoglycans are also upregulated after injury, potentially increasing HGF activity in the regenerating liver (Otsu et al., 1992; Gallai et al., 1996). Other ECM molecules are known to bind HGF with low affinity, possibly sequestering HGF in the ECM and preventing its activity (Schuppan et al., 1998). In fact, increased MMP expression during regeneration stimulates ECM degradation and hepatocyte proliferation. This increased proliferation is likely due to the proteolytic processing and release of matrix-bound HGF (Nishio et al., 2003; Mohammed et al., 2005). Increases in MMP production are followed by increased TIMP expression, which may prevent excessive hepatocyte proliferation (Rudolph et al., 1999; Mohammed et al., 2005). HGF, and thus hepatocyte proliferation, can also be activated by plasmin, suggesting a role for plasminogen activators in liver regeneration (Shimizu et al., 2001). Indeed, rapid increases in uPA activity after injury is followed by increases in plasmin activation and fibrinogen cleavage and a rapid loss of fibronectin, laminin, and entactin via proteolysis, although the levels of these latter proteins increase at later stages of healing (Kim et al., 1997). The importance of plasmin activation is underscored by studies in which the livers of uPA and tissue plasminogen activator (tPA) single and double knockout mice or plasminogen knockout mice were injured chemically (Bezerra et al., 1999, 2001). It was found that the plasminogen and uPA single knockouts, as well as the uPA/tPA double knockouts experienced significant liver regenerative problems accompanied by excessive fibrin and fibronectin, with a lesser effect seen in the tPA knockout. The observed disruption of regeneration may be due to a reduction of hepatocyte proliferation resulting from decreased HGF activity. Differentiation Myofibroblast differentiation can also occur from the stellate cells of the liver, which can then stimulate excessive ECM deposition, leading to fibrosis and cirrhosis rather than regeneration. Thus, myofibroblast differentiation must be very limited to allow appropriate liver regeneration. Plasma fibronectin levels are increased in the liver regenerating tissue, but are reduced in cirrhotic tissue (Kwon et al., 1990a; Chijiiwa et al., 1994). In addition, myofibroblast differentiation appears to require the ED-A domain of fibronectin (Serini et al., 1998; Kato et al., 2001), which is lacking in plasma fibronectin. These results, when taken together, suggest the possibility that plasma fibronectin may limit myofibroblast differentiation and fibrosis in the liver. This may be particularly important, given the increased quantity and activation of TGF-β1, TGF-β2, and TGF-β3 in the regenerating liver, which would otherwise promote differentiation and fibrosis (Jakowlew et al., 1991). In contrast, the stellate

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cell differentiation state may be maintained by the basement membrane, which appears to both maintain the differentiation state of stellate cells and, in vitro, promote myofibroblast de-differentiation back to stellate cells (Friedman et al., 1989; Sohara et al., 2002). Apoptosis In liver regeneration, prevention of hepatocyte apoptosis is critical for regeneration, while increased apoptotic rates are associated with impaired regeneration. Indeed, extensive cell death following a large liver resection leads to liver failure rather than regeneration (Panis et al., 1997). Liver ischemia–reperfusion injury can also promote apoptosis and liver failure rather than regeneration (Takeda et al., 2002). In the latter case of ischemia–reperfusion injury, prevention of apoptosis can significantly reduce the incidence of liver failure, underscoring the relationship between apoptosis and impaired regeneration or failure (Vilatoba et al., 2005b). The lack of regeneration in such cases is associated with the upregulation of pro-apoptotic gene expression and the downregulation of pro-survival genes (Morita et al., 2002), and may thus be related to the inability of hepatocytes to proliferate under such pro-apoptotic conditions (Iimuro et al., 1998). This hypothesis is supported by studies indicating that apoptosis and liver failure resulting from extensive liver resection or ischemia–reperfusion injury can be largely prevented by treatment conditions that promote cell proliferation (Longo et al., 2005; Vilatoba et al., 2005a). The prevention of apoptosis may thus require ECM molecules that are important in promoting hepatocyte proliferation, including laminin (Hirata et al., 1983; Kato et al., 1992), plasma fibronectin (Kwon et al., 1990b), and HGF-binding proteoglycans (Matsumoto et al., 1993; Kato et al., 1994; Lai et al., 2004). Different MMPs are activated after ischemia–reperfusion injury when compared with forms of injury that regenerate (Cursio et al., 2002), perhaps leading to the degradation of a different profile of ECM proteins; the activation of specific MMPs is thought to promote hepatocyte proliferation by releasing matrix-sequestered HGF (Nishio et al., 2003; Mohammed et al., 2005). The activation of different MMPs and cleavage of different substrates may alter HGF release and subsequent proliferation, leaving these cells more susceptible to apoptosis. This idea is supported by a study in which liver with ischemia–reperfusion injury was treated with an MMP inhibitor, which decreased apoptosis and necrosis in the injured liver (Cursio et al., 2002). Although apoptosis of hepatocytes disrupts the regenerative process, apoptosis of myofibroblastic hepatic stellate cells may be critical in preventing fibrosis and scarring during regeneration (Issa et al., 2001). These myofibroblastic hepatic stellate cells disappear via apoptosis (Saile et al., 1997; Issa et al., 2001), and also potentially by de-differentiation back to stellate cells (Friedman et al., 1989; Sohara et al., 2002). The apoptosis of these myofibroblastic cells seems to be dependent upon the activation of specific proteases and the subsequent degradation of matrix components. Mice expressing a collagen I gene that is resistant to proteolysis had decreased stellate cell myofibroblast apoptosis and increased fibrosis, and thus impaired regeneration, relative to wild type (Issa et al., 2003). These myofibroblasts also persist in plasminogen-deficient mice, and are associated with a general accumulation of non-degraded matrix components (Ng et al., 2001), further supporting a role for matrix degradation in the observed apoptosis. The matrix degradation important in apoptosis also likely involves the activation of MMPs, as inhibition of MMP activity using synthetic inhibitors or TIMP-1 (Murphy et al., 2002; Zhou et al., 2004) prevents apoptosis of myofibroblastic stellate cells in vitro, whereas MMP-9 activity promotes apoptosis of these cells (Zhou et al., 2004). In in vitro models of cutaneous wound healing, a release of mechanical tension within the collagen matrix (Fluck et al., 1998; Grinnell et al., 1999; Bride et al., 2004) can promote myofibroblast apoptosis. It is possible that a similar release of mechanical tension, perhaps via cleavage of collagen I, is critical for myofibroblast apoptosis in the liver. Proteolysis of ECM components may also contribute to stellate cell apoptosis by abolishing integrin signaling downstream of binding to these components. Experimental disruption of ECM–integrin binding via an RGD-containing peptide (Iwamoto et al., 1999) or

Cell–ECM Interactions in Repair and Regeneration

various αvβ3 antagonists (Zhou et al., 2004) induce stellate cell apoptosis in vitro, further supporting a role for integrin-mediated signaling in this apoptotic event.

IMPLICATIONS FOR REGENERATIVE MEDICINE One primary goal of studies comparing differences in cell–ECM interactions, and thus changes in signaling, that accompany regenerative and non-regenerative healing is to determine what types of interactions promote and which inhibit tissue regeneration (for an example, see Figure 5.3). After elucidating the functions of particular interactions, it may be possible to increase the regenerative response through (1) the induction of proregenerative ECM molecules or signaling events in the wounded area combined with (2) the antagonism of anti-regenerative/scarring interactions or signaling events using specific inhibitors. This discussion of regenerative medicine will focus upon possible strategies to promote regeneration in adult scarring wounds, thus causing adult wounds to more closely resemble fetal scarless wounds. Such an increased regenerative response would be particularly useful in the treatment of wounds that heal abnormally with increased scar formation, such as keloids and hypertrophic scars, ischemic reperfusion injury, and chronic inflammatory responses. Different types of approaches may be used to increase pro-regenerative ECM levels in the wounded area, including the introduction of these molecules via direct application of the molecules themselves, through the addition of agents that increase their expression, or through the addition into the wound sites of cells producing these types of ECM that have been prepared to minimize immunogenicity. Several different ECM molecules are present at higher levels in fetal wounds than in adult wounds, including hyaluronic acid, tenascin, fibronectin, and collagen III (Krummel et al., 1987; Hallock et al., 1988; Longaker et al., 1989; Whitby and Ferguson, 1991; Whitby et al., 1991; Sawai et al., 1997), and may play important roles in the regeneration process. Thus, altering the levels of these molecules in a scarring wound may improve regeneration. Indeed, preliminary experiments in rat wounds suggest that hyaluronic acid treatment decreases both the time required for healing and the amount of scar formation (Hu et al., 2003), underscoring the potential for this molecule in therapeutics. It is possible that treatment with tenascin, fibronectin, or collagen III in addition to hyaluronic acid could yield even more favorable outcomes. When attempting to promote regeneration, it is also imperative to inhibit events associated with scarring, including excessive ECM deposition, fibrosis, and contraction. During the adult healing process, these scarassociated processes are primarily controlled by the myofibroblast, a differentiated cell type that arises during the adult healing process but that is largely absent throughout fetal wound healing. As such, inhibition of myofibroblast differentiation or function along with the addition of pro-regenerative molecules may facilitate a stronger regenerative response. Inhibition of differentiation could be accomplished by blocking the factors that normally stimulate this process, such as TGF-β1 (Lin et al., 1995; Lanning et al., 1999) and IL-8 (Feugate et al., 2002a), or by preventing fibroblast–ECM interactions that facilitate myofibroblast differentiation, such as ED-A-containing fibronectin (Serini et al., 1998; Kato et al., 2001). Hyaluronic acid and fibromodulin appear to decrease TGF-β1 levels and activity, respectively, thus treatment of normally scarring wounds with these matrix components may thus decrease TGF-β1-mediated scarring (Hildebrand et al., 1994; Soo et al., 2000; Hu et al., 2003). IL-8, on the other hand, is a chemokine that activates G-protein linked receptors, which are highly amenable to inhibition by small molecules, which could be used to reduce the effects of this chemokine on myofibroblast differentiation. In summary, the recent surge in research regarding the ECM molecules themselves and their interactions with particular cells and cell-surface receptors has led to the realization that such interactions are many and complex, and that they are of the utmost importance in determining cell behavior during such events as wound repair and tissue regeneration. As such, the manipulation of specific cell–ECM interactions has the potential to modulate particular aspects of the repair process in order to promote a regenerative response.

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6 Developmental Mechanisms of Regeneration David L. Stocum

INTRODUCTION All the cell types of the adult vertebrate body are derived from stem cells of the early embryo. In mammals, these embryonic stem cells (ESCs) constitute the inner cell mass of the pre-implantation blastocyst (Rossant, 2001). ESCs are pluripotent, as demonstrated in vivo by their ability to make contributions to all tissues after injection into host blastocysts, their ability to form teratomas containing ectodermal, mesodermal and endodermal derivatives when implanted into immunodeficient mice, and their ability to differentiate spontaneously, or as directed, into multiple cell types in vitro. They express species and stage-specific embryonic antigens (SSEAs), alkaline phosphatase, and high levels of telomerase (Smith, 2001; Rippon and Bishop, 2004). ES cell lines have been established from a variety of vertebrate early embryos, including fish, birds, mice (Smith, 2001; Rippon and Bishop, 2004) and humans (Thomson et al, 1998; Shamblott et al, 1998). Several transcription factors have been implicated in the acquisition and maintenance of mouse and human ESC self-renewal and pluripotency. These are OCT4 (Smith, 2001), SOX2 (Avilion et al 2003), Fox D3 (Hanna et al, 2002), all activated by LIF (mouse) or FGF-2 (human) through STAT-3, and the LIF/STAT3independent transcription factors Nanog (Mitsui et al, 2003; Chambers et al, 2003), Tbx3, Esrrb, and Tcl1 (Ivanova et al, 2006). BMPs also play a role, by inducing the expression of inhibitor of differentiation (Id) genes via Smad transcription factors (Ying et al, 2003). ESCs give rise during cleavage to prospective ectoderm, endoderm and mesoderm cells. Once these cell types are established, they undergo the morphogenetic movements of gastrulation to position the mesoderm between the ectoderm and endoderm. Nanog is down-regulated as prospective mesoderm cells exit from the primitive streak during gastrulation, while OCT4 is still expressed. Subsequently all the pluripotency genes are down-regulated except in the germ cells (Hart et al, 2004). Cell interactions among these three embryonic tissue layers determine patterns of gene activity that establish the boundaries of organ and appendage fields. The distinct patterns of growth, tissue differentiation and morphogenesis that characterize the organs and appendages emerge as a result of further cell interactions within these fields. The cell and tissue interactions of development take place via autocrine, paracrine and juxtacrine signaling molecules that bind to receptors and activate intracellular signal transduction pathways leading to specific patterns of gene activity. Seven major signaling pathways have been identified: Notch, Wnt, hedgehog, JAK-SAT, RTK (receptor tyrosine kinase), TGF-β and the apoptotic, or cell death, pathway, which is important for eliminating excess cells in developing tissues and for morphogenesis (Gilbert, 2006).

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Regeneration is a developmental process that maintains (in the face of normal cell turnover) and restores (after damage) tissue integrity in the fetus, juvenile and adult. In most cases, it involves a partial recapitulation of tissue embryogenesis. The same signal transduction pathways and transcription factors are used in regeneration as in embryonic development, although they may not be deployed in exactly the same way. This chapter examines the mechanisms of regeneration, examples of each mechanism, and the strategies of regenerative medicine that are being developed from our knowledge of these mechanisms.

MECHANISMS OF REGENERATION There are three mechanisms of regeneration: compensatory hyperplasia, activation of resident adult stem cells (ASCs), and production of stem cells by the dedifferentiation of mature cells (Figure 6.1, Table 6.1). In all of these mechanisms, the regeneration-competent cells of adult tissues reside in three-dimensional environmental “niches” consisting of specific combinations and concentrations of soluble factors and extracellular matrix (ECM) that promote their survival, precisely regulate their proliferation, and determine the phenotypic direction and histological pattern of their differentiation (Scadden, 2006; Engler et al., 2006). Comprehending the elements and interlocking pathways of this “molecular ecology” (Powell, 2005) is one of the most important tasks of regeneration research today. Compensatory Hyperplasia Compensatory hyperplasia is defined as the mitosis of differentiated cells to maintain or restore tissue mass. New cells thus are derived solely from pre-existing differentiated cells. This is the only mechanism of regeneration, that does not recapitulate part of the embryonic developmental program. The classic example of regeneration by compensatory hyperplasia is the mammalian liver (Michalopoulos and De Francis, 1997; Fausto, 2004). Individual hepatocytes have an enormous capacity for replication, up to at least 70 times. They are maintained in a non-proliferative state by C/EBPα inhibition of cyclin-dependent kinases (cdks). Partial hepatectomy triggers the appearance of TNF-α, IL-6, HGF and EGF, mitogenic signals that prime the hepatocytes for entry into the cell cycle by activation of the transcription factors STAT3,

(a)

(b)

(c)

(d)

Figure 6.1 Mechanisms of regeneration. (A) Compensatory hyperplasia, the division of differentiated cells to restore tissue mass. (B) Activation and proliferation of adult stem cells. The mother stem cell self-renews while also giving rise to a transit amplifying cell that proliferates and gives rise to single or multiple types of terminally differentiated cells. (C) Dedifferentiation of muscle (left) by cellularization and loss of contractile apparatus to produce mesenchymal-like stem cells (right). (D) Epithelial (left) to mesenchymal (right) transformation and mesenchymal to epithelial transformation.

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Table 6.1 Mechanisms of regeneration and types of cells involved Compensatory Hyperplasia Hepatocytes • Liver Beta cells • Pancreas Activation of Adult Stem Cells Epithelial Stem Cells • Digestive tract (including canals of Hering and pancreatic ductules) • Respiratory tract • Interfollicular epidermis • Hair follicle (bulge) • Neural (olfactory epithelium, lateral ventricles of brain, hippocampus, hair cells of cochlea) • Kidney proximal tubules Endothelial Stem Cells • Bone marrow • Capillaries and venules • Epicardium? Hematopoietic Stem Cells • Bone marrow Mesenchymal Stem Cells • Bone marrow • Periosteum, endosteum

• Dental pulp, periodontal ligament • Adipose tissue • Connective tissue compartments Muscle Stem Cells • Skeletal muscle (satellite cells) • Myocardium (cardiac stem cells) Dedifferentiation Amphibian • Tail • Limb • Jaws • Lens, retina • Myocardium • Intestine, fish fins Fish • Fins • Retina • Myocardium Lizard • Tail Epithelial   Mesenchymal Transformation • Amphibian spinal cord • Capillaries and venules • Kidney proximal tubules

PHF/NF-κB, AP-1 and C/EBPβ. These transcription factors induce the activity of sets of “early immediate” and “delayed immediate” genes that encode proteins involved in entering and progressing through the G1 phase of the cell cycle. HGF appears to play a central role in this process. Pro-HGF is released by liver matrix degradation and its synthesis by sinusoidal endothelial cells is promoted by VEGF (Le Couter et al, 2003). Pro-HGF is activated by urokinase plasminogen activator (uPA) and triggers entry into the cell cycle by binding to its receptor, c-met. Once the original mass of the liver is attained, proliferation ceases and the original histological architecture of the liver is restored. Beta cells and acinar cells of the pancreas also appear to regenerate in vivo by compensatory hyperplasia. Genetic marking experiments have revealed that during growth of the mouse pancreas or during its injuryinduced regeneration, new β-cells and acinar cells are derived from pre-existing β and acinar cells (Dor et al, 2004; Desai et al, 2007). Beta cell regeneration can be initiated by a number of proteins: β-cell regeneration protein (Reg), islet neogenesis associated protein (INGAP, a 15 amino acid fragment of Reg), betacellulin (a member of the EGF family), and GLP-1 (Risbud and Bhonde, 2002; Bonner-Weir and Weir, 2005). Activation of Adult Stem Cells Adult stem cells (ASCs) are arrested in a pre-terminal differentiation phase of their developmental program, within their tissues of residence. Differentiation of the tissue of residence results in the creation of niche conditions that balance quiescence and activation of ASCs. The mechanisms that sequester small subpopulations of stem cells as other cells differentiate around them are not well understood. When activated, ASCs divide

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asymmetrically so that one daughter remains a stem cell (self-renewal) and the other gives rise to a transit amplifying population that undergoes terminal differentiation. Epithelia, dental tissues, tissues of the nervous system, musculoskeletal tissues, and hematopoietic cells regenerate via ASCs (Table 6.1). The ASCs of different tissues are commonly maintained in a quiescent state either directly by the Notch signal transduction pathway, or indirectly by absence or inhibition of signaling molecules for other signaling transduction pathways. Epithelial Stem Cells Nearly all of the epithelial tissues of the body undergo continual self-renewal throughout life and have a high capacity for injury-induced regeneration. Interfollicular epidermis and hair follicles are among the best studied. The stem cells that regenerate interfollicular skin epidermis are integrin-expressing cells located in the stratum basale. During maintenance regeneration, they give rise to transit amplifying cells that detach from the basement membrane and differentiate into keratinocytes as they move upward to the stratum corneum (Jones et al, 1995; Jensen et al, 1999). Gaps in the epidermis of an excisional wound are filled in by the division of stem cells at the edges of the wound whose progeny migrate laterally through the provisional fibrin matrix of the wound. Migration is initiated by TGF-α and EGF produced by macrophages. Cell division at the wound edges is promoted by KGF and GM-CSF secreted by fibroblasts of the granulation tissue (Werner et al, 1994; Woodley, 1996). Once the wound is re-covered, these same factors promote vertical division to thicken the epidermis. The trigger for stem cell proliferation after wounding may be the binding of EGF family ligands on the apical cell surface to EGF receptors on the basolateral surface (Vermeer et al, 2003). Normally, tight junctions separate the apical and basolateral domains, but the cell separation that occurs upon wounding allows ligands and receptors of the two domains to interact. The basal epidermal cells are continuous with the basal cells of the outer root sheath of the hair follicle. Hair follicle stem cells are located in a special region of the outer root sheath called the bulge. Marking studies have shown that these stem cells divide asymmetrically to self-renew and produce transit-amplifying cells that feed upward toward the surface of the epidermis where they differentiate into epidermal keratinocytes, and downward to the matrix of the hair follicle where they proliferate and differentiate as the hair shaft (Morris et al, 2004; Tumbar et al, 2004). Hairs go through a three-stage maintenance cycle of catagen (follicle regression), telogen (follicle rest) and anagen (regeneration of the follicle and new hair growth) that is regulated by growth factor signals from the dermal papilla at the base of the hair follicle (Hardy 1992; Messenger, 1993). Transcriptional repression by Lef-1/Tcf-3 maintains the stem cells of the epidermis and hair follicles in a quiescent state. The cells are activated to proliferate by Wnt signaling, which stabilizes β-catenin, allowing it to translocate to the nucleus and complex with Lef1 to form a transcriptional activating complex leading to proliferation. Corneal epithelium is regenerated continuously or after injury by epithelial stem cells located in the limbus, the region where the cornea undergoes a transition into the sclera of the eye (Cotsarelis et al., 1989). Limbal stem cells divide asymmetrically to produce transit amplifying cells that migrate centripetally to replace corneal epithelial cells lost by turnover or injury. The epithelia of the digestive, respiratory and urogenital systems have extensive capacity for regeneration that is regulated by overlapping sets of growth factors (Stocum, 2006, for review). Small intestinal epithelial stem cells are located in the crypts of Lieberkuhn (Potten, 1997; Brittan and Wright, 2004). Liver stem cells are located in the epithelium of the canals of Hering and are activated when the ability of hepatocytes to proliferate by compensatory hyperplasia is compromised (Dabeva and Shafritz, 2003). The ductules of the pancreas may also harbor stem cells that can differentiate into β-cells, and non-β-cells of the islets have been reported to transform in vitro into epithelial cells that differentiate into β-cells (Bonner Weir and Weir, 2005; Jamal et al., 2005). Overexpression of the Arx gene (which determines the embryonic differentiation of α and PP cells of the islets) in β-cells converts them to α and PP cells (Collombat et al., 2007). A subset of type I pneumocytes

104 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

in the alveolar epithelium regenerates injured type I pneumocytes (Reddy et al., 2004). The ciliated epithelium of the trachea and bronchial tree is constantly renewed by stem cells in the basal epithelium, as is the epithelium of the bladder, ureters and urethra (Ham and Cormack, 1979). Some types of neurons regenerate from epithelial stem cells. The sound and motion-sensing auditory hair cells of the sensory epithelium of the cochlea and vestibular apparatus in birds are regenerated by support stem cells interspersed with these sensory neurons (Stone and Rubel, 2000; Stone et al., 2004). Mammalian hippocampal NSCs exhibit a low level of maintenance regeneration that in mice is enhanced by environmental enrichment (Kempermann et al., 1998; Gage, 2000, for review). Neurons of the olfactory nerve and olfactory bulb turn over on a regular schedule. The olfactory nerve neurons are regenerated by NSCs in the nasal epithelium, while NSCs in the walls of the lateral ventricle replenish olfactory bulb neurons (Schwob, 2002, for review). Killing thalamic projection neurons in the cortex of the mammalian brain or granule neurons in the dentate gyrus of the hippocampus by focal or global ischemia results in a low level of NSC proliferation in the lateral ventricle walls and hippocampal ventricle walls. Intraventricular injection of a combination of FGF-2 and EGF elevates the number of regenerated hippocampal neurons to 40% of the number lost; these neurons are functionally integrated into the hippocampal circuitry. The small number of regenerated neurons in the absence of these growth factors may be due to inadequate output of growth factors by astrocytes, since neonatal hippocampal astrocytes induce hippocampal NSCs to differentiate into neurons in vitro, whereas adult astrocytes have only half the effect (Nakatomi et al., 2002; Song et al., 2002). Larval salamanders regenerate the spinal cord after amputation of the tail. Muscle, cartilage and connective tissue dedifferentiate to form a blastema. A tube of dividing ependymal cells extends from the cut end of the spinal cord into the blastema. As the ependymal tube grows distally, the cells closest to the amputation plane extend end feet. The endfeet form channels that promote the regeneration of axons from above the level of transection, while other ependymal cells differentiate into new motor neurons, interneurons and glia (Chernoff et al., 2003). FGF-2, Wnt, BMP and Notch signaling pathways all appear to be involved in regulating this regeneration. In order to migrate during regeneration, many epithelial stem cells, as well as the endothelium of blood vessels, undergo an epithelial to mesenchymal transformation (EMT), followed by the reverse mesenchymal to epithelial transformation (MET) to reconstitute the epithelium or endothelium (Fig. 6.1, Table 6.1). Well-studied examples of these transformations are the regeneration of wounded epidermis, the ependyma of the transected thoracic or lumbar spinal cord of urodele amphibians (Chernoff et al., 2003) and proximal tubule epithelial cells of the mammalian kidney (Bonventre, 2003). The cells of wounded epidermis are induced to migrate by macrophage-produced EGF and TGF-α. Kidney tubule epithelial cells are induced by TGF-β1 to undergo EMT to cover denuded areas of the basement membrane. Once having filled the gap, the mesenchymal cells are induced to undergo MET by BMP-7 (Zeisberg et al., 2003). In the case of urodele spinal cord regeneration, EMT produces a mass of cells that bridge the gap, followed by MET to restore the ependyma. The ependymal cells form endfeet that project to the glia limitans and form channels that support axon regeneration. In all cases, intermediate filament expression alternates between epithelial markers (cytokeratins) and mesenchymal markers (actin locomotory filaments, vimentin). The bone marrow harbors endothelial stem cells that circulate in the blood. Circulating EnScs have the phenotype [CD133 VEGFR2]+ and express the receptor for the chemoattractant, stromal cell derived factor 1 (SDF-1). These cells are recruited to sites of injury by SDF-1 and angiogenic factors such as VEGF-A and placental growth factor (PLGF). There, they are incorporated into regenerating blood vessels. EnSCs make only a minor contribution to the construction of the new vessels, which takes place primarily by sprouting from existing vessels (Stocum, 2006, for review). Endothelial cells in the injured vessel wall are induced by FGF-2, TGFβ1, IL-8 and TNF-α to undergo EMT. The activated cells may be a subpopulation in the vessel wall similar to EnSCs (Ingram et al., 2004). They lose their intercellular junctions and express proteases that break down their

Developmental Mechanisms of Regeneration 105

basement membrane. In the presence of VEGF synthesized by the epidermis of a healing wound, the cells proliferate and migrate out as mesenchymal cords into the fibrin matrix of the wound. PD-ECGF and TNF-α are chemotactic for endothelial cells; TGF-α, TGF-β, FGF 1 and 2, and PDGF-B stimulate proliferation (Madri et al., 1996; Tomanek and Shatteman, 2000). The proliferating cells then undergo MET and rearrange themselves into endothelial tubes, a process mediated by laminin and a fibroblast-secreted protein, Egfl7 (Parker et al., 2004). Mesenchymal Stem Cells Mesenchymal stem cells were first isolated from the bone marrow as an adherent cell capable of differentiating into fibroblasts, chondrocytes, osteoblasts and adipocytes. They are responsible primarily for the regeneration of bone, tendon and ligament, but also for a limited regenerative capacity of dentin in adult teeth. Bones regenerate after fracture by the proliferation of MSCs residing in the bone marrow, endosteum and periosteum. In fractured membrane (flat) bones, the MSCs differentiate directly to osteoblasts that secrete the bone matrix. The MSCs of fractured endochondral (long) bones first differentiate into a chondrocyte template that is subsequently replaced by osteoblasts (Ham and Cormack, 1979). The molecular mediators of fracture repair appear to be identical to those involved in embryonic bone development. BMPs, TGF-β, FGF-1 and 2, PDGF and IGF-1 expressed by the MSCs regulate chondrocyte differentiation (Bostrom, 1998; Einhorn, 1998; Trippel, 1998). The transcription factor Sox-9 activates the expression of type I, IX and collagen genes and the gene for aggrecan protein. Ihh signaling pathway components are expressed in a population of cells on the periphery of the soft callus that will reform the periosteum, indicating that the same mechanism used to regulate the rate at which chondrocytes mature during the embryonic development of long bones is operative during fracture healing. As the cartilage template is replaced by osteoblast invasion, the expression of genes involved in osteoblast differentiation, such as Runx2 and osteocalcin, is detected (Ferguson et al., 1998). Teeth contain two types of stem cells in the pulp (Gronthos et al., 2000; Shi et al., 2001; Miura et al., 2002). One type has been isolated from adult teeth. It is similar to the bone marrow mesenchymal stem cell and differentiates into odontoblasts that make new dentin to counter the loss of odontoblasts destroyed by trauma or bacterial invasion (Murray and Garcia-Godoy, 2004). The other has been isolated from normally exfoliated deciduous incisors and is associated with capillaries. Mesenchymal stem cells also reside in the periodontal ligament (Seo et al., 2000). These cells continually maintain the ligament, which is under constant stress, but they can also regenerate injured alveolar bone. Adult mammals cannot regenerate lost teeth, but adult urodele amphibians, sharks and crocodilians can do so, and may thus be valuable research models for learning how to regenerate human teeth. Mesenchymal stem cells are also found in adipose tissue (Zuk et al., 2002) and in most of the connective tissue compartments of the body (Young and Black, 2004). Like MSCs of the bone marrow, these cells can differentiate into fibroblasts, chondrocytes, osteoblasts and adipocytes. Whether they actually have a regenerative function in vivo is unknown, and their nature is unclear. It is possible that these cells, as well as the stem cells isolated from deciduous incisors, are pericytes, which are ubiquitous as cells that stabilize capillaries and venules, and have long been known to histologists as multipotential cells. Hematopoietic and Endothelial Stem Cells Blood cells are regenerated by hematopoietic stem cells (HSCs) in the bone marrow and blood vessels are regenerated both by endothelial cells in the walls of venules and circulating endothelial stem cells (EnSCs) from the bone marrow. HSCs are dependent on associated stromal cells of the marrow for their survival, proliferation and differentiation. They are small cells with the surface phenotype [CD34 c-Kit Sca-1 VEGFR2]+ Thy-1lo Lin and express the transcription factor Runx-1 (Spangrude et al., 1998; North et al., 2002). They divide asymmetrically to self-renew, while spawning a common erythroid/myeloid progenitor that gives rise to the blood cell lineages and a common lymphoid progenitor that gives rise to the cells of the immune

106 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

system (Adolfsson et al., 2005). HSCs are maintained in a quiescent state by interaction with a subset of N-cadherin+ CD45 stromal cells in the marrow, mediated by the Tie-2 receptor on HSCs and its ligand Ang-1 on the stromal cells. Wnt3a and Notch signaling are necessary for proliferation and self-renewal, as is Bmi, a protein that represses the expression of the p16 and p19 genes, which suppress proliferation and promote apoptosis, respectively (Calvi et al., 2003; Park et al., 2003; Arai et al., 2004). Stem Cells of Skeletal and Cardiac Muscle Skeletal muscle is regenerated by satellite cells (SCs) expressing the surface phenotype [CXR β1-integrin CD34 c-met]+ [CD45 Sca-1 Mac-1] (Sherwood et al., 2004). These cells reside between the sarcolemma and the basement membrane of the myofibers. SCs are held in a quiescent state by Notch signaling and do not express muscle regulatory factors (MRFs). Free-grafted muscle degenerates, followed by a typical inflammatory response. Quiescent SCs are activated, detach from their basement membranes, and proliferate within them, using the anaerobic pentose phosphate metabolic pathway (Hansen-Smith and Carlson, 1979; Carlson, 2003). The proliferating SCs strongly up-regulate Pax7 and MRFs and subsequently fuse and differentiate to form new myofibers. HGF released from muscle ECM and the growth hormone (GH)-stimulated upregulation of the IGF-IEc isoform (mechano growth factor, MGF) by myofibers are the major growth factors that stimulate SC proliferation, augmented by PDGF, FGF-2, LIF and TGF-β (Allen et al., 1995; Tatsumi et al., 1998; Pastoret and Partridge, 1998; Hill and Goldspink, 2003; Goldspink, 2005). Mammalian cardiac muscle initiates a regenerative response to ischemic injury that is not sustained. Stem cells in the myocardium proliferate, but fibroblast proliferation is faster, suppressing the regenerative response and creating a scar. There are three distinct cardiac stem cell phenotypes that can differentiate in vivo and in vitro into cardiomyocytes: [c-Kit Sca-1]+, Sca-1+ [c-Kit Lin], and Isl-1+ [Sca-1 c-Kit] (Beltrami et al., 2003; Oh et al., 2003; Laugwitz et al., 2005). The Isl-1+ cells are found only in those parts of the heart that have an embryonic contribution from the secondary heart field. The relationship between these three subpopulations with regard to cardiac regenerative potential is not clear. Heart muscle regenerates in the MRL/MpJ mouse after cryogenic infarction (Leferovich et al., 2001). The frequency of mitosis in the injured MRL hearts is 10–20%, compared to 1–3% in wild-type animals. This animal model offers the opportunity to investigate how stem cell populations and/or injury environments differ in regenerating vs. non-regenerating mammalian heart tissue. Dedifferentiation Dedifferentiation is a mechanism for making mature cells into mesenchymal-like stem cells by the loss of phenotypic specialization. Dedifferentiation is not observed during embryogenesis, but the cells derived by dedifferentiation of adult cells do recapitulate part of the embryonic developmental program. The divas of dedifferentiation are the larval and adult urodeles (salamanders and newts) and anuran (frog and toad) tadpoles. These animals can regenerate many complex structures by dedifferentiation, including lens and neural retina of the eye, spinal cord, intestine, heart muscle, upper and lower jaws and limbs and tails. The major difference between dedifferentiated cells and ASCs is that dedifferentiated cells do not self-renew in the conventional sense. One could argue, however, that because these structures can regenerate repeatedly, each cycle of regeneration represents a self-renewal. Amphibian structures known to regenerate by dedifferentiation are the lens, neural retina, intestine, upper and lower jaws, heart muscle, limbs and tails. We know the most about dedifferentiation from in vivo and in vitro studies on muscle of regenerating amphibian limbs (Brockes and Kumar, 2005; Stocum, 2006, for reviews). Limb regeneration in amphibians is achieved by the formation of a blastema derived from the satellite cells of muscle (Morrison et al., 2006) and by the dedifferentiation of dermis, muscle, skeletal and Schwann cells local to the amputation surface (Brockes and Kumar, 2005; Figure 6.2). Dedifferentiation is accomplished

Developmental Mechanisms of Regeneration 107

A Improved regeneration of young muscle

Young

Old

Depressed regeneration of young muscle

B Reg index ISO = 100% Hetero= 90% Delta+ SC ISO = 100% Hetero = 74%

Young

Reg index ISO = 17% Hetero= 100% Delta+ SC ISO = 19% Hetero = 93%

Old

C Delta+ SC= 97% Notch+ SC = 83%

Delta+ SC= 78% Notch+ SC = 89% Young serum old cells

Old serum young cells

Figure 6.2 Experiments demonstrating that the regenerative capacity of old rat muscle is restored by providing the muscle with a young environment. (A) Reciprocal transplantation of leg muscle between young and old rats. Old muscle regains regenerative capacity, while the capacity of young muscle for regeneration is reduced. (B) Parabiosis of young and old rats followed by cryoinjury to leg muscle. ISO  same age parabionts; HETERO  old/young parabionts. The regeneration index (RI) is the number of regenerated myofibers and Delta and Notch indicate the number of activated satellite cells. All HETERO values and the ISO value for old/old parabints are measured against the ISO value for young/young parabionts. Left, values for regeneration of young muscle in HETERO parabionts are depressed to 90%, and 74%, respectively of the control (ISO) value. Right, values for regeneration of old muscle are improved from 17% and 19% of the young control (ISO) value to 100% and 93%. (C) Left, old satellite cells in young serum. Delta and Notch are expressed at 97% and 83% of the control value for young cells in young serum. Right, young satellite cells in old serum. Delta and Notch are expressed at 78% and 89% of the control value for young cells in young serum. by the proteolytic degradation of ECM and the loss of phenotypic specialization by the liberated cells. In the case of muscle, this also involves cellularization. Re-entry of blastema cells into the cell cycle is induced by an as yet unidentified thrombin-activated protein (Tanaka et al., 1997; 1999). We do not yet have a clear picture of the molecular mechanism of dedifferentiation. Destabilization of microtubules is involved, but does not lead to the complete program of dedifferentiation and re-entry into the cell cycle (Duckmanton et al., 2005). Elements of the Notch signal transduction pathway are expressed in blastema cells, but other pathways may be involved as well. The blastema cells require growth and trophic factors from the wound epidermis and regenerating limb nerves for their survival and proliferation. Both the wound epithelium and nerves provide FGFs for this purpose. In addition, the nerves sustain blastema cells by glial growth factor-2, substance P, and transferrin (Stocum, 2006, for review).

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The blastema is a self-organizing entity from its inception (Stocum, 2006, for review). The developmental fate and handedness of the blastema cannot be altered by grafting it to an ectopic location, even under conditions that force it to repeat the earliest stages of blastema formation. The mechanism of self-organization can be explained by local cell interactions that intercalate missing limb structures within boundaries established in the early blastema. The limb can be viewed as a three-dimensional “normal neighbor” map in which each cell knows its position relative to all other cells (Mittenthal, 1981). When a limb is amputated, dedifferentiated cells inherit a memory of their position on the circumference and radius of the limb. How the proximodistal positional identities are restored is not yet clear, but probably involves intercalary regeneration via local cell interactions between boundary positional identities established during early blastema formation. In vitro and in vivo adhesion assays, in conjunction with RA treatment, have shown that positional identity is encoded in the cell surface (Nardi and Stocum, 1983; Crawford and Stocum, 1988; Escheverri and Tanaka, 2005). One molecule that has been implicated in PD positional identity is Prod-1, a molecule related to mammalian CD59, whose expression is differentially regulated by RA and whose overexpression causes distal blastema cells to translocate proximally (Morais da Silva et al., 2002; Echeverri and Tanaka, 2005). Patterning genes that are activated by local cell interactions during self-organization are similar to those that have been identified in the developing embryonic limb bud. In the PD axis, Hoxa9, Hoxd10 and Meis1 and 2 are involved in specification of the stylopodium and zeugopodium, and Hoxa13 is involved in specification of the autopodium (Simon and Tabin, 1993; Gardiner and Bryant, 1996; Mercader et al., 2005). In the AP axis, Shh plays a role in establishing digit number and identity, and Lmx-1 in the development of dorsal tissue pattern (Imokawa and Yoshizato, 1997; Endo et al., 1997). Effects of Aging on Regenerative Capacity Aging clearly reduces the capacity of tissues for regeneration. A major controversy is whether this decline is due to a decline in number or quality of regeneration-competent cells, a deteriorating niche environment (local and/or systemic), or both. A good example is skeletal muscle. The gastrocnemius muscle of young rats regenerates well, but regenerates poorly in old rats (Carlson and Faulkner, 1989; Carlson et al., 2001). Verdijk et al. (2007) reported that the absolute number of satellite cells per type I myofiber and the cross-sectional area of these myofibers is similar in the vastus lateralis muscle of young and elderly humans, but that the cross-sectional area, absolute number of SCs, percentage of SC myonuclei per myofiber, and the number of SCs per myofiber area is significantly lower in the type II myofibers of elderly muscle. Collins et al. (2007) reported a significantly lower number of SCs in aged rat muscle, but identified a subset of aged SCs in vitro that regenerate myofibers as efficiently as SCs from young muscle. Reciprocal exchange of the gastrocnemius muscle between young and old rats, or parabiosis of old and young rats improves the regeneration of old muscle, while depressing the regenerative capacity of young muscle (Carlson and Faulkner, 1989; Carlson et al., 2001; Conboy et al., 2005). Parabiotic studies indicated that the decline in regenerative capacity of old muscle was associated with a lower percentage of proliferating SCs, not with a decrease in the number or quality of satellite cells (Figure 6.3). The serum of old rats appears to be deficient in factors that in young rats promote the proliferation of satellite cells by increasing the expression levels of Notch and Delta. These factors have not been specifically identified, but one of them may be growth hormone, since strength training in elderly humans significantly retards sarcopenia, and increases MGF production, particularly in combination with administration of GH (Goldspink, 2004, 2005). Similar results have been obtained with young vs. old liver (Conboy et al., 2005). Other ASCs exhibit age-related declines in regenerative capacity as well, but no age-reversal experiments of the type performed on liver and skeletal muscle have been done on these tissues (Stocum 2006, for review).

Developmental Mechanisms of Regeneration 109

4 days

6–7 days

9 days

21 days

Figure 6.3 Longitudinal sections of a regenerating axolotl limb amputated through the distal radius and ulna, 4–21 days post-amputation. By 4 days, dedifferentiation has created an accumulation of mesenchymal stem cell-like cells under the wound epidermis, which becomes a cone by 6–7 days due to mitosis. The first signs of differentiation emerge at 9 days, and by 21 days, a replica of the missing wrist and hand has been regenerated.

STRATEGIES OF REGENERATIVE MEDICINE Regenerative medicine uses three strategies based on the regenerative biology of regeneration-competent cells: cell transplants, implantation of bioartificial tissues and the chemical induction of regeneration (Figure 6.4). These strategies seek to reconstruct damaged tissues, organs and appendages by cell transplants or bioartificial tissues, or by inducing resident cells to reconstruct them in situ. Cell Transplants and Bioartificial Tissues Cell Transplants Fetal cells Fetal cells have been used primarily to treat Parkinson’s and Huntington’d disease. Mesencephalic cells from 6–8 week old fetuses appeared to differentiate into dopaminergic neurons, increase dopamine output, and make synaptic connections with host neurons (Bjorklund and Lindvall, 2000, for review; Bjorklund et al., 2003). The results are highly variable, however, due to the differential survival of the transplanted cells, and double-blind studies suggest that there is a large placebo effect of the treatment (Lazic and Barker, 2003). Fetal striatal tissue grafted to the striatum of marmoset or macaque monkeys with NPA-induced Huntington’s was reported to reverse the symptoms of the disease (Kendall et al., 1998; Palfi et al., 1998). Immunohistochemical studies indicated good survival and differentiation of the grafted neurons, with establishment of functional connections with host tissue. Preliminary clinical trials in human patients given grafts of human fetal

110 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

Strategies of regenerative medicine

Chemical induction

Cell transplants

Bioartificial tissue

Figure 6.4 The three strategies of regenerative medicine. Chemical induction can involve administration of combinations of cytokines, growth factors, natural or artificial ECM templates, or small natural or synthetic molecules, such as reversine (see text). Cell transplants can be used as bioreactors to provide host tissues with paracrine factors, to rebuild tissue, or to construct bioartificial tissues, such as artificial blood vessels.

striatal tissue indicated that the tissue survived and that the symptoms of the disease were alleviated to some extent, with persistent benefits to some patients three years post-grafting (Rosser et al., 2002). Adult Stem Cells

The most sophisticated and successful clinical adult stem cell transplants, begun in 1968, are those of bone marrow for hematopoietic malignancies or genetic disorders. Variable success has been had with other types of ASCs. Cultured keratinocytes have been applied to acute and chronic wounds (Liu et al., 2004) and cultured autogeneic limbal or oral epithelial stem cells have been used to replace the cornea in patients who have suffered corneal damage (Tsai et al., 2000; Nishida et al., 2004). Transplantation of cultured satellite cells for human Duchenne muscular dystrophy has not been successful, but in mdx mice, fresh satellite cells or satellite cells derived from cultured wild-type MSCs have successfully regenerated normal muscle (Montarras et al., 2005). Human MSCs converted to satellite cells by transfection with the DNA sequence for the Notch intracellular domain (NICD), followed by treatment with satellite cell conditioned medium, regenerated muscle after transplantation to dystrophic mdx mice (Dezawa et al., 2005). In a similar experiment, cultured rat and human MSCs were reported to become dopaminergic neurons (41% frequency) when transfected with the NICD sequence and treated with glial derived neurotrophic factor (GDNF) (Dezawa et al., 2004). Transplantation of these cells into Parkinsonian rats significantly increase dopamine production and decreased symptoms. Over the past decade, one of the great hopes of regenerative medicine has been that adult stem cells will prove to have a plasticity that allows them to be reprogrammed by foreign injury environments in vivo or defined chemical factors in vitro to cell types of other lineages (lineage conversion) for transplantation or

Developmental Mechanisms of Regeneration 111

X-irrad or SCID adult host Injured host tissue Donor test cell BrdU

Blastocyst (chimeric embryo assay) Co-culture with inducing cells

GFP Lac Z + Y

Culture in medium conditioned by inducing cells Culture in medium containing cell-specific differentiation agent

Figure 6.5 Assays for lineage conversion of adult stem cells. Labels to identify the donor cells include the Y chromosome, transgenes for green fluorescent protein (GFP) or β-galactosidase, or some combination thereof. BrdU is added to detect DNA synthesis. The cells are then injected or implanted into a variety of host in vivo environments, or cultured in vitro with inducing cells or chemical agents.

bioartificial tissue construction. Several types of lineage conversion assays (Figure 6.5) have been used to test the developmental plasticity of various types of ASCs. Bone marrow stem cells have been of the most interest, because they are easy to harvest and expand as autogeneic cells. The results of such assays have been inconsistent and in many cases difficult to repeat, because of differences in ever-evolving experimental protocols, fusion with host cells, and artifacts such as contamination of the donor cell population with other differentiated cell types, incorporation of host leukocytes into donor tissues, or incorporation of donor cells into host tissues without long-term survival or differentiation into authentic cell phenotypes of that tissue. Cells with high putative plasticity have also been isolated from long-term cultures of bone marrow cells and from connective tissue compartments. These cells share some characteristics with ESCs and differentiate in vivo and in vitro into a wide variety of cell types at frequencies of 5–90% (Jiang et al., 2002; Young and Black, 2004). However, these results have also been difficult to repeat. The overall evaluation to date is that the lineage conversion of adult stem cells is possible and a goal worth pursuing, but requires more consistent and rigorous proof (Wagers et al., 2002; Murry et al., 2004; Balsam et al., 2004; Laflamme and Murry, 2005). Embryonic Stem Cells

Most ASCs are difficult to harvest and expand. Furthermore, any allogeneic cell transplant or bioartificial tissue will be immunorejected, unless the cells are encapsulated. ESCs are viewed as a prime source of cells for transplant or bioartificial tissue construction because they can be expanded indefinitely in culture to provide the large numbers of cells required to produce derivatives, while retaining their pluripotency. Transplanted neural and glial precursors differentiated from ESCs have successfully reversed the lesions and symptoms of Parkinson’s disease and deymelinating disorders in rodents (Kim et al., 2002; Barberi et al., 2003; Brustle et al., 1999). Beyond this, the use of ESC derivatives for tissue regeneration is still in a nascent state. Like other allogeneic cells, ESC derivatives are subject to immune rejection (Rippon and Bishop, 2004). Immunorejection of ESC derivatives theoretically can be avoided by using autogeneic ESC lines from blastocysts derived by somatic cell nuclear transplant (SCNT). Proof of principle has already been demonstrated in experimental animals (Munsie et al., 2000; Wakayama et al., 2001), but autogeneic human blastocysts have not

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yet been created, and their derivation is bioethically controversial. This has sparked a search for ways to derive autogeneic ESCs in other ways. Adult mouse fibroblasts have been successfully reprogrammed to express the transcriptional profile of ESC by fusing them with ESCs (Cowan et al., 2005) or transfecting them with the pluripotency transcription factors Oct-4, Sox-2, c-Myc and Klf-4 (Takahashi and Yamanaka, 2006). NSCs have been reprogrammed in the same way by fusing them with F9 embryonal carcinoma cells (Do et al., 2007). Other cells

Chondrocytes cultured from biopsies of healthy cartilage have been successful in repairing articular cartilage damaged by trauma (Brittberg et al., 1994), and β-cell transplants from cadavers have temporarily reversed the symptoms of diabetes, though such transplants are severely limited by donor shortage (Shapiro et al., 2000). Umbilical cord blood cells, which are easily harvested and preserved, show great promise for hematopoietic regeneration (Takahashi et al., 2004). Stem cells that express both embryonic and adult stem cell markers have been isolated from amniotic fluid. These cells were induced in vitro and in vivo to differentiate into neuronal, hepatic, and osteogenic phenotypes (De Coppi et al., 2007) and may represent the best of both ASC and ESC worlds. Bioartificial Tissues Cell transplants are primarily useful for replacing small areas of tissue. Bioartificial homologues are necessary to replace larger tissue areas or whole organs. Tissue homologues have been successfully created in experimental animals for long bone segments by seeding ceramic scaffolds (“bone blanks”) with MSCs that differentiate into osteoblasts (Dennis et al., 2001; Cowan et al., 2004), for intestine, trachea, and urinary bladder by seeding biodegradable polymer meshes with epithelial and smooth muscle cells and for blood vessels by culturing endothelial, smooth muscle and fibroblast cells around a mandrel (Stocum, 2006, for review). Work is ongoing to bioengineer whole organs such as the liver, but success in this endeavor has so far been limited because of the difficulty in providing the tissue with vascular channels in vitro. The most spectacular bioartificial tissue made so far is a human mandible constructed of a titanium mesh cage filled with blocks of bone matrix, bone marrow cells (for MSCs) and BMPs (Warnke et al., 2004). This construct was prevascularized and differentiated by growing it for seven weeks in a pocket made in the latissmus dorsi muscle of a patient who had lost his mandible to cancer. The construct was then removed and transplanted successfully into the position of the original mandible. A major issue for bioartificial tissue construction (or for regeneration templates, see ahead), aside from vascularization, is mimicking the properties of the ECM. The ECM is a complex, three-dimensional assembly of macromolecules synthesized by cells as an adjacent acellular basement membrane, and/or as an interstitial tissue matrix surrounding the cells. Interstitial ECM is composed of fibrous proteins (primarily collagens) embedded in a highly hydrated gel of GAGs and proteoglycans that is also a repository for signaling molecules such as growth factors, proteases and their inhibitors (Voytik-Harbin, 2001). The natural matrix releases the appropriate biological signaling information at the right times and places to promote and maintain cell adhesion, proliferation, differentiation and tissue organization. Thus, processed natural biomaterials such as cadaver dermis and pig SIS have been a logical choice for use as regeneration templates and scaffolds for bioartificial tissues. The use of synthetic biomaterials is advantageous because they can be manufactured in virtually unlimited quantities to specified standards, with additional shape-shifting features built in, such as liquidity and small volume at room temperature, changing to gelation, expansion, and space-filling at body temperature within a tissue gap. The goal of biomaterials science is to make synthetic scaffolds that mimic the ECM in vivo, providing not only the geometry and physical/chemical properties to maximize the migration of cells throughout the scaffold, but also the capability to sequester and release biological signals essential for cell

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Cells

Biomaterials or biomaterials plus adherent molecules

Figure 6.6 High throughput screening of biomaterials (yellow) or biomaterials with adherent molecules (red, blue and green dots) for their effects on cells layered on top of the biomaterials.

proliferation and differentiation (Langer and Tirell, 2004). Hench and Polak (2002) have described the evolution of biomaterials from a first generation in the 1960s and 70s that mimicked the physical properties of replaced tissue with minimal toxicity, to a second generation in the 1980s and 90s that was bioactive and biodegradable as well. This second generation of biomaterials is the set from which most scaffolds are currently made (for example, polyglycolic acid and polylactic acid). Third generation biomaterials focus on micro- and nanofibrillar biomaterial gels, including self-assembling peptide and non-biological amphiphiles, and non-fibrillar synthetic hydrophilic polymer hydrogels that have the physical and chemical properties of natural ECM (Lutolf and Hubbell, 2005). A number of biologically important signaling and enzyme-sensitive entities can be incorporated into these hydrogels, including recognition sequences for cell adhesion proteins, soluble growth factors, and protease-sensitive oligopeptide or protein elements. Derivatized amino reactive polyethylene glycols (PEG) containing both peptide substrates for proteases and binding peptides for soluble factors or cell adhesion molecules appear to be particularly promising for creating mimics of ECM-cell interactive processes (West and Hubbell, 1999; Zisch et al., 2003; Tessmar et al., 2004). There are significant technical hurdles yet to be overcome in making interactive synthetic biomaterials that mimic the specific microniche environments of regeneration-competent cells (Lutolf and Hubbell, 2005). However, the development of new generations of biomaterials with interactive effects on cell behavior is being aided by high-throughput screening of biomaterials (Anderson et al., 2004) (Figure 6.6). Hubbell (2004) has pointed out that polymer biomaterials could be used as tethering platforms to screen combinatorial libraries of molecules that bind to the polymers for their effects on cell activity. If the activity of such molecules is dependent on their association with the polymers, they would not show an effect when presented to cells by themselves, but would reveal their effects if bound to a polymer. Chemical/Physical Induction of Repair and Regeneration Topical Agents for Skin Repair Various topical agents have been tested for their efficacy in accelerating repair of acute wounds and chronic skin wounds (Fu et al., 2005; Stocum, 2006, for reviews). The growth factors TGF-β1 and 2, FGF-2, EGF, and IGF-1, and growth hormone (GH) have been reported to accelerate the repair of acute wounds in experimental animals and FGF-2 and GH have this effect in human patients. Other agents reported to accelerate the repair of acute skin wounds are extract of the Celosia argentea leaf, vanadate, oxandrolone, the opoid fentanyl,

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ketanserin, oleic fatty acids, pig enamel matrix, and the peptide HB-107. These agents increase the rate and extent of re-epithelialization, angiogenesis, or cellularity of granulation tissue. Still other topical agents act by reducing scarring. Extract of Allium cepia (onion), chitosan, the COX-2 inhibitor celicoxib, HGF, and antiTGF-β1 and 2 antibodies all promote healing with less scarring, as do hydrogels composed of cross-linked hyaluronic acid and chondroitin sulfate. Combinations of topically applied PDGF-B plus IGF-1, EGF plus insulin, and TGF-β plus PDGF-B enhance chronic wound repair in experimental animals. PDGF-B, FGF-2, EGF, TGF-β, and rhKGF-2 all accelerate the closure of chronic wounds in patients. Currently, FGF-2, PDGFB and rhKGF-2 are approved for clinical use. Other topically applied agents that accelerate the repair of chronic wounds are angiotensin (1–7), thymosin β4, L-arginine, and pentoxifylline. These agents exert their effect through anti-inflammatory and angiogenesis-promoting activities. Regeneration Templates Natural or bioartificial scaffolds have been used as templates to encourage immigration of resident cells bordering lesions to repair dermis and other connective tissues, peripheral nerves, urinary conduit tissue, digestive tract, and bone (Yannas, 2001; Stocum, 2006, for reviews). Cadaver dermal matrix (Alloderm®) and fetal bovine dermal matrix (Primatrix™) promote repair of burns; porcine dermal matrix (Permacol®) and porcine small intestine submucosa (SIS, Surgisis™) are approved for hernia repair. Primatrix™ and another form of SIS, Oasis™ accelerate the healing of diabetic ulcers. The most widely used bioartificial dermal matrix is Integra®, which consists of bovine dermal collagen and chondroitin 6-sulfate. Clinical assessments of Integra® have reported results superior to those of other constructs for excisional wounds, including burns (Heitland et al., 2004). Epidermal coverings do not take well on dermal regeneration templates when the dermis is badly damaged, due to slow vascularization. Thus they are often applied in a two-step procedure in which the dermal template is put on the wound first and allowed to revascularize, after which keratinocytes or meshed split thickness skin grafts are added. Collagen tubes filled with a copolymer of type I collagen and chondroitin 6-sulfate with longitudinally oriented pores promoted the regeneration of transected peripheral nerve axons, while a collagen/laminin matrix, alginate gel and intercostal nerve sheath embedded in fibrin matrix have been reported to foster the regeneration of spinal cord axons (Yannas, 2001; Goldsmith and de la Torre, 1992; Cheng et al., 1996; Ramer et al., 2005). Urinary bladder matrix has proved effective as a template to promote the regeneration of bladder wall tissue in pigs and urethral wall in human patients (Reddy et al., 2000; El-Kassaby et al., 2003). SIS matrix promoted regeneration of small defects in the esophagus, intestine, bile duct, trachea, bladder and ureter in experimental animals, and polyester mesh has been used as a template to regenerate small defects in the trachea and bladder (Stocum, 2006, for review). A wide variety of scaffolds, including ceramics, polymer combinations, and bioactive glass, encourage the regeneration of small segments of bone by MSCs that migrate into the biomaterial (Seeherman et al., 2002). Plasmid or retroviral growth factor constructs (primarily BMPs) have been incorporated into polymers to promote the commitment of MSCs to osteoblasts (Goldstein and Bonadio, 1998; Bonadio, 2002). Extra bone for transplant has been made in rabbits by filling a subperiosteal space in the tibia with alginate. MSCs from the periosteum migrated into the alginate and formed new bone that was then transplanted to fill a defect made in the contralateral tibia (Stevens et al., 2005). Soluble Factors A number of soluble agents have been found to protect neurons of the damaged mammalian spinal cord and to neutralize or remove molecules inhibitory to regeneration (Ramer et al., 2005). Neuroprotectives include

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molecules such as methylprednisolone and gacyclidine. Myelin proteins inhibitory to axon regeneration in the cord can be neutralized with antibodies, and chondroitinase ABC has been shown to promote spinal cord axon regeneration by cleaving off the chondroitin sulfate side chains from proteoglycans of glial scar (Filbin, 2000; Bradbury et al., 2002). A significant discovery is that transplanted cells are the source of paracrine factors that promote the survival of host cells in the injured region, reduce scarring and even promote regeneration of host tissue. For example, bioartificial skin equivalents, comprised of allogeneic neonatal foreskin fibroblasts in a collagen or polyester scaffold, act as living wound dressings to enhance the healing of chronic wounds by providing growth factors to host cells of the wound (Ehrlich, 2004; Jimenez and Jimenez, 2004). The fibroblasts are eventually rejected and replaced with host fibroblasts. Transplanted NSCs or NSCs transfected with a lentiviral GDNF construct or injection of the construct itself into the striatum promoted the survival of host dopaminergic neurons in Parkinsonian rats (Kordower et al., 2000). Regeneration of spinal cord axons is promoted by the incorporation of Schwann cells and olfactory ensheathing cells into regeneration templates (Ramer et al., 2005). These cells provide soluble factors and adhesion molecules to cord axons that are used normally in the regeneration of spinal nerve and olfactory nerve axons, respectively. Improvements in the symptoms of ALS patients have also been reported in China after injection of autogeneic olfactory ensheathing cells into the forebrain, presumably by paracrine action, although these results have been criticized because they are uncontrolled (Watts, 2005; Curt and Dietz, 2005). MSCs improved cardiomyocyte survival when injected into the infarcted hearts of mice. The effect of these cells is due to paracrine action that activates the cell survival gene Akt, as shown by the fact that conditioned medium of hypoxic MSCs activates this gene and reduces infarct size by reducing apoptosis of cardiomyocytes when injected into the infarct region. Thymosin β-4, which plays a role in regulating the assembly of G-actin into F-actin filaments, also enhances cardiomyocyte survival and cardiac function by activation of Akt (Mangi et al., 2003; Gnecchi et al., 2005; Bock-Marquette et al., 2004), but also by an effect on the migration of epicardial cells and their differentiation into endothelial cells (Smart et al., 2007). Modest improvement in cardiac function was reported in random, double-blinded clinical trials of bone marrow cells injected into the infarct region of patients (Wollert et al., 2004; Lovell and Mathur, 2004; Mathur and Martin, 2004). However, in another controlled, random double-blind study, G-CSF induced mobilization of bone marrow stem cells had no effect on cardiac function in patients who had suffered myocardial infarct (Zohlnhofer et al., 2006). Identifying Constellations of Natural Regeneration Promoting and Inhibitory Molecules Tissues that normally undergo maintenance or injury-induced regeneration clearly possess the niche factors requisite for regeneration. Regeneration-permissive signals must also be present in the injury environments of tissues that fail to regenerate, because such tissues (for example, spinal cord and heart) have regenerationcompetent cells that often initiate a regenerative response, which is then suppressed by fibrosis. To regenerate these tissues, it might only be necessary to neutralize molecules that promote scarring. For other tissues that do not initiate a regenerative response, it may prove essential to provide additional regeneration-permissive or inductive signals to the injury site, particularly if the tissue does not contain regeneration-competent cells. Two strategies can be used to identify regeneration-permissive/inductive and inhibitory molecules (Figure 6.7). One is to identify the molecules secreted by cells known to enhance host cell survival and inhibit scarring after transplantation, and determine which combinations are active in these processes. The other is to compare fibrosis and regeneration in three types of in vivo models. The first model compares wild-type tissues to genetic variations that confer a gain or loss of regenerative capacity. Several strains of MRL mouse can regenerate ear and heart tissue (Heber-Katz et al., 2004).

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A

Conditioned medium

Analysis of cell-secreted proteins

B

Comparative in vivo models

Cellular, biochemical, genomic, proteomic analyses

1. WT incompetent versus mutant competent 2. Early stage competent versus late stage deficient or incompetent 3. Competent species versus deficient or incompetent species

Figure 6.7 Approaches to the identification of natural molecules that constitute the molecular difference between regeneration-competence and regeneration-deficiency or incompetence.

Molecular comparisons can be made between these and regeneration-incompetent wild-type strains to reveal molecules permissive and inhibitory to regeneration. The second model compares tissues at developmental stages when they are capable of regeneration versus stages when they are not. For example, fetal skin in many mammalian species regenerates perfectly, but late in gestation the injury response switches to scar tissue formation characteristic of the adult (McCallion and Ferguson, 1996), whereas the skin of the neonatal PU.1 null mouse retains the fetal capacity for regeneration (Redd et al., 2004). The frog limb bud regenerates perfectly at early tadpole stages, but becomes regenerationdeficient at late tadpole stages. The loss of regenerative capacity in fetal skin and the frog limb bud may be related to maturation of the immune system and the resultant greater inflammatory response after wounding, while the retention of regenerative capacity in the PU.1 null mouse may be due to failure of the immune system to mature (Mescher and Neff, 2005; Godwin and Brockes, 2006, for reviews). The third model compares tissues in regenerating vs. non-regenerating species, such as the regenerating axolotl or newt limb vs. the non-regenerating frog or mouse limb. For example, it has been shown that newt dorsal iris cells and myofibers have the ability to respond to a thrombin-activated protein (as yet unidentified) by entering the cell cycle, whereas axolotl lens and mouse myofibers do not (Tanaka et al., 1999; Imokawa and Brockes, 2003). Comparative genomic analyses using these models have revealed differences in the gene activity of regeneration-competent vs. deficient tissues. For example, subtractive hybridization analysis of regenerationcompetent vs. deficient limbs in the frog Xenopus laevis has revealed not only the upregulation and downregulation of many known genes, but also many novel genes (King et al., 2003). Proteomic analyses should prove even more revealing. Coupled with bioinformatics and systems biology approaches, such data will be invaluable in providing complete molecular descriptions of regeneration competence vs. deficiency, allowing us to potentially promote regeneration in regeneration-deficient or incompetent tissues by manipulating the environment and/ or cellular responses at the injury site. Proof of concept has already been shown by experiments in which antibodies to TGF-β1 and 2, or application of TGF-β3 to adult skin wounds reduce scarring (Ferguson and O’Kane, 2004), by modest improvements in the regeneration of late frog tadpole limb buds by administration of FGF-8 and 10, BMP-4, HGF (Suzuki et al., 2006, for review), and by experiments showing

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that mouse muscle cells, which do not normally cellularize and dedifferentiate in response to injury, are induced to do so when treated in vitro with a protein extract of regenerating newt limb (McGann et al., 2001). Small Molecules A new approach to the chemical induction of regeneration is the use of small molecules to initiate regenerative responses. One such molecule is retinoic acid (RA, MW 300, acid derivative of vitamin A). RA is a key player in tissue embryogenesis, particularly the nervous system (Maden, 2002). It induces alveolar regeneration in the lung (Maden, 2004) and a lentiviral construct of the β2 retinoic acid receptor promotes functional recovery of injured rat spinal cord (Yip et al., 2006). RA has profound effects on the positional identity of blastema cells in regenerating urodele limbs, causing their proximalization, posteriorization and ventralization (Niazi, 1996; Maden, 1998; Stocum, 2006, for reviews). The emerging field of chemical biology has developed methods to systematically identify synthetic small molecules that have developmental or regeneration-related effects on cells. Combinatorial libraries of molecules are generated from starting molecules, and are screened on cells for specific effects. Two such molecules have been identified that effect dedifferentiation of C2C12 mouse myofibers in vitro. Myoseverin (a tri-substituted purine) depolymerizes microtubules and upregulates growth factor, immunomodulatory and stress-response genes. Reversine (a di-substituted purine) interacts with protein kinases and initiates a full dedifferentiation program in C2C12 myofibers to create mesenchymal stem cell-like cells that can differentiate into muscle, osteoblasts and adipocytes (Rosania, 2004; Chen et al., 2004). Another synthetic purine derivative is puromorphine, which induces osteogenesis via the hedgehog signaling pathway (Wu et al., 2004). Neuropathiazol is a synthetic 4-aminothiazole that selectively induces neuronal differentiation of hippocampal neural stem cells (Warashina et al., 2006). Molecules generated in this way clearly have the potential to be useful for initiating regenerative responses and/or suppressing fibrosis in injured regeneration-deficient or incompetent tissues.

CONCLUSION Developing the potential of regenerative medicine will require wide multidisciplinary efforts in the biological, chemical, physical, engineering and information sciences. The first wave of regenerative medicine was the transplantation of adult stem cells, begun in 1968 with the first bone marrow transplants. Current research aims to expand this success to other kinds of adult stem cells and derivatives of embryonic stem cells. The back part of this wave will be the chemical induction of regeneration using transplanted cells as bioreactors to provide survival and regeneration-permissive factors to host tissues and/or suppress fibrosis. These efforts have not seen much success as yet. The second wave will be the chemical induction of regeneration by cell-interactive regeneration templates, the direct delivery of regeneration-promoting and/or fibrosis inhibiting molecules or genes encoding these molecules to a lesion site, or some combination of these. These types of treatments will not only be relatively simple to administer clinically, but will also be much less expensive than cell transplantation therapies. To make the chemical induction of regeneration feasible, we must understand the biology of regeneration and how it differs from fibrosis to a much greater depth than is currently available. Only then will we know the appropriate places, times and at what concentrations to intervene in the pathways of repair to choose regeneration over fibrosis. The third wave will be the in vitro construction of bioartificial tissues and organs that can be implanted in place of the originals. A single type of regenerative therapy is unlikely to fit all degrees of tissue damage. For example, it may not be possible to regenerate tissues much beyond a critical size defect using a cell transplant, chemical cocktail, or regeneration template. Larger defects may require a regeneration template seeded with cells to make a bioartificial tissue. Nor will success in understanding the biology of regeneration be achieved by a singular

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focus on mammalian research models. Non-mammalian models that are more powerful regenerators than mammals, such as amphibians, planarians and coelenterates can teach us much about the mechanisms of regeneration that we need to know in order to stimulate the latent regenerative powers of, or even confer such powers on, non-regenerating mammalian tissues.

ACKNOWLEDGMENT Supported in part by a grant from the W.M. Keck Foundation

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7 The Molecular Basis of Pluripotency in Principles of Regenerative Medicine Ariel J. Levine and Ali H. Brivanlou

INTRODUCTION TO PLURIPOTENCY The union of sperm and egg, two highly differentiated cell types, gives rise to the zygote – the totipotent cell. The zygote has the potential to form every cell type of the embryo and the adult organism through a series of sequential cell fate decisions that successively limit its range of potency. For example, the cells of the very early mammalian embryo divide, maintaining their totipotency until they reach 16–32 cells, at which point outer cells will give rise to extra-embryonic tissues such as the placenta, and inner cells are fated to give rise to the embryo proper. This, the choice between the outer trophoblast and the “inner cell mass,” represents the first restriction in cell fate potential and therefore the end of totipotency. The inner cell mass will give rise to the reproductive germ lineage and all three germ layers of the embryo in vivo while, in vitro, the inner cell mass of the mouse embryo can give rise to embryonic stem cells (ESC) that share this pluripotency (Martin, 1981). Of note, human ESC that are also derived from the inner cell mass can form extra-embryonic derivatives (Xu et al., 2002) and may also spontaneously form primordial germ cells (Clark et al., 2004) and thus may be totipotent. The molecular basis of pluripotency has been best studied in ESC, about which this review will focus. In addition, other pluripotent cells types include “multipotent” adult progenitor cells (Reyes and Verfaillie, 2001) derived after prolonged culture of bone marrow cells, primordial germ cells cultured as “embryonic germ cells” (Matsui et al., 1992; Shamblott et al., 1998), embryonic carcinoma cells derived from teratomas (Finch and Ephrussi, 1967), and “multipotent” adult male germline stem cells (Guan et al., 2006). It is interesting how many of the pluripotent cell types are related to germ cells, highlighting the developmental proximity between the gametes and the totipotent zygote and potentially between the embryonic epiblast and the origin of primordial germ cells. Potency, or cell fate potential, is a functional characterization of cell types and does not necessarily describe the range of genes expressed in these cells, their origin, or whether they represent an endogenous cell type in the organism. The hallmark of pluripotent cells is the potential to give rise to germ cells, endoderm (gut, liver, pancreas), mesoderm (muscle, blood, bone), and ectoderm (neurons, glia, skin). This potential is determined using cell type specific molecular markers (such as insulin, cardiac actin, and neurofilament heavy chain), morphological criteria (such as typical histology, beating foci of cardiomyocytes, and branched axons of neurons), and functionally (through secretion of appropriate hormones or neurotransmitters in response to stimuli). The potency of cells may be revealed experimentally in vitro, using “embryoid bodies” in

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culture, within a teratoma upon injection into immunocompromised mice, and ultimately, in vivo within an embryo upon injection of plurpotent cells into a blastocyst (Brivanlou and Darnell, 2002). In addition to the ability to differentiate into all of the cell types of the organism, these pluripotent cells possess the ability to self-renew. These two characteristics together endow these cells with “stemness.” This quality is also measured by molecular markers (such as Oct4 and nanog), typical morphological criteria, and functionally by the ability to self-renew indefinitely and to differentiate into the broad spectrum of cell fates, in vitro and in vivo. Other stem cell types, such as adult hematopoietic stem cells (HSC), intestinal stem cells (ISC), hair follicle stem cells (HFSC), and neural stem cells (NSC) are multipotent, meaning that they can give rise to a range of cell types restricted to a particular tissue type. For example, HSC, found in the bone marrow, can give rise to all of the cell types of blood including macrophages, erythrocytes, and leukocytes. Many reports in the past several years have claimed transdifferentiation of multipotent cells into other tissues but these findings are contested and the molecular basis for them is not well understood so they are not considered in this review. Based on the above definitions of pluripotency and multipotency, it is clear that cells could exist with intermediate potencies. For example, recent data has suggested the existence of mesoangioblasts: cells derived from the embryonic aorta that can self-renew indefinitely in culture and give rise to many mesodermal cell types such as blood, bone, and muscle (Minasi et al., 2002). Further, these terms are defined along a unidirectional undifferentiated-to-differentiated vector, within a given window of time for a cell and barring major changes to a cell’s state. A fully differentiated “unipotent” cell type may be used to support totipotent development through nuclear transplantation and cloning. And in tumors, a fully differentiated cell may “dedifferentiate” by losing its markers of differentiation while gaining factors that support self-renewal. While these are somewhat semantic matters, they raise the point that our current abilities are limited for describing, and therefore fully characterizing and understanding the multiple states of potency and stemness.

EXTRACELLULAR SIGNALING FACTORS AND SIGNAL TRANSDUCTION Pluripotent cells in vivo exist in communication with other cell types, or in a “niche” that help to regulate their cell fate through extracellular signaling factors that activate signal transduction cascades within the stem cells. In vitro, the first pluripotent cell types, embryonic carcinoma and ESC, were cultured on feeder cells or in media conditioned by these cells (Martin, 1981). However, the factors secreted by these feeder cells were not known, and only a few have been characterized to date, and the media used for maintaining pluripotency included serum, which itself is replete with many known and unknown growth and signaling factors. While mouse embryonic stem cells may be grown on defined substrates such as gelatin, human ESC are still grown on either feeder cells or on a complex, and not defined, tumor cell extracellular matrix. Despite these many unknown inputs on pluripotent cells, several major signal transduction pathways have been shown to be sufficient and/or required for pluripotency. These pathways are the coded information that pluripotent and support cells exchange with pluripotent cells. The first of these was leukemia inhibitory factor (LIF). In addition, the Wnt pathway, the fibroblast growth factor (FGF) pathway, TGF-β/activin/nodal pathway, and the bone morphogenetic protein (BMP) pathway have all been shown to regulate pluripotency. Importantly, all of these latter pathways are initiated by morphogens – proteins that can produce different cell fates at different doses so it is imperative to consider the dose range of each pathway. Complete inhibition of a pathway is different than low levels of signaling which is qualitatively different than moderate or high levels of signaling (Figure 7.1). LIF LIF was the first factor that was demonstrated to maintain mouse ESC in the pluripotent state. It was identified as a pluripotency factor secreted by feeder cells (Smith et al., 1988) and can now be added to cells in

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Figure 7.1 Signal transduction pathways implicated in the molecular basis of pluripotency in ESC. The STAT pathway is activated in mouse ESC by LIF, but is not required for pluripotency and is not active in human ESC. The MAPK pathway can be activated by FGF signaling in human ESC to maintain stemness but promotes differentiation in mouse ESC. BMP/GDFs promote differentiation in human ESC but can support pluripotency in mouse ESC (in the presence of LIF) through Smad1/5/8 or through signaling to the MAPK pathway. Activin/nodal ligands activate signaling through Smad2/3 to maintain pluripotency in human ESC; this pathway is active in mouse ESC but not required for stemness. Wnt signaling through the canonical pathway maintains pluripotency in both human and mouse ESC.

a recombinant form. LIF binds to the LIF receptor (LIFR) and these proteins then form a complex with gp130 that activates STAT3 through tyrosine phosphorylation (Heinrich et al., 2003). While LIF can also activate other signal transduction pathways, such as ERK–MAPK, STAT3 is the major factor that mediates the affects of LIF on pluripotency. STAT3 activation alone is sufficient to maintain pluripotency in the presence of serum (Matsuda et al., 1999), bypassing a requirement for LIF, while STAT3 inhibition forces differentiation of mouse ESC (Niwa et al., 1998). STAT activation (of STAT5) also plays a role in the multipotency of HSC (Bradley et al., 2002; Schuringa et al., 2004). Surprisingly, though, neither LIF nor STAT3 is sufficient to maintain human ESC in a pluripotent state and STAT3 is not even activated in human ESC (Thomson et al., 1998; Humphrey et al., 2004; Sato et al., 2004). Further, LIF signaling is not required in vivo within the embryo for either pluripotency or viability of the organism (Stewart et al., 1992). BMP/GDF Recently, it has been shown in mouse ESC, BMPs can substitute for serum in cooperating with LIF to support the undifferentiated state (Ying et al., 2003; Qi et al., 2004). In contrast, it has also been demonstrated that BMP inhibition can synergize with FGF signaling to support pluripotency in human ESC (Xu et al., 2005).

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BMPs are TGF-β superfamily members that bind to Type 1 TGF-β receptors Alk1, Alk2, Alk3, or Alk6 together with Type 2 receptors to activate phosphorylation and subsequent nuclear localization of Smad1/5/8 (Shi and Massague, 2003). In addition, BMPs signal through Smad-independent means to regulate other signal transduction pathways, such as mitogen-activated protein kinase (MAPK) (von Bubnoff and Cho, 2001). There are two proposed mechanisms for BMPs effects on mouse ESC. The first suggests that BMPs mediate their effect on mouse ESC through Smad1 induction of Id proteins (Ying et al., 2003), while the other proposes that BMPs cooperate with LIF to maintain pluripotency by inhibiting MAPK signaling (Qi et al., 2004). The observation that BMPs can support pluripotency through Smad activation is surprising because endogenously, ESC and early mammalian embryos do not have active BMP signaling through Smad1/5/8 (James et al., 2005) and even in this work, the authors found that high levels of BMP signaling promoted differentiation of mouse ESC, even in the presence of LIF (Ying et al., 2003). However, as BMPs are morphogens (Wilson et al., 1997), it is possible that very low levels of BMP signaling support pluripotency while higher levels push the cells to differentiate. In support of this model, reduction of levels of GDF-3, a stem cell-associated BMP inhibitor, precludes normal differentiation of mouse ESC (Levine and Brivanlou, 2006). In human ESC, BMPs promote rapid differentiation to extra-embryonic cell fates even when these cells are cultured in feeder conditioned media that normally maintains their pluripotent state (Xu et al., 2002), despite the fact that stem cells express both the BMP inhibitor GDF-3 and the inhibitor Lefty (Sato et al., 2003) and that feeder cells secrete a BMP inhibitor as well (Xu et al., 2005). However, human ESC can be maintained without conditioned media by an exogenous combination of FGF activation and BMP inhibition (Xu et al., 2005). These findings suggest that the normal inhibition of BMP signaling in stem cells and early embryos are required to suppress differentiation to extra-embryonic fates. A role for BMP inhibition in maintaining the potency of stem cell types is conserved in adult, mulitpotent cells such as ISC (where it limits self-renewal of stem cells and antagonizes Wnt signaling (Haramis et al., 2004; He et al., 2004)), HFSC (where it antagonizes the ability of Wnt signaling to maintain the stem cells (Jamora et al., 2003)) and hematopoietic stem cells (where BMP signaling through Alk3 regulates the stem cell niche) (Zhang et al., 2003). TGF-β/Activin/Nodal The other branch of TGF-β signaling, the classic TGF-βs, activins, and nodal, support the pluripotent state and are required for stemness in human ESC (James et al., 2005). The members of this branch of the TGF-β pathway bind to Type 1 receptors Alk4, Alk5, or Alk7 together with a Type 2 receptor to activate signal transduction through Smad2/3 (Shi and Massague, 2003). Activin/nodal signaling is active in early mouse embryos and in both mouse and human ESC, as revealed by phosphorylation and nuclear localization of Smad2/3 (James et al., 2005). This activation is significant for the pluripotent state as exogenous activin or nodal promote pluripotency in human ESC (Vallier et al., 2004; Beattie et al., 2005; James et al., 2005). Further, activin/nodal signaling is required for the maintenance of stemness in human ESC, such that abrogation of signaling through a small molecule inhibitor of Alk4/5/7 (SB431542) or through excess extracellular domains of the receptors forces differentiation of human ESC even in conditioned media, or downstream of Wnt or FGF activation that maintain pluripotency, as described below (James et al., 2005; Vallier et al., 2005). While inhibition of Alk4/5/7 signaling in mouse ESC does not affect pluripotency (Dunn et al., 2004; James et al., 2005; Vallier et al., 2005), mice lacking both Smads2 and 3 are deficient in maintaining the epiblast, the immediate derivative of the inner cell mass, and have significantly reduced levels of Oct4 (Dunn et al., 2004). These findings show that activin/nodal signaling is important for pluripotency in human ESC and in mouse embryos, and suggest that the signaling events in mouse ESC may not represent the in vivo scenario.

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FGF In many cell types, FGF signaling promotes survival and proliferation through its activation of ras and ERK/MAPK. These downstream mediators of FGF signaling have been shown to play roles in pluripotent stem cells, but, as with many signaling pathways in these cells, the results are somewhat contradictory. In mouse ESC, eRas has been shown to promote the proliferation, self-renewal, and tumorigenicity of stem cells (Takahashi et al., 2003); however, active ERK/MAPK has also been shown to promote differentiation to extraembryonic endoderm cell fates (Yoshida-Koide et al., 2004). In human ESC, non-physiological levels of FGF signaling can act independently (Levenstein et al., 2006) or combine with BMP inhibition (Xu et al., 2005) or nodal activation (Vallier et al., 2005) to support pluripotency of the cells without conditioned media. In fact, FGF signaling itself can act as a BMP inhibitor in these cells, perhaps through an inhibitory phosphorylation by MAPK on the linker region of Smad1/5/8 (Pera et al., 2003; Xu et al., 2005). Wnt All of the above pathways are required or sufficient in either mouse or human ESC, but not both, reflecting a curious degree of species-dependent differences in the molecular basis of pluripotency. However, Wnt signaling has been shown to support pluripotency or multipotency in mouse ESC, human ESC, HSC, HFSC, and ISC, suggesting that it plays a core role in the molecular basis of pluripotency. Wnt ligands signal to cells through multiple pathways including the “canonical” pathway in which Wnt binds to frizzled receptors, which signal through disheveled to relieve GSK3β inhibition of β-catenin (Reya and Clevers, 2005). In mouse and human ESC, Wnt signaling can be activated with a small molecule BIO, an inhibitor of the Wnt inhibitor GSK3β and thereby maintain the pluripotent state, as determined by marker gene analysis and chimera formation (Sato et al., 2004). This is in agreement with the observation that Wnt signaling is normally active in mouse ESC and is decreased upon differentiation (Sato et al., 2004). BIO is only able to sustain pluripotency in human ESC for a limited number of passages. A possible explantation for this phenomenon is that the primary input of Wnt activation is on the self-renewal aspect of stemness rather than the maintenance of pluripotency. In this case, a fraction of each passage would differentiate spontaneously and be lost upon further passage. Interestingly, Wnt signaling requires intact activin/nodal signaling as inhibition of Alk4/5/7 abrogates the ability of BIO to maintain pluripotency (James et al., 2005). While these data relied on inhibition of GSK3β (which has Wnt-independent targets), the role of Wnt ligands in supporting stemness has been demonstrated in mouse ESC, in experiments that show that Wnts secreted by feeder cells or Wntconditioned media maintain stemness in mouse ESC (Hao et al., 2006; Ogawa et al., 2006). Further, constitutive activation of β-catenin synergizes with LIF to maintain pluripotency in mouse ESC (Ogawa et al., 2006). Wnt signaling also plays important roles in maintaining the mulitpotency of adult intestinal and HFSC. Loss of the Wnt-responsive transcription factor Tcf4 allows normal development of the gut but results in complete loss of the stem cells such that instead of a normal arrangement of differentiated villi and crypts that contain progenitors, only differentiated cells are present (van de Wetering et al., 2002). Forced activation of Wnt signaling in skin cells allows formation of new hair follicles and, eventually, skin tumors (Gat et al., 1998); indeed, mutations in a Wnt transcription factor are found in many human skin tumors (Chan et al., 1999). In HSC, Wnt3a has been shown to promote expansion of stem cells and activated β-catenin promotes self-renewal and maintains the undifferentiated state of HSC. These cells normally have active Wnt signaling and blocking Wnt activity inhibits self-renewal and the ability of HSC to reconstitute bone marrow. Importantly, in these experiments, the authors examined potential targets of Wnt signaling that could act to mediate the effects of Wnts on stemness and found that HoxB4 and Notch1 were upregulated by Wnt signaling and could play this role (Reya et al., 2003).

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Oct Sox Nanog Pluripotency targets

Polycomb

Differentiation targets

Figure 7.2 Intrinsic factors implicated in the molecular basis of pluripotency. Oct4, Sox2, and nanog coordinately regulate transcription of pluripotency targets. Differentiation targets are repressed in ESC by polycomb group epigenetic mechanisms.

TRANSCRIPTIONAL NETWORKS The nuclear factors that regulate pluripotency and convert extrinsic signals into intrinsic cellular responses have been the subject of intense scrutiny. Three principal transcription factors coordinately regulate the pluripotency program: Oct4, Sox2, and nanog. Each of these genes is expressed in the early mammalian embryo and within the blastocyst stage, they are localized to the inner cell mass (Rosner et al., 1990; Avilion et al., 2003; Chambers et al., 2003; Mitsui et al., 2003) (Figure 7.2). Mutants for these factors cannot maintain the pluripotent epiblast but, interestingly, different outcomes result from this common deficiency. In Oct4 knock-out embryos and stem cells, the cells differentiate into extra-embryonic trophectoderm. Reduction of Oct4 levels in human ESC confirms these findings, as these cells upregulate markers of trophoblast (Zaehres et al, 2005). Sox2 mutant embryos have a similar phenotype but fail slightly later in development and Sox2 mutant outgrowths of blastocyst embryos divert to trophectoderm (Avilion et al., 2003). In contrast, nanog mutant embryos form extra-embryonic endoderm (Mitsui et al., 2003), a fate that is shared upon nanog reduction in human ESC (Hyslop et al., 2005), although these cells also express a marker of trophoblast (Zaehres et al., 2005). Mouse ESC that overexpress Oct4 become primitive endoderm (Niwa et al., 2000), suggesting a possible morphogen effect mediated by Oct4. In contrast, nanog overexpressing stem cells retain pluripotency cellautonomously and do not require LIF or other factors (Chambers et al., 2003; Mitsui et al., 2003). Nanog overexpression similarly frees human ESC of exogenous factors to support pluripotency but converts these cells into a type that more closely resembles epiblast rather than inner cell mass (Darr et al., 2006). These results highlight the need for a critical balance of these stem cells factors to achieve pluripotency. Recent work has analyzed the targets of Oct4, Sox2, and nanog on a genome-wide scale and has found that these three factors coordinately regulate the stem cell program through both positive and negative regulation of target genes. Of the promoters bound by Oct4, more than half are bound by all three factors and the binding sites for these proteins are often very close together (Boyer et al., 2005). Further, synergistic co-regulation of the FGF4 promoter by Oct4 and Sox2 has been well established (Yuan et al., 1995). EPIGENETIC AND ENVIRONMENTAL REGULATION When the differentiated sperm and egg are converted into the totipotent zygote, a process known as “nuclear reprogramming” plays a critical role. This process reprograms the chromatin structure characteristic of differentiated cells into a new conformation typical of the pre-implantation mammalian embryo. Nuclear reprogramming is also a critical step in animal cloning and is required for the nucleus of a differentiated cell to support complete embryonic differentiation when placed into a host egg.

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Recently, two groups have shown that in pluripotent ESC, many genes whose expression is associated with differentiation are kept in either a suppressed state (Boyer et al., 2006) or a particular “bivalent” state in which the genes are expressed at very low levels but are easy to activate (Bernstein et al., 2006). One group of chromatin regulators that may be important for these epigenetic modifications is the polycomb group proteins. Specifically, ESC mutant for the member Eed aberrantly express many genes typical of differentiated tissues. These findings imply that unless these genes are silenced, the pluripotent cells will differentiate, meaning that differentiation is a default phenotype in ESC and maintaining cells in an undifferentiated, pluripotent state is an active process. A reciprocal relationship between pluripotent cells and differentiated cells has also been demonstrated, showing that DNA methylation is required for differentiation. For instance, ESC mutant for the DNA methyl transferase proteins Dnmt3a and Dnmt3b (Chen et al., 2003) or for the CpG binding protein CGBP (Carlone et al., 2005) do not differentiate normally in vitro or upon formation of teratomas; instead, these cells maintain expression of the pluripotency markers Oct4 and alkaline phosphatase (another stemness marker). In this case, targets for suppression by DNA methylation are the proteins that mediate pluripotency. The Oct4 locus is DNA methylated very early in development (Gidekel and Bergman, 2002). Another potential target is a region of human chromosome 12p13 that contains several genes involved in stemness and early germ cells including nanog, GDF-3, and Stella. Interestingly, this cluster of genes is overexpressed in almost all male germ cell tumors and nanog and GDF-3 are specifically overexpressed within pluripotent embryonic carcinomas relative to seminomas and their expression is decreased upon differentiation of embryonic carcinomas (Korkola et al., 2006). A local concentration of pluripotency genes would allow their coordinate regulation by epigenetic mechanisms such that they could be silenced after early development to avoid undue proliferation or inhibition of normal differentiation. Another non-classical type of molecular regulation of pluripotency includes environmental factors such as oxygen concentration. Low oxygen levels, or hypoxia, have been shown to promote more pluripotent and multipotent cell types at the expense of their differentiated progeny. For instance, it has been shown that low oxygen decreases the differentiation of human ESC, enhances the multipotency of NSC, and expands hematopoietic stem cells (Morrison et al., 2000; Danet et al., 2003; Ezashi et al., 2005). A possible mechanism for these observations is the fact that HIF2α, a key regulator of the cellular response to hypoxia, directly activates Oct4 (Covello et al., 2006). Accordingly, HIF2α knock-in ESC form teratomas with an increased percentage of undifferentiated cells (Covello et al., 2005) and knock-in embryos die shortly after implantation and often contain an expanded epiblast (Covello et al., 2006). In adult tissues, damage may be sensed by hypoxia, triggering local stem cells to self-renew and differentiate to repair the damaged tissue. In the embryonic environment the inner cell mass, from which ESC are derived, could be located further from a source of oxygen so that low levels of oxygen support that internal, pluripotent fate.

SUMMARY AND PERSPECTIVES The molecular basis of pluripotency is a complex coordination of extracellular and environmental factors, intracellular signal transduction and transcriptional networks, and global regulation of transcription through epigenetic mechanisms. The output of all of these factors is “stemness:” the ability of these special cells to selfrenew and to differentiate into the cell types of the embryo proper. Several themes emerge from this review of our understanding of pluripotency. First, a delicate balance of instructive and inhibitory signals maintain pluripotency. Second, Wnt activation and BMP inhibition are shared signaling characteristics of several types of pluripotent and multipotent cells. Third, Oct4, Sox2, and nanog regulate the transcriptional program of both human and mouse ESC. Despite these important discoveries, many questions remain regarding the molecular basis of pluripotency. Among the priorities is the molecular basis for the differences in human and mouse ESC. These differences

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may be artifacts of cell culture, may reflect differences in the endogenous cell type that each represents, or may be true differences in the potencies (“default” or otherwise) of human and mouse ESC. Further, it is unclear how the signaling pathways implicated in stemness mediate their effects. What are the targets of these pathways and how do they converge onto the transcriptional regulators of pluripotency? Ultimately, in addition to understanding the basis of stemness for purposes of basic biological knowledge, it is important to determine how these pathways can be manipulated and controlled to provide the material for regenerative medicine.

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8 How Do Cells Change Their Phenotype Peter W. Andrews and Paul J. Gokhale

INTRODUCTION Embryonic development is characterized by the progressive appearance of different cell types in an ordered and, for the most part, apparently irreversible manner. As commonly conceived, and in an idea popularized by Holtzer et al. (1972), these different cell types appear by a series of binary decisions, suggested to accompany cell division, during which cells successively restrict the fates that their progeny can ultimately adopt. In a few well-known cases, this generation of new cells with distinct phenotypes is accompanied by changes in the genome. In the nematode, Ascaris Boveri (1887) noted that segregation of somatic cells from the germ line involves loss of chromosomal material, and this loss does appear to include loss of DNA encoding specific genes (Etter et al., 1991). Of more direct relevance to human biology is the rearrangement of the immunoglobulin genes during development of the lymphoid system (Hozumi and Tonegawa, 1976). However, such examples appear to be the exception, and a fundamental insight into the development of most higher organisms was given by the work of Briggs and King (1952), and Gurdon (1962), who showed that the nuclei of differentiated cells in amphibia appear to retain a full complement of genes capable of directing development of an entire organism if replaced in an appropriate environment. This conclusion has since been extended to mammals by the cloning of sheep (Wilmut et al., 1997) and mice (Wakayama et al., 1998). Thus, it is generally accepted that the adoption of new phenotypes by cells during embryogenesis depends primarily upon the activation and inactivation of specific sets of genes. Understanding the processes that control gene expression and the cues to which cells respond as they acquire new phenotypes is central to the development of techniques for regenerative medicine. STEM CELLS Two general approaches to regenerative medicine can be envisaged. One is to capture cell types that express phenotypes intermediate between those of the zygote and its final differentiated progeny, in order to persuade them follow specific pathways of differentiation. The other is to reprogram later stage cells so that they can adopt phenotypes associated with different lineages, and different from those expected from normal progression during development. Intermediate cells may be progenitor cells that retain some capacity for proliferation while being committed to eventual differentiation into particular terminal cell types. Or, they may be stem cells that exhibit a capacity for apparently indefinite “self-renewal,” while retaining the ability to differentiate along one or more lineages in the future in response to specific cues. Some stem cells are “pluripotent,” such as embryonic stem (ES) cells, which are capable of differentiating into many, if not all cell types found in the adult, but others may be multipotent, retaining the capacity for

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generating only a few cell types, or even unipotent with the capacity to yield only a single terminal cell type. The distinction between progenitor cells and the various types of stem cells is somewhat blurred: it is difficult to assess the ability of cells to undergo indefinite self-renewal, while potency may also be difficult to chara cterize in practice. Considering ES cells, these are best considered to be in vitro artifacts, since the cells to which they correspond in vivo, the inner cell mass cells of the blastocyst, only have a limited capacity for selfrenewal before progressing to other cell types with more restricted potency. Would they be better regarded as progenitors? Such discussions have a semantic flavor. However, whatever the definitions, the crucial element is that for eventual applications we need to understand the physiological processes by which cells are between self-renewal and differentiation, and how, once committed to differentiation, they select the particular lineages they will follow. The modern concepts of stem cells were developed from studies of homeostasis in adult tissues. In early studies it was observed that hematopoietic colonies formed in the spleens of lethally irradiated mice after transplantation of marrow from healthy mice suggesting that the bone marrow contains cells that have the ability both to differentiate and to “self-renew,” that is, the capacity for extensive proliferation while retaining an undifferentiated phenotype and the capacity for future differentiation (Siminovitch et al., 1963). The spleen colonies derived from bone marrow transplants contained varying numbers of cells, from which it was suggested that the decision between differentiation and self-renewal may be stochastic (Till et al., 1964). However, an alternative, deterministic model of stem cell behavior holds that the balance between selfrenewal and differentiation is maintained by asymmetric cell division, so that of two daughter cells of a stem cell, one always retains a stem cell phenotype and one initiates differentiation, a mechanism that has been well studied in gametogenesis in Drosophila melanogaster (Lin and Spradling, 1995). This model, based upon asymmetric cell division, has found favour in many studies of stem cells in the adult, partly because it provides a simple mechanism for the tight control necessary to balance self-renewal and differentiation: in any adult tissue in homeostasis precisely 50% of the progeny of a stem cell must differentiate and 50% must retain a stem cell phenotype. Any lower proportion retaining a stem cell phenotype would lead to eventual depletion of the stem cell pool, while a higher proportion would lead to excess stem cells, which may be the primary issue in development of cancers. However, in the hematopoietic system, at least, the regulation of stem cells may involve a stochastic mechanism (Enver et al., 1998), while in human ES cells in vitro, differentiation is reported to involve symmetric, rather than asymmetric cell divisions (Zwaka and Thomson, 2005).

PLASTICITY – TRANSDIFFERENTIATION AND TRANSDETERMINATION The second approach to regenerative medicine is to discover how the normal, unidirectional sequence of differentiation during embryogenesis may be made to operate in reverse – to induce one type of differentiated cell to undergo a phenotypic conversion that does not normally occur. Cells differentiate in response to a variety of cues. During embryonic development the process is intimately linked to the control of morphogenesis so that particular cells are formed in the correct place. Separate from differentiation itself is the specification of cell fate. Typically, cells may become restricted with respect to the particular fates they can adopt before they actually acquire those fates, a process known as determination. On the other hand, cells may be fated to acquire particular phenotypes because of their location, but be capable of acquiring different fates if moved to alternative locations – “prospective fate” contrasted with “prospective potency” (Weiss, 1939). For example, in the late cleavage stage of early mouse development, the outer cells are fated to become the trophectoderm (Hillman et al., 1972; Kelly, 1977). However, if they are moved to an inner location they contribute to the inner cell mass of the blastocyst, retaining pluripotency and the capacity to contribute to all somatic lineages of the later embryo, a capacity is lost by the trophectoderm. Since determination inevitably involves a change in gene

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expression, it is not clear whether, at a molecular level, determination involves processes that are fundamentally distinct from those that control differentiation itself, when cells typically express at high-level specific genes pertinent to their particular function, for example, hemoglobin in erythrocytes or myosin in muscle. Plasticity of the differentiated or determined states has been more recently invoked in the context of reports that stem cells from different tissues can apparently adopt fates quite distinct from that expected in their tissues of origin. For example, it has been reported that neural stem cells can generate hematopoietic derivatives (Bjornson et al., 1999), and that hematopoietic stem cells can generate neurons (Brazelton et al., 2000). The interpretation of many of these results remains controversial and various mechanisms have been proposed that could have led to these observations. Whether the proposed “plasticity” is in any sense physiological, perhaps providing for endogenous repair mechanisms, or whether it is an experimental artifact remains to be resolved. Nevertheless, the concepts underlying the current ideas about plasticity are old, and the phenomena of transdifferentiation and transdetermination have been widely studied in a variety of organisms and situations. Many lower organisms retain the capacity to regenerate tissues that have been lost or damaged and the processes of tissue regeneration have been extensively studied by developmental biologists seeking to understand the underlying principles that guide embryonic differentiation and morphogenesis. For example, cutting a hydra in half results in the head forming a new tail and the tail forming a new head (Newman, 1974). The cells that had adopted the fates of these structures revert to a cell type that can then adopt both fates. In another example, when a limb is cut from a newt, it regenerates, apparently involving the ability of specific differentiated cells to revert to an earlier state and then redifferentiate to new tissues of the redeveloping limb (Brockes and Kumar, 2002). Transdifferentiation, the ability of a fully differentiated cell to adopt the phenotype of another differentiated cell, also occurs in mammals including humans, recognized as metaplasia in a variety of pathological conditions. For example, metaplasia of the stomach mucosa with the appearance of glands more typical of the small intestine, or the appearance keratinized epithelia in the squamous epithelia of the mouth, is well known. A related phenomenon is that of transdetermination, particularly studied by Hadorn (1968, 1969) in larval development in Drosophila. In that species, adult structures arise from larval primordia, the imaginal disks. During larval development, the cells of these structures become determined to generate specific body parts, such as the wings or leg or eye, etc., but retain an undifferentiated phenotype until pupation when the adult structures are formed. Hadorn found that cells from imaginal disks can be cultured and maintained in their larval undifferentiated state for prolonged periods by serial transplantation to new larvae, while retaining the ability to differentiate into their originally specified body parts when their larval host is permitted to pupate. Nevertheless, at a significant frequency, these imaginal disk cells undergo a switch, called transdetermination, in which they acquire the capacity to generate a distinctly different body part. A striking feature of this transdetermination is that it follows a specific hierarchy, so that certain imaginal disk cells only transdetermine to those lower in the hierarchy, but not vice versa.

CELL FUSION The discovery that nuclear transfer to enucleated oocytes could result in reprogramming of the genome, led to a general interest in the ability of the cytoplasm to control gene function. Fusion of somatic cells using agents such as Sendai virus or polyethylene glycol, provided the means to extrapolate studies of reprogramming by oocytes to a wide variety of cell combinations. For example, hybrids between leukocytes or fibroblasts and hepatoma cell lines may express liver specific proteins from the genome contributed by the leukocytes or

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fibroblasts (Peterson and Weiss, 1972; Darlington et al., 1974, 1984; Rankin and Darlington, 1979). Extensive other experiments also demonstrated that embryonal carcinoma (EC) and ES cells are capable of reprogramming somatic cells back to a pluripotent state, offering the possibility that cytoplasm from these cells might be used instead of oocyte cytoplasm to generate new pluripotent stem cells of defined genotypes (Miller and Ruddle, 1976; Andrews and Goodfellow, 1980; Gmur et al., 1980; Tada et al., 1997; Flasza et al., 2003). These extensive experiments over many years have not yet provided specific insights into the detailed mechanisms of reprogramming. However, they have clearly shown that reprogramming is possible and that the capacity to reprogram the genome extends well beyond that of oocyte cytoplasm. The results have raised the possibility that cell fusion events could confound at least some of the attempts to demonstrate plasticity of stem cells in vivo. For example, neural stem cells or bone marrow-derived cells co-cultured with ES cells have been found to fuse and retain both adult markers and pluripotent potential (Terada et al., 2002; Ying et al., 2002). Consequently, when apparent plasticity of stem cell fate is observed in transplantation experiments it is necessary to demonstrate that the change in phenotype occurs without cell fusion.

CELL PHENOTYPE Phenotype may be the subject of both theoretical and pragmatic definitions. Fundamentally, the phenotype of a cell represents the complete constellation of molecules of which it is composed, and hence a consequence of the activities of all the genes that comprise its genome. However, although the assessment of a large part of the transcriptome or proteome of cells may be feasible when working with cell populations, assessment of a single cell at such a comprehensive level is currently beyond our technological capacity. In practice most researchers lean heavily upon the expression patterns of a selected, small set of “marker” genes, or their products. Nevertheless, there are few markers that are uniquely expressed in only a single cell type, so that the expression patterns of a single, or even a few markers may be misleading if assessed uncritically. For example, several surface antigens, such as SSEA3 and SSEA4 are widely used to define human ES cells. In practice these work well within the context of studies of cultures of ES cells, or their malignant equivalents, the EC stem cells of teratocarcinomas. Our own view is that SSEA3 expression is a particularly sensitive indicator of the undifferentiated state of human ES cells (Enver et al., 2005). However, these antigens are members of the P blood group system, and are expressed on other cell types, including erythrocytes (Tippett et al., 1986). An uncritical attempt to define stem cells by expression of these markers alone outside the context of known cultures of ES or EC cells could easily lead to markedly misleading conclusions. Another way to assess cell phenotype is to analyze cell function. Since function may depend on the coordinated expression of a wide variety of molecules, it can provide a measure that integrates the activity of a large number of genes, and so provide a more robust indicator of cell state – certainly in a way most directly relevant to applications in regenerative medicine. For example, in the identification of pancreatic beta cells differentiating in culture from ES cells or other progenitor cells, the ability to secrete insulin in a measured way in response to changing glucose levels or, better, to rescue a diabetic mouse model following transplantation provides a more certain indicator of the beta cell state than merely detecting the expression of insulin. In the latter case, a number of cells might express insulin in vitro, yet not exhibit properties of pancreatic beta cells (Sipione et al., 2004). Nevertheless, direct measurement of gene expression may be much more rapid and convenient than assessment of function. The approach is undoubtedly essential, but results must be interpreted with care. Assessment of cell phenotype also usually involves extrapolation from a “snapshot” of marker expression or functional activity. Rarely attention is paid to analysis of successive samples and measuring variation, although evidence exists that cell phenotypes may “wobble” over time. For example, gene translation may occur

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in bursts, generating phenotypic “noise” in individual cells that would otherwise be assessed to express a similar “time averaged” phenotype (Elowitz et al., 2002; Ozbudak et al., 2002; Raser and O’Shea, 2004). One could imagine that cells in a terminally differentiated state would need to suppress any phenotypic noise generated by gene transcription and translation variation to ensure a stable physiology. However, cells that need to change could well exploit phenotypic noise to elicit differentiation. Several processes in development appear to rely on stochastic fluctuation to create cellular diversity. For example, in Drosophila, random fluctuations in the cellular levels of the cell surface receptor encoded by Notch, and its cell-associated ligand, encoded by the Delta family, underlie the setting up of feedback signaling that causes segregation of neural precursors from a uniform field of apparently identical epidermal precursors (Simpson, 1997). More generally stochastic fluctuations have been described as being critical for early Xenopus embryo differentiation (Wardle and Smith, 2004), hepatocyte differentiation (van Roon et al., 1989) and hematopoietic differentiation (Enver et al., 1998; Deenick et al., 1999; Hume, 2000).

CONTROL OF GENE ACTIVITY Clearly cells may change their phenotype in fully reversible ways in response to changes in their environment, and such changes involve modulation of gene activity. Building over many years upon the original studies of the lac operon in E. coli by Jacob and Monod (Jacob et al., 2005), such cellular responses to environmental cues have provided a paradigm for exploring the mechanisms that control gene activity. These studies have provided a plethora of transcription factors that interact directly with DNA, or modulate the activity of other components of the transcription complex, to activate or inhibit transcription of specific genes. Certainly in multicellular organisms, as in bacteria, the function of some factors may be directly or indirectly conditional upon the environment of the cell, including the presence of specific signaling molecules. The presence of some factors may also be dependent upon the presence of other transcription factors that regulate the genes that encode them; some may interact with the transcription complexes of the genes that encode themselves. In general, this complex array of factors that control transcription provides the basis of dynamic regulatory loops that could maintain cellular phenotype by a combination of balanced positive and negative feedback, reflecting the history of the cell. Certain transcription factors are known to play key roles in maintaining the phenotype of specific cells. Most notable, perhaps, is the requirement for Oct4 (Matin et al., 2004; Zaehres et al., 2005), Nanog (Chambers et al., 2003; Mitsui et al., 2003; Hyslop et al., 2005; Zaehres et al., 2005) and Sox2 (Yuan et al., 1995; Avilion et al., 2003; Catena et al., 2004) expression in the maintenance of the pluripotent, undifferentiated stem cell state of both mouse and human ES cells. Strikingly there is evidence of positive feedback of these transcription factors on expression of the genes that encode them (Boyer et al., 2005; Chew et al., 2005; OkumuraNakanishi et al., 2005) Other transcription factors have also been explicitly linked to specific pathways of differentiation, for example the expression of the helix loop helix transcription factor, MyoD, is required to initiate the differentiation of myoblasts and its introduction into non-myogenic cells may sometimes be sufficient to force them into a pathway of myogenic differentiation (Tapscott et al., 1988). However, for the most part we have only fragmentary knowledge of the cues that lead to the setting up of specific patterns of transcription factors that could maintain dynamically stable regulatory loops. Moreover, differentiation of cells may be maintained long after the cues that initiated them have disappeared. It seems unlikely that the stability of the differentiated state and its usual irreversibility can be adequately maintained by dynamic systems of transcription factors alone. Heritable changes in the organization of the genome through chromatin structure and DNA methylation, while retaining a constant DNA sequence, seem to play key roles (van Driel et al., 2003; Khorasanizadeh, 2004; Margueron et al., 2005).

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The histones, initially considered to be too simple to contribute to the specificity of gene regulation, nevertheless do appear to play crucial roles in controlling the activity of specific regions of the genome pertinent to particular cell types. Thus, one of the major biochemical modulators of gene activity is the covalent modification of histones, particularly by acetylation and deacetylation under the control of a family of histone acetyltransferases (HAT) and histone deacetyltransferases (HDAC). The deacetylation function of HDAC results in packaging of histones into nucleosomes and consequent transcriptional silencing. The reverse process, whereby acetyl groups are added to histones, reduces the ability of the histones to interact with the DNA backbone and thus “unwinds” the nucleosome complex allowing transcriptional activators to gain access to the DNA and elicit transcription. Several co-repressor complexes have been identified, such as N-CoR and mSIn3A/B (Hassig et al., 1997; Guenther et al., 2000); N-CoR and mSIn3A/B contain HDACs as part of the complex. In non-neuronal cells, neural specific genes containing the RE-1 response element bind a transcriptional repressor called REST, which can recruit HDAC-containing complexes, such as N-CoR and mSI3A/B to RE-1 response elements and induce silencing of the neural genes. Within the developing nervous system itself, HDAC activity plays a role in controlling lineage specification and terminal differentiation (Marin-Husstege et al., 2002; Cunliffe and Casaccia-Bonnefil, 2005). Similar systems no doubt also operate to control differentiation in other lineages. DNA methylation, occurring primarily at CpG dinucleotide palindromes, also plays a role in the regulation of gene activity in a heritable manner (Bird, 1986). It is evident that embryogenesis is associated at particular times with waves of DNA methylation and demethylation, and the heritability of DNA methylation once established could provide a basis for stable repression or activation of gene expression (Reik et al., 2001). However, the precise relationship of methylation to gene expression is certainly complex, and it may be that DNA methylation is associated with stabilization of gene expression patterns and thus stability of cell phenotypes, rather than their initial induction.

EXTRINSIC CONTROLS Whatever the internal mechanisms that control cell phenotype, cell differentiation during development is directed by external cues, whether they be direct cell–cell interactions, or interactions of cells with their substrate, or responses to diffusible factors. Such cues are essential for correct patterning during embryonic development, so that cells with appropriate phenotypes appear in the correct spatial and temporal relationships to one another. An enduring concept is that of the morphogen, an external cue that exhibits differing levels of activity across part of an embryo so that cells that are able to detect differences in its activity can identify their location and respond appropriately. For many years the concept remained hypothetical until a potential patterning role for retinoic acid in the developing limb bud was identified (Thaller and Eichele, 1987). Gradients of retinoic acid were also postulated to pattern the anterior–posterior axis of the embryo, such that higher concentrations are associated with a more posterior cell identity (Durston et al., 1989). A possible molecular mechanism for this was provided by the finding that expression of the HOX genes is induced in human EC cells by retinoic acid in a concentration-dependent manner consistent with their expression patterns along the anterior–posterior axis of the developing embryo (Simeone et al., 1990). However, the view that a gradient of retinoic acid is responsible for patterning the limb bud proved too simplistic and now interactions of several signaling molecules, including Sonic Hedgehog and Bone Morphogenetic Proteins (BMP), are believed to be involved (Drossopoulou et al., 2000). Indeed, the interplay among several signaling molecules seems to provide a common way by which different domains are established in the developing embryo. For example, in the gastrulating mouse embryo, proteins encoded by the Wnt and Nodal genes are produced by the posterior epiblast (Conlon et al., 1994;

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Varlet et al., 1997; Liu et al., 1999), while the opposing anterior viseral endoderm produces the Nodal inhibitors encoded by Cerberus-like and Lefty1 and the Wnt inhibitor encoded by Dkk1 (Belo et al., 1997; Glinka et al., 1998; Perea-Gomez et al., 2002). The Wnt and Nodal proteins induce primitive streak formation, while their inhibitors produced on the anterior side of the embryo create a gradient of signaling activity, thereby restricting primitive streak induction. A second embryo axis develops as a result of BMP4 signaling from the extraembryonic endoderm to the proximal epiblast (Coucouvanis and Martin, 1999); again, antagonists to BMP signaling are produced in the distal regions to create a gradient that patterns the proximal:distal axis (Brennan et al., 2001). In the context of regenerative medicine, understanding the cues that operate to guide cell differentiation during embryogenesis is crucial to developing protocols that will permit appropriate differentiation of stem cells in vitro, or, indeed, if the possibility of activating putative endogenous stem cells for tissue repair is ever realized. Many of the signaling molecules that do play a role in embryogenesis do also influence the behavior of stem cells in culture. For example, retinoic acid has found wide use for inducing differentiation in EC and ES cells, since Strickland and Madhavi (1978) first found that it can induce differentiation of the mouse EC cell, F9. Thus, retinoic acid also induces the differentiation of both human EC and ES cells (Andrews, 1984; Draper et al., 2002). Similarly, members of the BMP, Nodal and fibroblast growth factor (FGF) families of signaling molecules have been found to influence either self-renewal or differentiation of human EC and ES cells in culture (Andrews et al., 1994; Pera et al., 2004; Itsykson et al., 2005; Levenstein et al., 2005; Vallier et al., 2005).

CONCLUSIONS The development of regenerative medicine depends upon learning how to manipulate the phenotypes of cells. Whether, ultimately, the techniques used reflect normal physiological processes that operate during embryogenesis, or whether they are to a greater or lesser extent artifactual, is probably of little consequence provided that the terminal cells produced exhibit the required functions. Nevertheless, to develop rational approaches to the manipulation of cell phenotype does depend upon a detailed understanding of the mechanisms that do operate in vivo. At present, our understanding is limited and a pertinent question is whether our current concepts are adequate to the task? Most current studies of cell differentiation focus upon the effects of one, or a few cues in isolation, for example the response of signaling pathways initiated by binding of a particular ligand, such as BMP, or FGF or Wnt, or retinoic acid, to its receptor and consequent changes in gene activity. However, cells are exposed to a very large number of signals and most of the signaling pathways interact with one another, so that cellular responses analyzed in isolation might be misleading and not adequately reflect responses in vivo. We wonder, therefore, whether a proper understanding of cell phenotypes and the ways in which they can be altered will depend on developing concepts that can embrace complex control networks. One notion (Andrews, 2002) builds on Waddington’s ideas of an “epigenetic landscape” (Waddington, 1956, 1962, 1966). In this view, a cell can be considered to be capable of adopting a vast array of possible “states” reflecting the consequences of all possible permutations of gene activity in a given environment; these states will be associated, in thermodynamic terms, with different levels of free energy. States associated with low free energy levels would be relatively stable and would correspond to cell phenotypes that we observe; those states with high free energy levels would be unstable and would correspond to cell phenotypes we do not observe. Nevertheless, to convert from one phenotype to another, a cell would need to pass through such high energy, unstable “transition” states: the analogy is with activation states in chemical transitions.

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In this model, the normally observed lineages of differentiation would correspond to successive transitions to lower energy states, but along pathways for which there would be the lowest energy barriers and hence highest probability of transition. Movement in the opposite direction, corresponding to transdifferentiation and transdetermination could also occur but depending upon the relative energy barriers, the probability of such changes could be considerably lower. Factors or conditions that promote differentiation could function by altering the landscape so that barriers between particular stable states are lowered, increasing the probability of those transitions. Whether this or other models that seek to conceptualize the mechanisms by which cells change their phenotypes are eventually useful, will depend upon more detailed understanding than is available now of the chemistry of the intrinsic and extrinsic factors that control cell fate, in the environment and at a scale pertinent to individual cells. We also need to establish whether cellular differentiation processes are essentially stochastic until a dominant cellular phenotype takes hold in a population of cells, or whether they are based on binary divisions driven by intrinsic and extrinsic cues. Coupling insights into phenotypic changes at the single cell level with developmental cues gleaned from developmental biology should eventually enable a more directed approach to differentiation of stem cells, and stabilization of desired differentiated phenotypes for use in cell therapeutic applications.

ACKNOWLEDGMENTS This work was supported in part by a grant from the Medical Research Council of the United Kingdom.

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9 Somatic Cloning and Epigenetic Reprogramming in Mammals Heiner Niemann, Christine Wrenzycki, Wilfried A. Kues, Andrea Lucas-Hahn, and Joseph W. Carnwath

INTRODUCTION: SHORT HISTORY OF CLONING More than 50 years ago Briggs and King (1952) showed that normal hatched tadpoles could be obtained after transplanting the nucleus of a blastula cell into the enucleated egg of the amphibian, Rana pipiens. However, while cloning with embryonic cells resulted in normal offspring, development became more and more restricted when cells from more differentiated stages of development were employed (Briggs and King, 1952). This led to the hypothesis that the closer the nuclear donor is developmentally to early embryonic stages the more successful nuclear transfer (NT) is likely to be. This concept prevailed for many years (Gurdon and Byrne, 2003). Cloning of mammals became possible when technology had been developed that allowed micromanipulation of the small mammalian egg (120 μm), which is only one-tenth the diameter of an amphibian egg. This equipment became available in the late 1960s and early 1970s. The first report of cloning an adult mammal was that of Illmensee and Hoppe (1981), who reported the birth of three cloned mice after transfer of nuclei from inner cell mass (ICM) cells into enucleated zygotes. Unfortunately, these results were not repeatable in other laboratories. Subsequently it was shown that development to blastocysts could only be obtained when the nucleus of a zygote or a 2-cell embryo was transferred into an enucleated zygote (McGrath and Solter, 1983) and no development was obtained when donor cell nuclei from later developmental stages were used (McGrath and Solter, 1984). The concept that NT was only successful when both donor and recipient were at the same developmental stage contrasted with the results of the amphibian experiments, which had demonstrated the use of unfertilized eggs as recipients of somatic donor cell nuclei. However, the contradiction did not withstand the test of time. Willadsen (1986) soon demonstrated the use of blastomeres from cleavage stage mammalian embryos (sheep) for transfer into enucleated oocytes. This formed the basis for the successful embryonic cloning in rabbits (Stice and Robl, 1988), mice (Cheong et al., 1993), pigs (Prather et al., 1989), cows (Sims and First, 1994), and monkeys (Meng et al., 1997). Eventually in 1996, the full potential of somatic cloning in mammals became evident for the first time. Campbell et al. (1996) had success in using cells from an established cell line derived from a day 9 ovine embryo and maintained in vitro for 6–13 passages. These cells had been blocked into a quiescent state by serum starvation prior to fusing them with enucleated sheep oocytes. Transfer of these NT derived embryos resulted in two healthy cloned sheep and formed the basis for the birth of “Dolly,” the first mammal cloned from an adult donor cell, reported a year later by the same laboratory (Wilmut et al., 1997). Somatic NT has been successful in a total of 11 species, including sheep (Wilmut et al., 1997), cattle (Kato et al., 1998), mouse (Wakayama et al., 1998), goat (Baguisi et al., 1999), pig (Onishi et al., 2000; Polejaeva et al., 2000), cat (Shin et al., 2002), rabbit

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(Chesne et al., 2002), mule (Woods et al., 2003), horse (Galli et al., 2003), rat (Zhou et al., 2003), and dog (Lee et al., 2005). Worldwide research efforts have been undertaken to unravel the underlying mechanisms for successful somatic NT. The most critical factor is epigenetic reprogramming of the transferred somatic cell nucleus from its differentiated status into the totipotent stage of the early embryo. This involves erasure of the gene expression program of the respective donor cell and the re-establishment of the well orchestrated sequence of expression of an estimated 10,000–12,000 genes regulating embryonic and fetal development. Somatic NT holds great promise for basic biological research and for various agricultural and biomedical applications. The following is a comprehensive review of the present state of somatic cell cloning, including potential areas of application, with emphasis on the epigenetic reprogramming of the transferred somatic cell nucleus.

TECHNICAL ASPECTS OF SOMATIC NT Common somatic cloning protocols involve the following major technical steps (Figures 9.1 and 9.2): (1) enucleation of the recipient oocyte, (2) preparation and subzonal transfer of the donor cell, (3) fusion of the two components, (4) activation of the reconstructed complex, (5) temporary culture of the reconstructed embryo, and finally (6) transfer to a foster mother or storage in liquid nitrogen. Compelling evidence indicates that oocytes at the metaphase II stage rather than any other developmental stage are the most appropriate recipient for the production of viable cloned mammalian embryos. These oocytes possess high levels of maturation promoting factor (MPF), which is thought to be critical for development of the reconstructed embryo (Miyoshi et al., 2003). In many domesticated species, oocytes can be obtained from abattoir ovaries. These need to be matured in vitro but provide a potentially unlimited source of material for cloning experiments. In vitro maturation protocols have advanced to the extent that in vitro

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Figure 9.1 Sequence of steps in somatic cloning of pigs: IVM and enucleation of porcine oocytes. (a) Porcine cumulus oocyte complexes after isolation from abattoir ovaries. (b) Porcine oocyte after 42 h of IVM, note the expansion of the cumulus cells. (c) Microsurgical removal of the polar body plus adjacent cytoplasm containing the metaphase II chromosomes. (d) Microsurgical enucleation after labeling the DNA with a specific stain; note the fluorescence of the DNA within the cytoplasm indicating the metaphase plate and the polar body located in the enucleation pipette.

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(a)

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Figure 9.2 Sequence of steps in somatic cloning: from donor cell production to cloned blastocysts. (a) Porcine fetus from day 25 after insemination. (b) Outgrowing fibroblasts from minced fetal tissue, cultured as adhesive cells. (c) Isolated fibroblasts ready to be sucked up by the transfer pipette. (d) Transfer of a porcine fetal fibroblast into the perivitelline space of the enucleated recipient oocyte. (e) Fusion of the donor cell with the cytoplast in the electric field; note the great difference in size between donor cell and recipient. (f) Successful fusion of both components within 15 min. The donor cell has been completely integrated into the cytoplasm and is not further visible. (g) Cloned porcine blastocyst after 7 days of culture, image taken during the hatching process.

matured (IVM) oocytes can be used for somatic cloning without major losses and are comparable to their in vivo matured counterparts. Oocytes are enucleated by sucking or squeezing out a small portion of the oocyte cytoplasm, specifically the portion closely apposed to the first polar body, where the metaphase II chromosomes are usually located. The oocyte is treated with a mycotoxin, cytochalasin B, to destabilize the cytoskeleton, but this is washed out immediately after microsurgical removal of the chromosomes. In the second step, an intact donor cell (i.e. nucleus plus cytoplasm) is isolated from a cell culture dish by trypsin treatment and is inserted under the zona pellucida in intimate contact to the cytoplasmic membrane of the oocyte with the aid of an appropriate micropipette. These two components are then fused, usually by short, high voltage pulses through the point of contact between the two cells. In mice, instead of electrofusion, the piezoelectric microinjection tool is used. The donor cell membrane is disrupted and removed through repeated suction into thin glass pipettes and the remaining nucleus is injected into the oocytes’ cytoplasm (Wakayama et al., 1998). Activation of NT complexes is achieved either by short electrical pulses or by brief exposure to chemical substances such as ionomycin or dimethylaminopurin (DMAP), regulating the calcium influx into the complexes and/or the cell cycle. Cloned embryos can be cultured in vitro to the blastocyst stage (5–7 days) to assess the initial developmental competence prior to transfer into a foster mother. Another approach, which is frequently used in pigs, is the immediate transfer of the activated NT complexes into the oviducts of the recipient. Various somatic cells, including mammary epithelial cells, cumulus cells, oviductal cells, leukocytes, hepatocytes, granulosa cells, epithelial cells, myocytes, neurons, lymphocytes, and germ cells, have successfully been used as donors for the production of cloned animals (Brem and Kuhholzer, 2002; Hochedlinger and Jaenisch, 2002; Miyoshi et al., 2003; Eggan et al., 2004). In experiments with mice, nuclei from various cancer cells, including leukemia, lymphoma, and breast cancer, could be reprogrammed by NT and yielded apparently normal blastocysts; however, embryonic stem (ES) cells could not be derived from such blastocysts (Hochedlinger et al., 2004). In contrast, ES cells could be derived from blastocysts cloned from melanoma cells, and ES cells were subsequently able to differentiate into various cell types. Chimeras obtained from these ES cells showed a high incidence of tumor formation (Hochedlinger et al., 2004) suggesting that the tumorigenic potential of the

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donor cells was not fully erased by the reprogramming process. It is unclear which cell type is the most successful for NT into oocytes. No differences were found when the efficiency of cloning was compared using various somatic cell types, including those of adult, newborn or fetal, female or male donor cattle (Kato et al., 2000). Although initial experience suggested that cloning with adult somatic cells was only successful when cells were from the female reproductive tract, including mammary epithelial, cumulus, granulosa, or oviductal cells, male mice were eventually cloned from tail-tip cells (Wakayama and Yanagimachi, 1999) and subsequently similar developmental rates were observed for embryos cloned from either male or female nuclei in cattle and mice (Kato et al., 2000; Wakayama and Yanagimachi, 2001). Nevertheless, current experience in our laboratory still shows a bias for female donor cells in bovine NT (Lucas-Hahn et al., 2002). Fetal cells, in particular fibroblasts, have frequently been used in somatic cloning experiments, because they are thought to have less genetic damage and a higher proliferative capacity than adult somatic cells. Cells from early passages are most often chosen for somatic cloning, but high rates of development have also been obtained when donor cells from later passages of adult somatic cells were employed (Kubota et al., 2000). Whether donor cells need to be forced into a quiescent state by either serum starvation or treatment with cell cycle inhibitors is still a matter of debate. In most experiments, donor cells are induced to exit the cell cycle by serum starvation, which holds cells at the G0/G1 cell cycle stage (Campbell et al., 1996). Specific cyclindependent kinases, such as roscovitin, have been reported to increase the efficiency of the cloning process, although final evidence in the form of healthy offspring is lacking (Miyoshi et al., 2003). Nevertheless, unsynchronized somatic donor cells have been successfully used to clone offspring in mice and cattle (Cibelli et al., 1998; Wakayama et al., 1999). There is currently a great need to develop reliable methods to select or produce donor cells which are more efficient for somatic NT. The successful cloning of mice from terminally differentiated cells such as B- and T-lymphocytes or neurons demonstrated unequivocally that a fully differentiated nucleus can be returned to a genetically totipotent stage (Hochedlinger and Jaenisch, 2002; Eggan et al., 2004). Nevertheless, current results cannot yet completely rule out that at least some of the cloned offspring may have been derived from less differentiated adult cells, such as adult stem cells, present in low numbers in primary cell cultures. The chromatin of adult stem cells might to a large extent resemble that of ES cells, which have been shown to be significantly more efficient with regard to cloning in mice (Hochedlinger and Jaenisch, 2003). One option to improve cloning efficiency is to use less differentiated cells (fetal or stem cells) to minimize the complicated and error prone reprogramming process. Current results using such cells in various species are not yet conclusive, but somatic stem cells have been used successfully to give porcine blastocyst development and the birth of live piglets (Zhu et al., 2004; Hornen et al., 2006).

SUCCESS RATES OF SOMATIC CLONING AND THE QUESTION OF NORMALITY OF CLONED OFFSPRING The typical success rate (live births) of mammalian somatic NT is low and usually is only 1–2%. Cattle seem to be an exception to this rule as levels of 15–20% can be reached (Kues and Niemann, 2004). Pre- and postnatal development is often compromised and a variable proportion of the offspring shows aberrant developmental patterns and increased perinatal mortality. These abnormalities include a wide range of symptoms, summarized as “large offspring syndrome” (LOS). These include extended gestation length, oversized offspring, aberrant placenta, cardiovasculatory problems, respiratory defects, immunological deficiencies, problems with tendons, adult obesity, kidney and hepatic malfunctions, behavioral changes, and a higher susceptibility to neonatal diseases (Renard et al., 1999; Tamashiro et al., 2000; Ogonuki et al., 2002; Perry and Wakayama, 2002; Rhind et al., 2003). These pathologies have most often been observed in cloned ruminants

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and mice. The incidence is stochastic and has not been correlated with aberrant expression of single genes or specific pathophysiology. A new term “abnormal offspring syndrome” (AOS) with subclassification according to the outcome of such pregnancies has recently been proposed to better reflect the broad spectrum of this pathological phenomenon (Farin et al., 2006). The general assumption is that the underlying cause for these pathologies is insufficient or faulty reprogramming of the transferred somatic cell nucleus. However, a critical survey of the published literature on cloning of animals revealed that, most cloned animals are healthy and develop normally (Cibelli et al., 2002). This is consistent with the finding that mammalian development is rather tolerant of minor epigenetic aberrations in the genome and subtle abnormalities in gene expression do not interfere with the survival of cloned animals (Humphreys et al., 2001). It has become clear that once cloned offspring have survived the neonatal period and are approximately 6 months of age (cattle, sheep), they are not different from age matched controls with regard to numerous biochemical blood and urine parameters (Lanza et al., 2001; Chavatte-Palmer et al., 2002), immune status (Lanza et al., 2001), body score (Lanza et al., 2001), somatotrophic axis (Govoni et al., 2002), reproductive parameters (Enright et al., 2002), and yields and composition of milk (Pace et al., 2002). No differences were found in meat and milk composition of bovine clones when compared with age matched counterparts; all parameters were within the normal range (Kumugai, 2002; Tian et al., 2005). Similar findings were reported for cloned pigs (Archer et al., 2003). According to expert committees of the National Academy of Science of the USA and from the Japanese Ministry of Agriculture, Forestry and Fisheries (MAFF), and the Food and Drug Administration (FDA) of the USA. There is no scientific basis for questioning the safety of food derived from cloned animals. However, due to the limited experience with somatic cloning which has only been in general use since 1997 and the relatively long generation intervals in domestic animals, specific effects of cloning on longevity and senescence have not yet been fully assessed. Preliminary data indicate no pathology in second generation of cloned mice and cattle (Wakayama et al., 2000; Kubota et al., 2004).

EPIGENETIC REPROGRAMMING Basic Epigenetic Mechanisms DNA Methylation Normal development depends on a precise sequence of changes in the configuration of the chromatin which are primarily related to the acetylation and methylation status of the genomic DNA. These epigenetic modifications control the precise tissue-specific expression of genes. It is estimated that the mammalian genome with its 25,000 genes contains 30,000–40,000 CpG islands (i.e. areas which are rich in CG dinucleotides). These CpG islands are predominantly found in the promoters of housekeeping genes, but are also observed in tissue-specific genes (Antequera, 2003). The correct pattern of cytosine methylation in CpG dinucleotides is required for normal mammalian development (Li et al., 1993). DNA methylation is also thought to play a crucial role in suppressing the activities of parasitic promoters and is thus part of the gene silencing system in eukaryotic cells (Jones, 1999). Usually, methylation is associated with silencing of a given gene, but an increasing number of genes are found to be activated by methylation marks, specifically tumor suppressor genes (Bestor and Tycko, 1996; Jones, 1999). DNA methylation critically depends on the activity of specific enzymes, the DNA methyltransferases (Dnmts) (Figure 9.3). DNA-methytransferase1 (Dnmt1) is a maintenance enzyme that is responsible for restoring methylation to hemi-methylated CpG dinucleotides after DNA replication (Bestor, 1992). The oocyte-specific isoform, Dnmt1o, maintains maternal imprints. Dmnt3a and Dmnt3b catalyze de novo methylation and are thus critical for establishing DNA methylation during development (Hsieh, 1999; Okano et al., 1999). DNA methyltransferase 3-like protein (Dmnt3L) co-localizes with Dnmt3a and -b and presumably

Cloning and Reprogramming 153

De novo methylation Dnmt3b

Dnmt3a

Active demethylation

Dnmt1 Maintenance methylation

Passive demethylation

Figure 9.3 Methylation and de-methylation of DNA. The drawing shows DNA modifications by methylation and the involvement of various DNA-methyltransferases (Dnmts) and their function during methylation, de- and remethylation of a DNA strand.

Embryonic lineage

Maternal genome “Cloned” genome

Extraembryonic lineage Paternal genome IVP

NT

Figure 9.4 Methylation reprogramming of the genome during early bovine development. The paternal genome is rapidly and actively demethylated after fertilization, while the maternal genome becomes passively demethylated over time during cleavage. The embryonic genome is remethylated starting at the morula stage; the two cell lineages of the bovine blastocyst are methylated to different levels. In cloned embryos the methylation pattern may be completely different. Adapted from Dean et al. (2001), PNAS 98, 13734–13738.

is involved in establishing specific methylation imprints in the female germline (Bourc’his et al., 2001b). Dnmt activities are linked with histone deacetylases (HDACs), histone methyltransferases (HMTs), and several ATPases and are part of a complex system regulating chromatin structure and thus gene expression (Burgers et al., 2002). During early development, reprogramming of the DNA is observed shortly before and shortly after formation of the zygote (Figure 9.4). Paternal DNA is actively demethylated after fertilization, while the female DNA undergoes de novo methylation in several species, including murine, bovine, porcine, rat, and human zygotes (Mayer et al., 2000; Oswald et al., 2000; Dean et al., 2001; Santos et al., 2002; Beaujean et al., 2004; Xu et al., 2005). The maternal genome is then passively demethylated and the embryonic DNA begins to be remethylated at species-specific cell stages (Figure 9.4; Dean et al., 2001).

154 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

Imprinting Imprinting represents a specific function of DNA methylation. A typical feature of genomic imprinting is that the two alleles of a given gene are expressed differently. Usually one allele, either the maternal or the paternal, is silenced throughout development by covalent addition of methyl groups to cytosine residues in CpG dinucleotides (Constancia et al., 2004). This DNA methylation occurs in imprinting control regions (ICRs) of DNA and is established by the de novo methyltransferase Dnmt3a. A typical feature of imprinted genes is that they are found in clusters and the ICRs exert regional control of gene expression (Reik and Walter, 2001). In the mouse no more than 50 and in humans no more than 80 imprinted genes have been identified (Dean et al., 2003; Constancia et al., 2004). Imprinting is a genetic mechanism that regulates the demand, provision, and use of resources in mammals particularly during fetal and neonatal development. Usually genes expressed from the paternally inherited allele increase resource transfer from the mother to the fetus, whereas maternally expressed genes reduce this transfer to secure the mother’s well-being (Constancia et al., 2004). Imprints are established during development of germ cells into sperm and eggs. The germ line resets imprints such that mature gametes reflect the sex of a specific germ line due to the sequence of erasure and establishment (Reik and Walter, 2001). Histone Modifications Histones are the main protein component of chromatin and the core histones form the nucleosome. Covalent post-translational modifications of histones play a crucial role in controlling the capacity of the genome to store, release, and inherit biological information (Fischle et al., 2003). Numerous histone and chromatin related regulatory options are available including histone acetylation, phosphorylation, and methylation. Binary switches and modification cassettes have been suggested as new concept to understand the enormous versatility of histone function (Fischle et al., 2003). Specific HMTs catalyze methylation at specific positions of the nucleosome in mammalian cells. Deacetylation of histones is carried out by isoforms of HDACs. Histone acetyltransferases are involved in diverse processes including transcriptional activation, gene silencing, DNA repair, and cell-cycle progression and thus play a critical role in cell growth and development (Carrozza et al., 2003). Reprogramming can be divided into the pre-zygotic phase, which includes acquisition of genomic imprints and the epigenetic modification of most somatic genes during gametogenesis. X-chromosome inactivation and adjustment of telomere lengths are prominent examples of post-zygotic reprogramming (Hochedlinger and Jaenisch, 2003). Pre-zygotic Reprogramming Imprinted Gene Expression in Cloned Embryos and Fetuses The majority of imprinted genes are involved in fetal and placental growth and differentiation which makes them promising candidates for unraveling the developmental aberrations found after somatic NT. Disruption of imprinted genes has been observed in cloned mouse embryos (Mann et al., 2003). Knowledge about imprinted genes in bovine development is limited; only one out of eight genes known to be imprinted in mice appeared to be imprinted in bovine blastocysts (Ruddock et al., 2004). The normally imprinted H19 gene was expressed bi-allelically in bovine stillborn cloned calves, suggesting that aberrant imprinting is associated with abnormal development (Zhang et al., 2004). When calves survived, faulty H19 imprinted expression was corrected in the offspring showing that germline development was normal (Zhang et al., 2004). Genomic imprinting can be disrupted at the Xist (X-chromosome inactive specific transcript) locus in cloned fetuses, whereas IGF2 and GTL2 are properly expressed in fetal and placental tissue (Dindot et al., 2004). As in other species, the bovine IGF2 gene is controlled by an extremely complex regulatory mechanism based on multiple promoters, alternative splicing, and genomic imprinting which can be severely perturbated in cloned fetal, calf, and adult tissue (Curchoe et al., 2005). Recently, a differentially methylated region (DMR) has been

Cloning and Reprogramming 155

discovered in exon 10 of the bovine IGF2 gene (Gebert et al., 2006). This gene is critically involved in fetal and placental development and known to be imprinted in mice (Constancia et al., 2002). Thus, the basis for in-depth studies on imprinted expression in bovine development has been established. NT and Embryonic Gene Expression Patterns Cloning typically uses the unfertilized matured oocyte as the recipient cell. Reprogramming must occur within the short interval between the transfer of the donor cell into the oocyte and the initiation of embryonic transcription, the timing of which is species specific. In the mouse, embryonic transcription begins at the 2-cell stage, in the pig at the 4-cell stage and in sheep, cattle, and human at the 8–16-cell stage (Telford et al., 1990). The effects of somatic cloning on mRNA expression patterns have mostly been analyzed in bovine morula and blastocyst stages and numerous genes related to specific physiological functions have been identified as aberrantly expressed in cloned embryos as compared to their in vivo derived counterparts (Wrenzycki et al., 2005b). This group includes genes related to stress susceptibility, growth factor signaling, imprinting, trophoblast formation and function, sex-chromosome related mRNA expression, X-chromosome inactivation, DNA methylation, and histone modifications (Wrenzycki et al., 2005b; Nowak-Imialek et al., 2006). Expression of the transcription factor Oct4 within a certain range is crucial for maintaining toti- and pluripotency in early embryos. Oct4 is a transcription factor for a panel of developmentally important genes (Niwa et al., 2000; Pesce and Schöler, 2001). Aberrant spatial expression of Oct4 was found in murine embryos cloned from cumulus cells (Boiani et al., 2002). In a high proportion, up to 40% of cloned mouse embryos, Oct4 regulated genes were aberrantly expressed due to faulty reactivation of Oct4 (Bortvin et al., 2003). These findings indicate dysregulation of the pluripotent state in embryonic cells which could contribute to developmental failures in cloned embryos. Data from our laboratory have shown that Dnmt1 mRNA expression was significantly increased in cloned bovine embryos compared to in vivo derived controls. Similar observations have been made for Dnmt3a, while Dnmt3b expression did not differ between cloned, in vitro produced and in vivo produced bovine embryos (Figure 9.5; Wrenzycki and Niemann, 2003). Similarly, mice cloned from cumulus cells showed aberrant Dnmt1 localization and expression (Chung et al., 2003). These findings suggest perturbation of the normal wave of deand remethylation in early development, which could lead to developmental abnormalities in cloned animals. The pattern of aberrations in mRNA expression was extremely variable in embryos derived by in vitro production and/or cloning. Embryo production methods thus cause significant up- or downregulation, de novo induction, or silencing of the genes critically involved in embryonic and fetal development (Niemann and Wrenzycki, 2000). Some of the aberrant expression patterns found in cloned blastocysts could be the result of aberrant allocation of cells to the ICM and trophectoderm (Koo et al., 2003). But in most cases faulty expression patterns seem to be related to epigenetic errors rather than morphological deviations. Extended in vitro culture of mammalian embryos alone is known to result in aberrations in mRNA expression patterns, affecting imprinted and non-imprinted genes (Wrenzycki et al., 2001a; Young et al., 2001). In the case of cloning, it is difficult to discriminate between the effect of in vitro culture and dysregulation due to the cloning process. A recent analysis using a bovine cDNA microarray with 6,298 unique sequences revealed that the mRNA expression profile of cloned bovine embryos was completely different from that of the donor cells and was surprisingly similar to that of naturally fertilized embryos (Smith et al., 2005). This is confirmed by previous reverse transcriptase polymerase chain reaction (RT-PCR) analyses (Wrenzycki et al., 2001b, 2005a, b). A greater number of genes were differentially expressed in comparisons of artificial insemination (AI) and in vitro fertilization (IVF) embryos (n  198) and between NT and IVF embryos (n  133) than between NT and AI embryos (n  50), indicating that cloned embryos had undergone significant nuclear reprogramming at the blastocyst stage (Smith et al., 2005). In this case, it was suggested that

Relative abundance

Relative abundance

156 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

4.00

b

Dnmt1

b b

3.00 2.00

Maintenance methylation

a

1.00 0.00

Donor cell

1-Cell stage Mat. oocyte parth IVP NT

8-Cell stage IVP NT

Blastocyst In IVP parth NT vivo

5.00 4.00

b

Dnmt3a

3.00

a a a

2.00 De novo methylation

1.00 0.00 Donor cell

1-Cell stage Mat. oocyte parth IVP NT

8-Cell stage IVP NT

Blastocyst In IVP parth NT vivo a:b p0.05

Figure 9.5 mRNA expression pattern determined by gene-specific RT-PCR of the two Dnmts (Dnmt1 and Dnmt3a). The donor cells do not show Dnmt expression. Dnmt expression increases throughout early development and shows significant differences between blastocyst stages of various origin. Blastocysts cloned from fetal fibroblasts have an increased mRNA expression for Dnmt1 compared to in vivo produced control embryos. For Dnmt3a mRNA expression is also elevated for parthenogenetic (parth) and in vitro produced (IVP) blastocysts. Wrenzycki and Niemann (2003), RBMOnline 7, 135–142.

aberrations cause effects later in development during organogenesis because small reprogramming errors are magnified downstream in development. We have developed the hypothesis that deviations from the normal pattern of mRNA expression which are observed in the early preimplantation embryo persist throughout fetal development up to birth and that the many effects of this period of culture only become manifest later in development (Niemann and Wrenzycki, 2000). Consistent with this hypothesis, genes aberrantly expressed in blastocysts were also aberrantly expressed in the organs of clones that died shortly after birth (Li et al., 2005). This is particularly true for Xist and heat shock protein (HSP) for which aberrant expression patterns had been found in cloned blastocysts (Wrenzycki et al., 2001b, 2002). DNA Methylation Patterns in Cloned Embryos and Fetuses With regard to cloning, it is critical to assess to what extent the chromatin changes required by an adult somatic donor nucleus are similar to the changes which take place in gametogenesis and fertilization (Jaenisch and Wilmut, 2001). The abnormalities in cloned fetuses and live offspring cannot simply be due to the source of the donor nuclei. The most likely explanation for the variability is that it reflects the extent of failure in genomic reprogramming of the transferred nucleus. Cloned embryos all show aberrant patterns of the global DNA methylation (Dean et al., 2001; Kang et al., 2001a, b). The maintenance of high methylation levels during cleavage is thought to be related to the presence of the somatic form of Dnmt, an enzyme brought by the somatic donor cell nucleus into the cloned embryo. This probably interferes with the genome-wide demethylation process that takes

Cloning and Reprogramming 157

place in a normal preimplantation embryo (Reik et al., 2001). Methylation reprogramming is delayed and incomplete in cloned bovine embryos (Bourc’his et al., 2001a). A high degree of variability is observed among individual embryos with regard to methylation levels (Dean et al., 2001). At present it is not fully clear whether the aberrant methylation stems from a defective demethylation of the transferred somatic nucleus or is a consequence of failed nuclear re-organization. Only cloned ovine embryos which show re-organized chromatin appear to survive the early embryonic phase (Beaujean et al., 2004). Attempts to improve the developmental capacity of bovine cloned embryos by either complete or partial erasure of DNA methylation/acetylation of the donor cell by treatment with specific inhibitors prior to use in NT have met with only limited success (Enright et al., 2003, 2005). In support of the hypothesis that aberrant mRNA expression patterns persist throughout subsequent development (Niemann and Wrenzycki, 2000), epigenetic analysis revealed that methylation errors produced early in preimplantation development are in fact maintained throughout development and these genome-wide epigenetic aberrations can be identified in cloned bovine fetuses (Cezar et al., 2003). The proportion of methylated cytosine residues is reduced in cloned fetuses compared to in vivo produced controls and survivability of cloned bovine fetuses was found to be closely related to the reduced global DNA methylation status (Cezar et al., 2003). Significant hypermethylation was detected in the liver tissue of cloned bovine fetuses and was found to be correlated with fetal overgrowth (Hiendleder et al., 2004a). These results show that developmental abnormalities can be associated with both hypo- and hypermethylation during fetal bovine development. Remarkably, the degree of demethylation of repetitive sequences in the donor genome seems to be determined by the recipient ooplasm and not by the donor cell. Ooplasm from different species may have different capacity to demethylate-specific genes (Chen et al., 2006). The cytoplasm of the bovine oocyte may be particularly advantageous in this respect. The use of defined sources of highly effective recipient oocytes could render somatic cloning more efficient and could give significant improvements in the cloned phenotype (Hiendleder et al., 2004b). While reprogramming is considered to be essential in successful NT, it may not be the only factor affecting cloning efficiency. Additional improvements have been made by technical modifications (Hiiragi and Solter, 2005). Altogether, it is apparent that there has been a steady increase in the efficiency of somatic mammalian cloning since it was first described in 1997. Post-zygotic Reprogramming X-Chromosome Inactivation After Somatic Cloning X-chromosome inactivation is the developmentally regulated process by which one of the two X-chromosomes in female mammals is silenced early in development to provide dosage compensation for X-linked genes. A single X-chromosome is sufficient as shown in XY males (Lyon, 1961). Although the mechanism of X-chromosome inactivation is not yet fully understood, the paternal X-chromosome is typically inactivated by DNA methylation and remains inactive in placental tissue, whilst in the embryo proper either the paternal or maternal X-chromosome can be randomly selected on a cell by cell basis for inactivation leading to a mosaic pattern in adult cells (Hajkova and Surani, 2004). Recent findings in the mouse revealed that the paternal imprint in the ICM (i.e. the pluripotent cells that give rise to the fetus) is erased from the paternal X-chromosome late in preimplantation development followed by random X-inactivation (Mak et al., 2004). The paternal X-chromosome is partly silent at fertilization and becomes fully inactivated at the 2- or 4-cell stage (Huynh and Lee, 2003; Okamoto et al., 2004). Female somatic NT derived embryos inherit one active and one inactive X-chromosome from the donor cell. mRNA expression analysis of bovine embryos cloned from adult donor cells at the blastocyst stage revealed a significant upregulation of Xist compared to in vitro and in vivo derived embryos. Expression of X-chromosome related genes is delayed in cloned as compared to in vivo derived embryos (Wrenzycki et al., 2002). Premature X-inactivation was observed for the X-chromosome linked inhibitor of apoptosis (XIAP) gene in in vitro produced bovine embryos compared with their

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in vivo counterparts (Knijn et al., 2005). These findings indicate that perturbation of X-chromosome inactivation has occurred by the blastocyst stage after somatic cloning or IVF and culture. In female bovine cloned calves, aberrant expression patterns of X-linked genes and hypomethylation of Xist in various organs of stillborn calves were observed. Random inactivation of the X-chromosome was found in the placenta of deceased clones, but skewed in that of live bovine clones (Xue et al., 2002). This aberrant expression pattern of X-chromosome inactivation initiated in the trophectoderm seems to have resulted from incomplete nuclear reprogramming. Similar findings were obtained in studies of cloned mouse embryos (Eggan et al., 2000). Telomere Length and Somatic Cloning Telomeres are the natural ends of linear chromosomes and play a crucial role in maintaining the integrity of the entire genome by preventing loss of terminal coding DNA sequences or end to end chromosome fusion. Telomeres are composed of repetitive DNA elements and specific DNA proteins, which together form a nucleoprotein complex at the end of eukaryotic chromosomes (Blackburn, 2001). Although the sequence of these terminal DNA structures varies between organisms, telomeres are generally composed of a concatamer of short sequences of the form 5–TTAGGG–3. Changes in telomere length are closely related to ageing and cancer (de Lange, 2002). As a general rule, some loss of telomeres occurs with each cell division as a result of the incomplete replication of the lagging strand. A specialized RNA-dependent DNA polymerase, the telomerase, is then required to maintain the natural length of telomeric DNA. This ribonucleoprotein enzyme is composed of two essential subunits: the telomerase RNA component (TERC) and the telomerase reverse transcriptase (TERT) component (Nakayama et al., 1998). Telomerase is critically involved in maintaining normal telomere length (Blasco et al., 1999). This enzyme is active in hematopoietic cells, cancer cells, germ cells, and in early embryos at the blastocyst stage. Telomeres of the cloned sheep (Dolly), derived from adult mammary epithelial cells, were found to be shortened when compared to age matched naturally bred counterparts and telomere length reduction seemed to be correlated with telomere length of the donor cells (Shiels et al., 1999). Subsequently, however, the vast majority of cloning studies reported that telomere length in cloned cattle, pigs, goats, and mice are comparable with age matched naturally bred controls even when senescent donor cells were used for cloning (Jiang et al., 2004; Betts et al., 2005; Jeon et al., 2005; Schaetzlein and Rudolph, 2005). Regulation of telomere length is to some extent related to the donor cells employed for cloning. Telomere length in cattle cloned from fibroblasts or muscle cells was similar to that of age matched controls while clones derived from epithelial cells did not have telomeres restored to normal length (Miyashita et al., 2002). A checkpoint for elongation of telomeres to their species determined length has been discovered at the morula to blastocyst transition in bovine and mouse embryos (Schaetzlein et al., 2004). Telomeres are at the level of the donor cells in cloned morulae (Figure 9.6), whereas at the blastocyst stage telomeres have been restored to normal length (Figure 9.7). The telomere elongation process at this particular stage of embryogenesis is telomerase dependent since it was abrogated in telomerase deficient mice (Schaetzlein et al., 2004). The morula/blastocyst transition is a critical step in preimplantation development leading to first differentiation into two cell lineages: the ICM and the trophoblast, which coincides with dramatic changes in morphology and gene expression (Niemann and Wrenzycki, 2000).

APPLICATION OF SOMATIC NT Reproductive Cloning of Transgenic Animals Somatic NT holds great potential in three major areas: reproductive cloning, therapeutic cloning, and in basic research (Table 9.1). Improved transgenesis is of special relevance to the field of reproductive cloning due to

Cloning and Reprogramming 159

p = 0.0001 20 Telomere length (kb)

p = 0.1396 15

p = 0.464

10 5 0

Mean:

n=7

n=4

n=6

n=8

In vivo

In vitro

12.42 kb

14.36 kb

NT (adult) 9.47 kb

NT (fetal) 8.69 kb

Figure 9.6 Telomere lengths in bovine morulae as determined by qFISH (quantitative fluorescent in situ hybridization). Telomeres in morulae produced in vivo from superovulated cows or in vitro have significantly longer telomeres compared to morulae cloned from either fetal or adult fibroblasts. Schätzlein et al. (2004), PNAS 101, 8034–8038.

Telomere length (kb)

p = 0.4337 30 25 20 15 10 5 0

Mean:

p = 0.3282

p = 0.0583

ab n=7

fb n=6

cb n=6

21.67 kb

17.26 kb

19.53 kb

Figure 9.7 (a) Telomere length in bovine blastocysts as determined by qFISH. (b) The blastocysts cloned from either fetal (fb) or adult (ab) fibroblasts have similar telomere length as the in vitro produced “control” embryos (cb). Telomere length is restored to physiological length at morula/blastocyst transition. Schätzlein et al. (2004), PNAS 101, 8034–8038. Table 9.1 Application fields for somatic cloning Reproductive cloning

Therapeutic cloning

Basic research

Genetically identical multiplets Transgenic animals (transfection, homologous recombination) Disease models Maintenance of genetic resources Animal breeding strategies (milk, meat, etc.)

Derivation of customized ES cells Targeted differentiation Regenerative cells and tissues (autologous, heterologous) Tissue engineering

Toti and pluripotency Reprogramming De-differentiation Re-differentiation Ageing Tumorigenesis Epigenetics Telomere biology Many other areas

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a number of significant advantages over the previously used microinjection technology (Kues and Niemann, 2004). The major advantage is that somatic donor cells can be transfected with various gene constructs and those cells with the most appropriate expression pattern can be selected in vitro as donor cells. Even targeted genetic modifications such as a gene knock-out by homologous recombination are compatible with primary cell cultures. The transgenic expression patterns render much more control than was possible with microinjection (Kues and Niemann, 2004). Pre-eminent areas of application include the production of recombinant, pharmaceutically valuable proteins in the mammary gland of transgenic livestock (pharming), and the generation of transgenic pigs for xenotransplantation research. Proteins from the mammary gland of transgenic livestock, including antithrombin III, tissue plasminogen action (tPA), or α-antitrypsin, have successfully passed clinical trials and are now subject to registration by national and supranational regulatory agencies. Phase III trials for anti-thrombin III (AT III) (ATryn® from GTC Biotherapeutic, USA) produced in the mammary gland of transgenic goats have been completed and the recombinant protein has been approved as drug by the European Medicines Agency (EMEA) in August 2006. With regard to xenotransplantation, the hyperacute rejection response (HAR), which was the major rejection mechanism, can now reliably be overcome and further immunological hurdles are being tackled by the production of multi-transgenic pigs (Kues and Niemann, 2004). Cloning is the only practical approach to produce multi-transgenic animals for this kind of research as it is the only way to select the genotype precisely. Agricultural applications include modifications of animal products for food consumption, enhanced disease resistance, and the production of environmentally friendly farm animals (Kues and Niemann, 2004). Therapeutic Cloning With regard to therapeutic cloning, the generation of histocompatible tissue by nuclear transplantation has been demonstrated in a bovine model (Lanza et al., 2002). Despite expression of different mitochondrial DNA haptotypes, no rejection responses were observed when cloned renal cells were retransferred to the donor animal. Skin grafts between bovine clones with different mitochondrial haplotypes were accepted long term whereas non-cloned tissues were rejected (Theoret et al., 2006). The feasibility of therapeutic cloning has also been shown in mice where correction of a genetic defect by cell therapy was demonstrated (Rideout et al., 2002). Mouse ES cells derived from cloned or fertilized blastocysts were similar with regard to their transcriptional profile and differentiation potential and thus have equal value as stem cells (Brambrink et al., 2006). Cells cloned from a patient have the advantage that they are accepted by that patient without permanent immune suppression. The production of customized ES cells will be invaluable in human medicine for the treatment of degenerative diseases because no immunosuppressive treatment is required. The concept of “therapeutic cloning” (Figure 9.8) is fascinating but application in human medicine is still in its infancy. Major practical problems include the limited availability of human oocytes for reprogramming of the donor cells, the low efficiency of somatic NT, the difficulty of inserting genetic modifications, the increased risk of oncogenic transformation, and the epigenetic instability of embryos, and cells derived from somatic cloning (Colman and Kind, 2000; Humpherys et al., 2001). Alternatives to NT for reprogramming of somatic cell nuclei for the production of autologous therapeutic cells are being explored (Dennis, 2003). In humans, only preliminary data are available on therapeutic cloning (Cibelli et al., 2001). The papers on human ES cell isolation and cloning (Hwang et al., 2004, 2005) were retracted after discovery of significant fraud (Kennedy, 2006). The long-term goal of therapeutic cloning is to provide data on ES cell growth and differentiation which may make it possible to stimulate proliferation and differentiation of endogenous stem cells and reparation of sick stocks.

CONCLUSIONS Since the birth of Dolly, the first cloned mammal, significant progress has been made in increasing the efficiency of cloning. At the time of writing, cloned animals have been born in 11 species. While the majority of offspring derived from somatic cloning are outwardly normal, cloning may be still associated with pathological

Cloning and Reprogramming 161

Enucleated human oocyte

NT-derived embryo

Biopsied cell (i.e. fibroblast)

Blastocyst

In vitrodifferentiation Therapeutic cells (i.e. cardiomyocytes)

Embryonic stem cells

Figure 9.8 Principle of therapeutic cloning for the production of autologous cardiomyocytes.

side-effects summarized as LOS, which appear to be due to incomplete and/or faulty reprogramming of the genome of the donor nucleus by the oocyte’s cytoplasm. Epigenetic reprogramming is essential for successful cloning and involves a series of critical steps to ensure the well orchestrated gene expression pattern associated with normal development in which DNA methylation and histone modifications play a critical role. X-chromosome inactivation and telomere length restoration are post-zygotic epigenetic tasks that need to be performed for successful cloning. Identification of the specific factors present in the ooplasm which are necessary for epigenetic reprogramming will give us a better understanding of the underlying mechanisms and would improve cloning efficiency. Somatic cloning has promising application potentials and is a useful tool in basic research.

ACKNOWLEDGMENTS The authors gratefully acknowledge the valuable support during the course of the experiments on somatic cloning and reprogramming by various members of the Mariensee laboratory, in particular Doris Herrmann, Erika Lemme, Klaus-Gerd Hadeler, Lothar Schindler, Karin Korsawe, Hans-Herrmann Doepke, Drs Bjoern Petersen and Michael Hoelker. We thank Christine Weidemann for her expert technical assistance in the production of this manuscript. The financial support of the research on which this review is based through various DFG grants is gratefully acknowledged.

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Polejaeva, I.A., Chen, S.H., Vaught, T.D., Page, R.L., Mullins, J., Ball, S., Dai, Y., Boone, J., Walker, S., Ayares, D.L., Colman, A. and Campbell, K.H. (2000). Cloned pigs produced by nuclear transfer from adult somatic cells. Nature 407: 86–90. Prather, R.S., Sims, M.M. and First, N.L. (1989). Nuclear transplantation in early pig embryos. Biol. Reprod. 41: 414–418. Reik, W. and Walter, J. (2001). Genomic imprinting: parental influence on the genome. Nat. Rev. Genet. 2: 21–32. Reik, W., Dean, W. and Walter, J. (2001). Epigenetic reprogramming in mammalian development. Science 293: 1089–1093. Renard, J.P., Chastant, S., Chesne, P., Richard, C., Marchal, J., Cordonnier, N., Chavatte, P. and Vignon, X. (1999). Lymphoid hypoplasia and somatic cloning. Lancet 353: 1489–1491. Rhind, S.M., King, T.J., Harkness, L.M., Bellamy, C., Wallace, W., DeSousa, P. and Wilmut, I. (2003). Cloned lambs – lessons from pathology. Nat. Biotechnol. 21: 744–745. Rideout III, W.M., Hochedlinger, K., Kyba, M., Daley, G.Q. and Jaenisch, R. (2002). Correction of a genetic defect by nuclear transplantation and combined cell and gene therapy. Cell 109: 17–27. Ruddock, N.T., Wilson, K.J., Cooney, M.A., Korfiatis, N.A., Tecirlioglu, R.T. and French, A.J. (2004). Analysis of imprinted messenger RNA expression during bovine preimplantation development. Biol. Reprod. 70: 1131–1135. Santos, F., Hendrich, B., Reik, W. and Dean, W. (2002). Dynamic reprogramming of DNA methylation in the early mouse embryo. Dev. Biol. 241: 172–182. Schaetzlein, S. and Rudolph, K.L. (2005). Telomere length regulation during cloning, embryogenesis and ageing. Reprod. Fert. Develop. 17: 85–96. Schaetzlein, S., Lucas-Hahn, A., Lemme, E., Kues, W.A., Dorsch, M., Manns, M.P., Niemann, H. and Rudolph, K.L. (2004). Telomere length is reset during early mammalian embryogenesis. Proc. Natl Acad. Sci. USA 101: 8034–8038. Shiels, P.G., Kind, A.J., Campbell, K.H., Wilmut, I., Waddington, D., Colman, A. and Schnieke, A.E. (1999). Analysis of telomere lengths in cloned sheep. Nature 399: 316–317. Shin, T., Kraemer, D., Pryor, J., Liu, L., Rugila, J., Howe, L., Buck, S., Murphy, K., Lyons, L. and Westhusin, M. (2002). A cat cloned by nuclear transplantation. Nature 415: 859–860. Sims, M. and First, N.L. (1994). Production of calves by transfer of nuclei from cultured inner cell mass cells. Proc. Natl Acad. Sci. USA 91: 6143–6147. Smith, S.L., Everts, R.E., Tian, X.C., Du, F., Sung, L.Y., Rodriguez-Zas, S.L., Jeong, B.S., Renard, J.P., Lewin, H.A. and Yang, X. (2005). Global gene expression profiles reveal significant nuclear reprogramming by the blastocyst stage after cloning. Proc. Natl Acad. Sci. USA 102: 17582–17587. Stice, S.L. and Robl, J.M. (1988). Nuclear reprogramming in nuclear transplant rabbit embryos. Biol. Reprod. 39: 657–664. Tamashiro, K.L., Wakayama, T., Blanchard, R.J., Blanchard, D.C. and Yanagimachi, R. (2000). Postnatal growth and behavioral development of mice cloned from adult cumulus cells. Biol. Reprod. 63: 328–334. Telford, N.A., Watson, A.J. and Schultz, G.A. (1990). Transition from maternal to embryonic control in early mammalian development: a comparison of several species. Mol. Reprod. Dev. 26: 90–100. Theoret, C., Dore, M., Mulon, P.Y., Desrochers, A., Viramontes, F., Filion, F. and Smith, L.C. (2006). Short- and long-term skin graft survival in cattle clones with different mitochondrial haplotypes. Theriogenology 65: 1465–1479. Tian, X.C., Kubota, C., Sakashita, K., Izaike, Y., Okano, R., Tabara, N., Curchoe, C., Jacob, L., Zhang, Y., Smith, S., Bormann, C., Xu, J., Sato, M., Andrew, S. and Yang, X. (2005). Meat and milk compositions of bovine clones. Proc. Natl Acad. Sci. USA 102: 6261–6266. Wakayama, T. and Yanagimachi, R. (1999). Cloning of male mice from adult tail-tip cells. Nat. Genet. 22: 127–128. Wakayama, T. and Yanagimachi, R. (2001). Mouse cloning with nucleus donor cells of different age and type. Mol. Reprod. Dev. 58: 376–383. Wakayama, T., Perry, A.C., Zuccotti, M., Johnson, K.R. and Yanagimachi, R. (1998). Full-term development of mice from enucleated oocytes injected with cumulus cell nuclei. Nature 394: 369–374. Wakayama, T., Rodriguez, I., Perry, A.C., Yanagimachi, R. and Mombaerts, P. (1999). Mice cloned from embryonic stem cells. Proc. Natl Acad. Sci. USA 96: 14984–14989. Wakayama, T., Shinkai, Y., Tamashiro, K.L., Niida, H., Blanchard, D.C., Blanchard, R.J., Ogura, A., Tanemura, K., Tachibana, M., Perry, A.C., Colgan, D.F., Mombaerts, P. and Yanagimachi, R. (2000). Cloning of mice to six generations. Nature 407: 318–319. Willadsen, S.M. (1986). Nuclear transplantation in sheep embryos. Nature 320: 63–65.

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Wilmut, I., Schnieke, A.E., McWhir, J., Kind, A.J. and Campbell, K.H. (1997). Viable offspring derived from fetal and adult mammalian cells. Nature 385: 810–813. Woods, G.L., White, K.L., Vanderwall, D.K., Li, G.P., Aston, K.I., Bunch, T.D., Meerdo, L.N. and Pate, B.J. (2003). A mule cloned from fetal cells by nuclear transfer. Science 301: 1063. Wrenzycki, C. and Niemann, H. (2003). Epigenetic reprogramming in early embryonic development: effects of in-vitro production and somatic nuclear transfer. Reprod. Biomed. Online 7: 649–656. Wrenzycki, C., Herrmann, D., Keskintepe, L., Martins, Jr. A., Sirisathien, S., Brackett, B. and Niemann, H. (2001a). Effects of culture system and protein supplementation on mRNA expression in pre-implantation bovine embryos. Hum. Reprod. 16: 893–901. Wrenzycki, C., Wells, D., Herrmann, D., Miller, A., Oliver, J., Tervit, R. and Niemann, H. (2001b). Nuclear transfer protocol affects messenger RNA expression patterns in cloned bovine blastocysts. Biol. Reprod. 65: 309–317. Wrenzycki, C., Lucas-Hahn, A., Herrmann, D., Lemme, E., Korsawe, K. and Niemann, H. (2002). In vitro production and nuclear transfer affect dosage compensation of the X-linked gene transcripts G6PD, PGK, and Xist in preimplantation bovine embryos. Biol. Reprod. 66: 127–134. Wrenzycki, C., Herrmann, D., Lucas-Hahn, A., Gebert, C., Korsawe, K., Lemme, E., Carnwath, J.W. and Niemann, H. (2005a). Epigenetic reprogramming throughout preimplantation development and consequences for assisted reproductive technologies. Birth Defects Res. C Embryo Today 75: 1–9. Wrenzycki, C., Herrmann, D., Lucas-Hahn, A., Korsawe, K., Lemme, E. and Niemann, H. (2005b). Messenger RNA expression patterns in bovine embryos derived from in vitro procedures and their implications for development. Reprod. Fert. Develop. 17: 23–35. Xu, Y., Zhang, J.J., Grifo, J.A. and Krey, L.C. (2005) DNA methylation patterns in human tripronucleate zygotes. Mol. Hum. Reprod. 11: 167–171. Xue, F., Tian, X.C., Du, F., Kubota, C., Taneja, M., Dinnyes, A., Dai, Y., Levine, H., Pereira, L.V. and Yang, X. (2002). Aberrant patterns of X chromosome inactivation in bovine clones. Nat. Genet. 31: 216–220. Young, L.E., Fernandes, K., McEvoy, T.G., Butterwith, S.C., Gutierrez, C.G., Carolan, C., Broadbent, P.J., Robinson, J.J., Wilmut, I. and Sinclair, K.D. (2001). Epigenetic change in IGF2R is associated with fetal overgrowth after sheep embryo culture. Nat. Genet. 27: 153–154. Zhang, S., Kubota, C., Yang, L., Zhang, Y., Page, R., O’Neill, M., Yang, X. and Tian, X.C. (2004). Genomic imprinting of H19 in naturally reproduced and cloned cattle. Biol. Reprod. 71: 1540–1544. Zhou, Q., Renard, J.-P., Friec, G., Brochard, V., Beaujean, N., Cherifi, Y., Fraichard, A. and Cozzi, J. (2003). Generation of fertile cloned rats using controlled timing of oocyte activation. Science 302: 1179. Zhu, H., Craig, J.A., Dyce, P.W., Sunnen, N. and Li, J. (2004). Embryos derived from porcine skin-derived stem cells exhibit enhanced preimplantation development. Biol. Reprod. 71: 1890–1897.

10 Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins William Gavin, LiHow Chen, David Melican, Carol Ziomek, Yann Echelard, and Harry Meade

INTRODUCTION Transgenic Production: An Alternative Approach for an Expanding Recombinant Protein Market The use of recombinant proteins as human therapeutic agents has increased dramatically over the last two decades and is still climbing (Fox et al., 2001; Cooke et al., 2004; Schellekens, 2004; Mather et al., 2005). Over the last 10 years, several human plasma-derived therapeutic protein products have been replaced by recombinant versions of these proteins. Additionally, second generation recombinant products that have been engineered for increased efficacy or longer half-life have also made their appearance. However, many clinical applications require large quantities of highly purified biopharmaceuticals that are sometimes administered over long-time periods or through repeat dosing regimes. Therefore, the development of very efficient expression systems is essential to the full exploitation of recombinant technology for production of human therapeutic products. Production of recombinant proteins in the milk of transgenic animals has been under development for many years as an alternative to traditional stainless steel bioreactors for the production of biopharmaceuticals (reviewed in Houdebine, 1994; Clark, 1998; Meade et al., 1998). Recombinant human antithrombin (rhAT, ATryn®) derived from the milk of transgenic goats was the first transgenically produced human therapeutic product to enter clinical trials (Echelard et al., 2005). Recently, the European Medicines Evaluation Agency’s CHMP recommended the approval of ATryn for human use in the European Union (EMEA CHMP, June 1, 2006). This approval and the planned market launch of ATryn® will further validate this transgenic production technology and herald a significant milestone and advancement in recombinant pharmaceutical production. Transgenic Production To express a recombinant protein in the milk of a transgenic animal, the gene encoding the protein of interest is linked to milk-specific regulatory elements to generate the transgene. These DNA constructs must then be introduced into the genome of an animal. The microinjection (MI) of the transgene into the pronuclei of fertilized embryos has historically been the methodology of choice to produce transgenic animals (Hammer et al., 1985; Bondioli et al., 1991; Ebert et al., 1991; Wright et al., 1991). However, with the introduction of the cloned

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sheep “Dolly” (Wilmut et al., 1997), produced by somatic cell nuclear transfer (SCNT), this new methodology was viewed as a potential improvement over MI for large animal transgenic production. Subsequently, many other species, including cows and goats, were cloned from somatic cells with varying degrees of success (Cibelli et al., 1998; Baguisi et al., 1999). Further improvements in nuclear transfer technology have increased the efficiency of the process and lead to the production of genetically characterized and phenotypically selected cloned animals. The widespread adoption of this technology for generating transgenic farm animals is rapidly revolutionizing the transgenics’ field. This chapter will briefly review the transgenic technology platform, discuss the science behind transgenic production, and highlight some of the developments over the years related to nuclear transfer or cloning. Insights will also be provided into some of the challenges that have faced this technology during its development and to the promising future that awaits the widespread acceptance and adoption of this technology.

GENERATION OF TRANSGENIC ANIMALS Pronuclear MI The introduction of transgenes into the germline of large animals has often proven challenging and very labor intensive. Historically, the pronuclear MI approach to gene transfer has been widely used successfully for mice, but proven to be of more limited efficiency with large animal or ruminant species. MI involves the insertion of a fine micropipette into the pronucleus of a fertilized embryo with the injection of a few microliters of the transgene. In some embryos, the transgene will integrate into the host DNA and a transgenic animal will be produced. Transgene integration into the genome of founder animals is typically low and the frequency of generating mosaic animals, containing both transgenic and non-transgenic cells, can be high (Wilkie et al., 1986; Burdon and Wall, 1992; Whitelaw et al., 1993). This has sometimes complicated the expansion of transgenic herds from individual founder transgenic animals (Williams et al., 1998, 2000). Furthermore, transgenic founders can often carry multiple transgene integration sites in their chromosomes, frequently with various degrees of mosaicism, further complicating the genetic makeup and expansion of founder lines (Williams et al., 2001). For the expression of multi-chain proteins such as recombinant antibodies (Pollock et al., 1999), the co-integration of multiple transgenes is necessary. However, some transgenic animals may be generated by the MI approach that only carry one of the two required transgenes. Alternatively, one of the transgenes may integrate on one chromosome, while the other transgene becomes part of a different chromosome. These two independent chromosomes may segregate in the next generation such that some offspring may only express one of the two required antibody chains. These examples illustrate situations that decrease the frequency of “useful” founders (Gavin et al., 1998). Conversely, MI has been used successfully to generate small and large transgenic animals. It does not require extensive upfront manipulations prior to implementation and can be performed as soon as the transgene is assembled. The performance of MI, however, does require some mastery of basic embryological techniques. These include superovulation, embryo retrieval, short-term embryo culture, micromanipulation, and embryo transfer. These basic procedures have been employed over the years to generate transgenic animals and the required MI procedure itself is not overtly technically challenging. The current transgenically produced biotherapeutics that are in the clinic currently were produced by animals that were generated primarily using the MI technique (ATryn in goats, C-1 esterase in rabbits). Additionally, since this technology has been available for more than 25 years, the intellectual property (IP) landscape is established and straightforward.

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SCNT One of the major shortcomings of MI has been the low level of transgene integration into the genome in large animals (1–5%), in particular cattle, sheep, and goats. The use of MI for these species has been further hampered by long generation intervals and the low number of offspring typically generated normally per embryo recipient. The discovery that cultured cell lines can efficiently function as karyoplast donors for nuclear transfer has subsequently expanded the range of possibilities for germline modification in large animals. First sheep (Campbell, K.H.S. et al., 1996; Wilmut et al., 1997), then cattle (Cibelli et al., 1998), goats (Baguisi et al., 1999; Keefer et al., 2001), and pigs (Onishi et al., 2000; Polejaeva et al., 2000; Betthauser et al., 2000) have been generated by cloning, with reports of success in additional species on a regular basis. SCNT has now dramatically increased the efficiency of transgenic animal production to nearly 100% of animals produced. This higher level of production efficiency is solely a function of the ability to pre-select the transfected cell line based on genetic pre-characterization prior to its use in the nuclear transfer procedure. The implementation of SCNT has also had a very significant and positive impact on overall animal utilization. For pharmaceutical production using goats, the use of slaughterhouse-derived oocytes is generally not an option. MI utilized a relatively high number of animals per transgenic founder generated, when considering embryo donors and recipients required. This was due to the large number of offspring that had to be produced as a function of the low transgenic rate in the MI process. With SCNT, the number of donors and recipients utilized has been reduced based on the near 100% transgenic rate seen in the offspring, thereby significantly decreasing overall animal usage on a per founder basis. Compared to the MI process, however, nuclear transfer is more challenging technically. It requires additional laboratory equipment and demands a higher operator skill level. Nonetheless, nuclear transfer with preselected transfected somatic cells allows control over both the sex and chromosomal integration pattern of the transgenic animal produced. It also overcomes the problem of founder mosaicism typically seen with MI. The ability to pre-select transgenic cell lines by analysis of transgene integration sites before the generation of cloned transgenic embryos is extremely valuable and decreases the subsequent elimination of “non-useful” transgenic animals typically generated through the MI process. This pre-characterization is particularly important for the transgenic production in milk of recombinant monoclonal antibodies, where more often than not, several transgenes have to be expressed at similar levels in the same secretory cells of the mammary epithelium. Therefore, co-integration of the transgenes in the same chromosomal locus is paramount to avoid segregation of heavy chain and light chain transgenes during herd propagation.

SCNT: DONOR CELL LINE DEVELOPMENT AND CHARACTERIZATION DNA Construct Development Over the past 15 years, GTC has produced over 100 different proteins in the milk of transgenic animals. The initial expression construct, which was based upon the goat beta-casein promoter (Roberts et al., 1992), has proven to be robust and consistent in expression. It was shown in the early 1990s that this promoter could efficiently express cDNAs (Ebert et al., 1994), unlike other mammary gland promoters that are more selective in their ability to express cDNA versus genomic sequences. The most recent improvement to the promoter construct has been the addition of insulator sequences from the 5 hypersensitive site of chicken beta-globin (Chung et al., 1993). This 2.4 kb DNA fragment was linked to the 5 end of the casein promoter to insure position independent expression of the transgene. Although developed for MI for production of transgenic founders, this promoter construct has allowed the generation of dozens of SCNT founder animals, almost all of which have expressed the desired recombinant product in their milk at significant levels.

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins 171

Figure 10.1 Diagram of the transgene constructs used to establish donor cell lines. (A) Neo resistance gene is linked to the gene of interest in cis ; (B) Neo gene is supplied in trans by co-transfection. Lines represent 5 and 3 regulatory sequences of goat-casein gene; boxes represent coding region of gene of interest or neo; dark boxes represent insulator sequences.

The MI process utilizes DNA from which the prokaryotic vector sequences are removed in preparation of the injection fragment. The same methodology is also used in the preparation of DNA for transfections into cell lines. However, to carry out selection in cell cultures for SCNT, a selectable marker is also required. The traditional marker has been G418/Neo, which is the phosphotransferase isolated from the neomycin (neo) drug resistance TN5 transposon (De Lorenzo et al., 1990). The neo marker has therefore been linked to the beta-casein expression vector to facilitate selection in the primary transfected cell lines. To prevent interference with the beta-casein promoter, the neo resistance expression cassette is flanked with insulator sequences (Figure 10.1). This organization allows each expression sequence to function independently, thus ensuring the desired high-level expression of the beta-casein promoter in the mammary gland. Cell Line Development Timeline From a recombinant protein production point of view, the fundamental difference between the MI and SCNT methods is that the former results in animals of unpredictable genetic composition with regard to the transgene, while the latter produces animals derived from selected single donor cells with a predetermined and homogeneous transgenic genotype. For animals generated using MI, genetic characterization is only possible after the birth of the animals. Many times it is discovered at this later stage of the process that the animals, if transgenic at all, are not suitable for future development due to transgene rearrangement, undesirable copy number, multiple transgene integration sites, mosaicism or unfavorable gender ratio to serve as founders. The SCNT process allows the genetic characterization to be carried out upfront on a large number of transfectants by the combination of PCR, Southern blotting, and fluorescent DNA in situ hybridization (FISH) analyses, thereby ensuring the selection of suitable donor cell lines for transgenic founder production. With this pre-characterization, the uncertainty of the genotypic outcome is therefore removed. However, these selection and characterization steps add significant additional work to the process initially, resulting in a lag time between the DNA construction and the SCNT process. For this cell line development and pre-characterization assessment, up to 3 months can be added to the timeline to generate a large transgenic animal. Fortunately, this is the same amount of time required to confirm the expression of a transgene construct in a transgenic mouse model. Since demonstration of expression in transgenic mouse milk is generally recommended before embarking on a large animal SCNT program, these activities can run concurrently. Therefore, once the DNA construct is completed, it can be microinjected into mouse embryos and at the same time the transgene transfection can be initiated on the cell lines. Results of the mouse study will show whether the transgene is functional. If results from the mouse model are favorable,

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then the characterized cell lines can be used for SCNT, with the assurance that the resulting transgenic animals will produce the desired recombinant proteins. Transfection Transgenes are introduced into primary fibroblast cells isolated from adult skin or fetal tissue of goats or other farm animals by lipid-mediated transfection or by electroporation. Neo resistant colonies are isolated, expanded, and screened by the combination of PCR, Southern, and FISH analyses for the transgenic genotype. Multiple vials of a candidate cell line are cryo-preserved at the earliest possible passages to minimize the time in tissue culture before use as SCNT donors. A separate aliquot of the cells from each candidate line are expanded in culture for characterization purposes. Since the primary cells have limited life span in cell culture, the process of genotyping imposes an additional selection on candidate lines to meet the minimal growth requirement. Fibroblasts or cells of epithelial origin can both serve as transfection recipients and give rise to animals by SCNT that express the recombinant product. However, the shorter doubling time in tissue culture makes fibroblasts the recipient of choice in our laboratory. Neomycin Selection and Stable Donor Cell Lines The neomycin (neo) resistance selection procedure (G418 selection) allows for isolation of stable transfectants and is introduced into cells either linked in cis to genes of interest as one DNA fragment, or it can be introduced in trans by co-transfection. G418 selection is applied 48 h post-transfection and maintained thereafter. It is interesting to observe that in the initial phase of G418 selection, the apparent rate of rearrangement of the neo resistance gene and the gene of interest can differ significantly. This could result in the occurrence of multiple integrations of the neo gene independent of the gene of interest within a single colony, or resistant colonies that carry the neo cassette but not the gene of interest, even if the two were linked in cis in the transgene construct. This makes careful genotyping of each individual candidate donor cell line essential. In our experience, once the stable integration has been established however, the rearrangement of the transgene seems to be greatly reduced either in the presence (donor cells grown in culture) or in absence of G418 selection (the transgenic animals themselves). For transgenic animal lines produced by SCNT to date, we have not observed transgene rearrangement either within the lifespan of a transgenic animal or from one generation to another. Copy Number and Impact on Expression Level With the establishment of position independent expression by using the insulator elements in the transgene constructs, promoter strength and the copy number of the transgene become the dominate factors in determining the level of transgene expression in the transgenic animal. Transgene copy number of every candidate donor cell line is determined by Southern blotting analysis, which also detects gross transgene rearrangements. In the case of monoclonal antibodies, an equal molar ratio of heavy and light chain genes can be assured. For every program, the aim is to select several donor cell lines encompassing a range of different copy numbers for nuclear transfer, in anticipation of other factors that might also influence protein expression. Since very high copy number transgene integration sites increase the likelihood of interrupting normal lactation due to overexpression of the exogenous protein, moderate copy number integrations (10) are preferred. To date, of over 20 different founder transgenic animal lines that we have generated by SCNT, all but one (still currently under investigation) have expressed the transgene product in their milk: expression levels have ranged from 1 to 40 g/l (Table 10.1). FISH Analysis and Integration Site An important consideration for the selection of a donor cell line for SCNT is the number of transgene integration sites. Ideally, the founder(s) should have a single integration site to facilitate herd expansion by eliminating transgene segregation from multiple sites in subsequent generations. This becomes critical when

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins 173

Table 10.1 Project summary of caprine founders produced by SCNT Project (Product)

Founder line

Number of animals

Transgene copy number

Milk expression range (mg/ml)

Malaria (MSP-1)

A B

1 1

2 2

1.0 1.5

Antibody: cancer (CD-137)

A B C

2 1 1

2/1 6/3 6/6

8.0 8.0 10.0

Undisclosed serum protein

A

2

2

0.3–0.4

Antibody: anti-TNF

A B C D E F

3 1 2 1 5 3

1/1 2/1 10/10 5/2 3/4 12/10

1.5–2.0 0.2 14.5–20 10 4.5–7.2 18.0–19.3

Antibody: IL-8

A B C

1 1 1

6/6 2/4 4/6

15 3.7 6.7

Antibody: cancer (SCLC)

A B

1 1

10/20 2/2

NA 20

Antibody: amyloid B-peptide

A B C D E

1 2 1 5 2

60/16 5/3 30/20 60/60 20/2

50.0 0.9–1.0 22.0 42.1 8.6–10.5

multiple transgenes are involved, such as for the production of monoclonal antibodies. Highly sensitive FISH protocols that are able to detect a single copy of the transgene are essential in the selection process to determine the number of integration sites. In addition, FISH analyses are used to detect gross chromosome abnormalities, to verify the homogeneity of the karyotype, and to guard against donor cell lines that have mixed cell populations. Here again, the labor intensive nature of the FISH method adds time and expense to the process of screening hundreds of candidate cell lines. In our experience, only about 2% of the cell lines are selected as donors for SCNT following genetic characterization (Figure 10.2).

CAPRINE SCNT Summary The current technique used for nuclear transfer in the goat has been previously described in detail (Melican et al., 2005) and is depicted in Figure 10.3. Briefly, the process begins by obtaining either in vivo- (Baguisi et al., 1999; Echelard et al., 2004) or in vitro-sourced (Chen et al., 2001; Reggio et al., 2001) unfertilized goat oocytes. The choice of oocyte source must take into consideration oocyte availability, efficiency, cost, and potential regulatory issues (Ziomek, 1996, 1998) if the resulting transgenic goat is to be used for human recombinant therapeutic protein production. The oocytes are enucleated thereby removing their haploid maternal genetic material (Figures 10.4 and 10.5). The desired characterized transfected goat cell or karyoplast is then inserted into the perivitelline space between the egg and its protective outer coat. This is followed in the goat by simultaneously fusion and activation of the enucleated oocyte or cytoplast. The reconstructed embryo is then

174 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

Detection of an integrated transgene using FISH C719

2006

2007

A

B

C

D

E

F

Interphase FISH shows the FITC detected transgene integration site in the donor cell line (A) and in lymphocytes from the two cloned offspring (B,C). Metaphase FISH shows the identical transgene integration on the chromosomes in the donor (D) and offspring (E,F). Nuclei and chromosomes are counterstained with DAPI.

Figure 10.2 Selection of transfected primary cell lines for the generation of transgenic animals by somatic cell nuclear transfer and analysis of offspring.

Goat beta casein Gene DNA of interest 

1 mo.

Target protein expression vector Transfect cells

Select & mate/Al founders

Isolate oocytes & enucleate Transfer reconstructed embryo into recipient female

Select Fuse cell transgenic cell to enucleated oocyte

Hormonally induce lactation

Transgenic milk production herd Measure target protein expression Verify presence of transgene MilkMilkMilkMilk Source material

Figure 10.3 Schematic representation of the process used to generate transgenic goats by somatic cell nuclear transfer. The gene to be expressed is linked to caprine mammary gland-specific regulatory elements. The resulting transgene is then transfected in goat primary cells. Following selection, cell lines are used in the nuclear transfer process using in vivo derived oocyte. Reconstructed couplets are then transferred to the oviducts of recipients does and carried to term. Offspring are tested for the presence of the transgene. The female transgenic founders are induced to lactate to evaluate target protein expression in milk. Selected founders, are mated to non-transgenic males to generate the production herd.

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins 175

Figure 10.4 Enucleation of goat oocytes.

Figure 10.5 Reconstruction: Transfected primary cells are introduced into the perivitelline space of enucleated goat oocytes. Couplets are subsequently submitted to electrofusion and activation.

cultured (Figure 10.6) for a short period of time prior to transfer into a suitable synchronized recipient goat. In other species, such as cattle and pigs, long-term in vitro culture of the reconstructed embryo is generally performed, followed by embryo transfer at the morula or blastocyst stage. Effect of Various Nuclear Transfer Parameters Cell Type The early successes in SCNT were achieved primarily through the use of cultured cells isolated from embryos (Campbell, K.H. et al., 1996). Thereafter, successful cloning was reported with fetal fibroblast and adult mammary

176 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

Figure 10.6 In vitro culture of goat embryos resulting from somatic cell nuclear transfer. epithelial cells (Wilmut et al., 1997). Subsequently, many additional cell lines were shown to be amenable to nuclear transfer with varying degrees of success, including embryonic stem cells (Wakayama et al., 1999), cumulus cells (Forsberg et al., 2002; Chesne et al., 2002), and leukocytes (Galli et al., 2002) to name just a few. In our laboratory, we routinely use both fetal and adult skin fibroblast cell lines (Butler et al., 2003; Behboodi et al., 2004), although the adult skin fibroblast cell lines are easier to obtain. However, when evaluating process efficiency, there is considerable inter-cell line and laboratory variation, making it necessary to evaluate each cell line on a case-by-case basis. Cell Cycle Initially, cell cycle stage at time of enucleation/reconstruction was thought to be of paramount importance to the success of SCNT (Campbell, K.H.S. et al., 1996). Subsequently, successful cloning was reported using cells in many different stages of the cell cycle with differing degrees of efficiency. Dolly, the first cloned sheep in the world, was produced through serum starvation and transfer of quiescent (G0) stage cells (Wilmut et al., 1997), while subsequent cloned animals such as the cow (Cibelli et al., 1998) were produced with actively dividing (G1) cells. In our laboratory, cloned transgenic goats were produced not only using actively dividing cells, but also with simultaneous fusion and activation (Baguisi et al., 1999, Memili et al., 2004). Thereafter, other laboratories have shown that cells of other stages, such as G2/M in the goat (Zou et al., 2002; Zhang et al., 2004) and in the pig (Lai et al., 2001), were of use, albeit with lower overall efficiencies. The method employed to establish a population of cells either at the G0, G1, or other cell cycle stage also has the potential to impact cloning efficiency. For G0 cells, serum starvation over a number of days was the successful procedure (Campbell et al., 1996; Wilmut et al., 1997). However, serum starvation does not have a uniformly beneficial effect on cell populations and may cause apoptosis, which may be detrimental to an individual cells efficiency in the overall cloning process (Yu et al., 2003). Growing the cells to confluence is another method for synchronization of cells into the G0 state (Melican et al., 2005). Our limited data to date on serum exposure (Table 10.2) does not support the proposition that cell line confluency is advantageous in the number of cloned animals born in nuclear transfer experiments. However, one must consider the limited numbers of animals born and its effect on the statistical analysis of this data set. Therefore, additional data is needed to truly determine which cell synchronization protocol, if any, is beneficial to SCNT efficiency. Number of Passages One aspect of the karyoplast population used as a nuclear donor that has not been investigated nearly as much as cell type or cell cycle stage is the cells actual age. This is typically reported as either cell passage number or

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins 177

Table 10.2 Effect of donor karyoplast culture condition on caprine NT efficiencies FBS (%)

# Couplets/# fused (% fusion)

# Cleaved (% cleaved)

# Embryos/ # recipients

# Pregnancies (%) Day 50

0.5 10 a

964/655a (68) 682/411a (60)

298a (45) 238a (58)

587/87 315/51

5a (6) 2a (4)

# Offspring (% embryo)

Term 4a (5) 1a (2)

4a (0.7) 1a (0.3)

Within columns differ significantly, P  0.05.

Table 10.3 Effect of donor karyoplast harvest method on caprine NT efficiencies Trypsinization

Partial Complete a

# Couplets/# fused (% fusion)

1069/726a (68) 577/340a (59)

# Cleaved (% cleaved)

385a (53) 151a (44)

# Embryos/ # recipients

633/96 269/42

# Pregnancies (%) Day 50

Term

6a (6) 1a (2)

5(5) 0

# Offspring (% embryo)

5(0.8) 0

Within columns differ significantly, P  0.05.

sometimes cell doubling number. To date, it has been reported that the age of the donor from which the cells were taken did not impact cloning efficiency when looking at the bovine species comparing fetal fibroblast versus adult fibroblasts (Kasinathan et al., 2001a). However, in another report (Bhuiyan et al., 2004) looking at actual age of the cell line once in culture, the early-passage cell lines were shown to be less efficient than late-passage cell lines when considering nuclear transfer and blastocyst development. In our experience (unpublished data), for non-transfected cell lines, it appears that there is an increased efficiency of live animal produced per nuclear transfer attempt when we have used late-passage cell populations of fetal-derived fibroblasts and, conversely, when we have used early-passage cell populations of adult skin-derived fibroblast cell lines. Cell Isolation: Complete versus Partial Trypsinization, “Shake-Off” Method Isolation of individual adherent cells from culture for use as karyoplasts in reconstruction of enucleated oocytes or cytoplasts has routinely been done using standard trypsinization protocols. However, a nonenzymatic “shake-off ” method of harvesting bovine fetal fibroblast cells for use as nuclear donors was reported (Kasinathan et al., 2001b) and linked to better isolation of G1 cycling cells. Our laboratory has also investigated the best method to isolate an optimal cell population from culture for use in SCNT. Partial trypsinization was used to isolate a minimally adherent population of cells also believed to be a more G1 predominant cell cycle population (summarized in Table 10.3). Although based on early fusion and cleavage information, there was a statistical benefit with partial trypsinization. However, due to the lower than expected overall number of cloned animals produced, this difference could not be confirmed. Therefore, additional work in this area is warranted to definitively determine if a statistically significant benefit could be achieved. Ultraviolet versus Polarized Light Enucleation As part of the enucleation procedure, it is necessary to illuminate the nuclear material of the metaphase II (MII) oocyte, typically referred to as the metaphase plate or spindle. This has traditionally been done by

178 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

Table 10.4 Comparison of enucleation methods Method # Enucleated # Reconstructed # Couplets (% survival) (% survival)

1310b (92) 1243b (92)

UV 1419 Polarized 1348

1223b (93) 1176b (95)

# Fused # Cleaved # Recipients/ # Pregnant (% fusion) (% fused) # transferred recipients 24–48 ha 24–48 h (%) 1029b (79) 384b (37) 960b (77) 334b (35)

101/687 93/646

3 (3.0) 7 (7.5)

Values are totals of 50 experiments. Data were analyzed by Chi-square test. Includes experiments at 24 h post-fusion and activation prior to couplet cleavage. b Within columns differ significantly, P  0.05. a

Table 10.5 Effect of cycloheximide on development to term for SCNT Treatment

Enucleated Reconstructed Fused (%)

Cycloheximide 1474 No cycloheximide 1328

1320 1164

1122b (85) 955b (82)

Cleavage Pregnancies/ Development (24  48 h)a recipients offspring/embryo (%) (%) (%) 440b (39) 261b (27)

18/104b (17) 11/84b (13)

15/741b (2) 12/591b (2)

Values total from 46 experiments. Data was analyzed by the Chi-square test. Cleavage includes development at 24 h (1–2 cell) and 48 h (2–8 cell) post-fusion activation. b Within columns differ significantly, P  0.01. a

ultraviolet (UV) illumination of the MII plate following staining using Hoechst 33342 dye. However, this dye is known to be embryo toxic, permanently binds DNA and, during UV illumination, causes DNA damage. Our laboratory investigated the use of polarized light microscopy to visualize the MII plate (Gavin et al., 2003) thereby eliminating the need for nuclear staining and UV illumination (Table 10.4). Although this data shows that polarized light can be used for enucleation, this work did not show a statistical improvement on the efficiency of clonally produced animals. However, additional data from this laboratory relative to viability of cloned offspring at 6 months of age (unpublished data) supports the increased efficiency using polarized light illumination for enucleation. This data also indicates the possible negative impact of Hoescht 33342 dye on efficiency of SCNT and live cloned animal production. Use of Cycloheximide MII stage oocytes that are typically used in SCNT traditionally have high levels of maturation promoting factor (MPF) to maintain MII stage arrest. Typically, protein synthesis inhibitors are used to downregulate the levels of MPF in SCNT couplets following fusion and activation. Cycloheximide is a broad based protein synthesis inhibitor that blocks the levels of cyclin B, a component of MPF. However, its potential effects on embryo and fetal development were unknown. Therefore, this laboratory investigated the effects of cycloheximide (Table 10.5). We showed that cycloheximide does not have a detrimental impact on embryo or fetal development to term. Therefore, this protein synthesis inhibitor can be added to the list of compounds used for decreasing MPF activity through the SCNT process without any apparent negative impact on future development of the embryo/fetus.

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins 179

Table 10.6 Effect of fusion and activation on caprine NT efficiencies Fusion/ activation

1 2 Re-fused a

# Couplets/ # fused (% fusion)

# Fused/ # cleaved (% cleaved)

# Embryos/ # recipients

1646/720a (44) ND 812/346a (43)

353/112a (32) 364/128a (35) 346/231a (67)

225/35 230/37 447/66

# Pregnancies (%) Day 50

Term

0 1a (3) 6a (9)

0 1a (3) 4a (6)

# Offspring (% embryo)

0 1a (0.4) 4a (0.9)

Within columns differ significantly, P  0.05.

Activation: Calcium Oscillation Activation of the reconstructed couplet is of key importance to the optimization of the SCNT process. An electrical pulse is the most common method used for fusing the membrane of a donor cell/karyoplast to an enucleated oocyte/cytoplast. However, electrical or chemical stimuli can be used for activating couplets produced by SCNT. Additionally, multiple activation events have been suggested to improve efficiencies for generating both parthenogenetic porcine blastocysts and live porcine nuclear transfer offspring (Alberio et al., 2001). One area of investigation is the calcium release patterns in the goat oocyte that occur during normal fertilization and also during activation in SCNT (Jellerette et al., 2006). To try and improve cloned animal production, our laboratory investigated the hypothesis that mimicking the normal oscillatory pattern of calcium release upon normal oocyte fertilization or activation might prove advantageous for the SCNT process. The use of multiple electrical pulses to create multiple calcium spikes was evaluated (Melican et al., 2005) for improvement of SCNT efficiencies (Table 10.6). Unfortunately, there was not a statistical difference between treatment groups. Again, low numbers of cloned animals produced negatively impacted the power of the statistical analysis. However, it is typical in large animal SCNT laboratories that if couplets do not fuse after the first electrical pulse, they are typically subjected to a second electrical pulse (re-fused) as these embryonic materials are too valuable to waste. If one considers that re-fused couplets are exposed to a second electrical pulsation (and hence a second calcium rise) and do give rise to cloned animals, one could argue that multiple electrical pulsations/calcium rises are beneficial to SCNT. Furthermore, additional recent data from this laboratory relative to multiple pulsations for oocyte activation (unpublished) has strengthened our support for this hypothesis. Fusion/Cleavage as a Screening Tool Due to the marked cell line associated variability in SCNT efficiency and the need for case-by-case assessment of each cell line used, parameters were investigated that would indicate whether any given cell line was superior to another. Table 10.7 presents some of our preliminary data assessing fusion and cleavage rate as indicators for cell line efficiency in the SCNT process. It appears feasible and beneficial to screen cell lines in vitro for fusion and cleavage rates to determine which cell line to utilize for SCNT to obtain optimal efficiency rates of live animal production. This work is still ongoing and warrants further investigation based on the preliminary positive results using these parameters as markers of cell line efficiency for production of cloned animals. Embryo Culture The decision to employ either short- (24–48 h) or long-term embryo culture prior to transfer of SCNT embryos into suitable recipients is driven by a number of factors. A principal factor is the availability of wellestablished long-term culture conditions that allow efficient embryonic development. A secondary factor is

180 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

Table 10.7 Summary of SCNT pregnancies by fusion and cleavage

# Recipients # Experiments # Cell lines # Fusion attempted # Fused (%) Fusion range (%) # Cleaved @ 48 h/# fused (%) (Range %) a,b

NT recipients US positive (day 50)

NT recipients US negative

26 17 13 826 686a (83) (57–100) 239/339 (71)a (57–92)

139 35 15 1424 1093b (77) (32–100) 376/721 (52)b (22–93)

Values within rows with different superscripts differ significantly (P  0.001).

Table 10.8 Summary of effect via delivery method Method of birth

# of does birthing

# of kids born

# of kids lost at birth

Natural birth Cesarean section

10 11

11 13

1 1

the availability of non-surgical embryo transfer procedures. Due to technical effort and costs, long-term embryo culture and non-surgical embryo transfer at the blastocyst (D7) stage are preferentially employed in cattle. Bovine long-term culture conditions have been developed over the years that produce viable offspring with acceptable efficiencies upon embryo transfer. In the goat, since non-surgical embryo transfer has not yet been established with high efficiency and reproducibility, a surgical procedure is required. Although efficient culture conditions have been established for the goat embryo, in vivo embryo development is still viewed as optimal in this species and therefore only short-term embryo culture is primarily utilized. In the pig, due to similar circumstances to the goat, short-term culture and surgical transfer of SCNT embryos is the preferred methodology for SCNT. Veterinary Management of Cloned Goats Early in the development of SCNT, it became apparent that there were increased embryo/fetal losses throughout the process. These losses started during early embryo development and continued with higher rates of pregnancies losses in recipient animals during the perinatal and neonatal period. Based on the initial higher rates of loss seen with SCNT, many laboratories moved to cesarean section for delivery of all cloned animals. In our laboratory, we investigated the effects of delivery modality on survival rate of cloned goats (Table 10.8). Our studies did not support the generalized implementation of cesarean sections for increased offspring survival rate. Additionally, it was our belief that a normal delivery through the birth canal resulted in healthier offspring that required less neonatal attention when compared to cesarean delivery. Based on knowledge from the human and veterinary medicine arena, there are physiological developmental events (corticosteroid release and initiation of full respiratory functionality) that are known to occur during the normal delivery process that support this hypothesis. One of the initial reports on the increased losses with SCNT (Hill et al., 1999) highlighted the clinical and pathological abnormalities that were documented in cloned calves. This clinical pathology seen in these cloned animals has been linked to the abnormalities that were also found at the level of placentation where there was

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins 181

Table 10.9 Mouse data from transgenic MSP-1 program Transgene modification

Milk expression level (mg/ml)

Efficacies of purified protein in monkey vaccination study

Secretion variant Modified codon usage for mammalian expression of glycosylated MSP-1 Modified to express non-glycosylated MSP-1

2–4 2–4

NA Protected vaccinated monkey against lethal malaria challenge Protected vaccinated monkey against lethal malaria challenge

1–2

Table 10.10 Goat data from transgenic MSP-1 SCNT program Founder line

# of animals produced

Copy #

Milk expression level (g/l)

1 2

2 2

20 4

1 1

evidence of anatomical abnormalities. This abnormal placental development was also observed in goats produced by SCNT (in house unpublished data). A lower number and abnormal distribution of cotyledons, larger cotyledonary size, and abnormal vascularity within the placenta were observed. Subsequently, numerous reports have documented the abnormal reprogramming occurring at the level of the genome as the possible origin of abnormalities seen at the level of the cloned animals (Dean et al., 2001; Jones et al., 2001; Rideout et al., 2001). However, it was later reported that if one could get SCNT animals beyond the initial period of clinical compromise, there was the possibility for a rather normal subsequent development (Chavatte-Palmer et al., 2002; Pace et al., 2002). Our goat data (Melican et al., 2005; Behboodi et al., 2005: additional unpublished data) also supports this finding of normal healthy cloned animals following passage through a potentially vulnerable neonatal and early developmental period. Furthermore, additional published reports in other species showed that as adults, cloned animals had normal reproductive characteristics (Enright et al., 2002) and also normal milk production capabilities (Walsh et al., 2003). Transgenic Production in Cloned Goats: Two Case Studies To illustrate how a transgenic goat program actually progresses, the following are two examples of projects that were initiated within GTC Biotherapeutics and are still ongoing. The first program has already been partially detailed in Table 10.1 and is aimed at producing a recombinant version of the malaria surface antigen merozoite surface protein-1 (MSP-1) for use in a human therapeutic vaccine. At GTC, a mouse feasibility model is usually produced prior to moving into a larger species such as the goat or cow. For the MSP-1 program, many versions of the transgene were constructed and tested for their expression in mammalian cells and for the secretion of the gene product in the milk of transgenic mice. A total of 36 founder mouse lines (Table 10.9) were produced that expressed the MSP-1 antigen in their milk at a concentration ranging from 2 to 4 g/l. Based on this successful expression of the transgene, a goat SCNT program was initiated. Since only a small volume of MSP-1 antigen would be needed for the world market supply of this product, the goat was elected as the optimal large animal species for the production platform. Table 10.10 summarizes

182 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

the successful goat SCNT program. Whereas in the mouse model, only very small quantities of milk can be generated from a natural lactation, the goat can produce significant quantities of milk by hormonal induction of lactation at a pre-pubertal age prior to breeding. This milk is used for analytical purposes and for making decisions on founder lines and breeding scenarios. Founder line #2 (Table 10.10) was chosen for further development in the MSP-1 program and both females were bred and brought into a natural lactation. The average expression level was 1 g of MSP-1 antigen per liter of milk, and the average milk yield was 3 liters per day per doe. With these yields, it is calculated that a single goat would supply enough antigen to vaccinate several million people annually. The second case study involves a recombinant monoclonal antibody with a therapeutic anti-cancer application. The antigen recognized by this antibody is CD137, also known as 4-1BB, a member of the tumor necrosis factor/nerve growth factor family of receptors and a surface glycoprotein found on certain cells of the immune system. This agonistic antibody binds to and stimulates CD137 resulting in strengthening of an otherwise traditionally weak immune response to tumors. Utilizing the mouse model as a feasibility tool, three separate founder lines were produced and analyzed. The first line did not express any detectable levels of the recombinant antibody in the milk but the remaining two lines expressed at very high levels of 10 and 15 mg/ml, respectively. A successful goat founder program was carried out shortly following the mouse effort and Table 10.11 details that data generated. The transgenic goats generated in this program were hormonally induced to lactate and were found to express the agonistic CD137 antibody at levels in excess of 5 g/l. Recently, one of the transgenic does gave birth and entered into a natural lactation. The average yield of monoclonal antibody was 6 g/l and the daily volume of milk produced was 1.5 g/l. The above two case reports represent examples of two successful transgenic founder goat programs in development at GTC. While the malaria program may not appear to be as robust as one would prefer, there were challenges in expressing this antigen in the milk of the mammary system, as well as any other recombinant expression system that was investigated. Additionally, this level of production is more than suitable for a large market that only requires a small amount of recombinant material to produce adequate quantities of a vaccine. As for the CD137 program, this is somewhat at the other end of the scale from the point of view of expression level coupled with lactational volume. This founder goat program is more in line with a high volume recombinant protein need where expression level and good lactational output are critical to the success of the program. Transgenic Production in Cloned Cattle: A Case Study of Cows Expressing Human Albumin Human serum albumin (hSA), the most abundant protein in human plasma, is one the first human blood protein that has been mass produced by plasma fractionation. It was initially used during World War II as a blood replacement product (Finlayson, 1980; Peters, 1996). Currently therapeutic uses of hSA with critically ill patients cover numerous acute and chronic conditions (Hennessen, 1980; Alexander et al., 1982; Erstad et al., 1991; Wilkes

Table 10.11 Goat data from transgenic CD137 SCNT program Founder line

# of animals produced

Copy #

Milk expression level (g/l)

1 2 3

2 1 1

2–3 5–6 6, 12*

5 6–8 6–8

* Copies of heavy chain and light chain respectively for the antibody construct.

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins 183

and Navickis, 2001). In addition hSA is used as a stabilizer for drugs and vaccines, for the coating of devices, in imaging, and as an ingredient of cell culture media (Peters, 1996). Although purified hSA is generally considered to be safe, the supply of plasma itself is threatened by known and emerging viral infections, as well as by prion diseases. As with other therapeutic proteins traditionally derived from plasma fractionation, this concern has motivated the search for a recombinant version of albumin. However, the technical challenges related to the development of a recombinant human albumin (rhA) are daunting. Therapeutically, hSA is used in large amounts (10 s of grams per dose) and the end-user cost is low ($1.00–4.00 per gram). In addition the high dose used in treatment requires that the levels of contaminating host proteins must be extremely low. In an effort to provide an abundant source of rhA, a herd of transgenic cows expressing high levels of hA in their milk has been developed. The first step in the generation of transgenic cows expressing high levels of hA was to determine which combination of regulatory elements and hA sequences would be most reliably expressed in the lactating mammary gland. Transgenes linking human albumin DNA sequences to the regulatory sequence of milk-specific genes were first tested in transgenic mice. A construct that contained goat beta-casein upstream of a non-coding sequence linked to the 17 kb DNA fragment containing all of the exons and introns of the hA gene was shown to consistently direct high-level hA expression to the lactating mammary gland (Behboodi et al., 2001). Furthermore, addition of the chicken globin insulator element appeared to increase the frequency of high-level expression from line to line and also served to isolate the beta-casein – albumin transcriptional unit from potential interference with the promoter of the neomycin resistance cassette. This construct was then transfected into bovine primary fetal fibroblasts. Thirty-four neomycin-resistant isolates were screened by Southern blotting with a radiolabeled hA-specific probe. Four cell lines were then selected and employed in an intensive nuclear transfer program (Table 10.12). Twenty calves were generated from this effort. A combination of PCR, Southern blotting, and fluorescence in situ hybridization (FISH), as described above, was used to characterize the transgenic integration for each cloned calf. Of the 20 surviving calves, only 16 were transgenic. For cell lines B–D, all resulting 14 offspring were transgenic as expected. However, for cell line A, only two out of six resulting calves were transgenic, even though they all exhibited the same phenotypic appearance. It appears that these non-transgenic offspring were a consequence of a mixed cell population within the donor cell line. A posteriori FISH showed that a high proportion of nuclei from cell line A scored negatively as compared to the B–D isolates. Apart from the negative animals derived from line A, analyses confirmed that animals derived from cell lines 57, 59 and 60 each carried identical transgene integrations within their line. All transgenic cows were bred and rhA expression in milk was evaluated (Table 10.13). The rhA expression was roughly proportional to the transgene copy number. This is evidenced by line A producing more than 40 g of rhA per liter of milk. Unfortunately these high levels of expression were incompatible with normal lactations. Lines B and C were found to express

Table 10.12 Summary of bovine nuclear transfers for the recombinant albumin program Cell lines

NT* attempts Blastocysts Transfers (# blastocysts transferred) Calves (%NT) * NT: nuclear transfer.

53

57

59

60

Totals

2879 87 32 (79) 6 (0.2%)

3654 154 47 (98) 9 (0.2%)

4696 267 89 (76) 1 (0.02%)

7921 278 89 (193) 4 (0.05%)

19,150 786 257 (546) 20 (0.1%)

184 BIOLOGIC AND MOLECULAR BASIS OF REGENERATIVE MEDICINE

Table 10.13 Albumin expression in milk of four transgenic cattle lines Cell line

Transgene copies

hA expression (g/l)

A B C D

250–300 4–5 4–5 1–2

40 (short lactations) 1.5–2 (normal lactations) 2–2.5 (normal lactations) 1 (normal lactations)

approximately 2 g of rhA per liter of milk. Line B was expanded by in vitro fertilization (IVF) and artificial insemination and is being further developed.

CONCLUSION Overall, the application of SCNT to the generation of large dairy animals has benefited the field of transgenic production. When compared to pronuclear microinjection, SCNT has generally increased the number of transgenic animals obtained for any given effort by significantly increasing the efficiency of producing founders. More importantly, it has improved the predictability of the process of introducing transgenes into the germline of large animals. In addition, the ability to initially generate several identical transgenic females can accelerate the production of large quantities of drug substance in the milk for pre-clinical and clinical studies. Furthermore, the pre-characterization of the transfected cell line increases the likelihood that the transgenic animals will be useful and commercially viable. For these reasons, SCNT has become the method of choice for the production of transgenic ruminants in the biopharmaceutical production arena. However, the use of SCNT is not without drawbacks. The generation of well-characterized cell lines to be used as donors for the nuclear transfer process is time consuming. Often several hundreds, sometimes thousands, of clonal cell lines have to be expanded and genotyped to obtain a dozen candidate cell lines for use in SCNT. In our experience, it is necessary to perform Southern blotting to look for transgene copy number and possible rearrangement events, and FISH to eliminate multiple integration site cell lines. The necessary use of primary cells also complicates the selection process, since the majority of cell lines over time will become senescent and will not be usable as donor karyoplasts. Pronuclear MI is much less resource-intensive, since once a DNA construct is obtained in can be immediately microinjected. Although it requires less time to begin generating the transgenic founders, it is only after they are born that the screening process can begin. The SCNT allows one to more closely monitor the success of the program, since the pregnancy state of the animals can be determined after 40 days. Even with a 50% rate of live progeny generated, they are all expected to be transgenic and genetically characterized. IP is an additional source of concern. Patents covering the MI technique have either expired or will soon expire. However, the SCNT patent landscape is still fragmented and unsettled. This situation requires the evaluation of these patents through costly “freedom to practice” opinions. Complex licensing strategies may be necessary to move a commercial program forward. These agreements can add a significant royalty burden to a therapeutic product obtained from transgenic animals derived from lines whose founder was generated by SCNT. So the cost of the improvement in generating founders using SCNT must be factored into each program. Of course, this situation is not unique to SCNT and is also observed with recombinant therapeutic proteins obtained from traditional large-scale cell culture. A source of anxiety relative to the use of SCNT for transgenic production is the ongoing debate relative to the health of cloned animals. It appears that the health and reproductive fitness of SCNT cows and goats is very satisfactory and that these animals can effectively be used in transgenic production. However, the

Transgenic Cloned Goats and Cows for the Production of Therapeutic Proteins 185

pregnancy or yield of offspring from embryo transfer is still low with SCNT and has not significantly improved in almost 10 years. It is postulated that this is due to improper reprogramming of the SCNT embryos and as yet there has been no technique developed to improve this process. The current yield of transgenic founders by SCNT appears to be sufficient for pharmaceutical development. In this case, the cost of producing transgenic animals is a small fraction of the total cost of developing a drug candidate. This might not be the case for business concerns that aim to use SCNT animals for agricultural purposes. Further research in improving the efficiency of SCNT, although laudable, is not necessarily a priority for those who plan to produce recombinant proteins in the milk of transgenic ruminants. In summary, SCNT has become the method of choice for the generation of large animal founders for transgenic production. The main advantage over pronuclear MI is improved predictability and overall efficiency. The drawbacks of SCNT are an initial increase in demand of laboratory resources and an uncertain IP landscape. However, the bar is now very high for a new transgenic method to produce founder animals. It will probably need to improve the yield of transgenic founders while still preserving the ability to pre-select genetic characteristics. On the other hand, one could improve yield of founders to such an extent that predictability will be irrelevant. Approaches aimed at culturing germ cells for eventual laboratory manipulations and transplantation to a recipient male (Brinster, 2002; Dobrinski, 2005) might one day provide a solution and warrant continued interest and monitoring of future achievements. Currently, however, the SCNT can be used to reproducibly generate the transgenic founders required for recombinant protein production.

ACKNOWLEDGMENTS The authors would like to thank the GTC Farm Operations, Veterinary Services, and Molecular Biology staff for their efforts in handling and analysis of the animals used in this work.

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Part III Cells and Tissue Development

11 Genetic Approaches in Human Embryonic Stem Cells and Their Derivatives Junfeng Ji, Bonan Zhong, and Mickie Bhatia

INTRODUCTION Human embryonic stem cells (hESCs) were first derived from the inner cell mass of blasto cyststage embryos in 1998 (Thomson et al., 1998). Isolation of hESCs opened up exciting new opportunities to study human development which is inaccessible in vivo and develop cell replacement approaches to the treatment for a broad range of diseases based on two unique properties: (1) self-renewal capacity; hESCs are able to proliferate for extended periods of time while maintaining their undifferentiated state and normal karyotypes in the proper culture conditions in vitro and (2) broad developmental potential; hESCs are pluripotent cells which can give rise to cell types representing ectodermal, mesodermal, and endodermal germ layers as assessed by in vitro embryonic bodies (EBs) formation and in vivo teratoma assay (Itskovitz-Eldor et al., 2000; Schuldiner et al., 2000; Dvash et al., 2004). Despite the promising prospect of hESCs as an invaluable system to model human development in vitro and as an unlimited source of cells for transplantation for a broad spectrum of human disease, the emerging hESCs field is still in infancy and fundamental questions regarding the biology of hESCs remain to be addressed. Optimization of culture conditions to maintain hESCs in the undifferentiated state for a prolonged time in vitro is the first crucial step prior to any means to explore the therapeutic potential of hESCs, success of which requires a thorough understanding of molecular pathways regulating the self-renewal, pluripotency, apoptosis, and differentiation of hESCs. Moreover, only upon elucidation of cellular and molecular events dictating lineage specification and commitment of hESCs that faithfully recapitulate early human development will it be feasible to develop protocols to efficiently differentiate hESCs into diverse cell lineages potentially used for transplantation in the clinic. Genetic approaches to manipulating mouse embryonic stem cells (mESCs) in studies during the past 20 years have provided invaluable insights into the understanding of molecular signals governing pluripotency and specification of mESCs (Boiani and Scholer, 2005). To date, there is mounting evidence demonstrating that genetic manipulations such as homologous recombination, RNA interference (RNAi), overexpression of genes by transient transfection and stable viral infection are applicable to hESCs and their derivatives, which will allow us to investigate genetic program regulating pluripotency maintenance versus differentiation of hESCs into diverse lineages (Gropp et al., 2003; Zwaka and Thomson, 2003; Menendez et al., 2004; Zaehres et al., 2005). In this chapter, we will review current

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protocols to maintain hESCs, genetic approaches to modifying undifferentiated hESCs, differentiation of hESCs into multiple lineages and transplantation of their derivatives, genetic manipulation of hESCs-derived progenies, and discuss the potential applications of genetic modifications of hESCs and their derivatives in the context of regenerative medicine.

MAINTAINING UNDIFFERENTIATED HESCS hESCs were originally established and maintained by co-culture with mouse embryonic fibroblast (MEF) feeder layer (Thomson et al., 1998). In an attempt to free hESCs from animal feeder layer, researchers have successfully used human feeder cells to derive and grow hESCs (Richards et al., 2002). Xu and colleagues went one step further to show that hESCs can be maintained in feeder-free condition where hESCs are cultured on Matrigel, laminin, or fibronectin in media conditioned by MEFs (Xu et al., 2001). However, culturing hESCs on either feeder cells or in conditioned media from supportive feeder cells adds additional difficulties to hESCs maintenance and propagation, because preparing feeder layer or feeder layer-conditioned media is time consuming in that feeder cells like MEFs undergo senescence after approximately five passages and different batches vary significantly in their ability to support hESCs growth. Moreover, presence of xenogeneic components derived from MEFs or their conditioned media in hESCs culture harbors a potential risk for transmission of animal pathogens into human if cells derived in such conditions are used for cell replacement therapies in the clinic. Recently, four groups have made significant progress in eliminating animal product from hESCs culture (Amit et al., 2004; Wang et al., 2005a; Xu, C. et al., 2005; Xu, R.H. et al., 2005). Amit et al. reported a feeder layerfree system where hESCs were cultured on fibronectin-coated plate in media supplemented with 15% serum replacement (SR), a combination of growth factors including basic fibroblast growth factor (bFGF), leukemia inhibitory factor (LIF), and transforming growth factor beta 1 (TGF-β1) (Amit et al., 2004). Xu and colleagues have successfully sustained undifferentiated proliferation of hESCs on Matrigel in unconditioned media supplemented with 20% SR plus high dose of bFGF (40 ng/ml) and bone morphogenetic protein (BMP) antagonist noggin (Xu, R.H. et al., 2005). Similarly, Wang et al. have been able to maintain hESCs by culturing them on Matrigel in media supplemented with 20% SR and high dose of bFGF (36 ng/ml) alone (Wang et al., 2005a). Finally, Xu et al. demonstrated that Matrigel and SR supplemented with bFGF alone or in combination with other factors such as stem cell factor (SCF) or fetal liver tyrosine kinase 3 ligand (Flt3L) were able to maintain the growth of hESCs. Although all the above groups used SR and/or Matrigel to substitute for MEFs or their conditioned media to support hESCs, both SR and Matrigel are undefined and still contain animal-derived product. Subsequent to the reports, two groups have further demonstrated the successful derivation and growth of hESCs in defined culture conditions that are solely consist of human materials (Lu et al., 2006; Ludwig et al., 2006). Ludwig and colleagues reported the generation of two new hESC lines in TeSR1 media that is composed of DMEM/F12 base supplemented with human serum albumin, vitamins, antioxidants, trace minerals, specific lipids, and growth factors of human origin including bFGF, LiCl, gamma-aminobutyric acid (GABA), pipecolic acid, and TGF-β (Ludwig et al., 2006). Derivation of hESC lines in TeSR1 also requires a combination of collagen, fibronectin, laminin, and vibronectin as supporting matrices, pH (7.2), osmolarity (350 nanoosmoles), and gas atmosphere (10% CO2/5% O2). Lu et al. developed a less complex hESC cocktail (hESCO) containing bFGF, Wnt3a, a proliferation-inducing ligand (April), B cell-activating factor belonging to TNF (BAFF), albumin, cholesterol, insulin, and transferin to support the self-renewal of hESCs (Lu et al., 2006). However, both of the two studies used incompletely defined albumin derived from human sources in their culture conditions, which may introduce human pathogens into the hESC culture to comprise their potential application in the clinic. In addition, one new hESC line derived in TeSR1 media, although originally normal, developed genetic abnormality as previously observed (Draper et al., 2004) after a relatively long-term culture in vitro (Ludwig et al., 2006).

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Therefore, other than the requirement to eliminate feeder cells, animal product, and undefined components from hESCs culture, an optimal culture condition for the growth of hESC must be able to prevent spontaneous differentiation and maintain genomic stability in the long-term culture. Maintained in the existing conditions, hESC culture consists of morphologically heterogenous populations of cells in which a subset of fibroblast-like cells that are spontaneously differentiated from hESCs usually surrounds colonies. Although hESC-derived fibroblast-like cells have been used as a feeder layer to support the growth of hESCs (Yoo et al., 2005), its cellular and molecular identity and heterogeneity as to the proliferation propensity and developmental potential between individual colonies within hESC culture remain to be determined. Furthermore, during long-term hESC culture in suboptimal conditions, hESCs have been shown to progressively adapt to the culture and select for clones with alterations in survival and proliferation capacity (Enver et al., 2005). Maitra et al. reported that eight of nine late-passage hESC lines acquired genetic and epigenetic abnormalities implicated in human cancer development (Maitra et al., 2005). In an attempt to develop measures to ensure the genetic normality of hESCs, a recent study has established differential expression of CD30, a member of the tumor necrosis factor receptor superfamily, in transformed versus normal hESC lines, implying that CD30 may serve as a biomarker for transformed hESCs (Herszfeld et al., 2006). However, examination of CD30 expression must be extended to a larger array of normal hESC lines and their variants with subtle genetic alterations. Determining the cellular and molecular bases of heterogeneity and transformation due to spontaneous differentiation and adaptation is important for devising improved culture conditions that minimize the selective advantage of variant cells and therefore help maintain genetically normal cells suitable for therapeutic applications. Molecular dissection of signals dictating pluripotency and specification of hESCs by means of genetic manipulation will facilitate the optimization of culture conditions to maintain and specify hESCs.

GENETIC APPROACHES TO MANIPULATING HESCS Gene Regulation Knock-In/Knock-Out Traditionally, knock-in/knock-out technologies based on homologous recombination are the most widely used methods to study gene function in most organisms. Homologous recombination in hESCs is important for modifying specific hESC-derived tissues for therapeutic applications in transplantation medicine. In vitro studies of hESCs involved in understanding the pathogenesis of gene disorder diseases such as Wiskott–Aldrich syndrome or cancer also need the loss-and-gain methods. Although homologous recombination was efficient in generating mESCs mutant and knock-out mice (Joyner, 2000), it is difficult to be applied to hESCs. Firstly, comparing to their murine counterparts, hESCs cannot be cloned efficiently from single cells, making it difficult to screen for rare recombination events. Secondly, since the size of hESCs (14 μm) is larger than mESCs (8 μm), the transfection strategies between human and mESCs are different. Based on an electroporation method, the first homologous recombination in hESCs succeeded in generating the hypoxanthine phosphoribosyltranferase-1 (HPRT-1) knock-out mutant and the oct-4 knock-in mutant (Zwaka and Thomson, 2003). The transfection rate was 5.6  105 and the frequency of homologous recombination itself in hESCs was comparable to that in mESCs (2–40% and 2.7–86%, respectively) (Mountford et al., 1994). Knock-Down In 1998, the same year that hESCs were derived, RNAi was discovered in Caenorhabditis elegans and gained intense investigations till now (Fire et al., 1998). The first application of RNAi in hESC was achieved in hESCs 6 years later, oct-4, the important gene keeping hESCs in undifferentiated state was efficiently knocked down

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(Hay et al., 2004; Matin et al., 2004; Zaehres et al., 2005). RNAi is a mechanism of post-transcription silencing which degrades mRNA transcripts through homologous short RNA species in two steps: (1) double-stranded RNAs (dsRNA) larger than 30 bp are recognized by the highly conserved RNAse III nuclease, named Dicer, and cleaved into 21–24 nucleotides small interfering RNAs (siRNA) and (2) siRNAs are recruited into “RNAinduced silencing complex (RISC),” which is a multi-protein complex (with endogenous RNase activity) that induces endonucleolytic cleavage of the target mRNA (recognized by hybridization with the RISC-bound siRNA antisense strand). While lower degree of sequence complementary to the target mRNA only leads the RISC to interfere the translational machinery, leaving mRNA intact. Previous studies have found that in mammalian cells, dsRNAs larger than 30 bp (usually ranging from 500 to 1,000 bp) can trigger an interferon response by activating the dsRNA-dependent kinase (PKR), resulting in a non-specific global inhibition of protein translation and mRNA/rRNA hydrolysis (Kumar and Carmichael, 1998). Synthetic siRNA or short hairpin RNA (shRNA) can be exogenously delivered into cells to induce RNAi of target genes specifically without the activation of interferon response, which made RNAi applicable to manipulate genes in the hESCs study (Amarzguioui et al., 2005). The screening of RNAi libraries is very useful to identify novel gene functions, especially in the study of hESC differentiation. Libraries of synthetic shRNAs or siRNAs against a specified gene have been reported (Berns et al., 2004). One or multiple siRNAs could be delivered into the target cells with various transfection methods to increase the chance of successful repression. Because of the transient transfection, this method does not offer long-term stability. However, it also reduces the chances which potential inhibition of unknown genes may occur in the long-term assay. In order to achieve long-term therapeutic aims, more stable knock-down is required; for instance, expression of miRNAs, siRNAs, or shRNA in vivo may rely on the chromosomal integration of viral vectors containing homologous and complementary DNA sequence under the control of RNA pol II or RNA pol III promoters (Denti et al., 2004; Stegmeier et al., 2005). Among all these promoters, H1 and U6 promoters were mostly widely used to drive shRNAs/siRNAs (Tiscornia et al., 2003; Kaeser et al., 2004; Schomber et al., 2004; Zaehres et al., 2005), while vectors containing U6 promoter gave a higher frequency of interferon response induction than comparable H1 promoter-containing vectors. To avoid interferon induction by U6 promoter-driven vectors, Pebernard and colleagues recommended preserving the wild-type sequence around the transcription start site, in particular a C/G sequence at positions 1/1 (Pebernard and Iggo, 2004). In addition, the promoters could be constructed under various chemical-regulated transcription systems to achieve inducible expression, which offers the option of inhibiting gene expression at certain steps during hESCs differentiation (Wiznerowicz and Trono, 2003; Gupta et al., 2004; Higuchi et al., 2004; Tiscornia et al., 2004; Szulc et al., 2006). Because the sequence of shRNA plays a critical role in the efficiency of gene knock-down, several factors need to be considered during the designing of shRNA: the length of sequences should be within19–23 bp; avoiding the first 75–100 nucleotides (possible protein binding site) of target mRNA; G/C component within 30–50%; low internal stability at 5 antisense sequence; high internal stability at 5 sense sequence; absence of internal repeats or palindromes; 3 end of sense and antisense sequences should have 2 “U”; preference of G/C, A, U, A at the first, the third, the tenth, and the nineteenth nucleotide positions respectively in the sense sequence; avoiding the appearance of G/C, G at the nineteenth and thirteenth positions (Bantounas et al., 2004; Gilmore et al., 2004; Pebernard and Iggo, 2004). The most popularly used sequence, CAAGAGA, was designed as the nucleotides loop to link the sense and antisense sequence (Brummelkamp et al., 2002; Anderson et al., 2003; Kunath et al., 2003). Alternatively, because the nucleotide size of shRNA or siRNA is usually very small and hard to be examined during siRNA vector cloning based on enzyme digestion, using a restriction enzyme recognizing sequence as the nucleotide loop will be more convenient to screen the positive shRNA-inserted clones. There are several advantages of RNAi over “antisense oligonucleotides” and “knock-out” strategies. To silence the same gene, siRNA strategy was much more efficient than antisense oligos, as well as higher stability and less toxic side effects (Miyagishi et al., 2003). Compared to the time- and cost-consuming

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“knock-out” strategy, RNAi can achieve a gene knock-down in hESCs within several months. By lowering the expression level of one gene instead of completely eliciting it, RNAi allows a molecular “turning dial.” However, knock-down based on RNAi can not simply replace traditional knock-out techniques, but works as a relatively complementary tool. Recently, Persengiev and colleagues detected changes of 1,000 genes during introduction of siRNA against a non-existing gene (Persengiev et al., 2004). This indicates that off-target effects versus target specific effects issue need to be pay more attention in future studies. Transfection Chemical Transfection Synthetic chemicals such as cationic lipids have been extensively used for the DNA delivery into hESCs. It is based on the neutralization of cationic lipids to negatively charged DNA followed by the formation of DNA/lipid complexes, which possess an excess of positive charges. These complexes bind to the negatively charged membranes of hESCs and are subsequently taken by the cells through endocytosis. Although various optimizations have been compared by combining different chemicals, such as ExGen500 (Fermentas) (Eiges et al., 2001; Matin et al., 2004), calcium phosphate (Darr et al., 2006), FuGENE (Boehringer Mannheim) (Liu et al., 2004), LipofectAMINE Plus (Life Technologies) (Vallier et al. 2004), or Lipofectamine 2000 (Invitrogen) (Hay et al., 2004), with different media, concentrations of DNAs and cells, the transfection efficiency was still not promising. The inefficient performance of chemical methods may be due to cell cycle phase, the degradation of DNA caused by phagocytosis, or other unidentified factors. Physical Transfection Oligos delivery through electroporation is based on transient permeabilization of cell membrane via reversible formation of pores. Electrophoretic and electro-osmotic forces drive DNA through the destabilized cell membrane. Pre-stimulation on target cells by cytokines have controversial transfection results from different groups (Wu et al., 2001; Weissinger et al., 2003). This may be due to the different use of plasmids and various electroporation conditions. Lots of evidences indicated that electroporation is an efficient gene delivery method in mESCs, hematopoietic stem cells (HSCs), and hESCs (Kunath et al., 2003; Oliveira and Goodell, 2003; Fathi et al., 2006), and CD34 HSCs were relatively tolerant to electric forces and exhibited a higher cell survival rate after transfection compared to other primary cells. The death of the electroporated cells is proposed caused by colloidal-osmotic swelling of cells as well as the uptake of exogenous DNA, which triggers the apoptosis. Optimized protocol showed obvious greater post-electroporation viability when the hESCs were electroporated in clumps and plated out at high densities in isotonic, protein rich medium instead of phosphate-buffered saline (PBS) (Zwaka and Thomson, 2003). More recent emergency of “nucleofection” yielded acceptable cell survival rates (70%), and 66% of the surviving cells showed transgene expression 24 h after nucleofection (Siemen et al., 2005; Levetzow et al., 2006). As the oilgo is delivered into the nucleus, the transfection rate is comparable to those of retroviral systems. Thus, this method holds a promising wider application in the near future. Some other methods such as molecular vibration-mediated transfection and microinjection had high gene transfer rate (upto 100%), these one-step efficient procedures attracted more attention in the stem cells research (Capecchi, 1980; Wakayama et al., 2001; Song et al., 2004). Overall, physical methods of transfection are more efficient methods for plasmid DNA delivery, and are free from biocontamination as well as less concerns about immune reaction. It has low cost, ease of handling and is highly reproducible, the most importantly, biosafety. However, transient transgene expression in hESCs colonies is difficult to retain for longer than five passages (Vallier et al., 2004). To achieve long-term transgene expression, especially in the fast replicating cells, viral vector delivery may be needed.

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Viral Transduction Retroviral Vector In the past two decades, retroviral vectors have been used for stable gene transfer into mammalian cells (Cone and Mulligan, 1984). The first vectors studied in a clinical trial (adenosine deaminase deficiency) were also retroviral vectors (Anderson, 1990). In 2000, the first successful treatment of a genetic disease was relied on retroviral vectors, demonstrating the concept of gene therapy (Cavazzana-Calvo et al., 2000). The most popularly used retroviral vectors were those derived from the Moloney murine leukemia virus, which was also widely reported in the transduction of HSCs for gene therapy. Relative simplicity of their genomes, ease and safe of use and the ability of integrating into the cell genome resulting in long-term transgene expression render them ideal vectors for a genetic alteration. Upon this, stem cells in general, especially HSCs, constitute the best targets for retroviral vector-mediated gene transfer. Transgenes could be long-term expressed in vivo and may give rise to a large progeny of gene-modified mature cells during the continuous amplification process. Retroviral vectors are derived from retroviruses. This family consists of seven genera: alpharetrovirus, betaretrovirus, gammaretrovirus, deltaretrovirus, epsilonretrovirus, lentivirus, and spumavirus. The first five genera were previously classified as oncoretrovirus. Strictly speaking, vectors based on lentivirus or spumavirus are also retroviral vectors. However, the name retroviral vector is often used to refer to vectors based on murine leukemia virus or other oncoretrovirus. All retroviruses share some common features: lipidenveloped particles containing two identical copies of liner single-stranded RNA; depending on specific cell membrane receptor for viral entry; the RNA is reverse transcribed and integrates randomly into the target cell genome upon infection. All retroviral vectors contain long terminal repeats at the 5 and 3 ends (5LTR and 3LTR), a packaging signal located 3of the 5LTR(ψ), and the three groups of structural genes, gag, pol, and env, coding for the capsid proteins, reverse transcriptase and integrase, and envelop proteins, respectively. For the production of retroviral vectors, the complete coding region for the pol and env genes, and the majority coding region of the gag are removed leaving a backbone of the 5 and 3 LTRs, part of the gag coding region and the packaging signal (ψ). The transgene is constructed between the LTRs, and the resulting RNA transcript can be packaged into a virus with co-transfection of other separate packaging vectors (coding gag/pol, env proteins) within a cell. Some features of retrovirus have been problematic in the retroviral vector designing. First, cells not expressing the appropriate receptor are resistant to certain retroviruses, which limits the application of retroviral vectors for host transduction. To obtain a broad host range, retroviral vectors have been pseudotyped with amphoteric envelope, gibbon ape leukemia virus (GALV) envelope (transduction in hESC-derived CD45negPFV hemogenic precursors) or vesicular stomatitis virus glycoprotein (VSV-G) by which retroviruses were able to be transduced into even non-mammalian cells derived from fish, Xenopus, mosquito, and Lepidoptera (Burns et al., 1993; Menendez et al., 2004). VSV-G envelope is also useful to stabilize retroviruses during viral particles concentration by ultracentrifugation. However, the expression of the VSV-G is toxic to cells, resulting in only transient production of vectors in producer cell line. Therefore, conditional expression system of VSV-G in retroviral vector has been developed (Yang et al., 1995). Second, the nuclear membrane is a physical barrier for most retroviruses to migrate their transcribed dsDNA into the cell nucleus. Therefore, targets of most retroviral vectors, such as those based on murine leukemia virus, are limited to actively dividing cells (Miller et al., 1990). To disrupt the nuclear membrane, addition of a variety of stimulatory cytokines to introduce cycling in the HSCs population is usually applied before retrovirus infection. Third, retroviral regulatory elements are repressed in ESCs and HSCs, and this makes long-term expression mediated by integrated retroviral vector difficult to achieve. Short-term silencing of recombinant gene is due to the binding of trans-acting transcriptional repressor on specific region within the promoter of retroviral vector (Gautsch, 1980). Modification of the sequences in LTR to decrease the affinity of negative regulators has been applied to

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solve this problem (Laker et al., 1998). By engineering the regulatory regions, generation of noval retroviral vectors was reported, such as Friend mink cell focus-forming virus/murine ES cell virus hybrid vectors (FMEV), and higher expression levels of transgene than conventional retroviral vectors were observed in HSCs (Baum et al., 1995). In contrast, long-term silencing of target gene is often observed in retroviral vectors based on murine stem cell virus. Because of the high cis-acting methylation activity of ES cells, effective DNA methylation leads to the silencing on integrated retroviral vectors, while this was not detected within differentiated cells showing low methylation activity. Alteration of the cis elements in LTR could decrease the DNA methylation and increase transgene expression in embryonic carcinoma cells (Challita et al., 1995). From the cell aspect, disruption of the methyltransferase gene Dmnt1 to alter the endogenous level of DNA methylation in target ESCs may lead to another potential solution. Due to the multiple defects of retroviral vectors, lentivirus-based vectors are more attractive in the genetic research of hESCs. Lentiviral Vector Lentivirus is one genus of retrovirus including the human immunodeficiency virus (HIV) type 1. Principally, lentiviral vectors are derived from lentiviruses in a similar way as retroviral vectors. Some features of lentiviruses make lentiviral vectors better alternatives for gene regulation within the hESCs. Because their pre-integration complex can get through the intact membrane of the nucleus within the target cell, lentiviruses can infect both dividing and non-dividing cells or terminally differentiated cells such as macrophages, retinal photoreceptors, and liver cells (Naldini et al., 1996). Lentiviral vectors are also promising gene transfer vehicles for HSCs, which reside almost exclusively in the G0/G1 phase of the cell cycle (Cheshier et al., 1999). The only cells lentiviruses cannot gain access to are quiescent cells in the G0 state which blocks the reverse transcription step (Amado and Chen, 1999). Lentiviruses can stably change the gene expression within hESCs for up to 6 months and are more resistant to transcriptional silencing (Pfeifer et al., 2002). High expression level of enhanced green fluorescent protein (eGFP) was achieved both in undifferentiated hESCs and their derivatives (Gropp et al., 2003). Overexpression of different genes, for instance, oct-4, nanog, eGFP has been reported under the control of various promoters, such as human cytomegalovirus (CMV) immediate early region enhancer–promoter, the composite CAG promoter (consisting of the CMV immediate early enhancer and the chicken β-actin promoter), human phosphoglycerate kinase 1(PGK) promoter, human elongation factor 1α (EF1α) promoter, and Ubiquitin (Ub) promoter (Ramezani et al., 2000; Salmon et al., 2000; Luther-Wyrsch et al., 2001; Gropp et al., 2003; Ma et al., 2003). Among these promoters, the CMV promoter does not perform well in HSCs (Boshart et al., 1985). Moreover, it is often subject to extinction of expression and silencing in vivo (Kay et al., 1992). In comparison, EF1α promoter was the most popularly used one and showed consistently better performance. Single transgene expression can shorten the length of lentiviral vector, leading to relatively higher transduction efficiency of the recombinant lentivirus in the hESCs. However, screening of the positive transduced cells from the polyclonal population cannot be achieved unless the overexpressed gene encodes a fluorescent or membrane protein, or an antibiotics cassette. Instead, to express two recombinant genes and one of them could work as integration reporter, internal ribosome entry sites (IRES) and double-promoters have been extensively studied in lentiviral vector designing. IRES are sequences that can recruit ribosomes and allow cap-independent translation, which can link two coding sequences in one bicistronic vector and allow the translation of both proteins in hESCs. The expression level of target gene by bicistronic vectors could be higher than that by single gene vectors; however, the percentage of positively transduced cells was relatively lower (Ben-Dor et al., 2006). Besides, the expression of downstream gene to IRES may inconsistently depend on the sequence of its upstream gene in an unpredictable manner (Yu et al., 2003). In comparison, lentiviral vectors containing double-promoters allow expression of reporter gene and target gene independently as well as the permission of transgene expression under tissue-specific promoter.

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Gene regulation based on the bacterial tetracycline repressor/operator (tetR/tetO) system has been applied into lentiviral vector design. To make the expression of a transgene inducible, the tetO cassette is inserted upstream of the transgene promoter and the tetR cassette can either be transcribed by the same gene expression vector or by a separate vector within the same hESC, binding to the tetO and inhibit gene expression. Conditional gene expression can be achieved when tetracycline or doxycycline is added to the cells, releasing the tetR binding and turning on the promoter (Szulc et al., 2006). Accompanied with various benefits using lentiviral vectors in hESCs, the obvious concern came up on the biosafety issues. The lentiviral vectors based on HIV could self-replicate and could be produced during manufacture of the vectors in the packaging cells by a process of recombination. Also a self-replicating infectious vector may transform hESC into a cancer stem cell by chromosome integration and activation of a neighboring proto-oncogene. Therefore, a number of modifications and changes were made over time leading to the safe production of high-titer lentiviral vector preparations. In addition to the structural gag, pol, and env genes common to all retroviruses, more complex lentiviruses, contain two regulatory genes, tat and rev, crucial for viral replication, and four accessory genes, vif, vpr, vpu, and nef, which are not critical for viral growth in vitro but are essential for in vivo replication and pathogenesis. The Tat protein regulates the promoter activity of the 5 BMPs’ LTR and is necessary for the transcription from the 5LTR. The Rev protein regulates gene expression at post-transcription level. It promotes the transport of unspliced and singly spliced viral transcripts into cytoplasma, allowing the production of the late viral proteins. The Tat and Rev are necessary for efficient gag and pol expression and new viral particles production. Understanding the functions of these genes leads to a 10-year path of lentiviral vector design. The first generation of HIV-derived vectors was produced transiently by transfection of plasmids coding for the packaging functions and the transgene plasmid into a suitable cell line mostly derived from 293 cells (Naldini et al., 1996). The ψ sequences and the env gene were removed from the HIV genome, the 5LTR was replaced by heterologous promoter, and the 3LTR was replaced by a polyadenylation signal. The envelope was replaced from another virus, and was most often VSV-G (Burns et al., 1993). In the second generation, to attenuate the virulence of the virus, all four accessory genes were removed and the HIV-derived packaging component was reduced to the gag, pol, tat, and rev genes of HIV-1 in the second version of the system (Zufferey et al., 1997). However, viruses can still be produced in vitro. In the third generation, constitutively active promoter sequences replaced part of the U3 region in the 5LTR in the transgene vector. The activity of the 5 LTR during vector production became independent of tat gene, which could be completely removed from the packaging construct. The rev gene, necessary for the gag/pol expression, was separately cloned into another plasmid to minimize the likelihood of recombination. In addition, a 299 bp deletion in the 3 LTR blocks the function of enhancer and promoter, resulting in the self-inactivation (SIN) of the provirus in the infected cells and minimizes the risk of insertional oncogenesis. Therefore, an internal promoter is needed for SIN vectors to drive transgene expression, allowing the use of tissue-specific or inducible promoters. The resulting gene delivery system, which conserves only three genes (rev, gag, pol) of HIV-1 and relies on four separate transcriptional units for the production of transducing particles, offers significant advantages for its predicted biosafety. Other modifications of lentiviral vectors were performed to satisfy different expression requirements. To enhance the ability of infection, the central polypurine tract (cPPT) is often included in the transgene vectors. Insertion of the woodchuck hepatitis virus post-transcriptional regulatory element (WPRE) was previously found to enhance transgene expression (Zufferey et al., 1999). However, inclusion of WPRE from certain lentiviral vectors showed lower transgene expression in human HSCs KG1a cell line (Ramezani et al., 2000). Besides stable gene expression, mutation of integrase protein itself and the integrase recognition sequences (att) in the lentiviral LTR could disable the integration of lentiviral vector and permitted transient gene

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expression (Nightingale et al., 2006). To lower the possibility of integration by LTR during lentiviral vector construction, E. coli Stbl3 and E. coli Stbl2 strains (Invitrogen) instead of DH5 α were developed, and optimization of culturing temperature under 30°C instead of 37°C reduced the possibility of LTR recombination. Adenoviral Vectors and Adeno-Associated Viral Vectors Adenoviruses are a group of non-pathogenic viruses that contain a linear double-stranded DNA genome without envelope. They have been developed as gene delivery vehicles due to the ability to infect non-dividing cells. Adenoviral vectors do not integrate into the genome of host cells providing a transient expression of the transgene. Adenoviruses are capable of transducing cells in vivo taking up to 30 kb exogenous DNA and adenovirusassociated viruses can express 4.8 kb transgene (Tatsis and Ertl, 2004; Volpers and Kochanek, 2004). Co-infection with helper viruses such as herpes simplex virus is required for adeno-associated viral vectors, which still needs to be optimized to achieve productive infection. Adenoviruses-derived vectors have been successfully used in mESCs studies (Mitani et al., 1995; Kawabata et al., 2005) and their applications as homologous recombination and gene transfer vehicle in the hESCs and/or their differentiating progenies are under investigation (Ohbayashi et al., 2005; Stone et al., 2005).

DIFFERENTIATION OF HESCS INTO TISSUE-SPECIFIC LINEAGES AND TRANSPLANTATION OF HESC-DERIVED CELLS To date, a large number of methods and protocols to drive the differentiation of hESCs into a broad spectrum of tissue-specific lineages in vitro representing three germ layers have been documented. However, hESC-based regenerative medicine largely relies on the generation of transplantable progenies from hESCs that will function in vivo. Therefore, in addition to identifying tissue-specific lineages derived from hESCs by morphological and phenotypic criteria and in vitro functional assays, hESC-derived progenies have to be functionally evaluated in vivo by transplantation into appropriate animal models. In this chapter, we review the approaches to generating diverse cell lineages from hESCs that have been functionally assessed in vivo by transplantation assays. Mesodermal Derivatives and Their Transplantation Mesodermal including hematopoietic, vascular, and cardiac differentiation from hESCs have been well characterized in great detail. Derivation of hematopoietic cells from hESCs is not only important for studying hematopoietic development in human but also is opening exciting opportunities to create an alternative cell source in addition to cord blood and bone marrow for transplantation in the clinic. Different methods have been used to induce hematopoietic differentiation from hESCs in vitro. The first report on derivation of hematopoietic cells from hESCs employed co-culture of hESCs with murine bone marrow cell line S17 or the yolk sac endothelial cell line C166 (Kaufman et al., 2001). An improvement on the production of CD34 hematopoietic progenitor cells has then been achieved by co-culturing hESCs with OP9 stromal cells, a bone marrow stromal cell line created from mice deficient in macrophage colony stimulating factor (M-CSF) (Vodyanik et al., 2005). Nevertheless, hematopoietic differentiation by the co-culture system is inefficient and hematopoietic cells derived from the system lack the expression of pan-leukocyte marker CD45. Our group has recently demonstrated that a combination of hematopoietic cytokines and BMP-4 efficiently augment hematopoietic differentiation from hEBs (Chadwick et al., 2003; Cerdan et al., 2004), and identified a rare subpopulation of cells lacking CD45 but expressing PECAM-1, Flk-1, and VE-Cadherin (termed CD45negPFV precursors) that are exclusively responsible for hematopoietic cell fate (Wang et al., 2004). Function of hematopoietic cells derived by either stromal co-culture or EB formation system has been evaluated in vivo by xenotransplantation repopulation assays that have been instrumental in measuring human somatic HSCs

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(Dick et al., 1997). However, generation of in vivo repopulating hematopoietic cells from hESCs has been proven to be difficult. Our laboratory has recently demonstrated that CD45 cells isolated from EBs cannot be successfully intravenously transplanted into immunocompromised mice due to the rapid aggregation upon exposure to mouse serum and the levels of reconstitution were still very low despite direct intra-femoral injection of hESC-derived hematopoietic cells to bypass the circulation and allow mice to survive (Wang et al., 2005b). Moreover, CD45negPFV precursors or their derived hematopoietic cells were unable to engraft even after transplantation into the liver of newborn immunocompromised mice (unpublished data), an assay more amenable to readout repopulating hematopoietic cells (Yoder et al., 1997). In addition to our studies, sorted CD34lineage cells or unsorted cells from hESCs differentiated on S17 stromal cells have recently been shown to engraft, but at a very low level, after transplantation into fetal sheep or adult non-obese severe combined immunodeficient NOD/SCID mice, respectively (Narayan et al., 2006; Tian et al., 2006). Taken together, these studies suggest that full understanding of molecular and cellular events dictating hematopoiesis from hESCs is required to improve means to generate HSCs with potent repopulating ability from hESCs. Initiation of vascular development has been shown to be closely associated with the emergence of hematopoiesis and a common precursor termed “hemangioblast” with both vascular and hematopoietic potential has been identified during hematopoietic differentiation of mESCs and in the primitive streak of the mouse embryo (Choi et al., 1998; Huber et al., 2004). In human, our laboratory has recently identified a subpopulation of primitive endothelium-like cells termed CD45negPFV precursors with hemangioblast properties during EB differentiation of hESCs in the presence of exogenous hematopoietic cytokines and BMP-4 (Wang et al., 2004). Cells expressing PECAM1/CD31, a marker associated with cells capable of early hematopoietic potential in the human embryo (Oberlin et al., 2002), first emerged at day 3 and significantly increased at day 7 through day 10 of EB development. Isolated subpopulation of CD45negPFV precursors contained single cells with both hematopoietic and endothelial capacity. After 7 days in culture condition conducive to endothelial maturation, the cells not only strongly expressed CD31, VE-cadherin and mature endothelium markers vWF and eNOS, but also possessed low-density lipoprotein (LDL) uptake capacity (Wang et al., 2004). However, the in vivo function of hESC-derived endothelial cells from our system has not been assessed. Levenberg et al. reported the first study to characterize differentiation of hESCs into endothelial cells during spontaneous EB differentiation without adding any exogenous growth factors by functionally evaluating hESCs-derived endothelial cell both in vitro and in vivo (Levenberg et al., 2002). Although the efficiency of endothelial differentiation is relatively low in the spontaneous system as opposed to our system, their differentiation kinetics are similar in that the expression of CD31, VE-cadherin, and CD34 appeared at days 3–5 and reached a maximum about 2% at days 13–15 during EB differentiation. CD31 cells isolated from day 13 EBs displayed endothelium characteristics by expressing endothelium-specific markers VE-cadherin and vWF, taking up acetylated LDL (ac-LDL) and forming tube-like structures (Levenberg et al., 2002). Furthermore, hESC-derived CD31 cells were able to form functional bloodcarrying microvessels after transplantation into SCID mice (Levenberg et al., 2002). A recent study from the same group has further shown that hESC-derived endothelial cells are able to vascularize skeletal muscle tissue construct using a three-dimensional multiculture system in vitro (Levenberg et al., 2005). More significantly, pre-endothelialization of the construct, by promoting implant vascularization, can improve blood perfusion to the implant and implant survival in vivo (Levenberg et al., 2005). In summary, these studies demonstrate that endothelial differentiation of hESCs likely recapitulate vasculogenesis during human development and hESCderived endothelial cells are able to vascularize tissue construct in vitro and implant in vivo. However, it remains to further determine potential therapeutic implications of embryonic endothelial cells generated from hESCs for treatment of vascular disease and repair of ischemic tissues. Methods from different laboratories to induce cardiac differentiation from hESCs have also been demonstrated (Kehat et al., 2001; Xu et al., 2002; Mummery et al., 2003). During spontaneous EB differentiation of

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hESCs, 8% of EBs contained contracting cardiomyocytes that displayed structural, phenotypic, and functional properties of early-state cardiomyocytes (Kehat et al., 2001). Treatment of cells with 5-aza-2-deoxycytidine increased cardiomyocyte differentiation in a time-dependent and concentration-dependent manner and Percoll density centrifugation could achieve a population containing 70% cardiomyocytes (Xu et al., 2002). In addition to spontaneous differentiation, co-culture of hESCs with visceral endoderm-like cell line, END-2 has also been shown to induce cardiac differentiation of hESCs (Mummery et al., 2003). The induction events for cardiac development in the hESCs remain to be further defined in detail as cardiomyocytes are generated in serum-containing conditions in most studies. Recently, hESC-derived cardiomyocytes have been functionally tested in a swine model of complete atrioventricular block as “biologic pacemaker” for the treatment of bradycardia and the transplanted cells survive, integrate, ad successfully pace the ventricle with complete heart block (Kehat et al., 2004). However, long-term pacemaking function of grafted hESC-derived cardiomyocytes has not been evaluated in the study and it also raises the concern that transplanted cells could serve as a nidus for arrhythmia. Ectodermal Derivatives and Their Transplantation Most studies on derivation of ectodermal lineages from hESCs have focused on neuroectoderm and neural cells, aiming to create an unlimited source of neural cells for transplantation therapies. Differentiation of hESCs into neural lineages has been induced using different methods (Carpenter et al., 2001; Reubinoff et al., 2001; Zhang et al., 2001). hESC-derived neural progenitors that could differentiate into three neural lineages – mature neurons, astrocytes, and oligodendrocytes in vitro have been transplanted into neonatal mouse brain where they incorporated into host brain parenchyma, migrated along established brain migratory tracks, and differentiated into progeny of three neural lineages in vivo (Reubinoff et al., 2001; Zhang et al., 2001). Furthermore, enriched population of neural progenitors from hESCs that were grafted into the striatum of Parkinsonian rats induced partial behavioral recovery (Ben-Hur et al., 2004). The functional improvement is likely due to release of neurotropic factors from the graft to promote survival of impaired endogenous dopamine neurons as hESC-derived neural progenitors could not acquire dopaminergic fate in the host tissue. Despite recent availability of protocols to generated specific dopaminergic neurons from hESCs (Park et al., 2004; Perrier et al., 2004; Schulz et al., 2004; Zeng et al., 2004), only one of the studies has examined the in vivo functions of hESC-derived dopamine neurons after transplantation into the striatum of 6-hydroxydopamine treated rat and significance of the study is unclear, because only a few dopaminergic neurons survived 5 weeks after transplantation and no functional improvement has been demonstrated (Zeng et al., 2004). Future studies are required to determine the appropriate cell type for transplantation therapies by functionally evaluating hESC-derived dopamine neurons in comparison to neural progenitors in animal models of Parkinson disease. In addition to dopamine neurons, other specific neuronal subtypes like motoneurons that have also been recently generated from hESCs (Li et al., 2005) have to be functionally assessed in animal models of spinal cord injuries and motoneuronal degeneration. Endodermal Derivatives and Their Transplantation In contrast to mesodermal and ectodermal differentiation of hESCs, specification of hESCs into endodermal lineages, specifically insulin-producing cells, is less studied. Although differentiation of hESCs into insulinproducing cells have been demonstrated by either spontaneous system, exposure to inducing factors, or overexpression of Pdx1 or Foxa2, important transcription factors involved in pancreatic development (Assady et al., 2001; Segev et al., 2004; Brolen et al., 2005; Lavon et al., 2006), the frequency of these cells generated in the current differentiation conditions is too low to allow detailed characterization and functional analysis.

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GENETIC MODIFICATIONS OF HESC-DERIVED PROGENIES Successful derivation of diverse tissue-specific lineages from hESCs sets the stage to genetically manipulate hESC-derived progenies. However, in sharp contrast to the broad applications of genetic modifications to undifferentiated hESCs, very few studies have investigated genetic manipulations of specific lineages derived from hESCs possibly due to the difficulties in prospectively isolating a low frequency of lineage-specific progenies from the bulk population to allow detailed studies. To date, hESC-derived hematopoietic cells are the only cell type to which retrovirus-based gene transfer has been successfully applied (Menendez et al., 2004). Our laboratory has recently characterized and optimized a GALV-pseudotyped retroviral gene transfer strategy to stably transduce the hematopoietic progenitor cells derived from CD45negPFV hemogenic precursors that were prospectively isolated from hEBs (Menendez et al., 2004). We achieved 25% transduction efficiency using GALV-pseudotyped retrovirus into CD45negPFV precursors-derived hematopoietic cells and a proportion of transduced cells co-expressed CD34 and were able to give rise to hematopoietic colony-forming unit (Menendez et al., 2004). These studies are expected to provide a method to examine the functional effects of ectopic expression of candidate genes that may regulate primitive human hematopoietic development. Using the GALV-pseudotyped retroviral gene delivery method, we have very recently evaluated the role of HoxB4 overexpression in CD45negPFV precursors derived from hESCs (Wang et al., 2005b). In contrast to the generation of repopulating hematopoietic cells from mESCs by overexpressing HoxB4 in mESC-derived hematopoietic progenitors, ectopic expression of HoxB4 in hESC-derived hematopoietic cells does not confer engraftment potential (Kyba et al., 2002; Wang, Y. et al., 2005). Overexpression and knock-down of genes associated with lineage development in hESC-derived progenies is critical to further understand lineage specification and commitment from hESCs.

POTENTIAL APPLICATIONS OF GENETICALLY MANIPULATED HESCS AND THEIR DERIVATIVES Augmenting Differentiation of hESCs into Specific Lineages Once formed as EBs in serum-containing medium, hESCs will spontaneously differentiate into diverse lineages representing three germ layers, but at very low levels. Although many studies have demonstrated that adding growth factors or morphogens related to lineage development into the medium is able to significantly increase the differentiation of hESCs into specific lineages, the frequencies of lineage-specific cells are, in general, still low (Chadwick et al., 2003). In the setting of hematopoietic differentiation, our group has observed that 10–20% of EBs at days 10–13 still contained Oct-4 positive cells (unpublished observation), suggesting that the differentiation processes of cells within the EBs are not synchronized and some cells are reluctant to respond to differentiation clues in the culture. Very recent genetic mapping study has suggested that pluripotency-associated transcription factors Oct-4, Nanog, and Sox2 repress a set of developmental regulators of lineage specification to maintain the pluripotent status of hESCs (Lee et al., 2006). Therefore, RNAi-based genetic knock-down of Oct-4, Nanog, or Sox2 is expected to release the repression of differentiation and thereby facilitate the generation of tissue-specific progenies from hESCs with the induction of proper growth factors along the pathways of lineage development. Indeed, Oct-4 knock-down in hESCs has been shown to induce endoderm differentiation (Hay et al., 2004). On the other hand, enforced expression of lineage-specific genes in undifferentiated hESCs will likely promote the differentiation of hESCs into specific lineages. In the context of hematopoietic differentiation, overexpression of HoxB4, a transcription factor involved in hematopoietic development and self-renewal of HSCs, in undifferentiated hESCs by lipofection promotes a 6–20-fold increase in the frequency of hematopoietic cells derived from hESCs (Bowles et al., 2006). In line

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with the augmenting effect of constitutive expression of HoxB4 on the hematopoietic differentiation of hESCs, our group has observed that the mRNA expression profile of HoxB4 during EB differentiation is temporally correlated with hematopoietic development from hESCs (unpublished observation). A very recent study has evaluated the effect of transfection-based overexpression of Foxa2 and Pdx1, transcription factors involved in different phases of early endoderm and pancreatic development, on the differentiation of hESCs into pancreatic cells (Lavon et al., 2006). In contrast to the insignificant effect of overexpression of Foxa2 on the differentiation of hESCs into endoderm lineage, constitutive expression of Pdx1 promoted the differentiation of hESCs toward insulin cells as shown by induced expression of most transcription factors involved in pancreatic development (Lavon et al., 2006). However, expression of insulin gene was not induced by enforced Pdx1 expression, suggesting that differentiation signals that can further drive the specification into insulin cells is still missing in spite of constitutive expression of Pdx1. Future studies are required to investigate introduction of inducible gene expression system into hESCs, which will allow us to study the role of lineage-specific genes in lineage development from hESCs at specific stage of hESCs differentiation. Lineage Tracking and Purification In order to better understand temporal differentiation and spatial organization of specific lineages from hESCs, it is important to trace lineage specification and commitment within heterogeneous populations of cells during EB differentiation. Introduction of reporter/selection genes under the control of lineage-specific promoters will allow us to monitor the differentiation of hESCs toward specific lineages. Furthermore, it offers us the feasibility to select and purify specific lineages and eliminate undesirable cells from the bulk population based on reporter gene expression, which is critical for the potential use of these hESC-derived lineages in cell-based therapies, since any potential contamination by undifferentiated hESCs will likely result in the development of teratomas. Eiges et al. and Gerrard et al. introduced eGFP reporter gene under the control of ESC-enriched gene murine Rex1 or Oct-4 promoter into hESCs to select the undifferentiated hESCs from their spontaneously differentiated derivatives in the culture (Eiges et al., 2001; Gerrard et al., 2005). Lavon et al. have very recently traced the differentiation of hESCs into pancreatic cells by generating and differentiating hESC lines carrying eGFP reporter gene under the control of insulin promoter or Pdx1 promoter (Lavon et al., 2006). These studies paved the way for future endeavors to examine the molecular and cellular mechanisms governing lineage specification, which in turn will provide insight into better generation of lineagespecific cells from hESCs. Modifying the Immunogenicity of hESCs and Their Derivatives hESC-derived tissue-specific progenies represent an promising source for the potential transplantation therapies to a broad spectrum of diseases in the clinic. However, immune response launched by the host immune system to the graft may comprise the therapeutic potential of derivatives from hESCs. Although we and others have demonstrated that hESCs and their derivatives after a short period of differentiation in vitro express low levels of major histocompatability complex (MHC) class I and are less susceptible to immune rejection than adult cells (Li et al., 2004; Drukker et al., 2006), it remains unclear whether hESC-derived cells differentiated to a fully functional adult phenotype after successful engraftment will still possess immuno-privileged properties to permanently evade immune rejection. To overcome potential immune rejection, a few approaches have been proposed, which include somatic cell nuclear transfer to create hESCs lines with identical MHC to that of host tissue, collection of hESC banks representing the broadest diversity of MHC polymorphorisms, and induction of a state of immune tolerance to an hESC line using tolerogenic HSCs derived from it. Though promising, the feasibility of these strategies remains to be validated. Alternatively, strategies to genetically modify the immunogenicity of hESCs and their derivatives by targeting genes that encode and

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control the cell surface expression of MHC classes I and II molecules provide another theoretical means to circumvent the immune barrier. The deletion of both classes of MHC molecule has been achieved in mESCs by disruption of the genes critical for the correct assembly and membrane expression of MHC classes I and II (Zijlstra et al., 1990; Grusby et al., 1991). Although grafts deficient in the expression of either MHC class I or II target molecules do not completely avoid rejection by immunologically intact allogeneic hosts, MHC class Ideficient grafts are rejected more slowly than grafts from normal mice. Genetic modifications of similar target genes for MHC class I expression in hESCs and their derivatives remain to be fully explored in future studies, given the applicability of multiple genetic tools to manipulate hESCs and their progenies.

CONCLUSION Derivation of hESCs opens up a new era for human development biology and regenerative medicine. Almost one decade of research in the past has made considerable progress in defining culture conditions to grow hESCs and developing protocols to differentiate hESCs into tissue-specific lineages. However, formulated culture condition completely devoid of animal component and uncharacterized serum elements to maintain hESCs remains to be further optimized. Moreover, efficient generation of specialized derivatives from hESCs that are able to function in vivo after transplantation into animal models has not been achieved so far. Realization of hESCs as a model system to study human development and unlimited source for regenerative medicine relies on the dissection of molecular and cellular mechanisms dictating the pluripotency, selfrenewal, and lineage specification of hESCs. Genetic manipulations of hESCs and their derivatives are anticipated to provide invaluable insight into the understanding of fundamental biology of hESCs, which in turn will be instrumental in the optimization of protocols to either maintain hESCs or specify hESCs into functional tissue-specific lineages with potential use in the clinic. ACKNOWLEDGMENT We thank Dr. Marc Bosse in the Bhatia laboratory for his critical comments and insights during the preparation of this review.

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12 Embryonic Stem Cells: Derivation and Properties Junying Yu and James A. Thomson

INTRODUCTION Embryonic stem (ES) cells are derived from early embryos, and are capable of indefinite self-renewal in vitro while maintaining the potential to develop into all cell types of the body – they are pluripotent. With these remarkable features, ES cells hold great promise in both regenerative medicine and basic biological research. In this chapter, we will discuss how ES cells are derived and what is known about the mechanisms that allow these cells to maintain their pluripotency while proliferating in vitro. DERIVATION OF ES CELLS Embryonic Carcinoma Cells Teratocarcinoma is a form of malignant germ cell tumor that occurs in both animals and humans. These tumors comprise an undifferentiated embryonal carcinoma (EC) component and differentiated derivatives that can include all three germ layers. Although teratocarcinomas had been known as medical curiosities for centuries (Wheeler, 1983), it was the discovery that male mice of strain 129 had a high incidence of testicular teratocarcinomas (Stevens and Little, 1954) that made these tumors more routinely amenable to experimental analysis. Because their growth is sustained by a persistent EC cell component, teratocarcinomas can be serially transplanted between mice. In 1964, Kleinsmith and Pierce demonstrated that a single EC cell was capable of both self-renewal and multilineage differentiation, and this formal demonstration of a pluripotent stem cell provided the intellectual framework for both mouse and human ES cells. The first mouse EC cell lines were established in the early 1970s (Kahan and Ephrussi, 1970; Evans, 1972). EC cells exhibit similar antigen and protein expression as the cells present in the inner cell mass (ICM) (Klavins et al., 1971; Comoglio et al., 1975; Gachelin et al., 1977; Solter and Knowles, 1978; Calarco and Banka, 1979; Howe et al., 1980; Henderson et al., 2002), and this led to the notion that EC cells are the counterpart of pluripotent cells present in the ICM (Martin, 1980; Rossant and Papaioannou, 1984). When injected into mouse blastocysts, some EC cell lines are able to contribute to various somatic cell types (Brinster, 1974; Mintz and Illmensee, 1975; Papaioannou et al., 1975; Illmensee and Mintz, 1976;), but most EC cell lines have limited developmental potential and contribute poorly to chimeric mice, probably reflecting genetic changes acquired during teratocarcinoma formation (Atkin et al., 1974; McBurney, 1976; Bronson et al., 1980; Zeuthen et al., 1980). Mutations that confer growth advantages to EC cells are likely to accumulate during tumorigenesis, and EC cells in chimeras can result in tumor formation (Papaioannou et al., 1978). As a result, there are limitations in the application of EC cells to both regenerative medicine and research in basic developmental biology.

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Following fertilization, as the one-cell embryo migrates down the oviduct, it undergoes a series of cleavage divisions resulting in a morula. During blastocyst formation, the outer cell layer of the morula delaminates from the rest of the embryo to form the trophectoderm. The ICM of the blastocyst gives rise to all the fetal tissues (ectoderm, mesoderm and endoderm) and some extra-embryonic tissues, and the trophectoderm gives rise to the trophoblast. Although the early ICM can contribute to the trophoblast, the late ICM does not (Winkel and Pedersen, 1988), suggesting there is some restriction in developmental potential at this stage. In normal embryos, the pluripotent cells of the embryo have a transient existence, as these cells quickly give rise to other non-pluripotent cells through the normal developmental program. Thus, the pluripotent cells of the intact embryo really function in vivo as precursor cells and not as stem cells. However, if early mouse embryos are transferred to extra-uterine sites, such as the kidney or testis capsules of adult mice, they can develop into teratocarcinomas that include pluripotent stem (EC) cells (Solter et al., 1970; Stevens, 1970). These ectopic transplantation experiments result in teratocarcinomas at high frequencies, even in strains that do not spontaneously have elevated incidence of germ cell tumors, suggesting that this process is not the result of rare neoplastic transformation events. These key transplantation experiments led to the search for culture conditions that would allow the in vitro derivation of pluripotent stem cells directly from the embryo, without the intermediate need to form teratocarcinomas in vivo. Derivation of ES Cells In 1981, pluripotent ES cell lines were derived directly from the ICM of mouse blastocysts using culture conditions previously developed for mouse EC cells (Evans and Kaufman, 1981; Martin, 1981). ES cell cultures derived from a single cell could differentiate into a wide variety of cell types, or could form teratocarcinomas when injected into mice (Martin, 1981). Unlike EC cells, however, these karyotypically normal cells contributed at a high frequency to a variety of tissues in chimeras, including germ cells, and thus provided a practical way to introduce modifications to the mouse germ line (Bradley et al., 1984). The efficiency in mouse ES cell derivation is influenced by genetic background. For example, ES cells can be easily derived from the inbred 129/ter-Sv strain, but less efficiently from C57BL/6 and other mouse strains (Ledermann and Burki, 1991; Kitani et al., 1996), and these strain differences somewhat correspond with the propensity of mice of different strains to develop teratocarcinomas. These observations suggested that genetic and/or epigenetic components play an important role in the derivation of mouse ES cells. On the other hand, the efficiency of teratocarcinoma formation induced through extra-uterine mouse embryo transplantations appears to be somewhat less strain dependent (Damjanov et al., 1983). This indicates that the difference in the efficiency of ES cell derivation from different mouse strains might be due to suboptimal culture conditions. Indeed, mouse ES cells can be derived from some non-permissive strains using modified protocols (McWhir et al., 1996; Brook and Gardner, 1997). ES cell lines are generally derived from the culture of the ICM, but this does not mean that ES cells are the in vitro equivalent to ICM cells, or even that ICM cells are the immediate precursor to ES cells. It is possible that during culture, ICM cells give rise to other cells that serve as the immediate precursors. Some experiments suggest that ES cells more closely resemble cells from the primitive ectoderm, the cell layer derived from the ICM after delamination of the primitive endoderm. Isolated primitive ectoderm from the mouse gives rise to ES cell lines at a high frequency, and allows the isolation of ES cell lines from mouse strains that had previously been refractory to ES cell isolation (Brook and Gardner, 1997). Indeed, single primitive ectoderm cells can give rise to ES cell lines at a reasonable frequency, something not possible with early ICM cells (Brook and Gardner, 1997). Although these experiments do suggest that ES cells are more closely related to primitive ectoderm than to ICM, they do not reveal whether ES cells more closely resemble primitive ectoderm or another cell type (e.g. very early germ cells) derived from it in vitro (Zwaka and Thomson, 2005). As no pluripotent cell in the intact

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embryo undergoes long-term self-renewal, ES cells are in some ways tissue culture artifacts. It is surprising that even more than 20 years after their derivation, the origin of these cells is not completely understood. Given the dramatic improvement in molecular techniques since the initial derivation in the 1980s, there is considerable value in reexamining the origin of ES cells to better understand the control of their proliferative pluripotent state (Zwaka and Thomson, 2005). In addition to derivation from the ICM and isolated primitive ectoderm, mouse ES cells have also been derived from morula-stage embryos and even from individual blastomeres (Eistetter, 1989; Delhaise et al., 1996; Chung et al., 2005; Tesar, 2005). Again, although the ES cell lines were derived from morula, there may well be a progression of intermediate states during the derivation process. The frequencies of success were lower when starting with morula or blastomeres, but these results do suggest that it might be possible to derive human ES cells without the destruction of an embryo. Such cell lines could prove useful to the child resulting from the transfer of a biopsied embryo, as they would be genetically matched to the child. Derivation of Human ES Cells In 1978 the first baby was born from an embryo fertilized in vitro (Steptoe and Edwards, 1978), and without this event, the derivation of human ES cells would not have been possible. Although there were attempts to derive human ES cells as early as the 1980s, species-specific differences and suboptimal human embryo culture media delayed their successful isolation until 1998 (Thomson et al., 1998). For example, the culture of isolated ICMs from human blastocysts was reported (Bongso et al., 1994), but stable undifferentiated cell lines were not produced in medium supplemented with leukemia inhibitory factor (LIF) in the presence of feeder layers, conditions that allow the isolation of mouse ES cells. In the mid-1990s, ES cell lines were derived from two non-human primates: the rhesus monkey and the common marmoset (Thomson et al., 1995, 1996). Experience with these ES cell lines and concomitant improvements in culture conditions for human in vitro fertilization (IVF) embryos (Gardner et al., 1998) resulted in the successful derivation of human ES cell lines (Thomson et al., 1998). These human ES cells had normal karyotypes, and even after prolonged undifferentiated proliferation, maintained the developmental potential to contribute to advanced derivatives of all three germ layers. To date, more than 120 human ES cell lines have been established worldwide (Stojkovic et al., 2004b). Although most were derived from isolated ICMs, some were derived from morulae or later blastocyst stage embryos (Stojkovic et al., 2004a; Strelchenko et al., 2004). It is not yet known whether ES cells derived from these different developmental stages have any consistent differences or whether they are developmentally equivalent. Human ES cell lines have also been derived from embryos carrying various disease-associated genetic changes, which provide new in vitro models of disease (Verlinsky et al., 2005). Recently, and with a remarkably high efficiency, human ES cell lines have been derived through a process of somatic cell nuclear transfer (SCNT) (Hwang et al., 2004, 2005). By using the nuclear transfer technology, the nuclei of human somatic cells, such as skin cells, were transferred to donated human oocytes that were already stripped of their own genetic material. The oocytes were then activated and cultured in vitro to the blastocyst stage for ES cell derivation. Because such ES cells contain the genetic material present in the donor cell, it is hoped that they could provide immune-compatible ES cells for cell replacement therapies.

CULTURE OF ES CELLS Culture of Mouse ES Cells Mitotically inactivated feeder layers were first used to support difficult-to-culture epithelial cells (Puck et al., 1956), and were later successfully adapted for the culture of mouse EC cells (Martin and Evans, 1975; Martin

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et al., 1977) and mouse ES cells (Evans and Kaufman, 1981; Martin, 1981). Medium which is “conditioned” by co-culture with fibroblasts sustains EC cells (Smith and Hooper, 1983). Fractionation of conditioned medium led to the identification of a cytokine, LIF, that sustains ES cells (Smith et al., 1988; Williams et al., 1988). LIF and its related cytokines act via the gp130 receptor (Yoshida et al., 1994). Binding of LIF induces dimerization of LIF/gp130 receptors, which in turn activates the latent transcription factor STAT3 (Lutticken et al., 1994; Wegenka et al., 1994), and ERK mitogen-activated protein kinase (MAPK) cascade (Takahashi-Tezuka et al., 1998). STAT3 activation is sufficient for LIF-mediated self-renewal of mouse ES cells in the presence of serum (Matsuda et al., 1999). In contrast, suppression of the ERK pathway promotes ES cell proliferation (Burdon et al., 1999). In serum-free medium, LIF alone is insufficient to prevent mouse ES cell differentiation, but in combination with BMP (bone morphogenetic protein, a member of the TGFβ superfamily), mouse ES cells are sustained (Ying et al., 2003a). BMPs induce expression of Id (inhibitor of differentiation) proteins and inhibit the ERK and p38 MAPK pathways, thus attenuating the pro-differentiation activation of ERK MAPK pathway by LIF. Culture of Human ES Cells Mitotically inactivated fibroblast feeder layers and serum-containing medium were used in the initial derivation of human ES cells, essentially the same conditions used for the derivation of mouse ES cells prior to the identification of LIF (Thomson et al., 1998; Reubinoff et al., 2000). However, it now appears largely to be a lucky coincidence that fibroblast feeder layers support both mouse and human ES cells, as the specific factors identified to date that sustain mouse ES cells do not support human ES cells. LIF and its related cytokines fail to support human or non-human primate ES cells in serum-containing media that supports mouse ES cells (Thomson et al., 1998; Daheron et al., 2004; Humphrey et al., 2004; Sumi et al., 2004), and BMPs, when added to human ES cells, cause rapid differentiation in conditions that would otherwise support their self-renewal (Xu et al., 2002; Pera et al., 2004). Indeed, the LIF/STAT3 pathway has yet to be shown to have any relevance to the self-renewal of human ES cells (Thomson et al., 1998; Daheron et al., 2004; Humphrey et al., 2004). In contrast to mouse ES cells, fibroblast growth factor (FGF) signaling appears to be of central importance in the self-renewal of human ES cells. Basic FGF (bFGF or FGF2) allows the clonal growth of human ES cells on fibroblasts in the presence of a commercially available serum replacement (Amit et al., 2000; Xu et al., 2001). At higher concentrations, bFGF allows feeder independent growth of human ES cells cultured in the same serum replacement (Wang et al., 2005; Xu, C. et al., 2005; Xu, R.H. et al., 2005). The mechanism through which these high concentrations of bFGF exert their functions is incompletely known, although one of the effects is suppression of BMP signaling (Xu, R.H. et al., 2005). Serum and the serum replacement currently used have significant BMP-like activity, which is sufficient to induce differentiation of human ES cells, and conditioning this medium on fibroblasts reduces this activity (Xu, R.H. et al., 2005). At moderate concentrations of bFGF (40 ng/ml), the addition of noggin or other inhibitors of BMP signaling significantly decreases background differentiation of human ES cells. At higher concentrations (100 ng/ml), bFGF itself suppresses BMP signaling in human ES cells to levels comparable to those observed in fibroblast-conditioned medium, and the addition of noggin is no longer needed for feeder independent growth (Xu, R.H. et al., 2005). As more defined culture conditions are developed for human ES cells that lack serum products containing BMP activity, it is not yet clear how important the suppression of the BMP pathway will be, unless there is significant production of BMPs by the ES cells themselves. Also, the effects of BMP signaling could change depending on context. Even in mouse ES cells, BMPs are inducers of differentiation unless they are presented in combination with LIF, and it is entirely possible that in a different signaling context, the effects of BMPs on human ES cells could change. Suppression of BMP activity by itself is insufficient to maintain human ES cells (Xu, R.H. et al., 2005), thus bFGF must be serving other signaling functions. Human ES cells themselves produce FGFs, and in high-density

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cultures either on fibroblasts or in fibroblast-conditioned medium, it is not necessary to add FGFs. However, chemical inhibitors of FGF receptor-mediated phosphorylation cause differentiation of human ES cells under these standard culture conditions (Dvorak et al., 2005). The required downstream events are not yet well worked out, but some evidence implicates activation of the ERK pathway (Kang et al., 2005). Although FGF signaling appears to have a central role in the self-renewal of human ES cells, other pathways have also been implicated. When combined with low to moderate levels of FGFs, TGFβ/activin/nodal signaling has a positive effect on the undifferentiated proliferation of human ES cells (Amit et al., 2004; Beattie et al., 2005; James et al., 2005; Vallier et al., 2005), and inhibition of this pathway leads to differentiation (James et al., 2005; Vallier et al., 2005). However, one of the effects of inhibiting the TGFβ/activin/nodal pathway is a stimulation of the BMP pathway (James et al., 2005), which in itself would be sufficient to induce differentiation. Thus, it is not yet clear whether TGFβ/activin/nodal signaling has a role in human ES cell self-renewal independent of its effects on BMP signaling. Further studies directly inhibiting the BMP pathway in the context of inhibition or stimulation of the TGFβ/activin/nodal are needed to resolve this issue. The molecular components of the Wnt pathway are well represented in human ES cells (Sperger et al., 2003). In short-term cultures, activation of Wnt signaling by a pharmacological GSK-3-specfic inhibitor (6bromoindirubin-3-oxime (BIO)) has been reported to have a positive effect on human ES cell self-renewal (Sato et al., 2004), but in a different study, inhibition of Wnt signaling or stimulation of Wnt signaling by the addition of recombinant Wnt proteins showed no effect on the maintenance of human ES cells (Dravid et al., 2005). It is possible that the positive observed effect of BIO on human ES cells is mediated through other pathways (James et al., 2005). For human ES cells to be used in a clinical setting, it would be useful for these cells to be derived and maintained in conditions that are free of animal products. For example, human ES cells derived with mouse embryonic fibroblasts were shown to be contaminated with immunogenic non-human sialic acid, which would cause an immune reaction if the cells were used in human patients (Martin et al., 2005). Toward this goal, protein matrices including laminin and fibronectin, and different types of human feeder cells were developed to sustain human ES cells (Xu et al., 2001; Amit et al., 2003; Richards et al., 2003). New human ES cell lines have been derived in the absence of feeder cells, but in the presence of a mouse-derived matrix and a bovine-derived serum replacement product (Klimanskaya et al., 2005). Existing human ES cell lines have been grown in defined serum-free medium that included sphingosine-1-phosphate (S1P) and platelet-derived growth factor (PDGF) (Pebay et al., 2005), but this medium does not eliminate the need for feeder layers. Existing human ES cells lines have also been adapted to feeder-free conditions in which none of the protein components are animal derived, but it is not yet known whether these specific conditions will allow derivation of new lines (Li et al., 2005). Clearly, however, recent improvements in human ES cell culture suggest that the development of completely defined, feeder-free culture conditions are near at hand, and that such conditions will allow the derivation of new cell lines that will be more directly applicable to therapeutic purposes. During extended culture, genetic changes can accumulate in human ES cells (Draper et al., 2004; Maitra et al., 2005). The status of imprinted genes appears to be relatively stable in human ES cells, but can also change (Rugg-Gunn et al., 2005). Such genetic and epigenetic alterations present a challenge that must be appropriately managed if human ES cells are to be used in cell replacement therapy. The rates at which these changes accumulate in culture likely depend on the culture system used, and the particular selective pressures applied. For example, in all current culture conditions, the cloning efficiency of human ES cells is poor, typically 1% or less (Amit et al., 2000). If cells are dispersed into a suspension of single cells, there is a tremendous selective pressure for cells that clone at a higher efficiency, and indeed, such an increase in cloning efficiency is observed in karyotypically abnormal cells (Enver et al., 2005). Enzymatic methods of passaging ES cells can

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allow long-term passage without karyotypic changes if the clump size is carefully controlled (Amit et al., 2000), but if such methods are used to disperse cells to single cell suspensions or small clumps, karyotypic changes are more frequent (Cowan et al., 2004). This is a likely explanation for why mechanical splitting of individual colonies allows such long-term karyotypic stability (Buzzard et al., 2004). Understanding the rates at which genetic changes occur and the selective pressures that allow them to over-grow a culture in different culture conditions will be critical to the large-scale expansion and clinical use of human ES cells.

DEVELOPMENTAL POTENTIAL OF ES CELLS Differentiation of ES Cells Since ES cells have the ability to differentiate into clinically relevant cell types such as dopamine neurons, cardiomyocytes, and β cells, there is tremendous interest in using these cells both in basic biological research and in transplantation medicine. Both uses demand a great deal of control over lineage allocation and expansion. There are several experimental approaches to demonstrate the developmental potential of ES cells and to direct their differentiation to specific lineages. These approaches range in complexity and experimental control from allowing the ES cells to respond to normal developmental cues in a chimera within an intact embryo, to the addition of defined growth factors to a monolayer culture. Mouse ES cells reintroduced into blastocysts participate in normal embryogenesis, even after prolonged culture and extensive manipulation in vitro. In such chimeras, the progeny of ES cells contributes to both somatic tissues and germ cells (Bradley et al., 1984). When ES cells are introduced into tetraploid blastocysts, mice entirely derived from ES cells can be produced, as the teraploid component is out-competed in the ICMderived somatic tissues (Nagy et al., 1993; Ueda et al., 1995). Although mice entirely derived from ES cells can be generated, signals from the ICM of the blastocyst are likely necessary for mouse ES cells to contribute to offspring, as fetal development has not been reported when the ICM is completely replaced with ES cells. ES cells injected into syngeneic or immunocompromised adult mice form teratomas that contain differentiated derivatives of all three germ layers (ectoderm, mesoderm and endoderm) (Martin, 1981). This property is similar to both early embryos and EC cells, and is an approach now routinely used to demonstrate the pluripotency of human ES cells (Thomson et al., 1998). Very complex structures resembling neural tube, gut, teeth and hair form in these teratomas in a very consistent temporal pattern, and these teratomas do offer an experimental model to study the development of these structures in human material, but the environment of differentiation is complex and difficult to manipulate. Aggregates of EC cells or ES cells cultured in conditions that prevent their attachment form cystic “embryoid bodies” (Martin and Evans, 1975; Martin et al., 1977) that recapitulate some of the events of early development. Differentiated derivatives of all three germ layers form in these structures, and for ES cells, the temporal events occurring mimic in vivo embryogenesis. The formation of embryoid bodies has been used, for example, to produce neural cells (Bain et al., 1995; Zhang et al., 2001), cardiomyocyte (Klug et al., 1996; He et al., 2003), hematopoietic precursors (Keller et al., 1993; Chadwick et al., 2003), β-like cells (Assady et al., 2001; Lumelsky et al., 2001), hepatocytes (Hamazaki et al., 2001; Rambhatla et al., 2003), and germ cells (Hubner et al., 2003; Toyooka et al., 2003; Geijsen et al., 2004). The formation of a three-dimensional structure in embryonic bodies (EBs) is useful to promote certain developmental events, but the complicated cell–cell interactions make it difficult to elucidate the essential signaling pathways involved. A somewhat more controlled method to differentiate ES cells is to co-culture them with differentiated cells that induce their differentiation to specific lineages. For example, MS5, S2 and PA6 stromal cells have

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been used to derive dopamine neurons from human ES cells (Perrier et al., 2004; Zeng et al., 2004); bone marrow stromal cell lines S17 and OP9 support efficient hematopoietic differentiation (Kaufman et al., 2001; Vodyanik et al., 2005). The inducing activity provided by such stromal cells, while efficient in directing ES cell differentiation, contains many unknown factors, and such activity can change both between and within cell lines as a function of culture conditions. An even more controlled method is differentiation in monolayers on defined matrices in the presence of specific growth factors. Both mouse and human ES cells differentiate into neuroectodermal precursors in monolayer culture (Ying et al., 2003b; Gerrard et al., 2005), and human ES cells can be efficiently induced to differentiate into trophoblasts with addition of BMPs (Xu et al., 2002). This method eliminates many unknown factors provided by either EBs or stromal cells, thus allowing precise analysis of specific factors on the differentiation of ES cells into lineages of choice. With improved understanding of regulatory events governing germ layer and cell lineage specifications, more cell types will likely be derived from ES cells in increasingly defined conditions. Molecular Control of Pluripotency We remain remarkably ignorant about why one cell is pluripotent and another is not, although some of the key players important to maintaining this remarkable state have been identified. Oct4, a member of the POU family of transcription factors, is essential for both the derivation and maintenance of ES cells (Pesce et al., 1998). The expression of Oct4 in the mouse is restricted to early embryos and germ cells (Scholer et al., 1989; Okamoto et al., 1990), and homozygous deletion of this gene causes a failure in the formation of the ICM (Nichols et al., 1998). For mouse ES cells to remain undifferentiated, the expression of Oct4 must be maintained within a critical range. Overexpression of this protein causes differentiation into endoderm and mesoderm, while decreased expression leads to differentiation into trophoblast (Niwa et al., 2000). The expression of Oct4 is also a hallmark of human ES cells (Hansis et al., 2000), and its down-regulation also leads to differentiation and expression of trophoblast markers (Matin et al., 2004). Another transcription factor important for the pluripotency of ES cells is Nanog (Chambers et al., 2003; Mitsui et al., 2003). Similar to Oct4, the expression of Nanog decreases rapidly as ES cells differentiate. However, unlike Oct4, overexpression of this protein in mouse ES cells allows their self-renewal to be independent of LIF/STAT3, though Nanog appears not to be a direct downstream target of LIF/STAT3 pathway (Chambers et al., 2003). In both mouse and human ES cells, reduced expression of Nanog causes differentiation into extra-embryonic lineages (Chambers et al., 2003; Mitsui et al., 2003; Hyslop et al., 2005). The expression of genes enriched in ES cells has been extensively studied by several groups (see for example Rao and Stice, 2004 and references therein), and includes, for example, transcription factors Sox2, FOXD3, RNA-binding protein Esg-1 (Dppa5), and de novo DNA methyltransferase 3b. Deletion of some of them in mice does demonstrate a critical function in early development (Table 12.1). ES cells also express high levels of genes involved in protein synthesis and mRNA processing (Richards et al., 2004), and non-coding RNAs unique to ES cells (Suh et al., 2004). A surprisingly high percentage of genes enriched in ES cells have unknown functions (Tanaka et al., 2002; Robson, 2004 and references therein). A recent genome-wide location analysis of human ES cells showed that Oct4 and Nanog, along with Sox2, co-occupy the promoters of a high number of genes, many of which are transcription factors such as Oct4, Nanog and Sox2 (Boyer et al., 2005). These three proteins, in addition to regulating their own transcription as previously shown (Catena et al., 2004; Kuroda et al., 2005; Okumura-Nakanishi et al., 2005; Rodda et al., 2005), could also activate or repress the expression of many other genes. These genome-wide approaches hold great promise in elucidating the networks that control the pluripotent state.

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Table 12.1 Examples of genes with enriched expression in ES cells Genes

Protein features and functions

References

Sox2

HMG-box transcription factor; interacts with Oct4 to regulate transcription; Sox2–/– mouse embryos died shortly after implantation with loss of epiblast at E6.0.

Avilion et al. (2003)

FOXD3

Forkhead family transcription factor; FoxD3–/– mouse embryos died shortly after implantation with loss of epiblast (E6.5); no FoxD3–/– ES cells can be established.

Hanna et al. (2002)

Rex-1(Zfp-42)

Zinc-finger transcription factor; direct target of Oct4; Rex-1–/– EC cells failed to differentiate into primitive and visceral endoderm.

(Rosfjord and Rizzino, 1994; Thompson and Gudas, 2002)

Gbx2(Stra7)

Homeobox-containing transcription factor; Gbx–/– embryos displayed defects in neural crest cell patterning and pharyngeal arch artery.

(Byrd and Meyers, 2005)

Sall1

Potent zinc-finger transcription repressor; heterozygous mutations in humans cause Townes-Brocks syndrome; Sall1–/– mice died perinatally.

Kiefer et al. (2002); Kohlhase et al. (1998); Nishinakamura et al. (2001)

Sall2

Homolog of Sall1; Sall–/– mice showed no phenotype.

Sato et al. (2003))

Hoxa11

Transcription factor; Hoxa11–/– mice showed defects in male and female fertility.

Hsieh-Li et al. (1995)

UTF1

Transcriptional co-activator; stimulate ES cell proliferation.

Nishimoto et al. (2005)

TERT

Reverse transcriptase (catalytic component of telomerase).

Liu et al. (2000)

TERF1

Telomere repeat-binding factor 1; TERF1–/– mouse embryos died at E5-6 with severe growth defect in ICM.

Karlseder et al. (2003)

TERF2

Telomere repeat-binding factor 2.

Sakaguchi et al. (1998)

DNMT3b

De novo DNA methyltransferase; required for methylation of centrimeric minor satellite repeats; DNMT3b–/– embryos died before birth.

Okano et al. (1999)

DNMT3a

De novo DNA methyltransferase; DNMT3a–/– mice died at age of 4 weeks.

Okano et al. (1999)

Dppa2

Putative DNA binding motif SAP.

Bortvin et al. (2003)

Dppa3 (PGC7, Stella)

Putative DNA binding motif SAP.

Bortvin et al. (2003); Bowles et al. (2003); Saitou et al. (2002); Sato et al., (2002)

Dppa4 (FLJ10713)

Putative DNA binding motif SAP.

Bortvin et al. (2003); Sperger et al. (2003)

Dppa5 (Ph34, Esg-1)

Similar to KH RNA-binding motif.

Astigiano et al. (1991); Tanaka et al. (2002))

ECAT11

Conserved transposase 22 domain.

Sperger et al. (2003)

(FLJ10884)

CONCLUSION Progress in developmental biology has been dramatic over the last few decades, and one of the legacies of the derivation of human ES cells is that they provide a compelling link between that progress and the understanding and treatment of human disease. The derivation of mouse ES cells in 1981 and subsequent development of

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homologous recombination revolutionized mammalian developmental biology, as it allowed the very specific modification of the mouse genome to test gene function. Yet although the use of mouse ES cells as an in vitro model of differentiation was established soon after their initial derivation, it was only after the derivation of human ES cells in 1998, and their potential use in transplantation medicine was immediately appreciated, that there was an explosion of interest in the in vitro, lineage-specific differentiation of ES cells. Significant progress has been made in lineage-specific differentiation of human ES cells, and progress in this area is accelerating as new groups are now rapidly entering this field. An understanding of the basic mechanism controlling germ layer and lineage specification is rapidly unfolding through the interplay of knock-out mice, in vitro differentiation of ES cells, and conserved mechanisms identified in other model organisms. The basic biology of pluripotency is another area of research that the isolation of human ES cells rekindled. Even though significant differences exist between mouse and human ES cells, they share many key genes involved in pluripotency, such as Oct4 and Nanog. Global gene expression analysis of mouse and human ES cells reveals the existence of many novel genes unique to ES cells, but the challenge remains in identifying functions of those genes, and coming to understand how the proliferative, pluripotent state is established and maintained. Indeed, although certain genes have been identified that are required to maintain the pluripotent state, it remains a central problem in biology to understand why one cell can form anything in the body and another cannot. Such a basic understanding has implications for regenerative medicine that go far beyond the use of ES cells in transplantation, and may lead to methods of causing tissues to regenerate that fail to do so naturally.

ACKNOWLEDGMENT James A. Thomson is a co-founder and shareholder of Cellular Dynamics, International, Madison, Wisconsin.

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13 Stem Cells Derived from Amniotic Fluid and Placenta Paolo De Coppi, Shay Soker, and Anthony Atala

INTRODUCTION Amniotic fluid and placenta have been recently taken into consideration as potential sources of progenitor cells. Amniocentesis and chorionic villus sampling (CVS) are widely accepted methods for prenatal diagnosis. Minimal or no ethical concerns would be present if embryonic and fetal stem cells would be taken from amniotic fluid and placenta before or at birth. In the last few years our and other groups have described the presence of stem cell with various differentiative and proliferative potential in the amniotic fluid and placenta. We will briefly describe the techniques in use for amniocentesis and CVS, and we will examine the different progenitors that have been described. CVS AND AMNIOCENTESIS The first reported amniocentesis took place in 1930 when attempts were being made to correlate the cytologic examination of cell concentration, count and phenotypes in the amniotic fluid to the sex and the health of the baby. Since then, the development technique of karyotype and the discovery of reliable diagnostic markers, such as alpha-fetoprotein, as well as the development of ultrasound-guided amniocentesis, have greatly increased the reliability of the procedure as a valid diagnostic tool as well as the safety of the procedure (Milunsky, 1979; Hoehn and Salk, 1982; Gosden, 1983; Crane and Cheung, 1988). One of the primary uses of amniocentesis is a safe method of isolating cells from the fetus that can be karyotyped and examined for chromosomal abnormalities. In general, the protocol consists of acquiring 10–20 ml of fluid using a transabdominal approach. Amniotic fluid samples are then centrifuged, and the cell supernatant is resuspended in culture medium. Approximately 104 cells are seeded on 22  22 mm cover slips. Cultures are grown to confluence for 3–4 weeks in 5% CO2 at 37°C, and the chromosomes are characterized from mitotic phase cells (Brace and Resnik, 1999). Amniocentesis is performed typically around 16 weeks of gestation, although in some cases it may be performed as early as 14 weeks when the amnion fuses with the chorion and the risk of bursting the amniotic sac by needle puncture is minimized. Amniocentesis can be performed as late as term. The amniotic sac is usually noticed first by ultrasound around the 10-week gestational time point. With the introduction of CVS in the 1980s, first-trimester diagnosis became a reality. A small sample of chorionic villi (tissue from the developing placenta) is obtained from the mother’s uterus under ultrasound guidance, either transvaginally or transabdominally. Sampling of chorionic villi from the fetus is performed from 10 weeks of gestation. The biopsy is usually taken under ultrasound guidance via a transabdominal approach. Alternatively, the cervical approach may be utilized. Each biopsy yields 5–30 mg of tissue that can be used for fetal sexing, karyotyping, biochemical studies and DNA analysis. A direct fetal chromosomal

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analysis on cultured cells is possible within 24 h. However, given the problem with mosaicism in CVS samples, this should always be followed by chromosomal analysis on cultured cells from the sample 2–3 weeks later. An advantage of CVS is that termination can be completed in the first trimester when it is technically easier.

DIFFERENTIATED CELLS FROM AMNIOTIC FLUID AND PLACENTA Amniotic fluid cell culture consists of a heterogeneous cell population displaying a range of morphologies and behaviors. Studies on these cells have characterized them into many shapes and sizes varying from 6 μm to 50 μmm in diameter and from round to squamous in shape. Most cells in the fluid are terminally differentiated along epithelial lineages and have limited proliferative and differentiation capabilities. Previous studies have noted an interesting composition of the fluid consisting of a heterogeneous cell population expressing markers from all three germ layers. (Sarkar et al., 1980; Cousineau et al., 1982; Medina-Gomez and Johnston, 1982; von Koskull et al., 1984). The source of these cells and of the fluid itself underwent a great deal of research. Current theories suggest that the fluid is largely derived from the urine and pulmonary secretion from the fetus as well as from some ultra filtrate from the plasma of the mother entering though the placenta. The cells in the fluid have been shown to be overwhelmingly from the fetus and are thought to be mostly cells sloughed off the epithelium and digestive and urinary tracts of the fetus as well as the amnion (Lotgering and Wallenburg, 1986; Underwood et al., 2005). MESENCHYMAL CELL FROM PLACENTA AND AMNIOTIC FLUID Preliminary studies have been published a few years ago describing very simple protocols for the isolation of a nonspecific population of cells with “mesenchymal” characteristics from amniotic fluid and placenta (Haigh et al., 1999; Kaviani et al., 2001, 2002, 2003). These cells were able to proliferate in vitro, to be engineered in a threedimensional structure and used in vivo to repair a tissue defect (Kaviani et al., 2003). A few years later In’t Anker et al. were able to prove for the first time that both amniotic fluid and placenta were abundant sources of fetal mesenchymal stem cells (MSCs) that exhibit a phenotype and multilineage differentiation potential similar to that of postnatal bone marrow (BM)-derived MSCs (In’t Anker et al., 2003). They described a simple and repeatable protocol for their isolation and expansion. Briefly, amniotic fluid samples were centrifuged for 10 min at 1,283 rpm. Pellets were resuspended in Iscove’s modified Dulbecco’s medium containing 2% fetal calf serum (FCS) and antibiotics (defined as washing medium). Similarly, for the placenta, approximately 1 cm3 was washed in phosphate-buffered saline (PBS) and single-cell suspensions were made by mincing and flushing the tissue parts through a 100 μm nylon filter with washing medium. Single-cell suspensions of amniotic fluid and placenta were plated in six-well plates and cultured in M199 supplemented with 10% FCS, 20 μg/ml endothelial cell growth factor, heparin (8 U/ml), and antibiotics. After 7 days, non-adherent cells were removed and the medium was refreshed. When grown to confluence, adherent cells were detached with trypsin/EDTA and expanded in culture flasks pre-coated with 1% gelatin and kept in a humidified atmosphere at 37°C. The expansion potency of fetal MSCs was higher compared with adult BM-derived MSCs. As a result, they were able to expand amniotic fluid MSCs to about 180  106 cells within 4 weeks (three passages). The phenotype of the culture-expanded amniotic fluid-derived cells was similar to that reported for MSCs derived from second-trimester fetal tissues and adult BM. They were able to show that amniotic fluid-derived MSCs showed multilineage differentiation potential into fibroblasts, adipocytes, and osteocytes (In’t Anker et al., 2004). Furthermore, amniotic fluid-derived MSCs were successfully isolated, cultured, and enriched without interfering with the routine process of fetal karyotyping. Flow cytometry analyses showed that they were positive for SH2, SH3, SH4, CD29 and CD44, low positive for CD90 and CD105, but negative for CD10, CD11b, CD14, CD34, CD117, and EMA (Tsai et al., 2004). Most importantly, immuno-phenotypic analyses demonstrated that these cells expressed HLA-ABC, class I major histocompatibility complex (MHC-I), but they did not express HLA-DR, DP, and DQ (MHC-II molecules) (Li et al., 2005a). Li et al.

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have extensively investigated their immunological role and described that when mononucleated cells recovered from placentas by density gradient fractionation were added to umbilical cord blood (UCB) lymphocytes stimulated by human adult lymphocytes or potent T-cell mitogen phytohemagglutinin, a significant reduction in lymphocyte proliferation was observed. This immunoregulatory feature strongly implies that they may have potential application in allograft transplantation. As it is possible to obtain placenta and UCB from the same donor, they suggested the placenta as an attractive source of MSCs for co-transplantation in conjunction with UCB-derived hematopoietic stem cells (HSCs) to reduce the potential graft-versus-host disease (GVHD) in recipients (Li et al., 2005b). Other methods have been described for the isolation of mesenchymal cell from human placenta. Dissection and proteinase digestion are alternative techniques to harvest high numbers of viable mononuclear cells from human placenta at term, and a mesenchymal cell population with characteristic expression of CD9, CD29, and CD73 could be obtained in culture. The in vitro growth behavior of such placenta-derived mesengenic cells was similar to that of human BM mesengenic progenitor cells. Differentiation experiments showed differentiation potential along osteogenic, chondrogenic, adipogenic, and myogenic lineages (Figure 13.1). However, after in vitro propagation for more than three passages, the cells were exclusively of maternal origin (Wulf et al., 2004). Similar cells isolated from term placenta were described by Yen et al. They exhibited many markers common to mesenchymal stem cells – including CD105/endoglin/SH-2, SH-3, and SH-4 – and they lack hematopoietic-, endothelial-, and trophoblastic-specific cell markers. In addition, they exhibit embryonic stem (ES) cell surface markers of SSEA-4, TRA-1-61, and TRA-1-80. Adipogenic, osteogenic, and neurogenic differentiation were achieved after culturing under the appropriate conditions (Yen et al., 2005). Mesenchymal cells were also isolated from placentas collected after neonatal delivery (38–40 weeks of gestation). The cells expressed CD13, CD44, CD73, CD90, CDIO5, and HLA class I as surface epitopes, but not CD31, CD34, CD45, and HLA-DR, differentiated into osteocytes, chondrocytes, and adipocytes under specific culture conditions, and were also induced to form neural-like cells (Fukuchi et al., 2004; Igura et al., 2004). Different types of tissue were obtained by in vivo implantation of the cells. Hepatic

Endothelial

Myogenic

Progenitor

Neuronal

Adipogenic

Osteogenic

Figure 13.1 The isolated progenitor cells were capable of differentiation into multiple cell types, including muscle, liver, endothelial cells, adipocytes, osteoblasts, and neurons.

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Several studies suggested that human amniotic fluid and placenta-derived mesenchymal cells can be one of the possible allogeneic cell sources for tissue engineering of cartilage. In particular, Zhang et al. (2005) showed the possibility to make a cartilage-like tissue embedding mesenchymal stem cells derived from human placenta, into an atelocollagen gel with chondrogenic induction media. The in vitro pre-induced cells were implanted into nude mice and also into nude rats with osteochondral defect, and they were able to form chondrogenic structures. Ovine mesenchymal amniocytes have also been cultured and engineered into a collagen hydrogel in order to replace partial diaphragmatic loss or absences. The authors showed that diaphragmatic repair with an autologous tendon engineered from mesenchymal amniocytes leads to improved mechanical and functional outcomes when compared with an equivalent acellular bioprosthetic repair, depending on scaffold composition (Fuchs et al., 2004). Different groups have claimed that mesenchymal cells from placenta and amniotic fluid could have more plasticity than what initially thought. Phenotypic and gene expression studies indicated mesenchymal stem cell-like profiles in both amnion and chorion cells that were positive for neuronal, pulmonary, adhesion, and migration markers. In addition, transplantation in neonatal swine and rats resulted in human microchimerism in various organs and tissues, suggesting that amnion and chorion cells may represent an advantageous source of progenitor cells with potential applications in a variety of cell therapy and transplantation procedures (Bailo et al., 2004). Similarly, Zhao et al. have reported that human amniotic mesenchymal cells (hAMC) may also be a suitable cell source for cardiomyocytes. He showed that freshly isolated hAMC expressed cardiac-specific transcription factor GATA4, cardiac-specific genes, such as myosin light chain (MLC)-2a, MLC-2v, cTnI, and cTnT, and the alphasubunits of the cardiac-specific L-type calcium channel (alpha1c). After stimulation with basic fibroblast growth factor (bFGF) or activin A, hAMC expressed Nkx2.5, a specific transcription factor for the cardiomyocyte and cardiac-specific marker atrial natriuretic peptide. In addition, the cardiac-specific gene alpha-myosin heavy chain was detected after treatment with activin A. Co-culture experiments confirmed that hAMC were able to both integrate into cardiac tissues and differentiate into cardiomyocyte-like cells. After transplantation into the myocardial infarcts (AMI) in rat hearts, hAMC survived in the scar tissue for at least 2 months and differentiated into cardiomyocyte-like cells (Zhao et al., 2005). However, we have recently shown that this potential does not belong to mesenchymal progenitor cells in bigger animals, such as pigs. Amniotic fluid-derived mesenchymal cells (AFC) autotransplanted in a porcine model of AMI were able to transdifferentiate to cells of vascular cell lineages but failed to give origin to cardiomyocytes (Sankar and Muthusamy, 2003; Sartore et al., 2005). Regarding neuron regeneration, it has been shown that rat amniotic epithelial (RAE) cells were positive in vitro for both neuronal and neural stem cell markers, neurofilament microtubule-associated protein 2, and nestin. RT-PCR revealed that these cells expressed nestin mRNA. The RAE cells were also transplanted into the hippocampus of adult gerbils that were subjected to temporal occlusion of bilateral carotid arteries. Five weeks after transplantation, grafted cells migrated into the CA1 pyramidal layer that showed selective neuronal death, and survived in a manner similar to CA1 pyramidal neurons (Okawa et al., 2001). Different reports suggest that human amniotic epithelial cells (HAEC) also possess certain properties similar to that of neural and glial cells (Tsai et al., 2005). When transplanted into the transection cavities in the spinal cord of bonnet monkeys, HAEC were able to survive, support the growth of host axons through them, prevent the formation of glial scar at the cut ends and may prevent death in axotomized cells or attract the growth of new collateral sprouting (Okawa et al., 2001). Amniotic epithelial cells isolated from human term placenta express surface markers normally present on ES and germ cells. In addition, they express the pluripotent stem cell-specific transcription factors octamer-binding protein 4 (Oct-4) and nanog. Under certain culture conditions, amniotic epithelial cells form spheroid structures that retain stem cell characteristics. Amniotic epithelial cells did not require other cell-derived feeder layers to maintain Oct-4 expression, did not express telomerase, and are non-tumorigenic upon transplantation. Based on immunohistochemical and genetic analysis, amniotic epithelial cells had the potential to differentiate to all three germ layers – endoderm (liver, pancreas), mesoderm (cardiomyocyte), and ectoderm (neural cells) in vitro (Miki et al., 2005). Sarkar et al.

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(2003) have also shown that HAEC obtained from human placenta were able to survive into the transection cavities in the spinal cord of bonnet monkeys, to support the growth of host axons through them, to prevent the formation of glial scar at the cut ends and may prevent death in axotomized cells or attract the growth of new collateral sprouting. They speculated that HAEC may be having certain properties equal to the beneficial effects of neural tissue in repairing spinal cord injury. Apart from this speculation, there are two more reasons for why HAEC transplantation studies are warranted to understand the long-term effects of such transplantations. First, there was no evidence of immunological rejection probably due to the non-antigenic nature of the HAEC. Second, unlike neural tissue, procurement of HAEC does not involve many legal or ethical problems (Sakuragawa et al., 1996, 2000; Elwan and Sakuragawa, 1997; Takahashi et al., 2002).

PLURIPOTENT STEM CELLS FROM PLACENTA AND AMNIOTIC FLUID In the midgestation murine embryo, several major vascular tissues contain pluripotent stem cells, mainly defined for their HSC activity. These include the aorta-gonad-mesonephros (AGM) region, yolk sac, and fetal liver. Recently, different reports have shown that the mouse placenta functions as a hematopoietic organ that harbors a large pool of pluripotent HSCs during midgestation (Alvarez-Silva et al., 2003; Ottersbach and Dzierzak, 2005). The onset of HSC activity in the placenta parallels that of the AGM region starting at E10.5–E11.0. However, the placental HSC pool expands until E12.5–E13.5 and contains 15-fold more HSCs than the AGM (Gekas et al., 2005). Placental HSC activity starts before HSCs are found in circulation or have colonized the fetal liver. Moreover, hematopoietic cells in midgestation mouse placenta are not instructed for differentiation along the myeloerythroid lineage, as in the fetal liver. These findings suggest that the placenta provides a supportive niche where the definitive HSC pool can be temporarily established during development. Furthermore, if the stem cell-promoting properties of the placental niche can be harnessed in vitro to support HSC formation, maturation, and/or expansion in culture, these assets may greatly improve HSCbased therapies in the future (Mikkola et al., 2005). A part of HSCs, the presence of pluripotent stem cells, similar to ES cells, have been pointed out by us and others. Oct-4 is a marker for pluripotent human stem cells so far known to be expressed in embryonal carcinoma cells, ES cells, and embryonic germ cells. Performing RT-PCR, Western blot, and immunocytochemical analyses it has been evident that in human amniotic fluid in the background of Oct-4-negative cells, a distinct population of cells can be found, which express Oct-4 in the nucleus. Oct-4-positive amniotic fluid cell samples also expressed stem cell factor, vimentin, and alkaline phosphatase mRNA. The Oct-4-positive amniotic fluid cells were actively dividing, proven by the detection of cyclin A expression. They suggested that human amniotic fluid could represent a new source for the isolation of human Oct-4-positive stem cells without raising the ethical concerns associated with human embryonic research (De Coppi et al., 2001, 2002; Prusa et al., 2003; Karlmark et al., 2005). Established cell lines derived from human placenta by cloning technique using alpha-MEM culture medium containing 10 ng/ml of EGF (epidermal growth factor), 10 ng/ml of hLIF, and 10% FBS (fetal bovine serum) appeared to maintain a normal karyotype indefinitely in vitro and expressed markers characteristic of stem cells from mice and human, namely alkaline phosphatase. These cells contributed to the formation of chimeric mouse embryoid bodies and gave rise to cells of all germ layers in vitro (Tamagawa et al., 2004). Koegler et al. have also described a new pluripotent human somatic stem cell pluripotent, CD45-negative population from human cord blood, termed unrestricted somatic stem cells (USSCs). This rare population grows adherently and can be expanded to 1015 cells without losing pluripotency. In vitro USSCs showed homogeneous differentiation into osteoblasts, chondroblasts, adipocytes, and hematopoietic and neural cells including astrocytes and neurons that express neurofilament, sodium channel protein, and various neurotransmitter phenotypes. Stereotactic implantation of USSCs into intact adult rat brain revealed that human Tau-positive cells persisted for up to

Stem Cells Derived from Amniotic Fluid and Placenta 231

3 months and showed migratory activity and a typical neuron-like morphology. In vivo differentiation of USSCs along mesodermal and endodermal pathways was demonstrated in animal models. Bony reconstitution was observed after transplantation of USSC-loaded calcium phosphate cylinders in nude rat femurs. Chondrogenesis occurred after transplanting cell-loaded gelfoam sponges into nude mice. Transplantation of USSCs in a non-injury model, the pre-immune fetal sheep, resulted in up to 5% human hematopoietic engraftment. More than 20% albumin-producing human parenchymal hepatic cells with absence of cell fusion and substantial numbers of human cardiomyocytes in both atria and ventricles of the sheep heart were detected many months after USSC transplantation. No tumor formation was observed in any of these animals (Kogler et al., 2004). We have recently described a pluripotent population of cells derived from both amniotic fluid and placenta. We will describe in the following paragraphs in detaiòs their isolation, characterization, and differentiation in vitro into different lineages (De Coppi et al., 2007). Isolation and Characterization of Chorionic Villi and Amniotic-Derived Stem Cells Chorionic villi samples and human amniotic fluid were obtained under informed consent at 12–18 weeks of pregnancy from a total of 300 women between 23 and 42 years of age. In all cases the karyotype evaluated from the cultured chorionic villi and amniotic fluid cells was normal. Samples were seeded in a 22  22 mm cover slip in a volume of 2 ml and grown to confluence for 3–4 weeks at 95% humidity and 37°C. Fresh medium was applied after 5 days of culture and every third day thereafter. The culture medium consisted of alpha-MEM (GIBCO/BRL, Grand Island, NY), 18% Chang medium B (Irvine Scientific, Santa Ana, CA), 2% Chang C (Irvine Scientific, Santa Ana, CA) with 15% ES cell-certified FBS (ES-FBS, GIBCO/BRL, Grand Island, NY), 1% antibiotics (GIBCO/BRL, Grand Island, NY), and L-glutamine (Sigma-Aldrich, St. Louis, MO). The cells were subcultured using 0.25% trypsin containing 1 mM EDTA for 5 min at 37°C. In order to test the hypothesis that placenta and amniotic fluid could contain stem cells that would be able to differentiate into multiple lineages, cell colonies derived from single cells were expanded. The cells were successfully isolated from 300 fetuses and maintained in culture in Chang medium. The presence of cells of maternal origin in placenta and amniotic fluid is extremely low. In order to evaluate for the presence of maternal cells, the studies were performed using cells from male fetuses. Karyotypic analyses of the ckit pos cells showed an xy phenotype in all the cells. Female fetuses were used as controls and they did not show any difference in their pluripotential ability. Cytofluorimetric analysis and immunocytochemistry showed that most of the amniotic cells were epithelial and stained positive for cytokeratins. Most of the stromal cells stained for alpha-actin, and only a few cells were positive for desmin or myosin. Fluorescence-activated cell sorter (FACS) analyses showed that between 18% and 21% of the cells expressed CD105, while approximately the same proportion of cells (between 0.8% and 3%) expressed ckit and CD34. The ckit pos cells were successfully isolated and maintained in culture in Chang medium. The ckit pos cells were shown to be pluripotent. They maintained a round shape when cultured in bacterial plates for almost a week while they had a very low proliferative capability. After the first week the cells started to adhere to the plate and changed their morphology, becoming more elongated, and they started to proliferate. The medium was changed every 3 days and they were passed whenever they reached confluence. If the cells were not passaged, they aggregated, forming embryoid-shaped tissue-like structures measuring 1  5 mm3. Serial sections of these structures showed specific markers for the three embryonic germ layers immunohistochemically. The embryo-shaped tissue, if disaggregated, was still able to differentiate into different lineages under appropriate growth conditions. The CD105, CD90, and CD34 immunoseparated cells, and the remaining non-immunoseparated cells did not show any pluripotential ability. No feeder layers were required, either for maintenance or expansion (Takeda et al., 1992; Mosquera et al., 1999).

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Telomerase Activity Telomerase activity was evaluated using the telomerase repeat amplification protocol (TRAP) assay, and the presence of telomerase was analyzed immunocytochemically. No telomerase activity could be detected with the TRAP assay, either in the ckit pos cells (lanes1) or in the control BM stem cells (lane 2). In contrast, the prostate cancer cell line PC3 and an epithelial tumor cell line (HeLa), as a control, showed high telomerase activity (lanes 4 and 5). Anti-telomerase antibodies positively stained the amniotic ckit pos cells, suggesting that the cells may express telomerase protein, but the levels were too low to be detected by the TRAP assay. Differentiation Potential Induction of Osteogenic Phenotype (Figure 13.2a) Light microscopy analysis showed that ckit pos cells, within 4 days in osteogenic medium, developed an osteoblasticlike appearance with finger-like excavations into the cells’ cytoplasm (Karsenty, 2000; Olmsted-Davis et al., 2003). At 16 days the cells aggregated in the typical lamellar bone-like structures and increased their expression of alkaline phosphatase. Ca accumulation was evident after 1 week and increased over time. To confirm the cytochemical findings, AP activity was measured using a quantitative assay, which measured p-nitrophenol, equivalent to AP production. The ckit pos cells showed more than a two hundred time increase in AP production in the osteogenic-inducing medium compared to cells grown in control medium at days 16 and 24. After that time the levels of AP decreased. No AP production was detected in ckit neg amniotic cells cultured in osteogenic medium at any time point. AP expression was confirmed at the RNA level. No activation of the AP gene was detected at 8, 16, 24, and 32 days in the ckit pos cells grown in control medium. In contrast, ckit pos cells grown in osteogenic medium showed an activation of the AP gene at each time point. Expression of cbfa1, a transcription factor specifically expressed in osteoblasts and hypertrophic chondrocytes, was highest in cells grown in osteogenic-inducing medium at day 8, and decreased slightly at days 16, 24, and 32. The expression of cbfa1 in the controls was significantly lower at each time point. Osteocalcin was expressed only in the ckit pos cells in osteogenic conditions at 8 days. No expression of ostecalcin was detectedable in the ckit pos cells in the control medium and in the ckit neg cells in the osteogenic conditions at any time point. A major feature of osteogenic differentiation is the ability of the cells to precipitate calcium. Cell-associated mineralization can be analyzed using von Kossa staining and by measuring the calcium content of cells in culture. Von Kossa staining of cells grown in the osteogenic medium showed enhanced silver nitrate precipitations by day 16, indicating high levels of calcium. Calcium precipitation continued to increase exponentially at 24 and 32 days. In contrast, cells in the control medium did not form silver nitrate precipitations after 32 days. Microscopic examination of stained cells showed no calcification in the osteogenic treated cells at day 4 or 8, but strong black silver nitrate precipitates were noticed in osteogenic-induced cells after 16, 24, and 32 days in culture. In cells cultured in control medium, no precipitates were noticed over the 32-day time period. Calcium deposition by the cells was also measured with a quantitative chemical assay, which measures calcium–cresolophthalein complexes. Cells undergoing osteogenic induction showed a significant increase in calcium precipitation after 16 days (up to 4 mg/dl). The precipitation of calcium increased up to 70 mg/dl at 32 days. In contrast, cells grown in control medium did not show any increase in calcium precipitation (1.6 mg/dl) by day 32. Induction of Adipogenic Lineage (Figure13.2b) Ckit pos cells cultured in a medium containing dexametasone, insulin, indomethacin, and 3-isobutyl-1methylxanthine, within 8 days, changed their morphology from elongated to round (Kim et al., 1998). This coincided with the accumulation of intracellular triglyceride droplets. After 16 days in culture, more than 95% of the cells had their cytoplasm completely filled with lipid-rich vacuoles, which stained positively with Oil-O-Red. The amniotic ckit neg cells that were induced with the same medium and the ckit pos cells cultured in control medium did not show any phenotypic change of adipogenic differentiation and did not stain with Oil-O-Red after 16 days of culture. Adipogenic differentiation was confirmed by RT-PCR analysis. The expression of peroxisome proliferation-activated receptor 2 (ppart(2)), a transcription factor that

Stem Cells Derived from Amniotic Fluid and Placenta 233

regulates adipogenesis and of lipoprotein lipase was analyzed. Expression of these genes was upregulated in the ckit pos cells under adipogenic conditions. Ckit pos cells cultured under control conditions and ckit neg cells in adipogenic medium did not express either gene at any time point. Induction of Myogenic Phenotype (Figure 13.2c) Ckit pos cells were cultured with myogenic medium on Matrigel-coated dishes (Rosenblatt et al., 1995; Ferrari et al., 1998). Induction with 5-azacytidine for 24 h promoted the formation of multinucleated cells over a 24–48 h period. After 16 days, the cells grown with myogenic medium formed myofiber-like structures that stained immunocytochemically with desmin and sarcomeric tropomyosin. Ckit pos cells grown in control medium and ckit neg cells cultured in myogenic medium did not lead to cell fusion or multinucleated cells. Only a few desmin cells were present in the ckit neg amniotic cells cultured in myogenic medium at 16 days. Expression of MyoD, Myf5, Myf 6 (MRF4), and desmin were analyzed using RT-PCR. MyoD and MRF4 were expressed by the ckit pos cells in culture at 8 days and suppressed at 16 days. Both these genes were not expressed either at 8 or 16 days in the controls. Desmin expression was induced at 8 days and increased by 16 days in the ckit pos cells cultured in myogenic medium. In contrast, there was no activation of desmin in the control cells at 8 and 16 days. Myf5 was present at 8 days and increased at 16 days in the ckit pos cells. Lower levels of the Myf5 gene were detected in the cells maintained in culture with the control medium at 16 days. Induction of Endothelial Phenotype (Figure 13.2d) Ckit pos cells were cultured with endothelial medium in PBS–gelatin-coated dishes. After 1 week in culture the cells started to change their morphology, and by the second week, were mostly tubular. The cells stained positively for FVIII, KDR, and P1H12. Ckit neg cells cultured in the same conditions and ckit pos cells cultured in Chang medium for the same period were not able to form tubular structures and did not stain for endothelial specific markers. The cells, once differentiated, were able to grow in culture for more than 1 month. Induction of Hepatocytes Phenotype (Figure 13.2e) When cultured in hepatic conditions cells exhibited morphological changes after 7 days showing a change in the morphology from an elongated to a cobblestone appearance (Dunn et al., 1989; Hamazaki et al., 2001). The cells showed positive staining for albumin at day 45 post-differentiation, and were also found to express transcription factor HNF4α, c-met receptor, multidrug resistance gene (MDR) membrane transporter, albumin, and alphafetoprotein. RT-PCR analysis further provided evidence of albumin production. The maximum rate of urea production for hepatic differentiation induced cells was 1.21  103 ng urea/h/cell as opposed to 5.0  101 ng urea/h/cell for control progenitor cell populations. Induction of Neurogenic Phenotype (Figure 13.2f) Ckit pos cells cultured in neurogenic conditions changed their morphology within the first 24 h (Black and Woodbury et al., 2001; Barberi et al., 2003). Responsive cells progressively assumed neuronal morphological characteristics; initially the cytoplasm retracted toward the nucleus, forming contracted multipolar structures. Over the subsequent hours, the cells displayed primary and secondary branches, and cone-like terminal expansions. Induced ckit pos cells stained positively for beta-III tubulin and nestin. Ckit neg cells cultured in the same conditions and ckit pos cells cultured in Chang medium for the same period were not able to form tubular structures and did not stain for endothelial specific markers. The cells, once differentiated, were able to grow in culture for more than 1 month. Clonal and Proliferative Analyses Ckit pos cells were able to be expanded clonally. After serial dilution we observed that most of the wells contained no cells, and only a few of the 96 wells contained a single cell. Cells from numerous clones showed

234 CELLS AND TISSUE DEVELOPMENT

(b)

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du 4 dced ays ind uc 4 d ed ays no t in du ce 6d d ay ind s uc 6 d ed he ays at ina ctiv ind ated u 6 dced ays uro the lium

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Figure 13.2 The differentiated cell types expressed functional and biochemical characteristics of the target tissue. (a) Osteogenic-induced progenitor cells showed a significant increase of calcium deposition starting at day 16 (solid line). No calcium deposition was detected in the progenitor cells grown in control medium or in the negative control cells grown in osteogenic conditions (dashed line). RT-PCR showed presence of cbfa1 and osteocalcin at day 8 and confirmed the expression of AP in the osteogenic-induced cells. (b) Gene expression of ppar and lipoprotein lipase in cells grown in adipogenic-inducing medium was noted at days 8 and 16 (lanes 3 and 4). (c) Myogenic-induced cells showed a strong expression of desmin expression at day 16 (lane 4). MyoD and MRF4 were induced with myogenic treatment at day 8 (lane 3). Specific PCR-amplified DNA fragments of MyoD, MRF4, and desmin could not be detected in the control cells at days 8 and 16 (lanes 1 and 2). (d) RT-PCR of progenitor cells induced in endothelial medium (lane 2) showed the expression of CD31 and VCAM. (e) RT-PCR revealed an upregulation of albumin gene expression. Western blot analyses of cell lysate showed the presence of the hepatic lineage-related proteins HNF-4, c-met, MDR, albumin, and alpha-fetoprotein. Undifferentiated cells were used as negative control. (f) Only the progenitor cells cultured under neurogenic conditions showed the secretion of glutamic acid in the collected medium. The secretion of glutamic acid could be induced (20 min in 50-mM KCl buffer).

a similar morphology and growth behavior. Clonal lineages from different patients were tested. All the cells underwent osteogenic, adipogenic, myogenic, neurogenic, and endothelial differentiation. Amniotic stem cells did not show any decrease in their growth ability after more than 100 cell divisions, and they maintained their ability to differentiate into different lineages.

Stem Cells Derived from Amniotic Fluid and Placenta 235

FUTURE DIRECTION Fetal tissue has been used in the past for transplantation and tissue engineering research because of its pluripotency and proliferative ability. Fetal cells maintain a higher capacity to proliferate than adult cells and may preserve their pluripotency longer in culture. However, fetal cell transplants are plagued by problems that are very difficult to overcome. Beyond the ethical concerns regarding the use of cells from aborted fetuses or living fetuses, there are other issues which remain a challenge. Previous studies have shown that it takes almost six fetuses to provide enough material to treat one patient with Parkinson’s disease. In this study we hypothesized that placental and amniotic cells, which have been used for decades for prenatal diagnosis, could represent a viable source of fetal stem cells that could be used therapeutically. SUMMARY It is well known that placenta and amniotic fluid contain a large variety of cells. Our aim was to try to identify and isolate cells that still maintained their pluripotential and proliferative abilities. The vast majority of the cells in the placenta and in the amniotic fluid are already differentiated, and, therefore, have a limited proliferative ability. In this study the ckit pos cells were induced to different lineages. The ability to induce specific differentiation was initially evident by morphological changes, and was confirmed by immunocytochemical and gene expression analyses. In conclusion, placenta and amniotic fluid could be an excellent cell source for therapeutic applications. Fetal stem cells have a better potential for expansion than adult stem cells and for this reason they could represent a better source for any therapeutic application where large numbers of cells are needed. When compared with ES cells, fetal stem cells are easily differentiated into specific cell lineages, do not need any feeder layer to grow, and avoid the current controversies associated with the use of human ES cells.

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Mikkola, H.K., Gekas, C., Orkin, S.H. and Dieterlen-Lievre, F. (2005). Placenta as a site for hematopoietic stem cell development. Exp. Hematol. 33(9): 1048–1054. Milunsky, A. (1979). Amniotic fluid cell culture. In: Milunsky, A. (ed.), Genetic Disorder of the Fetus. New York: Plenum Press, pp. 75–84. Mosquera, A., Fernandez, J.L., Campos, A., et al. (1999). Simultaneous decrease of telomerase length and telomerase activity with ageing of human amniotic fluid cells. J. Med. Genet. 36: 494–496. Okawa, H., Okuda, O., Arai, H., Sakuragawa, N. and Sato, K. (2001). Amniotic epithelial cells transform into neuron-like cells in the ischemic brain. Neuroreport 12(18): 4003–4007. Olmsted-Davis, E.A., et al. (2003). Primitive adult hematopoietic stem cells can function as osteoblast precursors. Proc. Natl Acad. Sci. USA 100: 15877–15882. Ottersbach, K. and Dzierzak, E. (2005). The murine placenta contains hematopoietic stem cells within the vascular labyrinth region. Dev. Cell 8(3): 377–387. Prusa, A.R., Marton, E., Rosner, M., Bernaschek, G. and Hengstschlager, M. (2003). Oct-4-expressing cells in human amniotic fluid: a new source for stem cell research? Hum. Reprod. 18(7): 1489–1493. Rosenblatt, J.D., Lunt, A.I., Parry, D.J. and Partridge, T.A. (1995). Culturing satellite cells from living single muscle fiber explants. In Vitro Cell Dev. Biol. Anim. 31: 773–779. Sakuragawa, N., Thangavel, R., Mizuguchi, M., et al. (1996). Expression of markers for both neuronal and glial cells in human amniotic epithelial cells. Neurosci. Lett. 209: 9–12, 23. Sakuragawa, N., Enosawa, S., Ishii, T., et al. (2000). Human amniotic epithelial cells are promising transgene carriers for allogeneic cell transplantation into liver. J. Hum. Genet. 45: 171–176. Sankar, V. and Muthusamy, R. (2003). Role of human amniotic epithelial cell transplantation in spinal cord injury repair research. Neuroscience 118(1): 11–17. Sarkar, S., Chang, H.C., Porreco, R.P. and Jones, O.W. (1980). Neural origin of cells in amniotic fluid. Am. J. Obstet. Gynecol. 136(1): 67–72. Sartore, S., Lenzi, M., Angelini, A., Chiavegato, A., Gasparotto, L., De Coppi, P., Bianco, R. and Gerosa, G. (2005). Amniotic mesenchymal cells autotransplanted in a porcine model of cardiac ischemia do not differentiate to cardiogenic phenotypes. Eur. J. Cardiothorac. Surg. 28(5): 677–684. Takahashi, N., Enosawa, S., Mitani, T., et al. (2002). Transplantation of amniotic epithelial cells into fetal rat liver by in utero manipulation. Cell Transplant. 11: 443–449. Takeda, J., Seino, S. and Bell, G.I. (1992). Human Oct-3 gene family: cDNA sequences, alternative splicing, gene organization, chromosomal location, and expression at low levels in adult tissues. Nucl. Acid Res. 20: 4613–4620. Tamagawa, T., Ishiwata, I. and Saito, S. (2004). Establishment and characterization of a pluripotent stem cell line derived from human amniotic membranes and initiation of germ layers in vitro. Hum. Cell 17(3): 125–130. Tsai, M.S., Lee, J.L., Chang, Y.J. and Hwang, S.M. (2004). Isolation of human multipotent mesenchymal stem cells from second-trimester amniotic fluid using a novel two-stage culture protocol. Hum. Reprod. 19(6): 1450–1456. Tsai, M.S., Hwang, S.M., Tsai, Y.L., Cheng, F.C., Lee, J.L. and Chang, Y.J. (2005). Clonal amniotic fluid-derived stem cells express characteristics of both mesenchymal and neural stem cells. Biol. Reprod. Underwood, M.A., Gilbert, W.M. and Sherman, M.P. (2005). Amniotic fluid: not just fetal urine anymore. J. Perinatol. 25(5): 341–348. von Koskull, H., Aula, P., Trejdosiewicz, L.K. and Virtanen, I. (1984). Identification of cells from fetal bladder epithelium in human amniotic fluid. Hum. Genet. 65(3): 262–267. Wulf, G.G., Viereck, V., Hemmerlein, B., Haase, D., Vehmeyer, K., Pukrop, T., Glass, B., Emons, G. and Trumper, L. (2004). Mesengenic progenitor cells derived from human placenta. Tissue Eng. 10(7–8): 1136–1147. Yen, B.L., Huang, H.I., Chien, C.C., Jui, H.Y., Ko, B.S., Yao, M., Shun, C.T., Yen, M.L., Lee, M.C. and Chen, Y.C. (2005). Isolation of multipotent cells from human term placenta. Stem Cell 23(1): 3–9. Zhang, X., Mitsuru, A., Igura, K., Takahashi, K., Ichinose, S., Yamaguchi, S. and Takahashi, T.A. (2005). Mesenchymal progenitor cells derived from chorionic villi of human placenta for cartilage tissue engineering. Biochem. Biophys. Res. Commun. 340(3): 944–952. Zhao, P., Ise, H., Hongo, M., Ota, M., Konishi, I., Nikaido, T. (2005). Human amniotic mesenchymal cells have some characteristics of cardiomyocytes. Transplantation 79(5): 528–535.

14 Stem Cells Derived from Cord Blood Julie G. Allickson

INTRODUCTION Cord blood was first seen as biological waste product post childbirth. One of the first publications reported on the colony-forming capacity of cord blood summarizing its cloning efficiency to be similar to bone marrow was reported in 1980 (Di Landro et al., 1980). In 1988, the first cord blood transplant took place in France for Fanconi’s anemia with the donor being an identical human leukocyte antigen (HLA)-matched sibling (Gluckman et al., 1989). The transplant was successful without graft versus host disease (GVHD) (Gluckman et al., 2005) and the patient is reported to be alive and well 18 years after the transplant (Kurtzberg, personal communication). In 1989, Broxmeyer reported on the colony-forming potential of umbilical cord blood assessing its growth for colony-forming granulocyte-macrophage (CFU-GM), burst-forming erythroid (BFU-E), and colony-forming capacity for multipotent progenitors (CFU-GEMM) the most immature assessed. In 1990, it was reported that three patients had been transplanted for Fanconi’s anemia and it was suggested that cord blood transplantation maybe applicable to other diseases with a possibility of also transplanting adults. The cord blood cellular product viewed as a source of hematopoietic progenitors cells coupled to the immaturity of the immune system at birth is one of the advantages of using these cells for transplantation (Gluckman et al., 1990). In 1990, Thierry et al. reported difficulty in processing the procured cord blood in regards to cell recovery. It was also discovered that the total stem cell content correlated significantly with the time of delivery; the earlier in gestation the cord blood was collected the higher the number of stem cells retrieved (Thierry et al., 1990). In 1991 was the first report of a Public Cord Blood Bank for unrelated cord blood transplants (Rubinstein et al., 1993). One of the first reports in 1992 on the characterization of cord blood by flow cytometry was reported by Dr. Gluckman’s Laboratory to demonstrate that the content of the cord blood graft represented both suppressive and naive cells. Naive cells were noted by the T-cell content and its ability to produce receptors for interleukin (IL)-2 and HLA-DR6 (Rabian-Herzog et al., 1992). In 1994, researchers investigated the incidence of maternal cell contamination in the cord blood. It was demonstrated that only rarely are they discovered at birth and at an extremely low percentage as displayed in the lymphocyte population which was less than 1% (Socie et al., 1994). In 2000, Rocha reported a lower risk of acute and chronic GVHD in cord blood as compared to bone marrow in HLA-matched identical sibling transplants (Rocha et al., 2000). He was able to demonstrate the colony-forming capacity of cord blood which would remain viable 3 days after procurement if stored at 4°C or at room temperature, but not at 37°C (Broxmeyer et al., 1989). In this chapter, cord blood procurement, processing, and storage are briefly reviewed. The pluripotent capabilities of the umbilical cord blood stem cells have recently demonstrated differentiation potential in all

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three germ cell layers referring to the ectoderm, the endoderm, and the mesoderm. Researchers have studied the potential of differentiating in vivo and in vitro to not only characterize the cell, but to test its proliferative and clonogenic capacities. The most recent investigations will be discussed and summarized relating to the current efforts in the field of umbilical cord blood transplantation as it relates to regenerative medicine.

PROCUREMENT AND PROCESSING OF CORD BLOOD Cord blood procurement is generally performed by a health care professional or trained staff of the cord blood bank where the product will be processed and stored. The postnatal collection occurs after the cord blood vein is disinfected similar to a process used for whole blood collection utilizing a combination of iodine and alcohol through a series of steps. The procurement most commonly is harvested in a bag with citrate phosphate dextrose (CPD), acid citrate dextrose (ACD), or heparin as the anticoagulant, but may also be collected via a syringe with anticoagulant. Procurement of cord blood may be initiated while the placenta is in utero or ex utero. Most publications to date have shown statistically similar results by comparing methodology (Lasky et al., 2002; Pafumi et al., 2002; Solves et al., 2006), but some studies demonstrated the in utero collection yielded a higher recovery of hematopoietic cell content (Solves et al., 2003a, b; 2005). The collection usually takes place in a closed system mimicking the collection procedure used for whole blood. One other method published that yielded a higher cell recovery used a saline wash after the routine collection to procure residual cells residing in the vein after collection (Elchalal et al., 2000) although feasibility at the bedside may be difficult. Processing cord blood to enrich for hematopoietic progenitor cells most frequently depends on hydroxyethyl starch (HES), which was the method first published by Dr. Rubinstein and others (Rubinstein et al., 1995; Alonso et al., 2001; Liu et al., 2003) demonstrating great success. Cord blood banks reached a consensus that HES sedimentation is a reliable method which can easily be adapted to process large quantities of cord blood products. This method incorporates HES at a 1:5 ratio with the cord blood and allows it to sediment after a centrifugation step. The enriched hematopoietic progenitor cell product is expressed from the concentrated cord blood product. This fraction is then further volume reduced prior to cryopreservation with dimethylsulfoxide (DMSO). Alonso et al. in 2001 reported on a modified method according to Rubinstein et al. (1995). The modified method includes an inverted positioning of the cord blood product in a refrigerated centrifuge during the HES incubation and to reduce red blood cells they are drained from the bottom of the bag. Both methods yield a high recovery of nucleated and hematopoietic progenitor cells. Rubinstein et al. reported a minimum of 91% leukocyte and progenitor cell recovery and Alonso et al. reported an 87% recovery for total nucleated cells and 97% recovery for CD34 positive cells. Other methods used for manual processing include density gradient separations (using Percoll™ or Ficoll™) (Sato et al., 1995; Rogers et al., 2001) and gelatin (Oldak et al., 2000). Automated devices that have been evaluated for cell processing include the Optipress II (Armitage et al., 1999; U-pratya et al., 2003), the Biosafe, and Sepax (Tiumina et al., 2005) and the AutoXpress™ Platform (AXP™) by Thermogenesis (Dobrila et al., 2006) which have demonstrated a high cell recovery post processing. In 1995, a Request for Proposal was solicited by the National Heart, Lung, and Blood Institute entitled, “Transplant Centers for Clinical Research on Transplantation of Umbilical Cord Stem and Progenitor Cells” (Fraser et al., 1998). The study was designed to determine if cord blood transplantation is a viable option for bone marrow transplant. The study would also help to build standard operating protocols for cord blood banking and focused on building an ethnically diverse unrelated donor pool to supply nationalities under-represented (http://spitfire.emmes.com/study/cord/sop.htm). The study was initiated in 1996 and conducted with the United States Food and Drug Administration (FDA) under an Investigational New Drug (IND). The study end point was survival at 180 days with other end points including engraftment, GVHD, relapse, and long-term survival (Cairo et al., 2005). In summary, the report in 2005 states that cord blood transplant should continue

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along with research focusing on cord blood transplantation. The study consisted of approximately 11,000 donors (64% of the total donors collected) stored for potential future use; with 79% meeting donor eligibility criteria stated for the study. The study defined protocols for collection and processing, characterized a unique cell population in the cord blood which includes CD34 positive/CD38 negative a highly proliferative cell population capable of forming early uncommitted blast-like colonies in culture and CD34 positive/CD61 positive demonstrated a correlation to rapid platelet recovery (Cairo et al., 2005). A population of cells CD34 positive/ CD90 positive demonstrated significant correlation with colony-forming capacity (Cairo et al., 2005). They also examined factors of the collection such as sex of the donor, ethnicity, type of delivery, and gestational age which illustrated a significant effect on the progenitor cell content and the lymphocyte subsets (Cornetta et al., 2005).

CORD BLOOD STORAGE Cord blood products when stored long term will either be in liquid nitrogen vapor phase or stored directly in the liquid nitrogen. Concerns in the past with liquid nitrogen vapor storage erupted from temperature changes occurring at the top of the tank when it was opened, but liquid nitrogen storage tank models are available that can retain a temperature of less than 190°C on a consistent basis at the top and the bottom of the tank which allows consistent storage in the vapor phase. Since 1995, when a reported case of hepatitis B transmission occurred (Tedder et al., 1995) in the liquid storage of a nitrogen tank many moved to an overwrap bag system to add a second layer of protection and/or storage in the vapor phase of liquid nitrogen. Since cord blood banking allows an indefinite time period for storage, studies have evaluated the stability of these products for transplantation post cryopreservation. The most common functional viability assay to evaluate clonogenic potential is the colony-forming assay which frequently assesses these four parameters: colony-forming unit-granulocyte, erythrocyte, macrophage, megakaryocyte (CFU-GEMM), colony-forming unit-granulocyte, macrophage (CFU-GM), burst-forming unit-erythrocyte (BFU-E), and colony-forming unit-erythrocyte (CFU-E). Current studies performed have assessed cryopreserved products that have been stored for 10–15 years (Broxmeyer et al., 1997, 2003; Kobylka et al., 1998; Mugishima et al., 1999). Eight samples were assessed after 15 years of storage in liquid nitrogen. At post-thaw the recovery of mononuclear cells averaged 80% proliferative capability and demonstrated cytotoxic response potential against foreign HLA antigens (Kobylka et al., 1998). Proliferative capacities were demonstrated by assessing colony-forming units and replating CFU-GEMM as described by Broxmeyer as a test of “self-renewal” for hematopoietic progenitor cells (Broxmeyer et al., 2003). An assay testing repopulation and engraftment capability in a non-obese diabetic/ severe combined immune deficiency (NOD/SCID) mouse was tested and exhibited similar results compared to using fresh cord blood CD34 positive cells (Broxmeyer et al., 2003). They were also able to demonstrate that an average of 83% of the total nucleated cells was recovered from the products after 15 years in storage. HEMATOPOIETIC AND TISSUE REGENERATION Cord blood cells are now considered a standard product for hematopoietic reconstitution and a potential product for regenerative medicine. Hematopoietic cell transplantation is now a standard of care worldwide for a long list of different diseases which includes but is not limited to leukemia, myelodysplastic syndrome, myeloproliferative and lymphoproliferative disorders, phagocyte disorders, inherited metabolic disorders, inherited immune disorders, inherited platelet disorders, and other malignancies. Transplantation of umbilical cord blood attributes includes low immunogenicity as illustrated by reduced acute GVHD (Rocha et al., 2000) with graft versus leukemia effect remaining intact (Howrey et al., 2000), ease of collection as described earlier, generally a biohazardous discard product with no alternative use, lower risk of infectious disease transmission, potential for ex vivo expansion and the possibility of use in gene therapy.

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Cord blood in the past was viewed as a product for transplantation only in children due to the number of cells that could be harvested from a single cord blood collection, but recently adult cord blood transplantation has been successfully studied along with double cord blood unit transplantation. Umbilical cord blood cells have been demonstrated by Laughlin and colleagues to provide long-term hematopoietic reconstitution in adults over 40 kg in weight, but this cellular product also demonstrated significant delay in the time to hematopoietic engraftment by delayed neutrophil, red blood cell, and platelet recovery similar to the delay seen in children (Laughlin et al., 2001, 2004). A recent study examined high-risk malignancy patients where the median day to engraftment was day 23 when they infused two cord blood units. Twenty-four percent of patients demonstrated engraftment from both units with 76% demonstrating unit dominance on day 100. They were able to demonstrate the safety of transplantation of two partially HLA-matched cord blood products and demonstrated the possibility of adequate cell dose for hematopoietic reconstitution from two cord blood units in adults (Barker et al., 2005).

PLURIPOTENT CELLS FROM UMBILICAL CORD BLOOD CELLS Umbilical cord blood stem cells are not only considered for hematopoietic stem cell reconstitution, but also for other uses demonstrated by its pluripotent stem cell capabilities. The source of umbilical cord cells is almost endless as globally the birth rate is approximately 130 million which would allow for a source of cells easily retrievable. These cells are able to differentiate and expand without a feeder layer and are generally a more ethically accepted cell source. Advantages of this cell source include its naive nature and relatively unshortened telomere length (McGuckin et al., 2005). Kogler and colleagues have identified a cell population in the cord blood which is CD45 negative that they refer to as unrestricted somatic stem cells (USSC). They have demonstrated the potential for this cell population to differentiate into osteoblasts, adipocytes, chondroblasts, hematopoietic, and neural cells in vitro and bone, cartilage, heart, and liver cells in vivo. They were also able to show a time frame greater than 40 population doubling without recombinant cytokines and a longer telomere length as compared to mesenchymal stem cells from bone marrow (Kogler et al., 2004). NEUROLOGICAL REGENERATION Stroke There is an enormous potential for cord blood stem cells to assist in the repair and regeneration of cells and tissues that are afflicted by neurological diseases. Currently there is a wide range of neurological disorders in which scientists are studying the effects of cord blood as a treatment modality in small animal models and there is also a significant amount of work being done on the characterization of these cells. Umbilical cord blood stem cells are one current source of adult stem cells involved in this research today. Many researchers have demonstrated cells co-transplanted with other cells such as sertoli cells (Sanberg et al., 2002) or growth factors (cytokines and chemokines) (Newman et al., 2005) to produce a synergistic response toward therapeutic benefit. Some of the neurological diseases that have been proposed to be treated with stem cells derived for cord blood include stroke, Alzhemier’s disease, Parkinson’s disease, Huntington’s disease, spinal cord injury, central nervous system (CNS) injuries, amyotrophic lateral sclerosis (ALS), cerebral palsy, and generalized brain injuries. In review of the literature for treatment of non-hematopoietic disorders with cord blood cells, stroke is one of the more widely studied disorders. Stroke is also one of the leading causes of death, later-life dementia, and adult disability world-wide today ranking third as the cause of death in the United States behind heart disease and cancer. Researchers in the field of neurological disorders are proposing that adult stem cells may be able to differentiate into neurological tissue or cells and assist with repair at the site by promoting neogenesis specifically by the release of factors that will stimulate the growth of cells already at the site (Borlongan et al., 2004).

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In a review of the historical literature on umbilical cord blood cells in the treatment of neurological disorders, one of the first reports was issued by researchers in 2000 (Erices et al., 2000) where they reported the production of an adherent cell population with potential to differentiate into osteoclasts and mesenchymal-like phenotyped cells. About the same time, Ende and colleagues were studying the effects of cord blood in SOD1 mice used as an animal model for the disease of ALS. They were able to demonstrate that large doses of human umbilical cord blood mononuclear cells were able to prolong the lifespan of SOD1 mice (Ende et al., 2000). In 2001, Sanchez-Ramos reported on the presence of molecular markers on umbilical cord blood cells that are generally associated with neurons and glia cells (Sanchez-Ramos et al., 2001). Ha and colleagues also reported on the current status of neural markers (neurofilament (NF), microtubule associated protein (MAP2), glial fibrillary acidic protein(GFAP)) demonstrated in cultured human umbilical cord blood cells along with the classic neural morphology (Ha et al., 2001). Both groups suggested these cells may be used in the future for therapeutic applications related to neurological disorders. Cord blood cells may be a viable option compared to the use of neural progenitors due to the ease of collection. Researchers compared stromal cells in bone marrow to cells with multilineage potential found in cord blood and were able to demonstrate these cells could differentiate into neural cells as identified by immunofluorescent labeling and Western blot analysis, but lacked some of the neural markers seen in bone marrow cells which alluded to a more immature cell population in cord blood (Goodwin et al., 2001). Researchers were able to demonstrate a neural stem-like cell population from human umbilical cord blood after cell isolation and fractionation that yielded a high-potency cell population expressing the surface marker nestin. When fractionated cells were placed in culture with growth factors or rat brain the researchers were able to demonstrate the three main neural phenotypes representing neurons, astroglia, and oligodendroglia at 30%, 40%, and 11% of the population respectively (Buzanska et al., 2002). Chen and colleagues studied the effects of human umbilical cord blood infused intravenously after stroke in a rat model. The rats were subjected to middle cerebral artery occlusion prior to cell infusion. Cord blood cells significantly improved function as demonstrated by two behavioral tests where as 7 days after occlusion improvement in only one of the behavioral tests was observed. The investigators were able to determine that the cord blood cells could enter brain tissue, survive, and improve neurological recovery in this model which demonstrates the potential of using these cells in the future for therapeutic applications related to stroke (Chen et al., 2001). Zigova and colleagues infused human umbilical cord blood cells into a developing rat brain to evaluate cell survival and phenotypic properties of the cells after infusion. They cultured the cells in retinoic acid (RA) and nerve growth factor (NGF) prior to infusion and then cell suspensions were injected into the anterior part of the subventricular zone. When the brain tissue was assessed for neural markers the cells were found to be positive for the following neural markers; GFAP and beta-III-tubulin. They determined 1 month post infusion into a rat brain that approximately 20% of the cells infused into the brain survived (Zigova et al., 2002). Other researchers measured the effects of human umbilical cord blood cells infused into a rat after traumatic brain injury and the cell migrated to the site of injury in the brain and expressed neural markers. The rat model demonstrated the potential of cord blood cells in treatment of traumatic brain injury. The cells not only expressed neural markers within the brain, but also integrated into vascular tissue surrounding the injured brain tissue (Lu et al., 2002). Taguchi in 2004 was able to demonstrate neurogenesis and angiogenesis in a mouse model after the infusion of CD34 positive cells selected from human umbilical cord blood cells. The infusion was given to immunocompromised mice 48 h after injury. They were able to demonstrate endogenous neurogeneration accelerated by homing neural progenitors to the site of injury. They proposed that the CD34 positive cord blood cells promote neovascularization either directly or indirectly providing the environment for neovascularization (Taguchi et al., 2004). Borlongan and colleagues investigated why human umbilical cord blood transplants in a rat stroke model exhibited neuroprotection. They infused cord blood with mannitol to

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permeabilize the blood brain barrier and this reduced the cerebral infarcts and improved behavioral function in the animals that received the cord blood cells. Other animals that received the cord blood without the mannitol did not have an effect on the cerebral infarct or behavioral tests. The researchers concluded that the neuroprotection was transferred via secreted chemical factors rather than cells homing to the site of injury (Borlongan et al., 2004). These research groups both discuss the potential of indirect neovascularization that occurs in an in vivo stroke model. Vendrame and colleagues examined the behavioral recovery and stroke infarct volume before and after infusion of human umbilical cord blood cells assessing cell dose in a rat model. Four weeks after the cord blood infusion, there was a significant recovery in behavioral performance when a minimum of 1 million cells were infused. When doses of cord blood cells increased, they were able to demonstrate behavioral improvement, and neuronal sparing correlating directly to the number of cells infused. The researchers discussed the large cell dose required for potential human infusion and possibly using ex vivo expansion in this situation (Vendrame et al., 2004). This same group also reported on studies in a rat model of stroke that human umbilical cord blood cells may be effective by decreasing pro-inflammatory cytokines to result in enhanced neuroprotection. Testing results demonstrated a decrease in mRNA and protein expression of pro-inflammatory cytokines and a decrease in nuclear factor kappa B DNA binding activity in the brain of stroke animals treated with cord blood cells (Vendrame et al., 2005). Newman and colleagues investigated the migration of human cord blood cells to ischemic tissue extracts which correlated with an increase in certain cytokines and chemokines. They were also able to illustrate that the time frame for treatment may be extended out from a suggested 3 h to 24–72 h after a stroke when using approved anticoagulant therapy and cord blood cell infusion (Newman et al., 2005). In summary the research on human umbilical cord blood cells in relationship to stroke has demonstrated that neuronal cell markers are present on cells and that some of these cells actually demonstrate a more primitive status than the cells found in bone marrow. The cells when infused not only could enter the brain but also vascular tissue. Neuroprotection has a strong correlation to chemical factors produced at the site of injury. Decreasing pro-inflammatory cytokines appears to be a major factor and it may be possible to extend the treatment after cord blood infusion greater than 3 h as previously considered. Huntington’s disease, Alzheimer’s disease, Parkinson’s disease and ALS In the examination of neurodegenerative diseases one publication was found in the literature related to Huntington’s disease reported by Ende and colleagues where they treated mouse models for the disease with mega-doses of human umbilical cord cells. They infused approximately 70–100 million cells to treat a mouse to increase their lifespan from 88 days to 97.8–103.4 days with the largest dose of cells (Ende et al., 2001). Ende and colleagues also examined Alzheimer’s and Parkinson’s disease in a small animal model with human cord blood cells. Their reports included considerable life extension in the mouse model for Alzheimer’s disease after umbilical cord blood mononuclear infusion. A high dose of 110 million cells per mouse infused compared to control animals demonstrated a longer lifespan (Ende et al., 2001). In the mouse model for Parkinson’s disease three groups were studied: infused with congenic marrow mononuclear cells with one out of ten alive, infused with cord blood mononuclear cells with four out of twelve alive, and a control group with one out of ten alive. The experiments were terminated at day 200 and results demonstrated a delay in the onset of symptoms and prolonged lifespan in the group infused with umbilical cord blood mononuclear cells (Ende and Chen, 2002). Researchers examined ALS which is characterized by motor neuronal degeneration. An ALS mouse model (G93A) was used to study the infusion of human umbilical cord blood mononuclear cells into systemic circulation. The researchers demonstrated that the infusion delayed disease progression 2–3 weeks and increased the lifespan of the mice. The infused cells migrated to the parenchyma of the brain and spinal cord

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where neural markers were expressed on these cells including nestin, beta-III-tubulin, and GFAP (Garbuzova et al., 2003). Other researchers utilizing a mouse model for ALS, the SOD1 mice, infused a mega-dose of umbilical cord blood mononuclear cells after irradiation. They demonstrated that doubling the mega-dose of cells further increased the lifespan of the mice. To produce a mega-dose for infusion donors were pooled and no negative effects were observed (Chen and Ende, 2000). Willing and colleagues examined different sites of infusion in the mouse model for ALS. They demonstrated through behavioral tests that intraspinal infusion did not demonstrate improvements but intravenous did demonstrate improvements not only by behavioral tests, but also demonstrated by life-span (Willing et al., 2002). Spinal Cord Injury and CNS injuries Researchers recently examined the functional effects from an umbilical cord blood infusion in a rat model for spinal cord injury. Three groups were assessed for treatment by infusion of umbilical cord blood cells, umbilical cord blood cells with brain-derived neurotrophic factor, and a control group with media alone injected directly into the spinal cord. Groups that included the infusion of cord blood demonstrated improvement weekly over the control group in the locomotor rating scale. They also demonstrated that the transplanted cells differentiated into various neural cells (Kuh et al., 2005). Other researchers examined the effects of cord blood cell infusion in a sex-mismatched mouse model which demonstrated cells were generated in the CNS but concerns arose surrounding the available cell dose in a product and HLA disparity of the cells (Korbling et al., 2005). Researchers in 2005 published a case study on a spinal cord-injured patient transplanted with umbilical cord blood cells. The cells were HLA-matched and transplanted directly into the spinal cord. The case study demonstrated an improvement in the sensory perception and movement in the patient’s hips and thighs. An MRI and CT scan also demonstrated regeneration of the spinal cord at the site of injury (Kang et al., 2005). These are potentially exciting applications that will need further investigation for the site of infusion and possible assessment of the minimal cell dose required prior to scale up studies in larger animal models. Neural Cell Surface Markers It is known that researchers are able to differentiate adult multipotent cells into neurons, astrocytes, and oligodendrocytes in the CNS, but the mechanisms involved in the differentiation are a critical component of current research. More recently human umbilical cord blood cells specifically have been assessed for potential to produce neural progenitors some of the markers used to identify the cell population are discussed. Buzanska and colleagues selected CD34 negative umbilical cord blood cells after density gradient and expanded the cells in media to support the growth of neurogenic cells. Post culture of the cells expressed nestin, which is a primitive marker for neural cells and in culture with selected growth factors 30% of the cell population was neuronal, 40% astrocytic, and 11% were oligodendrocytes (Buzanska et al., 2002). Although nestin is a well-known neural progenitor cell marker it is also associated with other cell types such as pancreas, kidney, hair follicle cells, and blood vessels in the skin (Amoh et al., 2005). Nestin is also a filament protein that has been shown to play a role in cytoskeleton regulation (Chen et al., 2006). Jang and colleagues isolated CD133 via magnetic cell sorting by bead sorting and fluorescence-activated cell sorter (FACS). After selection, umbilical cord blood cells were cultured in RA and cells expressed neuronal and glial phenotypes. Post culture, the cells demonstrated transcription factors important for early neurogenesis including Otx2, Pax6, Wnt1, Olig2, Hash1, and NeuroD1 (Jang et al., 2004). Other researchers also isolated CD133 positive progenitor cells from human umbilical cord blood and cultured the selected cells in media containing Flt3-ligand (FL), thrombopoietin (TPO), and stem cell factor (SCF). The cells post culture expressed pluripotent markers including Sox-1, Sox-2, FGF-4, Rex-1, and Oct-4. After cell culture with RA the cells demonstrated a neural morphology coupled to the expression of beta-III-tubulin (Baal et al., 2004).

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McGuckin and colleagues performed a negative hematopoietic lineage depletion where they were able to recover 0.1% of the starting cell population and adherent cells were capable of demonstrating neuroglial progenitor cell morphology. Neuroglial progenitor cell markers were identified by gene expression analysis including, beta-III-tubulin (McGuckin et al., 2004). In summary the markers associated with identification of the neural cell lineage in umbilical cord blood cells include nestin, CD133, GFAP, NF, and MAP2. Nestin is a marker for primitive neural tissue and CD133 is a known cell surface marker associated with the production of neural and glial cells. Glial fibrillary acidic protein more commonly known as GFAP marks the astrocytes which is a type of glial cell. Neurofilament known as NF is an important structural component of the neuron and MAP2 is a protein found in the dendritic branching of the neuron. Cardiac Function Treatment Cardiac disease is the number one killer of men and women worldwide. Approximately 70 million Americans have some form of the disease and this is one of the reasons many investigators are studying how to treat the disease with cellular therapy. Researchers as early as 2001 discussed potential therapeutic applications, such as genetic modulation, cell transplantation, and tissue engineering as a novel approach to myocardial regeneration and tissue repair after myocardial infarction (Etzion et al., 2001). Current animal studies and human clinical trials are evaluating infusion of cells directly into the damaged myocardium or infusion of cells via intravenously to repair damaged and infarcted tissue. These cells may have the potential in the future to replace whole organ transplants with cell transplants derived from umbilical cord blood cells. Myocardial Infarction Regenerative medicine after a myocardial infarction may include the replacement of the damaged cells by either an intravenous infusion or infusion at the site of ischemia. Recently several articles have been published on the phenotypic properties of umbilical cord blood stem cells used in cardiac repair. The most common phenotypic marker published in the literature is CD34 which is a cell surface glycoprotein, generally marking the hematopoietic progenitor cell. In 2004, Botta and colleagues discussed the production of hematopoietic and endothelial cells from the hemangioblast. They were investigating a progenitor cell in cord blood with a phenotype of CD34 positive/KDR positive. KDR is an endothelial growth factor receptor. They assessed the potential of these cells in a NOD–SCID mouse model and were able to demonstrate beneficial effects of cord blood cells CD34 positive/KDR positive illustrating improvement in cardiac hemodynamics by resistance to apoptosis and their angiogenic action (Botta et al., 2004). Looking at a different disease process Cogle et al. (2004) also demonstrated the functional hemangioblast potential of CD34 positive human umbilical cord blood cells. They assessed an NOD–SCID mouse model for retinal ischemia which resulted in human retinal neovascularization in the mouse model. Hirata and colleagues in 2005 studied the effects of human umbilical cord blood CD34 positive cells in a rat model for myocardial infarction produced by ligation of the left coronary artery. The CD34 positive cells survived and improved cardiac function (Hirata et al., 2005). Two groups examined the effects of CD133 positive cells which has been identified as a neural and hematopoietic cell marker and recently was published as a marker for embryonic stem cell-derived progenitors (Kania et al., 2005). Leor and colleagues discussed the possibility of human umbilical cord blood stem cells for use in repair of infarcted myocardium. They infused approximately 1.2–2 million cells intravenously 7 days after coronary artery ligation in a rat model and were able to demonstrate that the cell infusion produced functional recovery by preventing scar thinning and left ventricular systolic dilation (Leor et al., 2006). Wu and colleagues expanded CD133 positive cells from human umbilical cord blood stem cells to produce endothelial progenitor cells (EPC) (Wu et al., 2004).

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Ott and colleagues suggested that cell therapy for myocardial infarction may be limited by the number of cells available. They expanded CD34 positive cord blood and cultured the cells in endothelial medium where cells were expanded up to 46 population doublings. These cells were able to form vascular structures and improve left ventricular function after experimental myocardial infarction in an athymic nude mouse model (Ott et al., 2005). Delorme and colleagues demonstrated that EPC in umbilical cord blood are CD146 positive cells, which is an adhesion marker on endothelial cells. These cells were selected from the non-adherent cell population in umbilical cord blood cells. They were able to demonstrate proliferation in long-term culture while maintaining the same phenotypic properties and the colonization of a matrigel plug in immunodeficient NOD–SCID mice. This study suggests that the CD146 positive cells contain a subpopulation of circulating EPC that may be used in pro-angiogenic therapy (Delorme et al., 2005). Other scientists investigated the use of mononuclear cells from human umbilical cord blood cells for treatment in acute myocardial infarction. They infused 1 million cells in a rat model that underwent left anterior descending coronary artery ligation and the cells were injected directly into the infarct border. The results of the experiments demonstrated a reduction in the infarction size in the rat model. Left ventricular functional measurements and ejection fractions were greater in the cord blood infusion group (Henning et al., 2004). Kim and colleagues discussed the USSC potential to differentiate into myogenic cells and induce angiogenesis. A porcine model demonstrated regional and global function of the heart after a myocardial infarction. These cells have been proposed to be used for cellular cardiomyoplasty due to efficacy and safety of the cells (Kim et al., 2005). Ishikawa and colleagues tested the potential of human umbilical cord blood stem cells to give rise to cardiomyocytes in vivo. They infused cord blood lineage negative cells which generated cardiomyocytes following transplantation into immune deficient mice (Ishikawa et al., 2006). Chen and colleagues assess the potential use of human umbilical cord blood cells with gene therapy to enhance angiogenesis via a mouse model after acute myocardial infarction. The goal of the study was to improve myocardial infarction by new vessel formation. A mouse model for acute myocardial infarction was infused intramyocardially with purified CD34 positive cells. The mouse model demonstrated a reduction in the infarct size with increased capillary density which resulted in a reversal of cardiac dysfunction (Chen et al., 2005). Ma and colleagues isolated human umbilical cord blood CD34 positive cells to inject them into the tail vein of an NOD–SCID mouse model with ligation of the left anterior coronary artery. Post infusion they analyzed capillaries for chimerism, but only occasionally human and mouse endothelial cells were discovered with most new vessels displaying mouse cells only. Post analysis, it was determined that up to 70% of the cord bloodderived cells in the heart were CD45 positive. The cells did not appear to differentiate, but did demonstrate migration to the infarcted tissue selectively where they engrafted to assist in neogenesis (Ma et al., 2005). Umbilical cord blood cells appear to be an attractive target for cell therapy after myocardial infarction due to the low immunogenicity of the cells and the ease of collection and storage of the cryopreserved product which render it easily accessible. It has no ethical concerns as embryonic stem cells and is currently used as an alternative for bone marrow in hematopoietic reconstitution in standard treatment protocols. All the recent data are encouraging for the use of human umbilical cord blood stem cells to assist in the reversal of cardiac dysfunction in the above described applications. Stem cell expansion maybe a major limiting factor if the cell dose used in the animal model needs to be translated to humans. The cord blood cell infusion may in the future eliminate the need to procure tissue or blood vessels from the patient for cardiac reconstruction. Clinical Trials for Cardiac Disorders Several phase I clinical trials are active involving progenitor cells derived from bone marrow for the treatment of myocardial infarction. The trials include the use of mesenchymal stem cells infused intravenously and autologous bone marrow mononuclear cells infused directly into the coronary artery. One other study

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involves the assessment of the safety of the autologous skeletal myoblasts cells via catheter delivery. Clinical studies with bone marrow cells in the past have shown improvement in function and decreased infarct size although the treatment is still an area of debate, but none the less is able to proceed with clinical trials to determine the process is safe. Currently bone marrow stem cells appear to be the cell of choice for treatment of myocardial infarction. The optimal cell type may allow for the promotion of angiogenesis and myogenesis, but further studies are required to determine the best cell type. Treatment of Diabetes Pancreatic Cells, Insulin-Producing Cells, Treatment of Type I/Type II Diabetes and Diabetic Neuropathy Recently a lot of work has been done in the area of regenerative medicine for Type I diabetes including the use of hepatocytes, bone marrow, intestinal epithelial cells, and pancreatic stem cells. Yoshida and colleagues demonstrated the production of insulin-producing cells from human umbilical cord blood via a mononuclear cell preparation that was infused into an NOD–SCID mouse. They were able to demonstrate cord bloodderived cells resulted, in insulin-producing cells at a rate of 0.65% 0.64% in xenogeneic hosts by fusion dependent and independent functions (Yoshida et al., 2005). Ende and colleagues assessed the use of human umbilical cord blood mononuclear cells in pre-diabetic stage NOD mice with autoimmune Type I diabetes. The outcome of the experiments demonstrated significantly lower glucose levels and increased their lifespan. The mice that received the highest dose had the most significant response with the highest dose at 200 million cells. The researchers were able to demonstrate that cord blood mononuclear cells infused at the pre-diabetic stage in the NOD mouse model without any immunosuppression is able to lower glucose levels and increase lifespan (Ende et al., 2004a). They also examined the effects of human umbilical cord blood cells for the treatment of Type II diabetes. They assessed blood glucose levels, survival, and renal pathology. In the obese mice with Type II diabetes infused with umbilical cord blood improvement was seen not only in blood glucose levels and survival rate, but also normalization of glomerular hypertrophy and tubular dilation (Ende et al., 2004b). Pessina and colleagues discuss a panel of markers required for human umbilical cord blood cells to form multipotent progenitor cells of the pancreas. The markers included nestin, which is generally viewed as a neuronal or pancreatic progenitor cell marker; other markers listed include cytokeratin (CK)-8 and CK-18. Transcription factors associated with islet-derived progenitors are Isl-1, Pdx-1, Pax-4, and Ngn-3. They were able to demonstrate that human umbilical cord blood cells contain a population of phenotyped cells similar to endocrine cell precursors forming beta cells (Pessina et al., 2004). Naruse and colleagues have assessed the use of EPCs from human umbilical cord blood cells for use in the reversal of diabetic neuropathy. Cord blood mononuclear cells were cultured and EPCs were isolated and expanded. The EPCs were injected intramuscular into the hindlimb skeletal muscles of streptozotocininduced diabetic nude rat model. The study results demonstrated an increased number of microvessels in hindlimb skeletal muscles in the diabetic rats compared to the controls (Naruse et al., 2005). Clinical Trials for Type I Diabetes Currently there is one active clinical trial assessing autologous cord blood infusion for Type I diabetes in an attempt to regenerate pancreatic islet insulin-producing beta cells and therefore improving glucose control. The researchers will track migration of the infused stem cells and study changes in metabolism and immune function leading to islet regeneration. The study is a phase I/phase II clinical trial so that it will evaluate safety and efficacy.

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Hepatocyte-Like Cells Currently research has focused on the search for alternatives, such as liver progenitors, fetal hepatoblasts, embryonic, bone marrow, or umbilical cord blood stem cells to replace hepatocytes in a disease state. Several investigators have examined the specific potential of human umbilical cord blood cells assessing the in vitro and in vivo potential to form hepatic cells. This is a relatively new area of experimentation as a significant amount of the work completed appeared in the literature over the last year. Most of the researchers assessed the potential of in vivo generation of hepatic cells after injury in a mouse model and a few others examined the potential of in vitro differentiation to the hepatic cell lineage. In the assessment of in vitro differentiation Kang and colleagues assessed the capability of human umbilical cord blood cells to differentiate into hepatocyte-like cells. Cord blood mononuclear cells were collected and cultured with hepatocyte growth factor (HGF), fibroblast growth factor 4 (FGF4), both, and no growth factor. The authors were able to demonstrate that HGF- and FGF4-induced cord blood mononuclear cells were capable of differentiation into hepatocyte-like cells (Kang et al., 2005). Other investigators reported on a cell population; cord-blood-derived embryonic-like stem cells (CBE) positive for TRA-1-60, TRA-1-81, SSEA-4, SSEA-3, and Oct-4, which are also embryonic stem cell markers. CBE were also cultured with hepatocyte growth medium and post culture the cells expressed characteristic hepatic markers, CK-18, alpha-fetoprotein, and albumin (McGuckin et al., 2005). Researchers investigated the potential of human cord blood to be used as cell therapy for an injured liver in vitro and in vivo. Cord blood cells post culture expressed albumin and hepatocyte lineage markers. When investigating liver-injured severe combined immunodeficient mice infused with human umbilical cord blood cells, they were able to demonstrate the development of functional hepatocytes in the liver (Kakinuma et al., 2003). Researchers suggest that these cells may have potential for treatment of hepatic diseases. Other researchers examined the potential of CD34 selected cells from human umbilical cord blood cells for production of hepatocytes in vitro. They also assessed NOD/SCID mice for the in vivo studies where it was exposed to liver injury by a Fas ligand-carried adenoviral vector. As demonstrated by RT-PCR the cord cells were able to differentiate into hepatocyte-like cells in the mouse liver and it was demonstrated that liver injury was essential during this process. There were no differences between the use of CD34 positive and CD34 negative cells (Nonome et al., 2005). Kashofer and colleagues also evaluated hepatic in vivo differentiation from human cord blood mononuclear cells selected for CD34 positive cells or lineage negative cells. The cells were infused after liver damage in NOD/SCID mice. To identify the infused cells they transduced, the stem cell population, with lentivirus construct expressing enhanced green fluorescent protein (eGFP) and fluorescent in situ hybridization (FISH) analysis performed as the cells were sex mismatched. The results of the study revealed that very little human chromosomes were present in the hepatocyte-like cells and they may have fused with host hepatocytes (Kashofer et al., 2005). Other researchers examined the potential of inducing hepatic differentiation in human umbilical cord blood cells. They assessed for newly formed hepatocyte-like cells in the liver of NOD–SCID mice after transplantation of human cord blood or murine bone marrow. Liver injury was induced by carbon tetrachloride and they detected clusters of hepatocyte-like cells derived from cord blood cells. FISH demonstrated mostly host-derived hepatocyte-like cells with murine bone marrow infusion. They demonstrated that human cord blood in an NOD–SCID mouse model has contrasting differentiation potential from murine bone marrow cells (Sharma et al., 2005). Investigators assessed the efficacy of human umbilical cord blood cells to decrease histologic damage and the mortality rate of animals previously damaged by allyl alcohol. NOD/SCID mice were treated with allyl alcohol with and without intraperitoneal infusion of human cord blood cells. The cord blood cells infused were able to transdifferentiate into hepatocytes and demonstrate a significant decrease in mortality rate in the mouse model. Researchers believe that endogenous regeneration occurs for early stage of damage (Di Campli et al.,

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2005). Investigators studied human umbilical cord blood cells that were CD34 positive or CD45 positive and they were transplanted into NOD/SCID/beta-II-microglobulin null mice. The livers were examined for evidence of human hepatocyte engraftment. Analysis of the mouse bone marrow revealed that 21.0–45.9% of the cells were human and FISH analysis excluded spontaneous cell fusion for the generation of human hepatocytes. The researchers demonstrated that human cord blood cells can give rise to hepatocytes in an xenogeneic transplantation model (Ishikawa et al., 2003). Researchers investigated the potential of human cord blood to be used as cell therapy for an injured liver in vitro and in vivo. Cord blood cells post culture expressed albumin and hepatocyte lineage markers. When investigating liver-injured severe combined immunodeficient mice infused with human umbilical cord blood cells they were able to develop into functional hepatocytes in the liver (Kakinuma et al., 2003). All researchers were able to demonstrate the in vitro repopulating capability of hepatic cells derived from human umbilical cord blood cells with the appropriate growth factors. Others working with the mouse model were able to demonstrate the in vivo differentiation potential of cord blood cells when hepatic injury occurs due to carbon tetrachloride or allyl alcohol. McGuin and colleagues were able to demonstrate CBE that may have the potential in the future as a source of transplantable hepatic progenitor cells. Endothelial Progenitors Angiogenic therapy by using EPC is currently a topic of debate. These cells have been used to treat of ischemic diseases for revascularization. They may also be used in diagnosis to assess the disease state in the patient; cord blood is fairly new in this arena. These cells have not only been assessed phenotypically by markers, but they also need to be evaluated for the proliferative and clonogenic potential. Ingram and colleagues described a group of EPCs derived from replating colonies by a single cell method in culture from umbilical cord blood cells. This culture gave rise to a new cell population capable of at least 100 population doublings and was able to retain high levels of telomerase activity (Ingram et al., 2004). Murga and colleagues isolated CD34 negative cells including endothelial precursor cells from human umbilical cord blood cells. The CD34 negative cell population with angiogenic factors produced cells that express the endothelial cell markers: vascular endothelial-cadherin, vascular endothelial growth factor receptor1 (VEGFR-1) and VEGFR-2, Tie-1 and Tie-2, von Willebrand factor, and CD31 and can be expanded in vitro for over 20 passages. Researchers were able to demonstrate endothelial precursors in the CD34 negative cell population of cord blood (Murga et al., 2004). Salven and colleagues were able to demonstrate that human CD34 positive and CD133 positive cells expressing VEGFR-3 constitute a phenotypically and functionally distinct population of endothelial stem and precursor cells that may play a role in angiogenesis (Salven et al., 2003). Other researchers were able to identify a cell population of circulating endothelial precursors expressing VEGFR2, CD34, and CD133 from human cord blood which may have a role in neogenesis (Peichev et al., 2000). Shin and researchers examined the cytokines and culture conditions required for large amounts of endothelial cells that may be required for vasculogenesis. The CD34 positive cells from human cord blood were selected and cultured in various cytokine cocktails. The quantity of cells adherent and non-adherent was the greatest with use of SCF, FL, and TPO cytokines. When growth factors were added: VEGF, IL-1 beta, FGFbasic (FGF-b); endothelial cells were identified by morphology and endothelial-specific markers (Shin et al., 2005). Researchers are investigating autologous patches engineered from human umbilical cord-derived fibroblasts and EPCs as a ready-to-use cell source for pediatric cardiovascular tissue engineering. EPCs were isolated from umbilical cord blood by density gradient centrifugation and myofibroblasts were harvested from umbilical cord tissue. Cells were differentiated and expanded in vitro. The investigators believe that these cells may be used for autologous replacement materials for congenital cardiac interventions (Schmidt et al., 2005). The possibility exists in the future to be able to use the differentiated cells produced from umbilical

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cord blood for vascular cell therapy, but more in vivo studies are required to assess the homing capabilities of the cells. Chondrocytes One article was published on the differentiation of cord blood cells into chondrocytes. Researchers investigated human umbilical cord blood cells lineage negative, CD45 negative, CD34 negative for the potential of chondrocyte differentiation. They cultured the cells with mouse embryonic limb bud cells which demonstrated that cord blood cells have the potential to differentiate into chondrocytes (Jay et al., 2004). Ex vivo Expansion A limiting factor for the use of cord blood cells is the low cell dose harvested at procurement. In clinical transplantation the low dose is often associated with delayed engraftment of neutrophils and platelets. Cell expansion is being investigated, but concerns lie in the low quantity of primitive cells observed in the expanded cell population. Different methods currently in use for ex vivo expansion include CD34 or CD133 positive selection. These selected cells are cultured to proliferate primitive cells, they can also be co-cultured with mesenchymal stem cells with growth factors or cells can be cultured in a bioreactor with continuous perfusion (Robinson et al., 2005). Robinson and colleagues compared two cord blood expansion methods. They studied the effects of CD133 positive cells expanded in culture and cord blood unmanipulated co-cultured with bone marrow mesenchymal stem cells both supplemented with growth factors. They were able to conclude through analysis of the total nucleated count, CD133 positive and CD34 positive cells that the cord blood co-cultured with mesenchymal stem cells performed better than the CD133 selected cells (Robinson et al., 2006). Tetraethylenepentamine (TEPA) enables preferential expansion of early hematopoietic progenitor cells in human umbilical cord blood-derived CD34 positive cell cultures as reported by Peled and colleagues in 2004. The copper chelation appears to modulate the balance between self-renewal and differentiation of hematopoietic progenitor cells (Peled et al., 2004). CD133 selected cells were cultured and expanded. The authors reported CD34 cells increased by 89-fold, CD34 positive/CD38 negative increased by 30-fold and colony-forming unit cells by 172-fold over the number of cells seeded (Peled et al., 2004). Subsequently they were transplanted into NOD/SCID mice which demonstrated the CD133 expanded cells faired better compared to the unexpanded for engraftment in terms of CD45 positive and CD45, CD34 positive cells (Peled et al., 2004). Peled and colleagues have demonstrated the enhancement effect of TEPA when they examined human umbilical cord blood selected for CD133 positive cells, cultured in a closed system with cytokines (SCF, TPO, IL-6, and FL). The cell yield of CD34 positive population made a 89-fold and a 172-fold increase in colony-forming units. Infusion into an irradiated non-obese diabetic (NOD/SCID) mice demonstrated superiority with the expanded product (Peled et al., 2005). McNiece and colleagues investigated the potential of human cord blood mononuclear cells in co-culture with mesenchymal stem cells. The expansion demonstrated 10- to 20-fold increase in total nucleated cells, 7to 18-fold increase in committed progenitors, 2- to 5-fold expansion of primitive progenitors and 16- to 37fold increase in CD34 positive cells which may allow for significant expansion without the use of pre-cell selection (McNiece et al., 2004). Delany and colleagues elaborate on the effect of Notch ligand density on induction of Notch signaling and the effect on expansion of human CD34 positive, CD38 negative cord blood progenitors. Lower densities of Delta1 (ext-IgG) enhanced production of CD34 positive cells while higher densities induced apoptosis of these cells. The density of Notch ligands may be an important factor in expansion of cord blood cells (Delaney et al., 2005). Jang and colleagues also examined expansion of cord blood cells in co-culture with mesenchymal stem cells lacking cytokines which demonstrated CFU-GM, CFU-GEMM, BFU-E, and CFU-E increased to 3.46-, 9.85-, 3.64-, and 2.03-folds, respectively (Jang et al., 2006). It appears

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that cord blood expansion has made progress in increasing not only the CD34 positive and the CD34 positive/CD38 negative hematopoietic progenitor cell population, but also capable of expanding the colonyforming unit capacity of the cell product. Aldehyde Dehydrogenase Expressing Cells Jones and colleagues described aldehyde dehydrogenase as an enzyme observed in high amounts in primitive cells initially discovered in bone marrow (Jones et al., 1995). Storms and colleagues discussed the stem cell and progenitor potential of umbilical cord blood with aldehyde dehydrogenase positive cells. They were able to demonstrate progenitors were highly enriched within the aldehyde dehydrogenase bright and CD34 positive population, but when compared to the aldehyde dehydrogenase negative and CD34 positive population few primitive progenitors were identified. They suggest the use of aldehyde dehydrogenase to discriminate between stem cell and progenitor cell populations in umbilical cord blood (Storms et al., 2005).

CONCLUSION For many years bone marrow and mobilized peripheral blood were the leaders in reconstitution for hematopoietic disorders, but now umbilical cord blood is gaining speed and is viewed as an alternative to bone marrow for transplantation. Mesenchymal stem cells were first described in bone marrow and now there are several reports on the unrestricted pluripotent cells identified in cord blood. Several advantages of these cells may assist in its leadership in regenerative medicine as one cell source lacking ethical concerns. Advantages are vast reaching including its ease of procurement, its naive immune status, its low contamination potential of infectious disease, its relatively unshortened telomere length and its homing capabilities that have been demonstrated in small animal models and in humans for hematopoietic reconstitution. With approximately 130 million births a year worldwide this is a largely under-utilized precious source of stem cells. An abundance of work has been done in the area of cord blood transplantation since the first reported case in 1988 which includes transplantation of umbilical cord blood stem cells for children to treat leukemia, lymphoma, and certain cancers including genetic disorders that affect the blood and immune system. Cord blood cells have also been used in adult transplants and double cord transplants which have been able to treat patients over 40 kg in weight. Due to the limiting nature of the number of cells in a cord blood product a significant amount of work has been done on the expansion of these cells focusing on the retention of the primitive stem cells required for engraftment. Amazing strides have been made to demonstrate that cord blood does in fact contain pluripotential cells that have proven differentiation to the lineages within all three germ cell layers. Publications on differentiation capability included neural related cells, cardiac cells, pancreatic progenitor cells, hepatocyte-like cells, endothelial cells, and chondrocytes. Ex vivo expansion is an active area of research due to the quantity of cells available at time of procurement. Researchers are examining different cells to expand and different culture conditions including a study assessing the bioreactor for continuous perfusion culture. The current challenges in umbilical cord blood stem cells include the quantity of cells in the procured product to be used for children and adults. Also tied into the quantity is the fact that a number of studies performed in small animal models required a significant amount of cells to demonstrate effective treatment; to be able to translate this dose to large animal models or human clinical trials may require an optimal technique for the expansion of these critical cells. FUTURE DEVELOPMENTS In evaluation of the pluripotent cells from umbilical cord blood investigators will be able to demonstrate safety in the infusion of the cells. Umbilical cord blood has already had years of safety data with routine transplant for

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leukemia and lymphoma. Unlike embryonic stem cells cord blood may have an abbreviated path to clinical trials as demonstrated with bone marrow mesenchymal stem cells in limited cases. More studies will be assessing the efficacy of cord blood transplants in adults. Basic research will continue to thrive in an effort to fuel significant changes in this area. Pre-clinical trials and phase I clinical trials will continue to move transplantation into an era where it will immensely expand the number of diseases it will have potential to treat or cure.

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Peichev, M., Naiyer, A.J., Pereira, D., Zhu, Z., Lane, W.J., Williams, M., Oz, M.C., Hicklin, D.J., Witte, L., Moore, M.A., et al. (2000). Expression of VEGFR-2 and AC133 by circulating human CD34() cells identifies a population of functional endothelial precursors. Blood 95(3): 952–958. Peled, T., Mandel, J., Goudsmid, R.N., Landor, C., Hasson, N., Harati, D., Austin, M., Hasson, A., Fibach, E., Shpall, E.J., et al. (2004). Pre-clinical development of cord blood-derived progenitor cell graft expanded ex vivo with cytokines and the polyamine copper chelator tetraethylenepentamine. Cytotherapy 6(4): 344–355. Peled, T., Glukhman, E., Hasson, N., Adi, S., Assor, H., Yudin, D., Landor, C., Mandel, J., Landau, E., Prus, E., et al. (2005). Chelatable cellular copper modulates differentiation and self-renewal of cord blood-derived hematopoietic progenitor cells. Exp. Hematol. 33(10): 1092–1100. Pessina, A., Eletti, B., Croera, C., Savalli, N., Diodovich, C. and Gribaldo, L. (2004). Pancreas developing markers expressed on human mononucleated umbilical cord blood cells. Biochem. Biophys. Res. Commun. 323(1): 315–322. Rabian-Herzog, C., Lesage, S. and Gluckman, E. (1992). Characterization of lymphocyte subpopulations in cord blood. Bone Marrow Transplant. 9(Suppl 1): 64–67. Robinson, S., Niu, T., de Lima, M., Ng, J., Yang, H., McMannis, J., Karandish, S., Sadeghi, T., Fu, P., del Angel, M., et al. (2005). Ex vivo expansion of umbilical cord blood. Cytother. Rev. 7(3): 243–250. Robinson, S.N., Ng, J., Niu, T., Yang, H., McMannis, J.D., Karandish, S., Kaur, I., Fu, P., Del Angel, M., Messinger, R., et al. (2006). Superior ex vivo cord blood expansion following co-culture with bone marrow-derived mesenchymal stem cells. Bone Marrow Transplant. 37(4): 359–366. Rocha, V., Wagner Jr., J.E., Sobocinski, K.A., Klein, J.P., Zhang, M.J., Horowitz, M.M. and Gluckman, E. (2000). Graft-versus-host disease in children who have received a cord-blood or bone marrow transplant from an HLA-identical sibiling. Eurocord and International Bone Marrow Transplant Registry working committee on alternative donor and stem cell sources. N. Engl. J. Med. 342(25): 1846–1854. Rogers, I., Sutherland, D.R., Holt, D., Macpate, F., Lains, A., Hollowell, S., Cruickshank, B. and Casper, R.F. (2001). .Human UC-blood banking: impact of blood volume, cell separation and cryopreservation on leukocyte and CD34() cell recovery. Cytotherapy 3(4): 269–276. Rubinstein, P., Rosenfield, R.E., Adamson, J.W. and Stevens, C.E. (1993). Stored placental blood for unrelated bone marrow reconstitution. Blood Rev. 81(7): 1679–1690. Rubinstein, P., Dobrila, L., Rosenfield, R.E., Adamson, J.W., Migliaccio, G., Migliaccio, A.R., Taylor, P.E. and Stevens, C.E. (1995). Processing and cryopreservation of placental/umbilical cord blood for unrelated bone marrow reconstitution. Proc. Natl Acad. Sci. USA 92(22): 10119–10122. Salven, P., Mustjoki, S., Alitalo, R., Alitalo, K. and Rafii, S. (2003). GFR-3 and CD133 identify a population of CD34 lymphatic/vascular endothelial precursor cells. Blood 101(1): 168–172. Sanberg, P.R., Willing, A.E. and Cahill, D.W. (2002). Novel cellular approaches to repair of neurodegenerative disease: from Sertoli cells to umbilical cord blood stem cells. Neurotox. Res. 4(2): 95–101. Sanchez-Ramos, J.R., Song, S., Kamath, S.G., Zigova, T., Willing, A., Cardozo-Pelaez, F., Stedeford, T., Chopp, M. and Sanberg, P.R. (2001). Expression of neural markers in human umbilical cord blood. Expression of neural markers in human umbilical cord blood. Expression of neural markers in human umbilical cord blood. Exp. Neurol. 171(1): 109–115. Sato, J., Kawano, Y., Takaue, Y., Hirao, A., Makimoto, A., Okamoto, Y., Abe, T., Shimokawa, T., Iwai, A. and Kuroda, Y. (1995). Quantitative and qualitative comparative analysis of gradient-separated hematopoietic cells from cord blood and chemotherapy-mobilized peripheral blood. Stem Cells 13(5): 548–555. Schmidt, D., Mol, A., Neuenschwander, S., Breymann, C., Gossi, M., Zund, G., Turina, M. and Hoerstrup, S.P. (2005). Living patches engineered from human umbilical cord derived fibroblasts and endothelial progenitor cells. Eur. J. Cardiothorac. Surg. 27(5): 795–800. Sharma, A.D., Cantz, T., Richter, R., Eckert, K., Henschler, R., Wilkens, L., Jochheim-Richter, A., Arseniev, L. and Ott, M. (2005). Human cord blood stem cells generate human cytokeratin 18-negative hepatocyte-like cells in injured mouse liver. Am. J. Pathol. 167(2): 555–564. Shin, J.W., Lee, D.W., Kim, M.J., Song, K.S., Kim, H.S. and Kim, H.O. (2005). Isolation of endothelial progenitor cells from cord blood and induction of differentiation by ex vivo expansion. Yonsei. Med. J. 46(2): 260–267. Socie, G., Gluckman, E., Carosella E., Brossard, Y., Lafon, C. and Brison, O. (1994). Search for maternal cells in human umbilical cord blood by polymerase chain reaction amplification of two minisatellite sequences. Blood 83(2): 340–344.

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Solves, P., Mirabet, V., Larrea, L., Moraga, R., Planelles, D., Saucedo, E., Uberos, F.C., Planells, T., Guillen, M., Andres, A., et al. (2003a). Comparison between two cord blood collection strategies. Acta Obstet. Gynecol. Scand. 82(5): 439–442. Solves, P., Moraga, R., Saucedo, E., Perales, A., Soler, M.A., Larrea, L., Mirabet, V., Planelles, D., Carbonell-Uberos, F., Monleon, J., et al. (2003b). Comparison between two strategies for umbilical cord blood collection. Bone Marrow Transplant. 31(4): 269–273. Solves, P., Perales, A., Moraga, R., Saucedo, E., Soler, M.A. and Monleon, J. (2005). Maternal, neonatal and collection factors influencing the haematopoietic content of cord blood units. Acta Haematol. 113(4): 241–246. Solves, P., Fillol, M., Lopez, M., Perales, A., Bonilla-Musoles, F., Mirabet, V., Soler, M.A. and Roig, R.J. (2006). Mode of collection does not influence haematopoietic content of umbilical cord blood units from caesarean deliveries. Gynecol. Obstet. Invest. 61(1): 34–39. Storms, R.W., Green, P.D., Safford, K.M., Niedzwiecki, D., Cogle, C.R., Colvin, O.M., Chao, N.J., Rice, H.E. and Smith, C.A. (2005). Distinct hematopoietic progenitor compartments are delineated by the expression of aldehyde dehydrogenase and CD34. Blood 106(1): 95–102. Taguchi, A., Soma, T., Tanaka, H., Kanda, T., Nishimura, H., Yoshikawa, H., Tsukamoto, Y., Iso, H., Fujimori, Y., Stern, D.M., et al. (2004). Administration of CD34 cells after stroke enhances neurogenesis via angiogenesis in a mouse model. J. Clin. Invest. 114(3): 330–338. Tedder, R.S., Zuckerman, M.A., Goldstone, A.H., Hawkins, A.E., Fielding, A., Briggs, E.M., Irwin, D., Blair, S., Gorman, A.M., Patterson, K.G., et al. (1995). Hepatitis B transmission from contaminated cryopreservation tank. Lancet 346(8968): 137–140. Thierry, D., Traineau, R., Adam, M., Delachaux, V., Brossard, Y., Richard, P., Gerotta, A., Devergie, A., Benbunan, M. and Gluckman, E. (1990). Hematopoietic stem cell potential from umbilical cord blood. Nouv. Rev. Fr. Hematol. 32(6): 439–440. Tiumina, O.V., Savchenko, V.G., Gusarova, G.I., Pavlov, V.V., Zharkov, M.N., Volchkov, S.E., Rossiev, V.A. and Gridasov, G.N. (2005). Optimization of isolation of the concentrate of stem cells from the umbilical blood. Ter. Arkh. 77(7): 39–41. U-pratya, Y., Boonmoh, S., Promsuwicha, O., Theerapitayanon, C., Kalanchai, L., Chanjerboon, V., Sirimai, K., Visuthisakchai, S., Bejrachandra, S. and Issaragrisil, S. (2003). Collection and processing of umbilical cord blood for cryopreservation.. J. Med. Assoc.Thai. 86(11): 1055–1062. Vendrame, M., Cassady, J., Newcomb, J., Butler, T., Pennypacker, K.R., Zigova, T., Sanberg, C.D., Sanberg, P.R. and Willing, A.E. (2004). Infusion of human umbilical cord blood cells in a rat model of stroke dose-dependently rescues behavioral deficits and reduces infarct volume. Stroke 35(10): 2390–2395. Vendrame, M., Gemma, C., de Mesquita, D., Collier, L., Bickford, P.C., Sanberg, C.D., Sanberg, P.R., Pennypacker, K.R. and Willing, A.E. (2005). Anti-inflammatory effects of human cord blood cells in a rat model of stroke. Stem Cells Dev. 14(5): 595–604. Willing, A.E., Saporta, S., Sanberg, P.R., Justen, E.B., Haywood, A.N., Garbuzova-Davis, S.N., Dellis, J.T. and Cahill, D.W. (2002). Intravenous and intraspinal transplantation of umbilical cord blood cells in a mouse model of familial amyotrophic lateral sclerosis. Soc. Neurosci. (Abstract# 852.13). Wu, X., Rabkin-Aikawa, E., Guleserian, K.J., Perry, T.E., Masuda, Y., Sutherland, F.W., Schoen, F.J., Mayer Jr., J.E. and Bischoff, J. (2004). Tissue-engineered microvessels on three-dimensional biodegradable scaffolds using human endothelial progenitor cells. Am. J. Physiol. Heart Circ. Physiol. 287(2): H480–H487. Yoshida, S., Ishikawa, F., Kawano, N., Shimoda, K., Nagafuchi, S., Shimoda, S., Yasukawa, M., Kanemaru, T., Ishibashi, H., Shultz, L.D., et al. (2005). Human cord blood-derived cells generate insulin-producing cells in vivo. Stem Cells 23(9): 1409–1416. Zigova, T., Song, S., Willing, A.E., Hudson, J.E., Newman, M.B., Saporta, S., Sanchez-Ramos, J. and Sanberg, P.R. (2002). Human umbilical cord blood cells express neural antigens after transplantation into the developing rat brain. Cell Transplant. 11(3): 265–274. NCBP Diseases – Diseases and Demographics. (2005). http://www.nationalcordbloodprogram.org/patients/ncbp_ diseases.htm. https://web.emmes.com/study/cord/sop.htm

15 Multipotent Adult Progenitor Cells Catherine M. Verfaillie, Aernout Luttun, Karen Pauwelyn, Jeff Ross, Lepeng Zeng, Marta Serafini, Yuehua Jiang, and Fernando Ulloa Montoya PLURIPOTENT STEM CELLS: EMBRYONIC STEM CELLS Embryonic stem cells (ESCs) are pluripotent stem cells as they can be propagated indefinitely, and differentiate into cells of all three germ layers, shown by teratoma and embryoid body (EB) formation. Following blastocyst injection, mouse ESCs contribute to all somatic and germline lineages. ESCs are derived from the inner cell mass (ICM) of the blastocyst and are true pluripotent stem cells. Mouse ESCs express the cell surface antigen SSEA1 and human ESC SSEA4, and both are characterized by the expression of a number of relative ESC specific genes, including the transcription factors (TFs) Oct4 (Scholer et al., 1989), Rex1 (Ben-Shushan et al., 1998), Nanog (Chambers et al., 2003; Mitsui et al., 2003), and Sox2 (Avilion et al., 2003). Oct4 is expressed in the pre-gastrulation embryo, primordial germ cells, the ICM, and germ cells (Scholer et al., 1989; Rosner et al., 1990). While normal expression levels of Oct4 maintain mouse ESC self-renewal, a decrease in expression to 50% leads to trophectoderm differentiation, and an increase to levels 200% to primitive endoderm differentiation (Niwa et al., 2000). Oct4 promotes self-renewal by promoting transcription of genes such as Oct4 (Boyer et al., 2006) and Sox2 (Catena, 2004), and repressing genes such as Hand1 and Cdx2 that promote trophectoderm differentiation (Niwa et al., 2000). What regulates expression of Oct4 is still poorly understood although recent studies have shown that Sall4 (Zhang et al., 2006), Epas1 (Hif-2α) (Covello et al., 2006), SF1 (Botquin, 1998) and RAR (Botquin, 1998) activate the Oct4 promoter. The homeoprotein Nanog appears to be an equally essential component for early mouse development and ESC propagation. Nanog–/– mice do not develop an epiblast, and Nanog–/– ESCs differentiate into mesoderm and endoderm (Chambers et al., 2003; Mitsui et al., 2003). Nanog prevents ICM cells from differentiating into extra-embryonic endoderm by inhibiting genes such as Gata4 and 6 that promote primitive endoderm differentiation. Forced expression of Nanog in ESC results in LIF-independent proliferation, demonstrating its important role in maintaining ESC pluripotency (Chambers et al., 2003; Mitsui et al., 2003). Intricate TF binding networks involving Oct4, Sox2, and Nanog are involved in global transcriptional activation and repression in ESC. Using ChiP on ChiP assays, unique and overlapping promoter binding sites have been identified for Oct4, Sox2, and Nanog, that serve as positive or negative regulators of transcription (Boyer et al., 2006). These interactions are controlled by feed-forward loops, where initial regulators control other regulators with the option of converging and controlling downstream target genes. Others have used proteomics to identify Nanog partners (Wang et al., 2006). This technique identified Nanog-bound genes such as Oct4, as well as other TFs including Sall1 and Sall4. POSTNATAL TISSUE-SPECIFIC STEM CELLS: ARE SOME MORE THAN MULTIPOTENT? During gastrulation, the pluripotent cells in the ICM become restricted first to a specific germ layer and then to a specific tissue. The latter persist throughout adult life, and are termed multipotent stem cells.

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One of the surprising findings of the last 5–7 years is that classical adult stem cells, thought to be multipotent, may actually be more pluripotent. Since the late 1990s, likely more than 1,000 papers have been published wherein authors described that adult stem cells from a given tissue may under some circumstances be capable of becoming a cell of an unexpected tissue. Reports describing stem cell plasticity initially caused great excitement, as they challenged the concept that adult stem cells function solely to maintain the tissue of origin, and might therefore provide a source of easy accessible cells not marred in ethical considerations, that could be used to treat a number of degenerative and genetic diseases. For instance, hematopoietic stem cells (HSCs) have been reported to differentiate into a variety of cell types of endoderm (lung epithelium, intestinal epithelium, kidney epithelium, endocrine pancreas, liver, bile ducts) (Petersen et al., 1999; Lagasse et al., 2000; Theise et al., 2000; Krause et al., 2001; Wagers et al., 2002; Alvarez-Dolado et al., 2003; Ianus et al., 2003; Kale et al., 2003; Vassilopoulos et al., 2003; Wang et al., 2003), ectoderm (epidermis and neural cells) (Brazelton et al., 2000; Mezey et al., 2000; Krause et al., 2001; Priller et al., 2001; Wagers et al., 2002; Alvarez-Dolado et al., 2003; Weimann et al., 2003; Weimann et al., 2005) as well as into mesoderm derivatives other than blood cells (skeletal and cardiac muscle, endothelium) (Ferrari et al., 1998; Gussoni et al., 1999; Orlic et al., 2000; Jackson et al., 2001; LaBarge and Blau 2001; Orlic et al., 2001; Grant et al., 2002; Camargo et al., 2003; Corbel et al., 2003; Balsam et al., 2004; Murry et al., 2004; Kajstura et al., 2005). However, after the initial series of optimistic reports a number of reports have appeared that challenge the initial observation, or provide alternative explanations to the claim of greater potency of adult stem cells. For instance, there is evidence that stem cells, such as HSCs, may not only reside in the bone marrow (BM), but can also be present in other tissues (Jackson et al., 1999; Kawada and Ogawa 2001; McKinney-Freeman et al., 2002). A second explanation for the perceived plasticity of chiefly hematopoietic cells is fusion between the hematopoietic cells and certain host cells in vivo, a phenomenon known from hybridoma cell production, and also shown to occur in vitro between hematopoietic cells or neurospheres and ESC (Terada et al., 2002; Ying, et al., 2002). Indeed, a number of studies described fusion between cells of hematopoietic origin and hepatocytes, cardiomyocytes, skeletal muscle cells, and Purkinje cells in the brain (Wagers et al., 2002; AlvarezDolado et al., 2003; Balsam et al., 2004; Doyonnas et al., 2004; Weimann et al., 2005). In many instances the nucleus of the donor cell becomes partially reprogrammed with suppression of the hematopoietic program and activation of genes from which the donor cell fused (Wang et al., 2003; Cossu, 2004; Weimann et al., 2005). Others have presented relatively convincing evidence that not all apparent plasticity is due to cell fusion, including differentiation of hematopoietic cells to lung epithelial cells (Harris et al., 2004), and neuronal lineage cells into endothelial cells (Wurmser et al., 2004). However, the efficiency with which one stem cell appears to acquire the phenotype of a tissue cell different from the tissue of origin, whether via fusion or direct, is limited; and it remains to be determined if this would have clinical relevance. The two remaining possible explanations for the apparent ability of some adult stem cells to generate cells of a tissue lineage different from the tissue of origin are that stem cells with more pluripotent characteristics persist into adulthood, or that adult stem cells can be “reprogrammed,” via a process of de-differentiation and then re-differentiation to another lineage, or via a process of trans-differentiation. Since 2001, a number of papers have reported that cells with greater potency can be isolated in culture. These include the isolation of SKPs (skin-derived progenitors) (Toma et al., 2001), PMPs (pancreas-derived multipotent precursors) (Seaberg et al., 2004) and hFLMPCs (human fetal liver multipotent progenitor cells) (Dan et al., 2006) that can differentiate into cells of two germ layers. We isolated apparently more pluripotent stem cell from the BM of mouse, rat, human, and swine, as well as from brain and muscle tissue from mice (Reyes et al., 2001; Jiang et al., 2002a; Zeng et al., 2006), termed multipotent adult progenitor cells (MAPCs). Since the initial description of MAPCs, a number of other cell populations isolated by culture of BM, umbilical cord blood, placental tissue, and amniotic fluid have been described that have the ability to differentiate into cells of the three

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germ layers. They have been named marrow-isolated adult multilineage inducible cells (MIAMI cells) (D’Ippolito et al., 2004), human bone marrow-derived stem cells (hBMSCs) (Yoon et al., 2005), unrestricted somatic stem cells (USSCs) (Kogler et al., 2004), fetal stem cells from somatic tissue (FSSCs) (Kues et al., 2005) very small embryonic-like cells (VSELs) (Kucia et al., 2005, 2006; Kucia et al., 2007), pre-mesenchymal stem cells (preMSC) (Anjos-Afonso and, Blood 2007); multipotent adult stem cells (MASC) (Beltrami et al., Blood, 2007) and amniotic fluid stem cells (AFS) (De Coppi et al., Nat Biotech). Although the phenotype differs somewhat between these different cell populations, they have in common that they can be expanded extensively in vitro; that most of them express stem-cell specific genes such as Oct4; and that they can differentiate in vitro to cells with features of mesoderm, endoderm, and ectoderm. However, not all studies show this at the single cell level, and the proof of differentiation differs between publications. Moreover, few if any of the studies have shown that the more potent cells can also regenerate a tissue in vivo.

ISOLATION OF MAPCs In 2001 and 2002 we described the isolation of MAPC from BM of human, mouse, and rat. MAPC can be expanded in vitro without obvious senescence, and can at the single cell level give rise to cells of mesoderm, endoderm and ectoderm in vitro. We also demonstrated that a Rosa26 mouse-derived MAPC cell-line contributed to many somatic tissues of the mouse when injected in the blastocyst (Jiang et al., 2002b). Since the initial description of MAPC isolation, we have made changes to the culture method, with an initial aim to decrease the aneuploidy/polyploidy. We detect chiefly in mouse MAPC when maintained for prolonged periods of time in vitro (Breyer et al., 2006). Such aneuploidy/polyploidy is seen significantly less when MAPC from rat, swine, or human are cultured. MAPC isolation is now performed under hypoxic conditions: BM cells are plated at relatively high density on fibronectin coated plates in 5% O2 and 6% CO2. After approximately 1 month, cells are passed through a Myltenii column to remove CD45 cells and Ter119 cells, and cells subcloned at 10 cells/well. Clones of “MAPC” are identified based on morphology and Oct4 mRNA levels (q-RT-PCR), and expanded. This has led to the isolation of MAPCs that have significantly higher levels of Oct4, with ΔCTs compared with GAPDH of 6 for mouse MAPC and 2 for rat MAPC. For mouse ESC, the ΔCT compared with GAPDH is 3–4. In addition 90% of MAPCs thus isolated and maintained express Oct4 protein in the nucleus. The phenotype of mouse MAPC is B220, CD3, CD15, CD31, CD34, CD44, CD45, CD105, Thy1.1, Sca-1, E-cadherin, MHC classes I and II negative, epithelial cell adhesion molecule (EpCAM) low and c-Kit, VLA-6 and CD9 positive. For rat MAPC the phenotype is CD44, CD45, MHC classes I and II negative, but CD31 positive. By contrast cells isolated under the same 5% O2 conditions with a mesenchymal stem cells (MSCs)-like phenotype (MSC-like cells) have an Oct4 ΔCT compared with GAPDH of 15 for both mouse and rat cells, and cells express CD44 as well as MHC class I antigens. In mouse, such MSC-like cells do not express c-Kit but express CD34 whereas in rat, such MSC-like cells do not express CD31. To generate single cell-derived populations of MAPC, we subclone established MAPC lines at 0.8 cells/well. Such subcloning is not usually possible at the initial subcloning step, but has a 30% efficiency when cells initially subcloned at 10 cells/well are subsequently subcloned at 0.8 cells/well. It should be noted that isolation of MAPC from rodent BM, as well as human and swine BM is much more readily accomplished when young donors are used (6 weeks in mouse and rat; 40 days in swine; 10 years in humans). This is consistent with our observations as well as the observations from Anjos-Afonso and Bonnet (2006), and Kucia et al. (2005, 2006, 2007) that cells expressing Oct4 are much more frequent in the BM of young compared with older animals. We have used transcriptome analysis to compare MAPC with ESC and MSC. These studies (F UlloaMontoya et al., manuscript under revision) demonstrate that MAPCs cluster more closely to ESC than MSC or with cells isolated under MAPC conditions but with functional characteristics of MSC Ulloa-Montoya

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et al., Genome Biol, 2007. MAPCs express a number of genes identified to be relatively uniquely expressed in ESC (ES cell associated transcripts or ECATs) (Mitsui et al., 2003), including Oct4, Rex1, and eight other genes, whereas MSCs express only two of these ECAT genes at very low level. We previously reported that mouse and rat MAPCs need to be maintained at low cell densities for expansion in order to maintain pluripotent capacity. However, the clonal populations of rodent MAPC that express high levels of Oct4 can be maintained at higher cell densities. Even when seeded at 5000 cells/cm2 and passaged every 2 days, the growth rate of rat MAPC was not affected, Oct4 mRNA and protein levels remained stable, cell surface phenotype was unaffected and cell differentiation toward endothelium-like and hepatocytelike cells was unaffected by maintenance at higher cell densities for 48 days.

DIFFERENTIATION ABILITY OF MAPC IN VITRO We have demonstrated that mouse, rat, human, and swine MAPCs differentiate to mesenchymal type cells such as osteoblasts, chondroblasts, and adipocytes (Carmeliet et al., 2001; Reyes et al., 2001; Zeng et al., 2006). In addition, we have shown that MAPC can generate endothelial cells in vitro and in vivo (Reyes et al., 2001; Jiang et al., 2002b; Reyes et al., 2002; Zeng et al., 2006). Moreover, we have recently demonstrated that in contrast to human AC133 cells, human MAPC can be specified to arterial and venous endothelium (Aranguren et al., 2006). Similar results have also been obtained using mouse and rat MAPCs, where we have also shown specification to Prox1, VLA9, podoplanin, Lyve-1 and Mmr positive lymphatic endothelium (Luttun, A. et al., unpublished observations). MAPCs from human, swine, rat, and mouse BM can be induced to differentiate to a homogenous population of smooth muscle cells with phenotypic as well as many functional attributes of smooth muscle cells, including remodeling of extracellular matrix and contractile properties (Ross et al., 2006). Since the initial description of differentiation of MAPC to hepatocyte-like cells (Jiang et al., 2002b; Schwartz et al., 2002), we have performed additional studies demonstrating robust acquisition of phenotypic and functional characteristics of hepatocytes from rat MAPC. These culture conditions consist of initial induction of endoderm using Wnt3 and activin-A, induction of hepatic endoderm using sequentially the mesodermal derived cytokines BMP4 and FGF2 followed by FGF1, FGF4 and FGF8, and finally hepatocyte growth factor (HGF), follistatin and dexamethasone (Pauwelyn, K. et al., manuscript in preparation). This yields a population of cells wherein 10% express mature liver markers and that have several functional characteristics of hepatocytes including albumin and urea secretion, glycogen storage, bilirubin glucuronidation, and steroid metabolization. A similar protocol may also be effective at inducing differentiation of mouse and human ESC towards hepatic endoderm. ENGRAFTMENT OF MAPC IN VIVO We have transplanted MAPC in postnatal animals in a number of models. In 2002, we reported that grafting of MAPC in sublethally irradiated NOD-SCID mice results in low levels of engraftment in the hematopoietic system, even though no lymphoid reconstitution was seen (Jiang et al., 2002b). BM from primary recipients could also generate hematopoietic cells in secondary recipients. In those studies we also identified donor MAPC-derived epithelial cells in gut, liver, and lung (Jiang et al., 2002b). Anjos-Afonso and Bonnet (2006) also demonstrated that pre-MSC can generate hematopoietic cells in vivo when grafted in the femur. Since 2002, Tolar et al. demonstrated that engraftment of MAPCs that are MHC class-I negative is inhibited by natural killer (NK) activity (Tolar et al., 2006). This lead us to transplant two independent clones of MAPC expressing Oct4 at levels between 10% and 100% of mESC, derived from green fluorescent protein (GFP)transgenic mice, in sublethally irradiated NOD-SCID mice also treated with an anti-NK antibody for the first 3 weeks. We demonstrated that this results in multi-lineage hematopoietic reconstitution in 75% of animals, without evidence of fusion in the hematopoietic progeny. MAPC-derived KLS cells from primary recipients can rescue secondary C57Bl/6 mice from lethal irradiation and establish long-term hematopoiesis. The primary

262 CELLS AND TISSUE DEVELOPMENT

recipient mice have also evidence of presence of common myeloid progenitors (CMP) and common lymphoid progenitors (CLP) in the marrow (Serafini et al., 2007). MAPC-derived progeny cells that are CD45 negative can be found in multiple organs, although differentiation in a tissue-specific manner was not seen except for the hematopoietic system and the heart where GFP positive cardiac cells were detected following transplantation of high Oct4 MAPC, but not KTLS-HSC. One technical problem we encountered is that the most specific anti-GFP antibody (from Clonetech) did not stain all epithelial cells of the β-actin–GFP-transgenic animals from whom MAPC were isolated. Hence, the apparent lack of contribution to epithelial tissues following transplantation could be an indication that the MAPC used in Serafini et al. (2007) do not contribute to tissues other than blood and heart, or our inability to identify such contribution.

CONTRIBUTION OF MAPC TO CHIMERAS We evaluated the ability of MAPC to contribute to chimeras when injected in the blastocyst. Using the Rosa26 MAPC line described in Jiang et al. (2002b) we found chimerism in 80% of mice derived from blastocysts in which 10–12 MAPCs were injected and in 33% of mice derived from blastocysts microinjected with one MAPC. In both sets of animals, chimerism was low (in the 1–10% range for 76% and 71% of chimeras from 10–12 and 1 cell injections respectively). In 1 and 2 animals, respectively, from 10–12 and 1 cell injection, 40% chimerism was detected. Injection of additional cell lines in the blastocyst has yielded fewer chimeric animals, and chimerism was in general low. Nevertheless, injection of high Oct4 rat MAPC in 20 blastocysts yielded two embryos with chimerism in the 1–5% range at E10 gestational age (GFP positive by fluorescence microscopy); injection of a GFP-transgenic mouse MAPC line yielded embryos with 1–5% chimerism determined by q-PCR, and injection of 10–12 cells from a GFP-transgenic mouse MAPC line generated by Reyes, M. in the Chamberlain lab at the University of Washington in Seattle yielded three embryos with 1–10% contribution (determined by q-RT-PCR). MECHANISM UNDERLYING GREATER POTENCY OF MAPC AND SIMILAR ADULT STEM CELLS WITH GREATER POTENCY One question that has not been answered is whether the cell populations described above (SKPs, PMPs, hFLMPCs, MAPCs, MIAMI cells, hBMSCs, USSCs, FSSCs, AFS, MASCs, VSEL, and pre-MSCs) exist in vivo or are created in culture as the result of dedifferentiation. From all the cells described, SKPs have recently been isolated directly from skin without intervening culture step. Toma et al. showed that SKPs can also be derived freshly, without preceding culture, from fetal mice as well as from adult mice where they appear to reside in a niche in the hair papillae and whisker follicles (Fernandes et al., 2004). Anjos-Afonso and Bonnet (2006) found the SSEA1 antigen positive pre-MSCs that express high levels of Oct4 and can be expanded under MAPC conditions to generate cells capable of differentiating to the mesodermal, endodermal, and ectodermal lineage, and can contribute to hematopoiesis when grafted in vivo, can be isolated from mesenchymal cultures at passage 1. Compared with MAPCs, the cells isolated by Anjos-Afonso also expressed Nanog and Sox2. In addition, Kucia et al. (2006) demonstrated that a homogenous population of rare Sca-1 positive lineage negative, CD45 negative cells can be selected directly from BM of mouse and humans. These VSELs express like the cells identified by Anjos-Afonso and Bonnet (2006) and like ESC, SSEA-1, Oct-4, Nanog and Rex-1. The latter two studies suggest that rare cells exist in murine and human marrow with phenotypic features of MAPCs, MIAMI cells, hBMSCs, USSCs, AFS or FSSCs. Whether the differentiation ability ascribed to MAPCs and like cells (Jiang et al., 2002b; D’Ippolito et al., 2004; Kogler et al., 2004; Kues et al., 2005; Yoon et al., 2005; Anjos-Afonso and Bonnet 2006) is already present in the primary selected, uncultured BM cells isolated by Anjos-Afonso and Bonnet (2006) and Kucia et al. (2006), and hence represent cells with greater potency persisting in vivo into postnatal life, or whether the differentiation ability is acquired once cells are culture expanded in vitro, and therefore represent de-differentiation of a rare Oct4 positive cell, is not known. Interestingly, during the last year, 4 reports have been published demonstrating that

Multipotent Adult Progenitor Cells 263

mouse embryonic fibroblasts and tail clip fibroblasts can be reprogrammed towards cells with all ESC characteristics, by introduction of four transcription factors known to be expressed in ESC (Oct4, Sox2, Klf4 and c-Myc), and selecting for cells that start to express Nanog or Oct4. This provides proof of principle that adult cells can be reprogrammed. It should be noted that of the three of the four transcription factors used to reprogram fibroblasts are expressed in culture established mouse and rat MAPC (Ulloa-Montoya, 2007). Again, we do not know whether these genes were expressed in the fresh bone marrow cells prior to culture. The question as to whether MAPCs, and like cells, exist as such is not only of academic importance, but the answer may have profound biological implications as well as potential clinical applications. In vitro generated cells have tremendous potential clinical usefulness, as long as the cells can be generated in an efficient and reliable manner. If MAPCs exist as such in vivo it may one day be possible to manipulate their function in vivo, without the need for in vitro manipulation. Hence future studies should be aimed at determining whether MAPC and like cells exist in vivo, and if so what the optimal way of isolation and in vitro expansion is; and whether they could be mobilized and/or activated in vivo. If the answer is “No,” then it will be of the utmost importance to determine which cell population in a given tissue generate cells with greater potency in vitro, and develop strategies to select the precursor and induce with great efficiency the phenotype in vitro.

ACKNOWLEDGMENTS We acknowledge the support of the FWO (Odysseus fund) and the KUL COE funding.

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Theise, N.D., Badve, S., et al. (2000). Derivation of hepatocytes from bone marrow cells in mice after radiation-induced myeloablation. Hepatology 31: 235–240. Tolar, J., O’Shaughnessy, M.J., et al. (2006). Host factors that impact the biodistribution and persistence of multipotent adult progenitor cells. Blood 107: 4182–4188, (Epub January 12) Toma, J.G., Akhavan, M., et al. (2001). Isolation of multipotent adult stem cells from the dermis of mammalian skin. Nat. Cell Biol. 3: 778–784. Ulloa-Montoya, F., Kidder, B., Pauwelyn, K., Chase, L., Luttun, A., Crabbe, A., Sharov, AA., Piao, Y., Ko, MSH., Hu, W-S., Verfaillie, CM. (2007). Comparative Transcriptome Analysis of Embryonic and Adult Stem Cells with Extended and Limited Differentiation Capacity. Genome Biol. 8(8): R163. [Epub ahead of print] Vassilopoulos, G., Wang, P.R., et al. (2003). Transplanted bone marrow regenerates liver by cell fusion. Nature 422: 901–904. Wagers, A.J., Sherwood, R.I., et al. (2002). Little evidence for developmental plasticity of adult hematopoietic stem cells. Science 297(5590): 2256–2259. Wang, J., Rao, S., et al. (2006). A protein interaction network for pluripotency of embryonic stem cells. Nature 444: 364–368. Wang, X., Willenbring, H., et al. (2003). Cell fusion is the principal source of bone-marrow-derived hepatocytes. Nature 422: 897–901. Weimann, J.M., Charlton, C.A., et al. (2003). Contribution of transplanted bone marrow cells to Purkinje neurons in human adult brains. Proc. Natl Acad. Sci. USA 100: 2088–2093. Weimann, J.M., Johansson, C.B., et al. (2005). Stable reprogrammed heterokaryons form spontaneously in Purkinje neurons after bone marrow transplant. Nat. Cell Biol. 5: 959–966. Wernig, M., Meissner, A., Foreman, R., Brambrink, T., Ku, M., Hochedlinger, K., Bernstein, B.E., Jaenisch, R. (2007). In vitro reprogramming of fibroblasts into a pluripotent ES-cell-like state. Nature 448: 318–324. Wurmser, A.E., Nakashima, K., et al. (2004). Cell fusion-independent differentiation of neural stem cells to the endothelial lineage. Nature 430: 350–356. Ying, Q.Y., Nichols, J., et al. (2002). Changing potency by spontaneous fusion. Nature 416: 545–548. Yoon, Y.S., Wecker, A., et al. (2005). Clonally expanded novel multipotent stem cells from human bone marrow regenerate myocardium after myocardial infarction. J. Clin. Invest. 115: 326–338. Zeng, L., Rahrmann, R., et al. (2006). Swine bone marrow derived multipotent adult progenitor cells. Stem Cells 24: 2355–2366. Zhang, J., Tam, W.L., et al. (2006). Sall4 modulates embryonic stem cell pluripotency and early embryonic development by the transcriptional regulation of Pou5f1. Nat Cell Biol. 8: 1114–1123.

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16 Bone Marrow Stem Cells: Properties and Pluripotency Munira Xaymardan, Massimo Cimini, Richard D. Weisel, and Ren-Ke Li

INTRODUCTION The formation of new tissue in animals is generally confined to the embryonic and developmental stages. Regeneration of highly complex tissue does occur in some amphibians and reptiles. In mammals, however, healing of damaged tissue essentially results in the replacement of functional cells by highly fibrotic reparative tissue, which leads to diminished or even lost organ function. In the past 5 years, accumulating evidence has shown that multipotential stem cells are in fact present in many adult tissues. Bone marrow and tissue-specific stem cells can be induced to differentiate into adult cell types previously thought terminally differentiated, including cardiomyocytes, skeletal muscle cells, and neurons. Stem cells, by definition, have two characteristics: (1) the ability to self-renew and generate more stem cells through cell division and (2) under appropriate induction, the ability to give rise to clonal progency that continue to differentiate into one or more specialized cell types. In adults, bone marrow is a major reservoir for stem cells. Unlike totipotent stem cells (such as a fertilized egg), which can give rise to entire organism, bone marrow stem cells (BMSCs) are multipotent cells, which can give rise to most of the adult cell types, but not yet proven to be able to develop into a fetus (Stocum, 2001). Stem- and progenitor-based therapies are currently being developed for the treatment of cardiovascular diseases, which represent the major cause of death in the Western world (NIH, 2000). A number of clinical trials are underway to test the efficacy of local and systemic delivery of bone marrow-derived stem cells for the replacement of cardiomyocytes and vascular endothelial cells (Britten et al., 2003; Perin and Silva, 2004) (Table 16.1). Preliminary results are suggestive, but their widespread application necessitates a thorough understanding of the mechanisms of cellular replacement in order to optimize the efficient use of BMSCs for vascular repair and cardioprotection.

BONE MARROW STEM CELLs Bone marrow is hematopoietic tissue that lies within the trabecular bone. The trabecular and the bone marrow stroma are the elements that physically support and physiologically maintain the hematopoietic tissue. In adult humans, bone marrow is the site for production of all hematopoietic cells; the supporting stroma consists of reticular cells, osteocytes, adipocytes, vascular endothelium, and extracellular matrix. And together with the blood vessels, the bone marrow forms a hematopoietic inductive microenvironment that controls adult hematopoiesis, where five billion blood cells are produced every day. The vascular structure of bone marrow consists of sinusoidal vasculature in which the endothelial cells do not have subsequent encapsulation of other types of cells; this is highly permissive for the emigration and immigration of the bone marrow cells.

268

Table 16.1 Summary of clinical trials Author

Trial name Disease

Perin and Silva (2004)

Ischemic heart failure

Trial size Length of follow-up

Cell source and type

Delivery route

Outcome

21

4 months

BM, mononuclear TransIncrease cells endocardial LVEF

Kang et al. (2004)

MAGIC; 2 days randomizedcontrolled

27

6 months

G-CSF  CPC CPC

Schachinger et al. (2004)

TOPCAREAMI

AMI, 4.7 /1.7 days

59

6 months and 1 year

BM, CPC (contain IntraCD133; CD34) coronary

Improved global and regional contractility; decreased MI size

Wollert et al. BOOST; (2004) randomcontrolled

AMI, 4.5 days

60

6 months

BM, CD34

Intracoronary

Improved global LVEF

Lunde et al. (2005)

AMI, 5–8 days 100

1 year

BMC

Intracoronary

No benefit

18 months plus

BMC

Intracoronary

Benefit up to 18 months

ASTAMI Randomcontrolled

Cleland et al. REPAIR-AMI; 4 days after (2006) randomAMI controlled

204

Intracoronary

G-CSF induced restenosis

Cx-non-treatment control; AMI-acute myocardial infarction; BM-bone marrow; CPCs-circulating blood progenitor cells; LV-left ventricular; EF-ejection fraction.

Hematopoietic Stem Cells Hematopoietic stem cells (HSCs) are the stem cells from which all red and white blood cells develop. They are

entirely responsible for the development, maintenance, and regeneration of the blood forming tissue for life (Weissman, 2000). Because HSCs can reconstitute and restore the hematopoietic system of a myeloablated host, they have been traditionally used for treating hematologic disorders, starting in 1945 (Gengozian and Makinodan, 1956), when donor-derived HSCs were first used to protect a lethally irradiated civilian population. In adult mouse bone marrow, HSC activity has been shown in a cell population marked by c-kitpos, thy-1low, and sca-1pos (Bradfute et al., 2005). In adult humans, HSCs are marked by c-kitpos, thy-1pos, and CD34pos (Weissman, 2000). HSCs from mice and humans are being isolated, starting with a lineage depletion step in which all the lineage-specific cells (B220, CD3, 4, 8, 11b Mac-1, Gr-1 and Tcr-119 for mice and CD10, 14, 15, 16, 19, and 20 in human) are removed (Figure 16.1). The resultant population, referred to as Linneg, can be enriched 10–100-fold, and is able to re-populate of bone marrow of a lethally irradiated host. In vitro expansion of HSCs can be achieved by co-culturing them with stromal cells from bone marrow. Researchers have found several subpopulations within Linneg HSCs. One homogenous population is characterized as side population (SP) cells based on their unique ability to extrude hoechst dye. When examined by fluorescence activated cell sorter (FACS) analysis, SP cells fall within a separate population to the side of the rest of the cells on a dot plot of emission data. SP cells express the ABCG2 transporter, a transmembrane protein, which allows them to actively exclude hoechst dye and fluoresce in this specific manner. These cells are also able

269

270 CELLS AND TISSUE DEVELOPMENT

Hematopoietic stem cell

Positive for

Negative for

Mesenchymal stem cell

Thy-1 Thy-1Lo Sca-1 C-Kit CD34

Mouse

Human

B220 CD3 CD4 CD8 CD11b Mac-1 Gr-1 Tcr119

CD10 CD15 CD16 CD19 CD20

Sro-1

Human

Mouse

CD13 CD49a CD49b CD29 CD44 CD71

CD90 CD106 CD16 CD54 CD55 CD124

CD34 CD45 CD14 CD14

Isolation

HSCs are normally purified using a fluorescence cell sorting system or antibody conjugated magnetic beads to deplete all committed cell types by negative selection followed by positive selection of targeted cells

MSCs are typically isolated from the mononeclear layer of the bone marrow after separation by discontinuous gradient centrifugation. In some cases, further purification is performed based on MSC markers, such as STRO-1

Figure 16.1 Isolation of HSCs and MSCs from bone marrow. to home rapidly to the bone marrow of a lethally irradiated host (Goodell et al., 1996) and contribute progeny to the lung and liver in irradiated mice, and infiltrate into the infarcted heart (Abe et al., 2003). SP cells are also present in other tissues, including skeletal muscle and skin (Liadaki et al., 2005). Data are conflicting: some suggest that SP cells can be tissue-specific stem cells within these organs, and others suggest that are actually bone marrow-derived SP cells lodged within these tissues. Another group of highly plastic stem cells isolated from bone marrow are known as bone marrowderived stem cells (Leone et al., 2005). Several recent studies indicate that these cells are highly plastic, exhibiting tremendous differentiation activity in numerous non-hematopoietic organs. It is unclear whether these populations are enriched for pre-hematopoietic cells that maintain greater pluripotentiality than HSCs. An additional possibility is that a differentiated hematopoietic cell, such as a macrophage, may be able to assume the gene expression pattern of a different cell type by fusion (Ozturk et al., 2004). Mesenchymal Stem Cells Mesenchymal stem cells (MSCs) are stem cells found in bone marrow, from where they can generate bone,

cartilage, fat, and fibrous connective tissue. They are the non-hematopoietic, structural components of bone marrow that support hematopoiesis by providing extracellular matrix components, cytokines, and growth factors. MSCs represent 0.001–0.01% of the bone marrow cell population, and in culture as clonal, plastic adherent cells that assume a spindle cell morphology with a finite life span. Friedenstein first discovered MSCs in 1970s (1978), when his laboratory was able to culture these cells in media and induce them to differentiate into multilineage cell types, including osteoblasts, chondroblasts, and adipocytes, in response to

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appropriate stimuli. No specific constellation of surface markers has been agreed upon for these cells, but human MSCs are typically isolated from the mononuclear layer of the bone marrow after separation by discontinuous gradient centrifugation. In some cases, further purification is performed based on MSC markers, such as STRO-1 (Gronthos et al., 1994). Other surface antigens reported to exist on MSCs are: CD13 (aminopeptidase-N), CD49a and CD49b (integrins-alpha), CD29 (integrin-beta), CD44 (hyaluronate), CD71 (transferrin), CD90 (thy-1), CD106 (vascular cell adhesion molecule-1), CD166 (activated leukocyte cell adhesion molecule), CD54 (intercellular adhesion molecule-1), CD55 (decay accelerating factor), and CD124 (interleukin-4 (IL-4) receptor). MSCs uniformly lack antigens CD34, CD45, CD14, and CD31 that typically identify hematopoietic cells (Pittenger and Martin, 2004) (Figure 16.1). A wide array of cytokines, including fibroblast growth factor-2 (FGF-2), FGF-4, platelet-derived growth factor-BB (PDGF-BB), and leukemia inhibitory factor (LIF), have been used to expand MSCs (Gregory et al., 2005). Because MSCs are easily expandable in culture and differentiate into multiple tissue lineages, there has been much interest in their clinical potential for tissue repair and gene therapy. In particular, a population of highly plastic, adult-derived bone marrow cells, referred to as multipotent adult progenitor cells (MAPCs), can be grown in vitro from the postnatal marrow (and other organs) of mice, rats, and humans. These cells co-purify initially with MSCs and grow as adherent cells in vitro (Reyes et al., 2001). However, unlike MSCs, MAPCs can be cultured indefinitely in a relatively nutrient-poor medium (Jiang et al., 2002). Specific changes in growth factors induce differentiation of MAPCs into myoblasts, hepatocytes, and even neural tissue (Jiang et al., 2002; Schwartz et al., 2002). Endothelial Progenitor Cells Endothelial progenitor cells (EPCs) are a group of non-endothelial cells that can give rise to endothelial cells.

Stemness of the cells is not clear, but they can be expanded, and increasing evidence shows that EPCs play a major role in postnatal neovascularization. Bone marrow HSCs and MSCs, as well as other tissues (fat, cord blood, and circulating blood), are the sources of the EPCs, of which HSC-derived EPCs are perhaps the best characterized. HSC-derived EPCs are maintained in the BMSC niche and are released upon mobilization with cytokines such as vascular endothelial growth factor (VEGF) or stromal cell-derived factor-1 (SDF-1), which are synthesized by ischemic tissue (Leone et al., 2005). Indeed, Asahara et al. (1999) demonstrated that bone marrow-derived HPCs give rise to endothelial cells and contribute to endothelial recovery and new capillary formation after ischemia. EPCs have been subsequently defined as cells that express HSC markers such as CD34 or CD133, and an endothelial marker protein, VEGF receptor 2 (VEGFR2 or flk-1). Isolated cells express the classic HSC marker protein CD34 or the more immature HSC marker protein CD133. Both cell populations differentiate to endothelial cells in vitro under appropriate endothelial differentiation-promoting factors (Gehling et al., 2000). Most importantly, injection of CD34pos or CD133pos cells enhanced neovascularization in animal models after ischemia (Asahara et al., 1999). Likewise, MSCs can differentiate into endothelial cells (Oswald et al., 2004) and improve neovascularization in vivo (Pittenger and Martin, 2004). Because MSCs can release a variety of angiogenic growth factors, this cocktail of growth factors may also act in a paracrine manner to support angiogenesis and arteriogenesis. Verfailliea’s group (Reyes et al., 2002) reported that MAPCs that co-purify with MSCs isolated from postnatal human bone marrow can differentiate into cells that express endothelial markers, function in vitro as mature endothelial cells, and contribute to in vivo neoangiogenesis during tumor angiogenesis and wound healing (Reyes et al., 2002). Interaction of Bone Marrow Cells and Stem Cell Niches The stem cells in the bone marrow are not randomly distributed. They reside in specific compartments consisting of support cells known as niche, the microenvironment, which in turn controls the fate of the stem cells.

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The concept of a stem cell niche was first proposed for the human hematopoietic system in the 1970s (Schofield, 1978). At present, the hematopoietic niche is conceptually divided into two parts: an osteoblastic niche and a vascular niche. The osteoblastic niche located near the trabecular bone is a hypoxic environment that hosts quiescent state HSCs (slow cycling or G0), whereas the vascular niche located near sinusoids is an oxygenic niche, where stem/progenitor cells actively proliferate. The presence of osteoblasts not only sustains the bone but is required for the maintenance and expansion of HSCs through interaction of N-cadherin and betacatenin (Zhang et al., 2003). Other soluble and membrane bound proteins which are required for self-renewing within the niches are: mKirre, the Wnt proteins, stem cell factor (SCF), and bone morphogenic proteins (BMP) such as BMP-4 (Ueno et al., 2003; Zhang et al., 2003; de Boer et al., 2004). The limiting factor of HSC selfrenewal is perhaps the space within the niche (Zhang et al., 2003). As a niche is filled with stem cells, the excess cells are pushed into the adjacent vascular niche, which fosters the maturation of the HSCs and where HSCs finally mature and egress from marrow into the peripheral circulation via the bone marrow sinusoids. Hemotopoiesis in the vascular niche is partly regulated by growth factors (cytokines), particularly by ILs and colonal stimulating factor (CSF) (Barria et al., 2004), both of which stimulate the proliferation and maturation of the HSCs. The growth factor binding ligands are tyrosine kinase receptors such as c-kit, flt-3, and thrombopoietin, and all are expressed on primitive hematopoietic cells. The major inhibiting factors of hematopoiesis are perhaps transforming growth factor-beta (TGF-β), and tumor necrosis factor alpha (TNF-α).

BMSCS AND TISSUE REGENERATION BMSCs are multipotential in that they not only act as myelo-regenerative and supportive cells, but they also can differentiate into multilineage cell types. HSCs are capable of differentiating into endothelial cells and have also demonstrated an ability to differentiate into liver cells, skeleton muscle cells, and cardiac cells. Increasing evidence indicates that MSCs can differentiate into functional cells and repair damaged tissue. MSCs have been demonstrated to adopt osteoblasts, chondrocytes, and adipocytes in vitro (Friedenstein et al., 1978). When implanted in vivo, they are able to help repair multiple tissues including blood vessels, heart, liver, kidney, and muscle (Pittenger and Martin, 2004). Their ability to generate almost all the mesenchymal lineages of connective tissues has strengthened the idea that MSCs represent, or at least contain, a population of stem cells from which all mesenchymal lineages originate under the influence of different microenvironments. Understanding the molecular signals that underlie the process of bone marrow cell differentiation, and moreover, controlling the microenvironment, will help advance cell-based therapies (Figure 16.2). In cases where female animals or female human patients have received a male donor bone marrow transplant, tracing of a bone marrow cell that differentiated into multiple tissue types is achieved through fluorescent in situ hybridization (FISH) techniques to detect the Y-chromosome (Deb et al., 2003). Alternatively, wild-type animals may be transplanted with green fluorescence protein cells, which are easily detected using a fluorescence microscopy (Orlic et al., 2001b). BMSCs and Heart Regeneration Ventricular remodeling following an acute myocardial infarction leads to ventricular dilatation and progressive heart failure. The remodeling process is characterized by the removal of necrotic cardiac cells accompanied by granulation tissue formation with the simultaneous induction of neovascularization in the peri-infarct bed. The latter is a prerequisite for the survival of surrounding hypertrophic but viable cardiomyocytes, and the prevention of further cardiomyocyte loss by apoptosis. Ultimately, the remodeling process culminates in the formation of a non-contractile fibrous scar, which may expand, leading to further cardiac deterioration and heart failure (Chandrashekhar, 2005).

Bone Marrow Stem Cells: Properties and Pluripotency 273

Fat cell

Osteoblast

Osteocyte

Stromal cells

HSC Vessel

Hematopoietic stem cell

Blood cells

Liver cells

MSC Mesenchymal stem cell

Bone cells

Skeleton muscle cells

Cardiac muscle cells

Nerve cells

Skin cells

Figure 16.2 Pulripotency of BMSCs. Historically, the adult heart has been viewed as a terminally differentiated organ without the capacity of self-renewal or regeneration. But recent data challenges this doctrine, suggesting the existence of innate mechanisms for myocardial regeneration. Studies have shown evidence of low-level mitotic activity in the normal human myocardium, and proliferation of cardiomyocytes increases in the heart with end-stage ischemic disease (IHD) (Beltrami et al., 2001). The most intriguing finding perhaps is the data from Quaini et al. (2002), showing that cardiac regeneration following orthotopic heart transplantation. Using the Y-chromosome as a marker, this study found recipient-derived cardiomyocytes and vascular structure within the donor hearts of male patients who had received female donor hearts. BMSCs are considered to be the major contributors to the regeneration of cardiac tissue. Supporting this hypothesis are reports from female patients who received sex-mismatch bone marrow transplantation. The female hearts were examined for Y-chromosome and results confirmed the presence of bone marrow cells within the myocardium (Deb et al., 2003). Both HSCs and MSCs are reported to have the ability to repair damaged hearts. The establishment of a cardiomyogenic cell line from murine bone marrow MSCs marks a typical example of bone marrow differentiation into cardiomyocyte (Tomita et al., 1999; Fukuda, 2000). Fukuda’s group repeatedly passaged bone marrow cells until a single clone of immortalized homogenous fibroblast-like cells was obtained. After prolonged treatment with the DNA demethylating agent 5-azacytidine, the cells formed myotubes connected by intercalated disks, which beat synchronously after 3 weeks of culture. Ultrastructurally, the differentiated myotubes had well-organized sarcomeres, central nuclei, and contained atrial granules and mitochondria.

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The cardiomyocyte phenotype was further confirmed by both electrophysiological and cardiac-specific gene expression studies. These included the identification of both sinus node-like action potentials and ventricular cardiomyocyte-like action potentials. Differentiated myotubes expressed atrial natriuretic peptide (ANP), brain natriuretic peptide (BNP), low levels of a-MHC and a-cardiac actin, and high levels of b-MHC and a-skeletal actin, as well as MLC-2v, which was consistent with a fetal ventricular phenotype. The cells have also been shown to express functional adrenergic and muscarinic receptors, which mediate heart rate, conduction velocity, contractility, and cardiac hypertrophy (Fukuda, 2000). 5-xathydidine treated cardiomyocyte cell lines have been shown to establish stable cardiac engraftment and site-specific differentiation in myocardial scar tissue in the rat cryo-injury model of infarction (Tomita et al., 1999). These findings are also supported by rat myocardial infarction and porcine myocardial ischemia models (Bittira et al., 2002; Moscoso et al., 2005). Other experiments have shown that bone marrow cell interaction with neonatal cardiomyocytes or cellular extract may induce cardiomyogenic differentiation of the BMSCs. For example, bone marrow c-kit cells express cardiac markers when co-cultured with neonatal cardiomyocyte (Lagostena et al., 2005). Anversa’s group isolated Linneg/c-kitpos cells from adult bone marrow. When injected into an ischemic heart, these cells reconstituted well-differentiated myocardium formed by blood-carrying new vessels and myocytes with the characteristics of young cells (Beltrami et al., 2003). The results of some studies cast doubt on the ability of HSCs to adopt cardiac myocyte phenotypes in vivo. For example, Murry et al. isolated HSCs from mice carrying the alpha-cardiac myosin heavy chain promoter driving nuclear-localized enhanced green fluorescence protein (EGFP), and delivered these cells into mice after acute myocardial infarction. Unfortunately, neither systemic delivery nor direct injection of HSCs produced myocyte regeneration. Most studies have shown improved cardiac function by exogenous delivery of HSCs. This implies that an alternative mechanism of the HSC cardiac repair may be due to the paracine system: BMSCs secrete growth factors that augment angiogenesis, which in turn improves the remodeling process associated with cardiac regeneration. BMSCS and Skeletal Muscle Regeneration BMSCs are reported to differentiate into skeletal myoblasts. Human BMSCs are shown to differentiate into multinucleated myotubes in culture (Bossolasco et al., 2004). Moreover, direct injection of human whole bone marrow into the right tibialis anterior muscle of immunodeficient mice previously been treated with cardiotoxin to induce muscle degeneration showed a variable but significant level of human cell engraftment (Bossolasco et al., 2004). BMSCS and Bone Regeneration The osteogenic lineage is considered a default pathway of in vitro differentiation of the bone marrow stromal cells. Indeed, regeneration of the bone tissue has been successfully used in the clinical practice. The earliest studies used clonal forming unit fibroblasts (CFU-F) like cells to form bone structures in culture (Friedenstein et al., 1978). The phenomenon is also observed in the stro-1 population of MSCs (Gronthos et al., 1994). The bone forming ability of these cells has also been tested in diffusing wound chambers in rabbit models. When cells are isolated and expanded in the presence of FGF-2, the frequency of clones able to differentiate into the osteogenic, chondrogenic, and adipogenic lineages is greater than in the other lineages. BMSCS and Liver Regeneration BMSC engraftment to hepatocytes using male-to-female bone marrow transplantation in rats and mice (Fujii et al., 2002) was first demonstrated in response to liver damage, which may promote BMSC-to-hepatocyte transition. In rats, a combination of hepatotoxin, which induces widespread liver damage, and 2-acetylaminofluorine, which prevents endogenous liver repair, was used. A combination of Y-chromosome FISH and transgene expression was then employed to confirm that BMSCs were the source of the resultant hepatocytes. The

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effect on other forms of liver damage could be assessed in liver samples from men who received orthotopic liver transplants from female donors. In these patients, the degree of subsequent damage to the transplanted liver correlated with the extent of male (host-derived) hepatocyte engraftment (Theise et al., 2000). BMSCS and Nerve Cell Regeneration Two different systems show that bone marrow-derived stem cells can serve as progenitors of non-hematopoietic cells in the murine central nervous system (CNS). In one study, lethally irradiated adult mice that received whole marrow intravenously developed donor-derived brain cells bearing the neuronal antigens NeuN and class 3 b-tubulin (Brazelton et al., 2000). Similarly, adult rat and human BMSCs induced the stromal cells to exhibit a neuronal phenotype, expressing neuron-specific enolase, NeuN, neurofilament-M, and tau (Woodbury et al., 2000). Bone Marrow to Kidney, Pancreas, Lung, and Gastrointestinal Tract Similarly, BMSCs can also differentiate in vivo into pancreas islet cells, lung clara cells, and GI crypt cells, which are the functional stem cells of the gastrointestinal epithelium.

IMPORTANT FACTORS REGULATING BMSC HOMING AND DIFFERENTIATION As discussed in the previous sections, there is an accumulating body of evidence suggesting that stem and progenitor cells have the potential to regenerate and revascularize injured tissue. Stem cell-mediated cardiac repair involves three components: (1) the bone marrow as a stem cell reservoir; (2) the injured myocardium as the area where repair is required with the release of mediating factors; and (3) the circulation for transport of the signals and stem cells from the bone marrow to the injured myocardium (Vandervelde et al., 2005). Upon myocardial injury, molecular pathways are upregulated immediately, followed by streams of chemical mediators (cytokines and chemokines) that are released into circulation to help recruit BMSCs and allow for their homing to the distal injured myocardium (Vandervelde et al., 2005). Alternatively, exogenous BMSCs can be delivered directly to the injured site. Repair may have three different foci: (1) the vasculature, (2) the cardiomyoctes, and (3) the stability of the extracellular matrix. Together, these components orchestrate the signaling, mobilization, homing, incorporation, survival, proliferation, and differentiation of stem cells – a progression that involves a dynamic process of metalloproteinase activity, adhesion molecules, and remodeling of the extracellular matrix. The cytokines and chemokines involved may be classified according to function as mediators of homing and mobilization, inflammation (to aid in incorporation), survival and differentiation of stem cells (Figure 16.3). Granulocyte colony-stimulating factor (G-CSF) and VEGF are among the best-characterized cytokines for mobilization of BMSCs and EPCs to the site of injury. Vascular Endothelial Growth Factor VEGFs are a group of secreted proteins produced by almost every cell type; they appear to be the most prominent protein that guides vascular growth during vasculargenesis and angiogenesis (Carmeliet et al., 1996). However, the angiogenic capabilities of VEGF often overshadow its importance in the mobilization of BMSCs. Patients with high levels of plasma VEGF were found to have an increase in BMSCs in the heart, indicating the ability of VEGF to recruit stem cells post-myocardial infarction (Kamihata et al., 2001). In animal models, infusion of BMSCs was related to reduction in infarct size; the effect was attenuated with neutralizing antibodies to either VEGF and by increasing its soluble receptor VEGF-R1 (Flt-1) (Hiasa et al., 2004). Injection of VEGF plasmid DNA has also been documented to have mitogenic effects on porcine cardiomyocytes (Laguens et al., 2002). Furthermore, naked plasmid DNA directly injected into the ischemic myocardium

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5. Engrafting and regeneration/repair of myocardium

Myocardium

1. Cardiac injury

4. Homing of BMSCs to the myocardium

Circulation

2. Release of chemical mediators

VEGF G-CSF EPCs

SCF

Other factors MSCs

Bone marrow

C-kit

Other cell types

3. Mobilization of BMSCs

Figure 16.3 The stages of BMSC recruitment involved in regeneration of the damaged myocardium. of symptomatic myocardial ischemia patients led to a reduction in symptoms and improved myocardial perfusion (Laguens et al., 2002). Recent study shows that the VEGF is sufficient for organ homing of BMSCs to the perivascular area (Grunewald et al., 2006). VEGF may therefore enhance homing, mobilization of BMSCs, and augment cardiomyocyte proliferation. G-CSF G-CSF is a hematopoietic factor that stimulates neutrophils and BMSC mobilization through cleavage of

intercellular adhesion molecule-1, thereby disrupting the homing mechanism of the stem cells in the bone (Levesque et al., 2001). G-CSF is also involved in the proliferation, differentiation, and survival of bone marrow-derived stem and progenitor cells. The mobilization properties of G-CSF are widely utilized in clinical stem cell therapies with promising results seen in most of the trials (Kang et al., 2004; Valgimigli et al., 2005). Indeed, G-CSF has been shown to increase the number of in CD34pos cells in the circulation from 5- to 30-fold (Powell et al., 2005). When Kocher et al. isolated circulating human CD34pos cells released by G-CSF treatment and injected it into the infarcted hearts of nude rats, they found these CD34pos cells demonstrated phenotypic and functional properties of embryonic hemangioblasts stimulating neoangiogenesis in the infarct vascular bed (Kocher et al., 2001). G-CSF treatment was also found to increase the density of macrophages and neutrophils which may enhance the absorption of necrotic tissue in post-infarct myocardium, and coincide with proliferating cardiomyocytes and improved cardiac function; these results suggest additional pathways for G-CSF treatment (Minatoguchi et al., 2004). Further, G-CSF may act on non-hematopoietic cells that may potentially serve as the origin of the bone marrow-derived cardiomyocytes observed in the mouse myocardial infarction model (Kawada et al., 2004).

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Stromal Cell-Derived Factor-1 Stromal cell-derived factor-1 (SDF-1) (CXCL12) is produced by bone marrow stromal cells; its function is

to promote homing and engraftment of HSCs within the recipient bone marrow (Peled et al., 2000). SDF-1 transcription is partially controlled by hypoxia inducible factor-1 (HIF-1), upregulated by hypoxia during vascular injury (Ceradini and Gurtner, 2005). It therefore seems logical that SDF-1 forms a gradient from the hypoxic to the oxygenic bone marrow compartments. Studies have shown that blockage of SDF-1 binding to its receptor, CXCR4, inhibits stem cell homing to the infarcted heart, strongly suggesting that SDF-1/CXCR4 interactions play a crucial role in the recruitment of BMSCs to the heart after myocardial infarction (Abbott et al., 2004). Interestingly, a non-hematopoietic CXCR4pos population in the bone marrow has been found to also express early cardiac progenitor markers such as Nkx2.5/Csx, GATA-4, and MEF2C. This population can be mobilized into the peripheral blood after experimental myocardial infarction, providing a possible therapeutic target for myocardial regeneration (Kucia et al., 2004). Stem Cell Factor SCF, also known as c-kit ligand or steel factor, binds to its receptor c-kit and induces chemotactic properties in stem and progenitor cells (Chute et al., 2005). SCF is abundantly expressed in the normal bone marrow and heart, but it is downregulated following myocardial infarction (Woldbaek et al., 2002). Orlic et al. demonstrated that combined SCF and G-CSF treatment synergistically improved mouse cardiac function after myocardial infarction via mobilization of the BMSCs (Orlic et al., 2001); however, other groups failed to reproduce the described effect (Norol et al., 2003; Ohtsuka et al., 2004). Since SCF is produced by infiltrating macrophages in the ischemic myocardium and attracts mast cell precursors, SCF treatment for increased BMSCs may have a detrimental hyper-inflammatory effect (Frangogiannis et al., 1998). Interleukin-8 IL-8 is a member of small chemokine CXC family. It is upregulated by pro-inflammatory cytokines, like SCF-1, and is an important factor for stem cell proliferation in the bone marrow niche and promoting rapid mobilization of BMSCs. On the one hand, IL-8 promotes endothelial cell migration to sites of injury duration (Fibbe et al., 2000). On the other hand, it activates neutrophil adhesion to cardiomyocytes, subsequently promoting cardiomyocyte death (Kukielka et al., 1995). Similar to SCF, IL-8 may have detrimental effects on the cardiac tissue. Transforming Growth Factor-Beta and Bone Morphogenetic Proteins TGF-βs and bone morphogenetic proteins (BMPs) constitute a single morphogenic protein super family involved in cardiogenesis (Zaffran and Frash, 2002). Both have been demonstrated to induce embryonic stem cell differentiation into a cardiogenic phenotype (Behfar et al., 2002). Recently, it has been demonstrated that CD117pos cells partially positive in various fractions for Lin, CD34, and Sca-1, and negative for Lin, were able to undergo cardiomyogenic differentiation when treated with TGF-β1 (Li et al., 2005). Furthermore, transplantation of these cells into the infarcted region expressed ventricular heavy chain myosin, reduced fibrosis, and improved shortening. Unfortunately, angiogenesis, wall thickness, and LVEDD/LVESD did not significantly differ between the untreated and preprogrammed c-kit transplanted cells. Similar affects have also been reported in MSCs treated with BMP-2 and FGF-4 in a rat model of experimental myocardial infarction (Yoon et al., 2005), and hepatic growth factor (HGF) and platelet-derived growth factor-B (PDGF-B) were reported to stimulate BMSC survival and differentiation in experimental bone marrow transplantation models.

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CLINICAL APPLICATIONS OF BMSCS FOR CARDIAC REGENERATION Based on the ex vivo confirmation of bone marrow plasticity and the idea of possible cardiomyocyte regeneration, numerous clinical trails have been initiated to augment this process by transplanting exogenous bone marrow cells into damaged myocardium in patients with acute myocardial infarct or ischemic heart failure. Both HSC and MSC populations have been used in patients, with delivery methods including: (1) intra-coronary delivery (Britten et al., 2003; Kang et al., 2004); (2) direct injection into myocardium (Stamm et al., 2003; Perin and Silva, 2004; Pompilio et al., 2004); and (3) G-SCF-mediated BMSC mobilization (Kang et al., 2004). Most of the techniques described in these studies have been combined with conventional treatments, including surgical revascularization like angioplasty and stenting. Although some authors indicated that the cells used in these studies contained CD133 and/or CD34 populations, most of the cells and treatments were not clearly characterized, and their fates are undetermined. The results, however, (summarized in Table 16.1) have demonstrated that bone marrow transplantation in ischemic heart disease patients is safe and feasible, with the exception of a single report by Kang et al. (2004), which showed increasing restenosis in patients treated with G-CSF. While G-SCF release of smooth muscle progenitor cells may contribute to increased in-stent restenosis, transplantation of bone marrow cells into the ischemic heart augments angiogenesis and improves cardiac function. Most trials did not find significant risk in patients receiving bone marrow treatments. Indeed, most of the short-term trials have shown improvements in left ventricular ejection fraction and other functional parameters tested (Britten et al., 2003; Stamm et al., 2003; Kang et al., 2004; Perin and Silva, 2004; Pompilio et al., 2004; Kuethe et al., 2005). However, more recently, randomized, controlled clinical trials have produced controversial results. The REPAIR-AMI study (Cleland et al., 2006) (Germany and Switzerland) randomly assigned 204 patients to infusion of BMSCs or cell-free supernatant an average of 4 days after a myocardial infarction. By 4 months after treatment, left ventricular ejection fraction had improved in both groups, but the improvement was significantly greater in the patients who had received stem cells. In contrast, the ASTAMI study (Lunde et al., 2005) (Norway) randomly assigned 100 patients to stem cell implant or treatment after an acute anterior myocardial infarction. The investigators observed no benefit from stem cell implants, and indeed suggested that at 6 months, left ventricular ejection fraction had increased more in the control group. A recent update of the BOOST study (Wollert et al., 2004) suggested that the benefits of bone marrow transfer post-myocardial infarction were sustained at 18 months. However, there was a further improvement in global left ventricular function in the control group rendered the inter-group comparison non-significant.

CONCLUSION As discussed in this chapter, BMSCs provide a promising new arena for regenerative medicine. Although the challenging nature of the research raises some skepticism within the field, bone marrow studies still in their infancy are showing great potential for regeneration of various tissues, at least through the delivery of endothelium and paracrine factors improving revascularization and preventing apoptosis. Current debates addressing the therapeutic potential of this fundamental biologic process should encourage collaborative effort in defining the microenvironment that controls BMSC transdifferentiation. The characterization of such a factor could help harness the mechanism by which cellular repair is achieved. With an understanding of the mechanisms involved in BMSC activity, and with solutions to existing technical difficulties – through improved cellular tracking, improved imaging technology and research standards, and increased communication between international research groups – the mist surrounding BMSCs may soon be lifted.

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17 Hematopoietic Stem Cell Properties, Markers, and Therapeutics S.M. Chambers, William J. Lindblad, and M.A. Goodell

INTRODUCTION Hematopoietic stem cells (HSCs), which primarily reside in bone marrow (BM), maintain blood formation and replenish themselves throughout the adults’ lifespan. The activity of BM HSCs was discovered half a century ago when Ford et al. (1956) identified a robust contribution of donor BM cells in lethally irradiated recipient mice. After three decades of work, the contribution of donor hematopoietic cells in recipients had been demonstrated to originate from a few “clones,” suggesting the existence of HSCs (Becker et al., 1963; Lemischka et al., 1986). However, the isolation of HSCs was not achieved until 1988 when Weissman et al. enriched HSCs from the murine BM using a fluorescent-activated cell sorter (Spangrude et al., 1988). Since these seminal studies, researchers have been able to demonstrate that HSCs possess stem cell properties including the ability to give rise to daughter HSC (self-renewal) as well as to repopulate all of the hematopoietic lineages (differentiation, Figure 17.1). As one of the most investigated tissue stem cells, studies of HSCs have inspired the exploring of various stem cells, and will continuously provide insight of stem cell biology.

Self-renewal

HSC

Differentiation Lineage cells

Figure 17.1 Self-renewal and differentiation of HSC: The HSC-to-niche interaction influences the two definitive properties of HSC. HSC can expand to create more HSC (self-renew) and they can regenerate the hematopoietic system (differentiate).

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DEVELOPMENTAL ORIGIN OF HEMATOPOIESIS During embryonic development, two waves of hematopoiesis occur: primitive hematopoiesis and definitive hematopoiesis, which, respectively, give rise to embryonic and adult hematopoietic cells. Primitive hematopoiesis begins at day 7 of gestation in the mouse yolk sac and generates embryonic primitive erythroblasts (EryP) (reviewed in Lensch and Daley (2004)). However, the hematopoietic precursors from yolk sac are not able to reconstitute lethally irradiated adult recipients (Medvinsky et al., 1993; Muller et al., 1994), which is the gold standard for demonstrating functional HSC activity. The second wave of hematopoiesis, definitive, or adult hematopoiesis arises around day 10 of gestation in the aorta-gonad-mesonephros (AGM) region (Muller et al., 1994; Medvinsky and Dzierzak, 1996). The definitive embryonic HSCs are able to self-renew and give rise to mature hematopoietic lineages in adults. These cells seed the BM, where HSCs contribute to blood formation throughout the lifespan of the adult (Lensch and Daley, 2004). Several studies have identified molecules required for primitive and definitive hematopoiesis, bringing insight to HSC ontogeny and shedding light on mechanisms that regulate HSC self-renewal and differentiation (reviewed in Lensch and Daley (2004) and Medvinsky and Dzierzak (1998)). Deficiency in some of these genes is found to cause embryonic anemia due to inefficient hematopoiesis, indicating they are essential for HSC formation. For example, mutants of Runx1, a member of the runt transcription factor family, exhibit normal primitive hematopoiesis in the yolk sac but lack of hematopoietic clusters in the intra-aorta region at E10.5 of the AGM, and are embryonic lethal at E12.5 with anemic fetal liver, a temporary site of primitive hematopoiesis between E11 and E14 (Okuda et al., 1996; North et al., 1999). The evidence indicates that Runx1 is indispensable for the definitive hematopoiesis but not primitive hematopoiesis. Flk-1 (vascular endothelial growth factor receptor-2) null mice are also embryo lethal (at E8.5–E9.5) with defects in forming blood clusters and in developing vascular network in yolk sac region (Shalaby et al., 1995; Sakurai et al., 2005). Likewise, Scl/Tal1 null mice are found to be embryonic lethal (at E9.5–E11.5), and lack yolk sac vitelline vessels and primitive hematopoiesis (Robb et al., 1995; Shivdasani et al., 1995). Scl/Tal1 null cells also fail to contribute to definitive hematopoiesis of both the AGM and fetal liver in chimeric mice (Porcher et al., 1996; Robb et al., 1996), suggesting critical roles of Scl/Tal1 in both primitive and definitive hematopoiesis.

INTRINSIC REGULATORS OF SELF-RENEWAL Adult HSCs comprise only 0.02% of whole BM cells but possess abilities to self-renew and replenish the whole hematopoietic system. Self-renewal, a signature process of all stem cells, is the process by which one stem cell is able to give rise to at least one daughter stem cell via cell division. By using retroviral integration to mark HSCs, it has been demonstrated that HSCs undergo clonal expansion while repopulating the hematopoietic system in vivo (Lemischka et al., 1986). Moreover, as few as a single HSC is sufficient to establish long-term multi-lineage engraftment (Osawa et al., 1996; Camargo et al., 2005), which would only be possible through a self-renewal process. However, culture conditions have not been fully established to expand HSC, which has hindered our ability to manipulate HSC in vitro. Nevertheless, in the attempt to understand the mechanisms that intrinsically underlie HSC self-renewal in vivo, gene-targeted mice have been developed. HSCs lacking self-renewal mediators usually bear defects in maintaining homeostasis of stem cell population upon proliferation stimuli. Genes that are involved in intrinsically regulating HSC self-renewal can be roughly divided into several groups: cell cycle regulators, transcription factors, anti-apoptotic molecules, and development pathway regulators (Table 17.1). The majority of HSCs remain in G0 phase under normal homeostasis but are able to extensively expand upon receiving proliferation cues. p18INK4C, p21cip/waf1, and p27kip1 are G1-phase cyclin-dependent kinase inhibitors (CKIs) that tightly regulate the G0/G1 stage transition of cell cycle. By perturbing these cell cycle regulators, HSCs exhibit distinct self-renewal phenotypes. Surprisingly, although all three gene-targeted

285

286 CELLS AND TISSUE DEVELOPMENT

Table 17.1 Intrinsic factors of HSC self-renewal Gene Cell cycle regulators p21cip/waf1

KO

p27kip1

KO

p18INK4X

KO

Transcription factors Hoxb4

KO

Tg (retroviral)

Hoxa9

Tg (retroviral)

Pbx1

KO

c-myc

Tg (retroviral)

CBP

KO-Het

Bmi-1

KO

Tg (retroviral) Rae28

KO

Mel-18

KO

Tg-mice Gfi-1

KO

HSC phenotypes

Reference

Increased number of proliferating HSC (sensitive to 5FU treatment) Increased HSC cell number (CAFC assay) Decreased self-renewal ability (serial transplantation) Unaltered HSC pool Increased frequency of committed progenitor Increased multi-lineage repopulation due to outgrowth of progenitors (competitive transplantation) Increased frequency of self-renewing HSC

Cheng et al. (2000)

Lower proliferative rate in HSC compartment Decreased repopulation efficiency when competing with wild-type HSC Impaired embryo primitive hematopoiesis Normal hematopoiesis Extend the self-renewing ability of HSC after short-term in vitro cell culture (transplantation)

Bjornsson et al. (2003)

Expanded number of self-renewing HSC after short-term in vitro cell culture (competitive transplantation) Defects in fetal liver-derived HSC to engraft (competitive transplantation) Loss of adhesion molecule expression and long-term repopulation ability CBP/ HSC pool prematurely exhausted after birth HSC prematurely exhausted after birth, mice die from anemia Bmi1 null HSC fail to repopulate Enforced expression of Bmi-1 increased HSC engraftment Reduced repopulating activity (competitive transplantation, serial transplantation) Slightly increased repopulation activity of Mel-18 null HSC (competitive repopulation) Slightly decrease in HSC proliferation status Decreased repopulation efficiency Increased frequency of proliferative HSC Increased frequency of proliferative HSC and HSC pool Compromised repopulation ability of Gfi-1 null HSC (transplantation and competitive transplantation)

Cheng et al. (2000)

Yuan et al. (2004)

Sauvageau et al. (1995) and Thorsteinsdottir et al. (1999) Thorsteinsdottir et al. (2002) DiMartino et al. (2001) Wilson et al. (2004) Kung et al. (2000) and Rebel et al. (2002) Park et al. (2003)

Iwama et al. (2004) Ohta et al. (2002) Kajiume et al. (2004)

Hock et al. (2004)

Hematopoietic Stem Cell Properties, Markers, and Therapeutics 287

Table 17.1 (Continued) Gene Anti-apoptotic molecules Bcl2 Tg-mice

MCL-1

KO

Anti-oxidative stress ATM

KO

Developmental molecules Notch1 Tg (retroviral)

β-catenin

Tg (retroviral)

HSC phenotypes

Reference

Higher frequency of HSC compartment Increased repopulation ability of HSC (competitive transplantation) Decreased frequency of proliferative HSC Decreased engraftment after Mcl-1 gene deletion Mcl-1 gene deletion leads to loss of HSC and committed progenitors

Domen et al. (2000)

Decreased frequency of BM HSC compartments (KTSL and KSL compartments) Decreased long-term engraftment (competitive transplantation)

Ito et al. (2004)

HSCs are able to survive through a long-term cell culture, retain repopulation ability after in vitro cell culture Enforced expression leads to higher engraftment ability

Varnum-Finney et al. (2000)

Opferman et al. (2005)

Reya et al. (2003)

These genes have been demonstrated to alter aspects HSC self-renewal; Tg: transgenic, KO: knock-out.

HSCs exhibit proliferative phenotypes, p21 null HSCs exhaust over time (Cheng et al., 2000b), whereas p27 and p18 null HSCs possess increased engraftment in transplantation assays (Cheng et al., 2000a; Yuan et al., 2004). More interestingly, the increased multi-lineage engraftment in p18 and p27 null mice originates from different mechanisms; p18 null HSCs possess a higher proportion of self-renewing HSCs (Yuan et al., 2004), whereas p27 null progenitors preferentially outcompete the progenitors of wild type HSC (Cheng et al., 2000a). These data demonstrate how components of cell cycle regulation impinge on HSC cell fate decision. Gfi-1, a zinc-finger proto-oncogene, has been found to be involved in self-renewal with a regulatory role in HSC proliferation. In addition to the defects in generation of neutrophils which will be discussed later, Gfi-1 null HSC exhibited a higher proliferation rate but failed to engraft after serial transplantation, indicating Gfi-1 is a transcriptional regulator that restricts HSC proliferation and prevents exhaustion of HSC pool (Hock et al., 2004). Molecules that regulate cell apoptosis have been implicated to affect HSC function as well. Overexpression of Bcl-2, an anti-apoptotic factor, leads to not only an increased HSC number, but also elevated HSC repopulating ability (Domen et al., 2000). In addition, deletion of Mcl-1, a Bcl-2 family member, leads to a loss of HSCs and committed progenitors, and ultimately severe anemia in mice (Opferman et al., 2005). The largest group of genes that regulate HSC self-renewal are transcription factors. Homeobox genes (HOX) have been found to be involved in both hematopoiesis and leukemia ontogenesis (Abramovich et al., 2005). Wild-type HSCs quickly lose engraftment ability after in vitro culture. However, enforced expression of Hoxb4 with a retroviral target vector extends the self-renewing capability of cultured-mouse HSC. Hoxb4expressing HSC engraft as well as freshly isolated HSC while the control retroviral-transduced HSC showed 5–10% decline in engraftment over the course of cell culture (Sauvageau et al., 1995; Thorsteinsdottir et al.,

288 CELLS AND TISSUE DEVELOPMENT

1999). Similarly, overexpression of Hoxb9 enhances engraftment when they are directly competed against wildtype HSC (Thorsteinsdottir et al., 2002). The findings that Hoxb4 overexpression in HSC leads to a robust expansion of HSC in vitro (Antonchuk et al., 2002), but retained HSC under control of homeostasis in vivo (Thorsteinsdottir et al., 1999), suggests a therapeutic strategy to expand HSC in vitro prior to transplantation. Human HSCs with enforced expression of Hoxb4 have been shown to be capable of repopulating immunodeficient mice. Unfortunately, the repopulation which is derived from Hoxb4-transduced HSC preferentially differentiates into myeloid lineages at a loss of lymphoid cells (Sauvageau et al., 1995; Brun et al., 2003), limiting the clinical utility of this strategy. Polycomb genes that are involved in epigenetic regulation also regulate HSC self-renewal. Bmi-1deficient mice exhibit a hypocellular BM phenotype, and their HSCs exhaust 2 months after birth. Fetal liver and adult BM cells from Bmi-1-deficient mice failed to engraft after transplantation, which indicates Bmi-1 is critical for HSC self-renewal (Park et al., 2003). Moreover, enforced expression of Bmi-1 by a retrovirus in murine HSC has shown an increased repopulating activity, suggesting a therapeutic role of Bmi-1 (Iwama et al., 2004). Other polycomb genes such as Rae28 (Ohta et al., 2002) and Mel-18 (Kajiume et al., 2004) have also been shown to modulate HSC self-renewal. Interestingly, unlike Rae28 and Bmi-1, Mel-18 acts as a negative regulator of HSC self-renewal. Mel-18 null HSCs possess a higher repopulation activity, whereas Mel-18 transgenic HSCs have decreased ability to engraft into lethally irradiated mice (Kajiume et al., 2004). Pathways that trigger cell fate decisions during early development of vertebrates and invertebrates have been demonstrated to be involved in dictating cell fate decisions of HSC during cell division. Overexpression of constitutively activated Notch1 in HSC immortalizes HSC in long-term in vitro culture. A single Notch1transduced HSC clone is able to undergo multi-lineage repopulation in vivo (Varnum-Finney et al., 2000). The Wnt signaling pathway has also been implicated in HSC self-renewal. Extrinsic stimulation of Wnt3a or overexpression of beta-catenin in combination with the presence of Bcl2 leads to HSC with a high repopulating activity after a long period of in vitro cell culture while ectopic expression of Axin, an inhibitor of the Wnt pathway, decreased HSC proliferation in vitro and reconstitution function in vivo, suggesting a role of Wnt signaling in retaining HSC self-renewal (Reya et al., 2003).

MULTI-LINEAGE REPOPULATION During adult hematopoiesis, BM-HSCs generate both lymphoid and myeloid cells (Figure 17.2). Lymphoid cells are comprised of primarily T-cells, B-cells, and natural killer cells (NK cells). Myeloid cells include granulocytes, macrophages, megakaryocytes, and erythrocytes. Hematopoiesis is a gradual differentiation process that involves multiple decision points beginning with HSC and ending with terminally differentiated lineages (Figure 17.2). This concept of stepwise hematopoiesis has lead to the identification of several differentiation intermediates. From this concept, Morrison and Weissman described two populations within BM that possess transient engraftment ability when transplanted into lethally irradiated mice. These two populations are considered short-term HSC (ST-HSC, Mac-1loCD4) and multipotent progenitor (MPP, Mac-1loCD4lo), which are distinguished from long-term repopulating HSC (Morrison and Weissman, 1994; Morrison et al., 1997). Weissman et al. also first identified common lymphoid progenitors (CLPs) within BM which specifically give rise to lymphoid lineages (Tcells, B-cells, and NK cells) (Kondo et al., 1997), and common myeloid progenitors (CMPs), which give rise to granulocytes/macrophages and megakaryocytes/erythrocytes colonies in methylcellulose cell culture and in lethally irradiated recipients (Akashi et al., 2000). More recently, Adolfsson et al. have revised the role of MPPs as lymphoid-primed multipotent progenitors (LMPP). They discovered that MPPs (now LMPP) preferentially differentiate into lymphoid lineages, but retain some myeloid development capacity (dashed line in Figure 17.2), restricted to granulocytes and macrophages (Adolfsson et al., 2005). The CMP is the major generator of all

Hematopoietic Stem Cell Properties, Markers, and Therapeutics 289

Long term HSC Hematopoietic stem cell (HSC) Short term HSC

Lymphoid primed multipotent progenitor (LMPP)

Common myeloid progenitor (CMP)

Common lymphoid progenitor (CLP)

B-cells

T-cells

Granulocytes Monocytes

Erythrocytes Megakaryocytes

Figure 17.2 HSC differentiation: HSC regenerates the hematopoietic system, which is comprised of a myeloid and lymphoid branch, ultimately creating all the cells that comprise the blood. The multipotentprogenitor (MPPs), previously thought of as a bipotential progenitor, is now identified as a lymphoid-primed progenitor (LMPP).

myeloid cells, including megakaryocytes/erythrocytes lineages. In summary, these studies have suggested a hematopoietic hierarchy in which long-term HSCs give rise to ST-HSC that differentiates into CMP and CLP. The CMP and CLP then generate the myeloid and lymphoid lineages (Figure 17.2). Although the differentiation pathway in humans is not as well established as in rodent models, transplantation of HSC has been utilized for decades to treat patients with hematopoietic diseases, demonstrating the repopulation ability of human HSC to reconstitute the entire hematopoietic system.

PLAYERS IN HEMATOPOIESIS Transcription factors that regulate cell fate decisions during adult hematopoiesis have been identified mostly by phenotypes of gene-targeted mice. Deficiencies in genes controlling cell fate lead to defective, definitive hematopoiesis in embryos and lineage skewing in adult hematopoietic system. For example, PU.1, a member of the ETS transcription factor family, has been shown to dictate cell fate at the divergence point of myeloid and lymphoid differentiation in a dosage-dependent manner. Overexpression of PU.1 drives fetal liverderived hematopoietic precursors to differentiate into macrophages; whereas haploinsufficiency in PU.1/ cells leads to both pro-B-cells and macrophages (DeKoter and Singh, 2000). In humans, aberrant expression of several transcription factors was shown to cause diseases that result from defects in hematopoietic development or abnormal proliferation of a particular hematopoietic lineage. For example, mutations of a protooncogene Gfi-1, a zinc finger transcription factor, result in neutropenia in which patients lack neutrophils

290 CELLS AND TISSUE DEVELOPMENT

(Hock et al., 2003; Person et al., 2003). Chromosome translocation that results in overexpression of Scl/Tal, a basic helix-loop-helix gene that is also critical to murine primitive and definitive hematopoiesis, causes human acute T-lymphocytic leukemia (T-ALL) (Begley et al., 1989; Chen et al., 1990). Runx1-Eto, the fusion protein caused by chromosome translocation t(8;21), and Tel-Runx1, caused by t(12;21), result in AML and B-cell precursor ALL, respectively (reviewed in Izraeli (2004)). In addition, retroviral integration which results in overexpressing LMO2 has been found to cause T-cell malignancy in a retroviral-gene therapy that was originally designed to treat X-linked severe combined immunodeficient (SCID) patients (HaceinBey-Abina et al., 2003). These studies have shed light on therapeutic methodology to target hematopoietic disorders as well as defining roles of genes that control homeostasis of hematopoietic system.

IN VITRO DIFFERENTIATION OF HEMATOPOIETIC LINEAGES In vitro cell culture of blood precursors has been established to quantify the differentiation ability of hematopoietic precursors. The in vitro culture conditions were established with the goal to mimic the in vivo growth stimuli and maturation signals. A functional assay of hematopoietic precursors, the colony forming unit in cell culture (CFU-C) was first established in 1980s and measured generation of myeloid and erythroid cells (Metcalf, 1989). In vitro differentiation of lymphocytes was later found to require microenvironments that are provided by co-cultured cells. In vitro co-culture of stromal cell line with B-cell precursors (Cumano et al., 1990) and fetal thymic organ culture (FTOC) to generate thymocytes (Robinson and Owen, 1978) have been aimed at recapitulating the in vivo environment and identifying regulators of lymphocyte differentiation. Moreover, in vitro assays such as cobblestone area forming cells (CAFC) and long-term culture initiating cells (LT-CIC) were developed to detect earlier precursors, including the HSCs. A more comprehensive outline of in vitro HSC differentiation assays has been described by Ramos et al. (2003). In addition, differentiation of murine and human hematopoietic progenitor cells in vitro has been utilized to modulate immune responses. The best example is terminally differentiated dendritic cells (DCs), one of the professional antigen presenting cells in immune system. Antigen (Ag)-pulsed DCs have been utilized to modulate Ag-specific immune response. In clinical trials of immunotherapy, these in vitro differentiated DCs are able to stimulate tumor antigen-specific immune responses and to induce tolerance in autoimmune diseases (Figdor et al., 2004). IN VITRO EXPANSION OF SELF-RENEWING HSCs Expansion of HSC in vitro has been the most difficult challenge for decades due to a decline in repopulation capacity of HSC in long-term ex vivo culture. There are currently two main strategies for ex vivo HSC expansion: HSC stromal cell co-cultivation and HSC suspension culture. BM stromal cells support HSC maintenance, measured by repopulating ability, in the absence of additional growth factors (Fraser et al., 1990; Fraser et al., 1992). BM stromal cells are thought to mimic the microenvironment of the HSC niche. To identify the molecules in these stromal cells that retain HSC function, genome wide studies of cloned stromal cells have been reported (Moore, 2004), providing an emerging picture that underlies the HSC-to-niche interaction which we will discuss below. For suspension culture, growth factor cocktails have been utilized in attempt to expand human and murine HSC (Sauvageau et al., 2004), albeit with low recovery rate of repopulating HSC. The differentiation of HSC during in vitro culture has drawn into question whether or not HSC self-renewal occurs during cell expansion. To address the question, Glimm and Eaves labeled HSC with a fluorescent membrane-specific dye, carboxyfluorescein diacetate succinimidyl (CFSE), to track the proliferation history. By transplanting cells that had divided from an in vitro cell suspension culture, they discovered that human HSCs were still able to give rise to multi-lineage repopulation after a low number of cell divisions (Glimm and Eaves, 1999). Additionally, Nakauchi et al. have cultured highly purified murine single-cell HSC to track cell

Hematopoietic Stem Cell Properties, Markers, and Therapeutics 291

? Notch-1 N-cadherin Wnt 3a Frizzled

CXCR4 SDF-1 HSC Ang-1 Tie-2

Osteoblast

Figure 17.3 Hypothetical HSC-to-niche interactions: Through several cell surface molecules (markers) and cytokines, the HSC is thought to directly or indirectly interact with osteoblast cells. These are few of the molecules that provide instructions for self-renewal and differentiation to the HSC.

division as well as cell fate. They have been able to generate limited self-renewing HSC under the influence of various combinations of cytokines (Ema et al., 2000). However, after more than two cell divisions in vitro, the HSCs greatly lost their repopulating activity.

HSC NICHE In adults, HSCs reside in the BM cavity, closely associated with surrounding stromal cells. There is mounting evidence suggesting that the most primitive HSCs localize to the interior surface of bone (periosteum/endosteum border) on the basis of colony forming assays (Lord and Hendry, 1972) and and Brd-U label retention (Zhang et al., 2003), bringing them within close contact with osteoblasts. Murine osteoblasts have long been thought to provide essential cues for HSC, as they (or their transformed counterparts) express various cytokines known to influence hematopoiesis, including but not limited to G/M/GM-CSFs, Il-1, Il-6, SDF-1, and VEGF (reviewed in Taichman (2005)). In addition to expressing HSC-modulating cytokines, recent genetic evidence from mice has demonstrated that expanding the number of osteoblasts within the BM increases the relative percentage of HSC (Calvi et al., 2003; Zhang et al., 2003), and genetically ablating osteoblasts results in the failure of BM hematopoiesis (Visnjic et al., 2004), suggesting that osteoblasts provide a direct physical niche for the HSC that maintains their self-renewing capacity via various cell surface molecules (Figure 17.3). Researchers have also provided evidence of a second HSC niche provided by sinusoidal endothelial cells resident in the BM (Kiel et al., 2005). Molecules including N-cadherin, Notch-1, Tie2, and CXCR-4 have all been implicated in the HSC-to-niche interface. N-cadherin, a Ca2-dependent homophilic adhesion molecule expressed on osteoblasts, exhibits an asymmetrical localization on HSC, as determined by fluorescence microscopy (Zhang et al., 2003), and is found expressed on a fraction (10%) of hematopoietic progenitors that include the HSC, thus suggesting that it may be involved in self-renewal or niche retention. However, it remains to be seen whether or not N-cadherin is required for HSC-to-niche interaction, by examining if there is a functional difference between N-cadherin and Ncadherin HSC in vivo. Several studies have demonstrated that Notch-1 is expressed on HSC and its activation by incubating with Jagged-1 expressing cells (Varnum-Finney et al., 2000) or constitutive Notch-1 signaling (Varnum-Finney et al., 2000) results in in vitro expansion of self-renewing HSC with normal homeostasis while transplanting

292 CELLS AND TISSUE DEVELOPMENT

into lethally irradiated mice. In contrast to the evidence discovered from in vitro Notch1 stimulation, targeted disruption of Jagged1 and Notch-1 in mice does not result in reconstitution or self-renewal defects in vivo (Mancini et al., 2005). Therefore, Notch-1 stimulation may regulate HSC cell fate decision toward self-renewal during in vitro cell culture, although it is not essential for HSC homeostasis in vivo. Hence, stimulation of the signaling pathway is of interest in HSC expansion. A second such receptor–ligand interaction is that of Tie2, a tyrosine kinase receptor that is expressed on HSC, and its ligand Angiopoietin-1 (Ang-1), expressed by osteoblasts. These two molecules were demonstrated to be important retention factors for HSC (Arai et al., 2004). Incubation of HSC in Ang-1 or overexpression of Tie-2 in transduced BM cells resulted in expansion of the quiescent portion of the HSC compartment. A loss of self-renewal or rapid differentiation continues to be a hurdle for in vitro expansion of HSC. Exploiting Tie2 signaling with soluble Ang-1 may be a promising method for expanding long-term, quiescent HSC cultures. One such receptor that has been employed in the clinic is the chemokine receptor, CXCR-4, expressed on both human (Viardot et al., 1998) and mouse (Wright et al., 2002) HSCs. In addition, SDF-1, the ligand for CXCR-4, is a well-established HSC homing factor expressed by osteoblasts and BM fibroblasts (Ponomaryov et al., 2000). Antibodies against SDF-1 or CXCR-4 block human HSC engraftment in the non-obese severe combined immunodeficient (NOD/SCID) mouse model, and SDF-1 enhances transwell migration of human HSC (Peled et al., 1999). Therefore, the CXCR-4/SDF-1 axis is thought to provide a BM homing and retention mechanism for HSC in vivo. In a secondary transplantation assay, treatment of stem cell factor (SCF) and IL6 enhanced HSC engraftment correlating with elevated CXCR4 expression and increased in vitro migration activity to SDF-1 (Peled et al., 1999). However, they did not distinguish between a rescued migratory defect and enhanced self-renewal as the cause for increased engraftment in vivo. Therefore, it remains less clear whether CXCR-4 plays a role in HSC self-renewal and in vitro expansion of HSC. In summary, over the past decade it has become increasingly clear that components such as N-cadherin, Tie2/Ang-1, Notch-1, and CXCR-4/SDF-1 may be required in order to establish an ex vivo niche for HSC expansion and self-renewal, unfortunately, research has yet to elucidate the appropriate cocktail of soluble factors, cytokines, and/or niche support cells needed to stimulate faithful and prolonged in vitro HSC self-renewal and expansion.

PURIFICATION AND MOLECULAR SIGNATURE OF HSC In order to assess HSC function and identify novel molecular components, it is essential that HSCs are distinguished from the heterogeneous cell mixture that comprises the BM niche. Consequently, functional purity has been an important focus in the HSC field, since the purification method can drastically influence engraftment and hematopoietic reconstitution. Murine and human HSCs are purified by a combination of enrichment (by magnetic cell sorting) and fluorescence-activated cell sorting based on cell surface markers or vital dye staining (Table 17.2). Human HSCs have been shown to express CD34 and lack the expression of CD38, forming the basis of all human HSCs purification schemes. Positive selection for CD34 and Thy-1, as well as the removal of differentiated progeny by CD38 and a cocktail of antibodies that recognize lineage-specific cell surface molecules (Lin cocktail), has also been used to purify human hematopoietic progenitors from fetal liver (Muench et al., 2002), umbilical cord blood (Gluckman et al., 1989; Rocha et al., 2004), BM (Thomas, 2000), or patient periperal blood treated with an HSC mobilizing agent (Ho et al., 1996; Broxmeyer et al., 2005). In the mouse, HSCs are typically harvested from the medullar cavity of the leg bones (tibias and femurs). Similar to human HSCs, differentiated cells are excluded from the purification on the basis of expression of lineage-specific cell surface markers (Lineagenegative or Lin). To further purify mouse HSCs, cells that express two canonical HSC markers, stem cell antigen-1 (Sca-1) and the tyrosine receptor kinase c-Kit, are selected. Low expression of the marker Thy-1 is sometimes used to further subdivide this population but its utility is limited

Hematopoietic Stem Cell Properties, Markers, and Therapeutics 293

Table 17.2 HSC cell surface molecules that allow for purification of specific HSC subpopulation; UK:unknown, Gene

Function

Homology Mouse

Human

N-Cadherin

Cell surface marker; Wnt signaling



UK

Notch-1

Cell surface marker; Notch signaling





Tie2

Cell surface marker



UK

Endoglin



UK

CD34

Cell surface marker; TGF-b signaling Cell surface marker

ST-HSC



CXCR-4

Cell surface marker





CD150 (SLAM) CD48

Cell surface marker



UK

Cell surface marker

ST-HSC

UK

Bcrp-1

Cell property





Thy-1

Cell surface marker





Sca-1

Cell surface marker; HSC self-renewal





c-Kit

Cell surface marker; Tyrosine Kinase Receptor; expansion and self-renewal of HSC pool in vivo





Details/discovery

References

Homophilic adhesion molecule found to be expressed on both HSC and osteoblasts In vitro activation results in self-renewal/ expansion

Zhang et al. (2003)

Activation with Ang-1 results in enhanced quiescence Discovered by microarray, marks LT-HSC Purification marker in humans; ST-HSC marker in mouse Homing and retention of HSC to niche Discovered by microarray, marks LT-HSC Discovered by microarray, marks ST-HSC Thought to be responsible for HSC SP phenotype ST-HSC marker

Purification marker, null HSC exhibit reduced in vivo repopulation (competitive transplantation and serial transplantation) HSC from W/Wv mice fail to engraft (transplantation, parabiosis)

Varnum-Finney et al. (1998) and VarnumFinney et al. (2000) Arai et al. (2004) Chen et al. (2002) Civin et al. (1996) Peled et al. (1999) Kiel et al. (2005) Kiel et al. (2005) Zhou et al. (2002) Uchida and Weissman (1992) Ito et al. (2003) Gardner et al. (1988)

HSC cell surface molecules: Both mouse and human HSC express cell surface molecules (markers) that allow for purification of specific HSC subpopulations.

to strains bearing the Thy-1.1 allele. This isolation scheme has been collectively referred to as “KTSL” in the literature (c-Kit, Thylow, Sca-1, Lin). Two purification schemes have used to simplify the bewildering number of antibodies needed in order to obtain highly purified HSC. The first isolation scheme exploits a property of HSC to efflux the fluorescent DNA binding dye Hoechst 33342. Cells that posses this efflux ability are referred to as side population (SP) cells (Goodell et al., 1996) and have been shown to have a KTSL phenotype (Camargo et al., 2005). This efflux property is thought to be conferred by the ABC transporter Bcrp1 (ABCG2) (Zhou et al., 2002). In mouse BM, SP cells comprise approximately 0.03–0.07% of all nucleated cells, express low to no lineage marker expression, and greater than 95% express c-kit and Sca-1 (Figure 17.4). Details about the Hoechst staining

294 CELLS AND TISSUE DEVELOPMENT

104

4000 95%

103

3000

2000

1000

0 0

1000

2000

3000

Hoechst-Red

4000

97%

4 c-Kit

0.33%

Histogram

Hoechst-Blue

6

2

0 100

102 101

101

102 Lineage

103

104

100 100

101

102 Sca-1

103

104

Figure 17.4 HSC SP phenotype: Mouse HSC can be purified by their ability to efflux Hoechst dye. When whole BM is excited with a UV laser and viewed by two wavelengths (Hoechst-Red, Hoechst-Blue), HSCs are found in a SP (left). Less than 5% of these SP cells express differentiation markers (lineage, middle), and greater than 95% express the two canonical stem cell markers Sca-1 and c-Kit (right).

procedure can be found in a recent review (Goodell, 2002; Goodell et al., 2005). Finally, a new scheme may simplify the purification scheme further. Antibodies against a combination of three markers, all members of the SLAM family proteins, have been demonstrated to recognize high-purity HSC comparable to KTSL (Kiel et al., 2005). Although these results have yet to be verified by other labs, this method reduces the markers needed to purify HSC to three (CD150, CD48, and CD244). Now that HSCs have been purified to near homogeneity, microarrays have been employed to examine HSC transcriptome under steady state and activated conditions. Initial analyses were aimed at identifying transcriptional similarities shared between embryonic, hematopoietic, and neuronal stem cells (“stemness”) (Ivanova et al., 2002; Ramalho-Santos et al., 2002). While some similarities were observed, the implications remain unclear (Evsikov and Solter, 2003; Fortunel et al., 2003; Vogel, 2003). Fortunately, transcriptional profiling of both mouse and human HSCs under steady state conditions has lead to a number of observations about cell surface markers and potential regulators of HSC function. For instance, CD150 was found to be a nearly exclusive marker of LTHSC, after comparing the expression profiles of ST- and LT-HSC (Kiel et al., 2005). The expression profile of mouse HSCs under activated conditions has been studied by chemotherapeutic drug 5-fluorouracil (5-FU) treatment, which destroys the majority of hematopoietic cells prompting hematopoietic regeneration. This stressful, anemic environment stimulates the normally quiescent HSC (1% S-phase) into a transient proliferative phase (20% S-phase) lasting approximately 12 days. The expression profile of HSC was monitored at 2–3 day intervals after 5-FU injection (Venezia et al., 2004). The data not only elucidated a model of HSC activation but identified CD48, a molecule which marks proliferating HSC and has been shown to distinguish ST-HSC from LT-HSC (Kiel et al., 2005). Venezia, et al. also employed a refined gene ontology (GO) analysis (Young et al., 2005), which quantified fold enrichment (over random chance) of all GO biological processes, thus conferring more biological relevance to the global microarray results. Microarrays are an extremely powerful emerging technology, with continued bioinformatic developments that can augment and standardize the results, microarrays will become increasingly prevalent in the HSC field as a tool for expression profiling, molecular phenotyping, and screening.

PLASTICITY AND THERAPEUTICS The central tenet to embryonic development states that HSCs become committed to a multipotent program of differentiation, in that they exclusively create all of the cells of the hematopoietic hierarchy. However, reports

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began to arise that HSCs could give rise to non-autologous tissues leading to the idea the HSCs may be pluripotent (Ferrari et al., 1998; Brazelton et al., 2000). Although provocative, many of these reported cases of HSC “plasticity” have been called into question, due to inadequate assays (Jackson et al., 2004), cell fusion (Terada et al., 2002; Ying et al., 2002), and extramedullary HSC (McKinney-Freeman et al., 2002). Through stringent genetic studies it has become clear that certain hematopoietic progeny, most likely the fusogenic and invasive macrophage, are able to contribute at subtherapeutic levels to skeletal muscle (Camargo et al., 2003) and liver (Camargo et al., 2004) regeneration. Clinical trials of HSC transplantation for the purpose of replenishing hematopoiesis have been widely performed for treating leukemia (Armitage, 1994), severe combined immunodeficiency (Fischer et al., 2004), and severe autoimmune disease (Tyndall and Saccardi, 2005). In addition, hematopoietic treatment has been used in adjunct to chemotherapy for non-hematopoietic cancer such as breast cancer, neuroblastoma, and testicular cancer (reviewed in Armitage (1994)). Currently, HSC transplant research is investigating ways to improve the overall success by reducing graft-versus-host-disease, infection during recovery, and accelerating robust hematopoiesis after transplantation.

CONCLUSIONS The HSC field continues to be at the forefront of regenerative medicine with therapeutic potential to cure a wide range of diseases. Developmental origin, self-renewal, differentiation, molecular signature, and therapeutic potential of HSC are currently being established. The next horizon for the HSC field is to create an ex vivo niche. Identifying the essential support cells and their secreted cytokines required for HSC self-renewal and expansion would open the gateway for unfettered genetic modification of HSC to aid the continued efforts in gene therapy. Equally important, HSC expansion could drastically reduce the number of HSCs needed to provide adequate reconstitution in human BM transplant therapy. Therefore, further characterization of components that contribute to regulation of HSC self-renewal is needed in order to co-ordinate HSC expansion.

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18 Neural Stem Cells Yang D. Teng, Filipe N.C. Santos, Peter M. Black, Deniz Konya, Kook In Park, Richard L. Sidman, and Evan Y. Snyder

INTRODUCTION – HISTORICAL BACKGROUND In 1913, Santiago Ramon y Cajal hypothesized that neurons are generated exclusively during the prenatal phase of development. Although some investigators held different opinions, it was impossible, with the resources and methods available at that time, to prove that postnatally born cells were actually neurons and not glia. The idea that no new neurons are formed after birth became doctrine in neuroscience. Only in the late 1950s did new methodologies make alternative hypotheses possible (Ming and Song, 2005). In 1959, a new method of marking dividing cells in the mammalian brain was developed (Sidman et al., 1959). The method was based on the already known fact (Hughes et al., 1958) that [H3]-thymidine injected systemically in mammals would be incorporated selectively into DNA replicating during S-phase of the cell cycle (Howard and Pelc, 1953), and their positions and numbers ascertained by autoradiography. Using this technique, patterns of neuron genesis and migration were demonstrated in several developing areas of the mouse brain (Angevine and Sidman, 1961; Miale and Sidman, 1961; Sidman, 1961) and in the adult mouse brain (Smart, 1961). Soon after, it was suggested that not only astrocytes and microglial cells, but also oligodendrocytes (Altman, 1962a) and neurons proliferate (Altman, 1962b) in adult mammals. Subsequently, Angevine (1965) reported that granule cell neurons were still forming in the mouse at least until postnatal day 20, the oldest age he examined, in and near the dentate gyrus of the hippocampal formation, and Altman et al. reported evidence of new neurons forming in various regions of the adult rat brain including the dentate gyrus of the hippocampus (Altman and Das, 1965), the neocortex (Altman, 1966) and the olfactory bulb, the latter after cell proliferation in the wall of the anterior horn of the lateral ventricle and migration via a pathway he named the “rostral migratory stream” (Altman, 1969). A decade later, Nottebohm and collaborators demonstrated telencephalic neuronal replacement in the adult avian brain related to seasonal song learning (Graziadei and Graziadei, 1979; Goldman and Nottebohm, 1983; Nottebohm, 2004). Around that same time, it was observed that newborn neurons in the hippocampus appeared to receive synaptic inputs (Kaplan and Bell, 1983), and later it was established that they also extended axon projections to their target area (Stanfield and Trice, 1988). Parallel to these studies, a new cell tracer was being explored that would accelerate research in the stem/progenitor cell field (Gratzner, 1982). The marker, bromodeoxyuridine (BrdU), is a synthetic thymidine analog that also becomes incorporated into a cell’s DNA during the S-phase of the cell cycle, much like [H3]-thymidine, but, in contrast to thymidine, could be easily, rapidly, inexpensively, and safely detected by simple immunocytochemistry. By immunostaining a given cell for dual expression of BrdU and a cell type-specific marker, the phenotypic fate of a newly divided cell could be readily analyzed and quantified by unbiased stereological methods. This technique became a mainstay in stem cell research. Only recently has there been concern that BrdU and other thymidine

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analogs might be passed from labeled grafted stem cells to dividing host cells, allowing for an element of misidentification unless adequate controls are invoked (Burns et al., 2006). Also, labeling might represent DNA repair in a non-mitotic cell (Bauer and Patterson, 2005), and cell cycle activity with labeling by thymidine analogs can be a prelude not to cell division, but to cell death of neurons (Herrup et al., 2004). In 1992, adult and fetal neural stem cells (NSCs) from the central nervous system (CNS) of rodents were isolated (Reynolds and Weiss, 1992; Snyder et al., 1992) and, in 1998, human NSCs (Flax et al., 1998; Kukekov et al., 1999). Also in 1998, taking advantage of the fortuitous use of a BrdU-like chemotherapeutic agent in the treatment of a head and neck cancer, the presence of neurogenesis in the hippocampus of adult humans was confirmed, further suggesting that concepts regarding stem cell biology being formulated in rodents may be evolutionarily conserved (Eriksson et al., 1998; Curtis et al., 2007). Also in the late 1980s and early 1990s, complementary methods for identifying and tracking newly mitotic cells were being developed and employed for tracing the lineage of cells. The most powerful of these was the use of replication-incompetent retroviral vectors that could transduce a reporter transgene to the genome of cells passing through S-phase such that this marker would become permanently integrated in a unique chromosomal site and passed to all progeny and generations (Sanes et al., 1986; Price et al., 1987). Unlike BrdU, which becomes diluted 50% with each cell division, retroviral-transduced genes are not diluted. Combining such marked cells with electrophysiological analysis could provide convincing evidence that newborn neurons in the adult mammalian CNS are functional and synaptically integrated (van Praag et al., 2002; Belluzzi et al., 2003; Carleton et al., 2003). Yet, though it is now known that neurogenesis persists in some regions of the human adult CNS and that these new neurons are functional and connected, it remains unclear what biological roles are served by the newborn cells in these few neurogenic “hot-spots” (Teng et al., 2006).

THE NEURAL STEM CELL Definition Neural Stem Cells (NSCs) are the most primordial and uncommitted cells of the nervous system, and are believed to give rise to the vast array of more specialized cells of the CNS and peripheral nervous system (PNS). To be considered a “neural stem cell,” in contrast to a “progenitor” cell (i.e., cells that have already become lineage committed to give rise to only one category of neural component, e.g., glial cells versus neurons), that cell must be capable of (1) generating all neural lineages (neurons, astrocytes, and oligodendrocytes) throughout the nervous system, (2) having some capacity for self-renewal, and (3) being able to give rise to cell types in addition to themselves through asymmetric cell division (Gage, 2000). Stem cells are defined according to their repertoires. A “totipotent” stem cell, if implanted in the uterus of a living animal, can give rise to a full organism and all its organ systems, including CNS and PNS. A “pluripotent” stem cell, in the simplest definition, is similar to the totipotent cell, except that it cannot give rise to trophoblasts of the placenta. Such cells have been convincingly affirmed to exist only in the inner cell mass of the blastocyst, although a series of recent controversial papers have suggested that such pluripotent cells may be harbored in the amniotic fluid, placenta, umbilical cord, and bone marrow mesenchyme (Ortiz-Gonzalez et al., 2004). The next developmental stage in the progressive restriction of potency is the “multipotent” somatic stem cell, which is capable of differentiating into all cell types of a given organ or tissue and to only cells of that organ or tissue (Martinez-Serrano et al., 2001). Even though that is the traditional and most wellaccepted definition of the multipotent stem cell, recent findings have forced a reexamination of this biological concept, based on experiments showing possible cross-differentiation, although some such instances have

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been found to be artifacts caused by cell fusion, as discussed in the next section. Hence, while an active area of inquiry at the time of this writing, the distinctions above remain the lexicon of the field. Importantly, a common belief among stem cell biologists is that the distinction between totipotence, pluripotence, and multipotence is not discrete, but rather a continuum in development. Cross-differentiation and Cell Fusion As previously stated, somatic stem cells – sometime erroneously called “adult” stem cells in the lay press – have been viewed as exceptionally plastic but nevertheless restricted to the production of cells from the tissue of origin but not cells of non-related tissues. For example, NSCs give rise to the three main types of nervous system cells (neuron, oligodendrocyte, and astrocyte), while hematopoietic stem cells (HSCs) produce only blood-derived cells, etc. However, various reports in the last few years have challenged this central dogma by demonstrating that adult stem cells, under certain microenvironmental (often very non-physiological) conditions, generate cell types besides those in the tissue of origin, possibly indicating that they can switch cell fate. HSCs, for instance, in addition to forming blood cells, have been reported to develop into liver cells (Petersen et al., 1999) and NSCs to give rise not only to nerve cells but also to early hematopoietic precursors (Bjornson et al., 1999). These cell behaviors have been termed “cross-differentiation” or “stem cell plasticity.” Such reports have generated excitement as well as skepticism in the field of stem cell biology, as the concept of plasticity defies the developmental biology principle that lineage restriction is imparted during morphogenesis. However, if correct, the ability of adult stem cells to change fate also holds immense therapeutic potential as well as circumventing the concerns by some of having to obtain pluripotent cells from human embryos (Lakshmipathy and Verfaillie, 2005). The controversy comes when more recent studies suggest that such stem cell plasticity might not actually exist. Rather, the illusion of cross-differentiation/transdifferentiation might actually result from the fusion of donor cells with host cells – conferring on these host cells the transplanted cell’s reporter gene and creating the misperception that these host cells are donor derived. In other words, the markers used to visualize the grafted cells in vivo or in vitro have simply been taken up by the host cells but no differentiation of donor cells into differentiated organ-specific cells had actually occurred. It is also possible that cell fusion may cause a change in the phenotype and/or the function of cells (through nuclear reprogramming) and may also explain the apparent inherent plasticity of committed cells. While cell fusion may have created, in many instances, the misleading appearance of cross-differentiation, the prevalence of this process itself has caused stem cell biologists to speculate whether fusion may actually play a natural role in some normal biological as well as pathophysiological functions. This process, as suggested by some investigators, could provide a strategy for de-differentiating committed cells that might then be reprogrammed for tissue reconstitution and reversal or repair of injury (Ogle et al., 2005). Resolution of the “cross-differentiation” controversy will obviously have a major impact on our current definitions of stem cells.

ANALYSIS OF NEUROGENESIS In Vivo As briefly mentioned, the analysis of neurogenesis in vivo was largely established by the use of [H3]-thymidine in 1959 and BrdU after 1982 as markers of cell division. Those labeling techniques made it possible to track stem/progenitor cells and do quantitative analysis of proliferation, differentiation, and survival of all cells born after the injection of such a marker (Sidman, 1970; Miller and Nowakowski, 1988; Kempermann et al., 1997).

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Because all marking techniques carry caveats and require rigorous controls, a good rule of thumb is to use at least two independent markers and marking techniques for identifying cells. For example, a strategy for marking newly born cells might include both BrdU incorporation and the tagging of a cell with a genetic marker carried in a replication-incompetent, helper-virus-free retroviral vector that lacks nuclear import mechanisms. With the latter technique, viral integration occurs only when the nuclear membrane breaks down during mitosis. Also the transgene becomes transduced only when the vector adheres to a specific cell surface receptor. Therefore, they are good, non-diffusible, non-dilutable markers of dividing cells (Lewis and Emerman, 1994; Ming and Song, 2005). Other methods for distinguishing donor cells from host cells in transplantation studies might include mismatching the species or sex between the two; for example, transplanting human cells into a rodent organ or implanting a male cell (recognized by its Y chromosome) into a female host. Some research might employ cell type- or developmental stage-specific markers to track the development and differentiation of a stem cell. For example, a developmental neuron is defined as immature when it expresses immature markers but lack mature ones. Therefore, antibodies directed against those immature markers would stain only newborn, still immature cells that are undergoing neuronal differentiation. Oligodendrocyte (Olg) development can be used to illustrate the use of this technique. Even though Olgs express proteolipid protein (PLP) and DM20 (myelin proteins) during most of their development, the ratio between these two proteins varies during differentiation (Yu et al., 1994; Timsit et al., 1995). While DM20’s expression precedes that of PLP and is kept high during the first stages of differentiation, PLP is expressed in higher levels only at the latter stages of differentiation and, therefore, defines a mature myelin-forming oligodendrocyte (Meng et al., 2005). Multiple steps at multiple CNS sites have been defined for this cell type (Miller, 2002, 2005). Similarly, other markers for other cell types can be used to track the development of a particular cell type. It is important to reiterate that this method is best used in combination with some or all of the others described above, because antibodies to developmental markers are not always ideally specific (Brandt et al., 2003; Kempermann et al., 2004). In Vitro The isolation and culture of NSCs in vitro is of great value, providing an opportunity to scrutinize individual cells and their differentiation more closely. It also allows one to exclude confounding variables (e.g., intervention by the immune system) and to manipulate the microenvironment in systematic and prescribed ways. Stem cells can be propagated as clonal populations in vitro by many effective and safe means that include both epigenetic and genetic strategies. Epigenetic The identification of cytokines and growth factors that affect survival and proliferation of NSCs has been very important for cell isolation and culture. In the late 1980s and early 1990s it became evident that most regions of the CNS contain a small population of individual cells that could be isolated and could give rise to clonally related populations consisting of multiple neural cell types that heretofore had not been regarded as deriving from a common progenitor. These regions included olfactory bulb and cerebellum (Ryder et al., 1990; Snyder et al., 1992), postnatal (Altman, 1966) as well as embryonic neocortex (where 18% of the cells could generate both neurons and oligodendrocytes in vitro) (Williams et al., 1991), hippocampus (Renfranz et al., 1991; Gage et al., 1995), and septum (Temple, 1989). From these studies, however, it became clear that neural progenitors or stem cells, when isolated from the CNS, had a predisposition to exit the cell cycle and differentiate unless an intervention was imposed upon them to remain cycling and hence hold commitment in abeyance. For example, cells cultured from the embryonic rat septum were incapable of more than one cycle of cell division unless they were co-cultured with fetal striatal tissue (Temple, 1989). When they were prodded to continue

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proliferating, some of the cells appeared to be multipotent. Subsequent studies were launched to identify exogenous growth factors and/or internal genes that could be used to maintain multipotent cells in a proliferative state that might allow them to be expanded and used for developmental or therapeutic studies. The principal mitogens studied and most commonly used to date are basic fibroblast growth factor (bFGF) (Richards et al., 1992), epidermal growth factor (EGF) (Reynolds and Weiss, 1992), and leukemia inhibitory factor (LIF) (Smith et al., 1988). As described above, one of the first difficulties encountered in attempting to culture NSCs was that isolated single cells tended to cease dividing and differentiate. Only cells surrounded by other cells remained mitotic, suggesting that cell-to-cell contact was important for cell proliferation (Temple and Qian, 1995). Further investigation suggested that a likely candidate for mediating this intercellular interaction was bFGF. Besides being known for its association with extracellular matrix and cell membranes, bFGF was surprisingly found to be present throughout CNS development (Baird, 1994; Kilpatrick et al., 1995). Previous studies had already shown that the addition of exogenous bFGF to cultures of cortical neuroectoderm cells appeared to stimulate their proliferation and lead to increased numbers of neurons (Gensburger et al., 1987), but, it was not until the 1990s that the issue was addressed of which type of cortical cell was the target of bFGF. It was found that bFGF would stimulate multipotent stem cells from many regions of the CNS (Temple and Qian, 1995; Gritti et al., 1996). In addition to bFGF, studies in the 1980s had already shown that EGF bound to CNS cells and stimulated their proliferation (Simpson et al., 1982; Anchan et al., 1991). EGF-induced proliferation of progenitor cells that began to divide and would come to form a cluster of undifferentiated cells that expressed nestin, an intermediate filament present in neuroepithelial stem-like cells. These cells, when they exited the cell cycle, would differentiate into neurons and astrocytes (Reynolds et al., 1992). These EGF-expanded cells, when then stimulated by bFGF would yield two progenitor cell subtypes: one giving rise to neurons and astrocytes, the other generating only neurons (Vescovi et al., 1993). Addition of both bFGF and EGF simultaneously in the same cell culture yielded better cell survival than either agent alone, and would also better maintain their ability to differentiate into neurons, oligodendrocytes, and astrocytes (Vescovi et al., 1999). The notion emerged that the most immature – and hence most multipotent – stem/progenitor cells bore receptors for both mitogens. Selecting for the presence of both receptors (and hence responsiveness to both mitogens) became the basis of a quick and easy screening technique for selecting NSC populations from a heterogeneous CNS culture (including from human material; Flax et al., 1998). The addition of LIF to cell cultures yielded better long-term growth than bFGF and EGF alone. No difference was noted in early stage cultures, but, after 50–60 days in vitro, the cultures without LIF consistently showed slower expansion while those with all three mitogens continued to expand (Carpenter et al., 1999). Therefore, the method most used now for expanding and maintaining NSC cultures by epigenetic means involves addition of bFGF, EGF, and LIF to a serum-free basal medium. Many studies (as well as our own experience) suggest that (at least for short periods) bFGF alone (at 20 ng/ml) is sufficient. As noted above, in addition to stimulating proliferation of NSCs, responsiveness to these factors may be one technique for selecting NSCs from a mixed population of cells at various degrees of maturation, potency, and lineage commitment. Culturing cells with mitogens alone is a double-edged sword. Because one does not by intent genetically manipulate the cells, approval by regulatory agencies for translation to the clinic may be easier to obtain. On the downside, however, obtaining a sufficient number of cells while avoiding senescence (particularly of human cells) can be quite challenging. Furthermore, one may be selecting for cells with particular avidity for these factors and hence may be more prone to neoplastic transformation. Genetic modifications that might be

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imposed by tonic exposure to high concentrations of a mitogen are mainly unknown and less controllable than some of the well-studied self-regulated or regulatable genetic augmentations described in the next section (Snyder et al., 1992; Roy et al., 2004). Genetic The genetic strategy for culturing NSCs in vitro seeks to augment and prolong the expression of “stemness” genes. Myc represents such a gene; it is becoming recognized as essential to the proliferative state of stem cells. When its expression is prolonged, cells remain in the cell cycle and hence their differentiation is held in abeyance. When they exit the cell cycle, they respond to regional epigenetic cues and differentiate appropriately – spontaneously and inherently downregulating myc translation. In other words, the exogenous extra copy of myc is regulated by the cell as it does its cellular myc. Interestingly, embryonic stem cells (ESCs) are colloquially said to be “naturally immortalized.” The goal in augmenting certain genes in somatic stem cells is to temporarily and controllably co-opt that quality – safely and effectively. The NSCs are usually transfected using a retroviral vector encoding the stemness gene. These clonal NSCs have the advantage of being easily expanded to large numbers that typically remain stable and homogenous over long periods of time from experiment to experiment and recipient to recipient without senescing. Such cells have not only proven to be an ideal model for understanding fundamental aspects of stem cell biology, but, because they have been used safely and effectively as transplant material for a wide variety of disease models over the past two decades, they have established the “gold standard” of what therapeutic benefit should be safely achievable by a cell with stem-like qualities (Parker et al., 2005). In fact, whatever expansion technique is elected for use by an investigator should be judged in part by its ability to attain at least this degree of safety and efficacy in models of disease and injury. Of course, when contemplating clinical strategies, a complete knowledge of the fate of these cells – as for any implanted cell – must be vouchsafed. As noted above, while such cells have proven exceptionally safe, regulatory agencies generally feel most comfortable approving the use of non-genetically modified cells. Ultimately, however, a well-characterized, uniform, stable, readily available, inexhaustible supply of clonally related NSCs may prove to be most efficacious and practical (Vescovi and Snyder, 1999). Other genes (e.g., telomerase) have been explored and proven valuable (Roy et al., 2004). To maximize comfort with stem cells, their “stemness” genes may be placed under regulatable control or put in tandem with suicide genes.

THERAPEUTICS AND CLINICAL APPROACHES Although the field of NSC biology is very young, strategies by which these cells and their emerging properties might be harnessed to ameliorate a range of neurological disorders have already captured the imagination of clinicians and helped fuel the new field of regenerative medicine. Broadly, there are two fundamental interlocking strategies for using NSCs therapeutically. The first seeks to use the stem cell to provide replacement cells for those that have become dysfunctional or lost. The second exploits the observation that stem cells (particularly in their non-neuronal state) constitutively produce neuroprotective, immunoregulatory, and homeostatic molecules that serve to reduce host cell loss and inhibit the formation of barriers to self-repair. We will review both lines of research as they apply to the most frequently studied diseases and where NSCs seem to be most promising. Spinal Cord Injury Experimental treatment strategies aimed at the acute stage of the injury process currently include neuronal protection and preservation of residual axons and white matter. Several approaches have been proposed and

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applied to this end, most notably anti-secondary injury therapy using high-dose methylprednisolone (MP) (Hall and Springer, 2004). Ideally, with effective neuroprotection in place, treatments to promote axonal regeneration can be started simultaneously or immediately thereafter. In pursuing this possibility, experimental use of neurotrophins or neurotrophic factors to increase neuroprotection and axonal growth, and neutralizing antibodies to Nogo and other oligodendrocyte-related inhibitory molecules have been investigated (Blesch and Tuszynski, 2002; Schwab 2004). While some produced encouraging outcomes, evidence to the contrary has also been reported. The conflicting findings do not negate the potential for this therapy but suggest that this strategy, when used in isolation may not be sufficient to promote functionally meaningful axonal regrowth in the injured spinal cord. For example, enzymatic or genetic antagonism to chondroitin sulfate proteoglycans (CSPGs) has offered some encouraging results. This approach was strengthened, however, when it was combined with maneuvers to prompt adult CNS neurons to reenter “growth mode” (Silver and Miller, 2004). Hence, combined therapies are likely to be most effective. Despite the fact that various treatment strategies have shown benefit in experimental animal models, there is still no effective therapy for chronic spinal cord injury (SCI). This frustrating situation, in our opinion, is attributable to the following realities. First, there has been no conclusive evidence favoring one process as the predominant pathophysiological mechanism which can account for all the spinal dysfunction seen following SCI. Most of the pathophysiological processes (e.g., secondary molecular events: glutamate toxicity, sodium and calcium influx, free radical insult, cytochrome c release; secondary pathophysiological events: ischemia, anoxia, apoptosis, etc.) apparently exist either simultaneously or sequentially in an interlocked manner throughout the evolution of the injury and represent different facets of this complicated disease entity (Tator and Fehlings, 1991; Teng et al., 2004). Most interventions reported to date target solely one facet of the injury process which, in isolation, is doomed to have limited benefit. To further complicate the situation, a given approach that may be useful when used alone, may become ineffective or even detrimental when used in combination with other interventions, perhaps working at cross purposes. Hence, it is critical to understand the intricate interactions between these options and identify the underlying mechanisms of their actions so that they may be orchestrated in a safe, synergetic, and clinically feasible fashion. Apropos to this point, we have come to view NSCs as the “glue” that can bind and integrate multiple approaches. Several studies have suggested that NSCs, when transplanted into the injured brain or spinal cord of rodents, migrate preferentially to and become integrated within the damaged areas. A subpopulation of the transplanted NSCs is redirected to differentiate into cell types that might replace the diseased or degenerated host cells (Rosario et al., 1997; Snyder et al., 1997; Park et al., 1999, 2002b; Yandava et al., 1999). More intriguingly – and, ultimately, of probably greater importance and utility – is the observation that undifferentiated NSCs or NSCs that have pursued a glial lineage seem to recondition the host CNS microenvironment and promote functional recovery by protecting pre-existent but threatened host neurons and circuitry (Teng et al., 2002). The impact of this action is probably greater than neuronal replacement. The precise mechanism by which NSCs exert this homeostatic pressure is unclear, though it is likely attributable to a large degree by intrinsic ability of NSCs to secrete neurotrophic factors, and/or immunomodulators (as demonstrated by Teng et al. (2002) and subsequently reported by many others, including Ourednik et al. (2002), Lu et al. (2003), Llado et al. (2004), Li et al. (2004, 2005), Yan et al. (2004), Bjugstad et al. (2005), and Pluchino et al. (2005)). Thus, preserving the multipotency of these cells – as opposed to attempting to direct them invariantly down the differentiation pathway of a single cell type – might offer the greatest chance for cell-based therapies of the different interlocked stages of SCI but in a parsimonious fashion. Harnessing the potentially broad therapeutic capacity of the NSC for use in an intelligent and rationale manner requires learning the principles

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that can govern interaction between the pathologic target and host environmental components, the NSC, and other therapeutic reagents. We believe that the innate biology of NSCs (i.e., their default production and secretion of various neurotrophic factors and other molecules in a differentiation stage-dependent fashion) enables them to interact with the surrounding environment, including releasing trophic factors in an appropriate, regulated, stimulus-appropriate manner. These factors, in our view, are components of the stem cell’s inherent developmental program which “calls upon it” to exert homeostatic forces upon a dynamically growing nervous system which, otherwise, could become dysequilibrated. The result of this inherent “program” – a dividend, so to speak, from developmental biology – is to promote, enable, induce, or catalyze the host to attempt to reconstitute its own tissue, to minimize barriers to this process, and to protect endangered cells from cell death or other toxic influences. Methods to optimize this process (i.e., to work in concert with normal developmental tendencies) is undoubtedly desirable for optimizing repair. An example of harnessing and exploiting such inherent stem cell programs will be presented here. To direct neural repair more effectively following SCI, we cultured NSCs ex vivo upon a biosynthetic scaffold that mimicked the general structure of a healthy spinal cord (Figure 18.1a). It had an inner section, engineered to

(a)

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Inner scaffold

Outer scaffold 1.5 mm

3 mm

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4 mm

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Figure 18.1 (a) Schematic of the scaffold design showing the inner and outer scaffolds. (b and c) Inner scaffolds seeded with NSCs (scale bars: 200 and 50 μm, respectively). The outer section of the scaffold was created with long, axially oriented pores intended for support axonal growth as well as radial pores to allow fluid transport and inhibit the ingrowth of connective and astrogliosis tissue: (d) scale bar, 100 μm. (e) Schematic of surgical insertion of the implant into the spinal cord (adapted from Teng et al., 2002).

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emulate the gray matter with an isotropic pore structure of 250–500 μm in diameter to facilitate seeding of the NSCs (Figure 18.1b and c). The outer section of the scaffold, modeled to emulate the white matter, had long, axially oriented pores for axonal guidance and radial porosity to allow fluid transport while inhibiting the ingrowth of scar tissue (Figure 18.1d) (Teng et al., 2002). Implantation of the scaffold–NSC unit into an adult rat hemi-section model of SCI promoted long-term improvement in function (persistent for 1 year in a group of animals specifically designated for long-term studies) relative to control groups. At 70 days postinjury, animals implanted with scaffold-plus-NSCs exhibited coordinated, weight-bearing hindlimb stepping. Histological and immunocytochemical analysis suggested that this recovery was not caused by neural cell replacement; rather, it is attributable predominantly to a reduction in host tissue loss from secondary injury processes as well as diminished glial scarring. This work is the first to demonstrate explicitly the “chaperone” neuroprotective effects of the NSC. Tract tracing demonstrated host corticospinal tract fibers passing through the injury epicenter to the caudal side of the lesion epicenter with classic growth cone-type varicosities at their leading ends, a finding not observed in untreated groups. Together with evidence of enhanced local GrowthAssociated Protein (GAP)-43 expression by axons, a marker associated with regenerative processes (also not seen in controls), these findings suggest a host-derived neuroregenerative component that might have also contributed to functional recovery. These results, besides suggesting a novel approach to SCI, may more broadly serve as a prototype for the use of NSCs to anchor multidisciplinary strategies in regenerative medicine, including gene therapy, material science, growth factor delivery, anti-inflammation and anti-scarring strategies, and pharmacological interventions against secondary injury. The use of scaffolds and cellular bridges are well suited for lesions in which there is large parenchymal loss (as was created experimentally with the hemi-section model described above) or where a syrinx might otherwise form because of extensive cell death following contusion. For injuries where the amount of tissue loss is less dramatic, the implantation of cells alone (without a template) may be useful. Murine ESCs were observed to survive and promote recovery in the contused spinal cord (McDonald et al., 1999). Although this recovery was at first attributed to the few neurons that appeared to emerge, a more detailed study showed that functional impact may, in fact, have been more plausibly due to oligodendrocytes that myelinated some traumatized host axons (Liu et al., 2000). Indeed, it has been suggested that the injured cord offers a microenvironment that is not favorable to the differentiation of multipotent NSCs into neurons (Cao et al., 2002). Rather, it has been proposed that transplanting neuronal- and glial-restricted precursors (NRP/GRP) is more tractable, that is, precommitting the cells to a particular lineage ex vivo rather than letting the in vivo environment direct their differentiation. The transplantation of NRP/GRPs into the postcontusion spinal cord did improve motor and sensory function. Histological analysis showed that a subset of the NRP/GRPs survived, filled the lesion site, differentiated into neurons and glia, and migrated selectively (Cao et al., 2005; Mitsui et al., 2005). Interestingly, the volume of spinal cord spared was increased in NRP/GRP recipients, suggesting that their action may nevertheless have been attributable in a large part to local neuroprotection. The actual role that donor-derived neurons played in recreating neurocircuitry is not clear. We are coming to learn that replacing lost neural tissue with its complex connections is more challenging than once assumed, even using a cell that can yield immunocytochemically- and electrophysiologically proven neurons in vivo. We do believe that, as we better understand the molecular milieu of the injured cord, some degree of reconstruction will be achievable – perhaps not of long projection neurons, but perhaps of interneurons with shorter axons. However, it is fortuitous that the very same exogenous cells that we hope will participate in circuit reconstitution may also concurrently be providing trophic and immunoregulatory factors that can preserve host circuits. Future research must clearly pay attention not only to how the donor cells are changed by the host but also how the host is changed by the donor cells (i.e., placing a greater research

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emphasis on dynamic donor–host interactions). Such an understanding may allow this dynamic to be optimized such that a greater number of endogenous stem cells may be effectively mobilized. Endogenous NSCs have been established to reside in a few well-characterized secondary germinal zones of the adult nervous system, most notably the subgranular zone (SGZ) of the dentate gyrus of the hippocampus and the subventricular zone (SVZ) of the forebrain. Endogenous NSCs may also reside in the ependymal region of the spinal cord. In response to injuries like stroke, adult hippocampal NSCs may proliferate and differentiate into new, functioning neurons (Schaffer and Gage, 2004). However, NSCs in the adult spinal ependymal region do not seem to differentiate into neurons when they reside in their normal spinal cord niche. Nevertheless, when these same NSCs are transplanted into the SGZ, they will yield neurons (Doetsch et al., 1999; Johansson et al., 1999). Hence, limitations to neurogenesis must emanate in large part from the microenvironment of the adult spinal cord rather than from the cells themselves. Therefore, it seems feasible that either altering the milieu or changing the cells to respond differently to that milieu may permit these endogenous spinal NSCs to play a more prominent role in neuronal reconstitution in the adult cord (Danilov et al., 2006). To overcome the normal impediments to adult neurogenesis will require a better understanding of the biological roles of spinal NSCs, especially their proliferation in response to injury, inflammation, and neuroactivity – all significantly unexplored (Teng et al., 2006). Neurodegenerative Diseases Neurodegenerative diseases are a prominent target for stem cell therapy in general and NSCs in particular. Although adult-onset Alzheimer’s disease (AD), Parkinson’s disease (PD), Amyotrophic lateral sclerosis (ALS), and Huntington’s Disease (HD) tend to receive the most attention, there is a growing recognition that some of the lysosomal storage diseases (LSDs) affecting the nervous system in childhood are attractive targets because each disease requires the replacement of a single well-characterized, usually diffusible, enzyme without a significant need for cell replacement if treated early. In the case of the LSDs, the enzymes to be replaced are typically produced constitutively by the NSC as part of its normal “housekeeping” (e.g., Sidman et al., 2004). The notion that using NSCs for molecular therapies is likely more tractable even for neurodegenerative diseases that do ostensibly require cell replacement has begun to spill over into the adult literature, as well. Although traditionally, for diseases like PD, this has entailed arming cells to provide substrates for dopamine production, such as tyrosine hydroxylase (TH) and dopa-decarboxylase (DDC) (e.g., Kim et al., 2000; Kim, 2004), we have come to view “molecular therapies” more broadly, as described above for SCI, using the NSC’s intrinsic propensity to supply homeostasis-promoting, neuroprotective, and neurotrophic agents that directly benefit host neural functioning. In preliminary studies, as proof of concept of this notion using human NSCs (hNSCs), we have approached a model of PD that most faithfully mirrors the actual human entity: non-human primates (African green monkeys, a macaque) exposed to the complex I mitochondrial toxin 1-methyl4-phenyl-1,2,3,6-tetra-hydropyridine (MPTP). hNSCs were implanted into the left and right caudate nucleus and the right substantia nigra (SN) of eight MPTP-treated, severely Parkinsonian monkeys. By 4 months (and clearly by 7 months) post-transplantation, the majority of hNSCs were found bilaterally along the nigrostriatal pathway and in the SN. In the presence of NSC, the number and size of host TH–and dopamine transporter (DAT)-expressing neurons in SN, typically “poisoned” by MPTP, returned to nearly non-MPTP-exposure control levels. Host TH neurons in the caudate, whose size-to-number ratio had become dysequilibrated, also returned to their normal ratio, as if forced toward equipoise by stem cell transplant (Bjugstad et al., 2005; Redmond et al., 2007). The host-nigral striatal pathway, normally attenuated and dysfunctional in the MPTPlesioned monkey (as in PD), was restored (or preserved). Taken together, these actions resulted in significant

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and long-term diminution in Parkinsonian symptoms. Therefore, a broader view of the role of stem cells – even for diseases characterized by degeneration of a particular neuronal subtype, as in PD – may make these diseases a more tractable therapeutic target than we initially presumed (Li et al., 2006a, 2006b; Sidman et al., 2007). Diseases characterized by glial dysfunction might be viewed in a similar manner. Although we had established long ago that NSCs could yield effectively remyelinating oligodendrocytes throughout the brains of poorly myelinated mouse mutants (Yandava et al., 1999), some of the most prevalent white matter diseases of adulthood, such as multiple sclerosis (MS), are characterized by an environment that is likely inhospitable to both host and donor-derived oligodendrocytes. For example, in the experimental allergic encephalomyelitis (EAE) model of MS (characterized by chronic CNS inflammation, multifocal demyelination, and axonal loss), syngeneic undifferentiated neural precursor cells injected intravascularly, transited from the vascular space into the inflamed intracerebral microenvironment (via constitutively activated integrins and functional chemokine receptors), accumulated and survived within perivascular areas (in the company of reactive astrocytes, inflamed endothelial cells, and encephalitogenic T-cells that produced neurogenic and gliogenic regulators) and helped to preserve host oligodendrocytes (more so than replacing oligodendrocytes) by exerting an anti-inflammatory effect (Pluchino et al., 2003, 2005), thus protecting against chronic neural tissue loss. We had previously observed this anti-inflammatory action of NSCs in a cerebral-ischemia model (Park et al., 2002a); now it was being recapitulated in a bona fide inflammatory disease. In the EAE model, the NSCs appeared to counteract the inflammation by inducing apoptosis of blood-borne CNS-infiltrating encephalitogenic T-cells. In addition, there was a significant reduction in astrogliosis as we had noted also in traumatic diseases and described above (Park et al., 2002a; Teng et al., 2002). Taken together, these actions resulted in a marked decrease in the extent of demyelination and axonal loss, and, in turn, disease-related disability, both clinically and neurophysiologically (Pluchino et al., 2003). Stroke The presumptive goal in using cell-based therapies in ischemic (stroke) or hypoxic–ischemic injury would be to replace infarcted CNS tissue in an appropriate organotypic manner. Given the data presented above, however, it seems plausible to expect that cell-based approaches might also provide trophic and neuroprotective support to tissue at risk in the penumbra surrounding the infarct, inhibit inflammation and scarring, promote angiogenesis, and help promote the mobilization, migration, survival, and differentiation of endogenous precursor cells (Hass et al., 2005). Stem cell therapy for stroke may be divided into two approaches: the first focuses on mobilizing endogenous NSCs and the second depends on providing exogeneous NSCs. Obviously, as suggested above, not only will both approaches likely be required for optimal restitution of function, but the two strategies likely act synergistically. Typically, the strategy for exploiting the population of endogenous NSCs has been to attempt to stimulate their proliferation and neuronal differentiation by administering exogenous growth factors and other pharmacological agents (e.g., bFGF, TGF-alpha, erythropoietin). Although some human studies suggest safety and efficacy of this approach (Ehrenreich et al., 2002), the misadventures of the neurotrophic factor field in the 1980s and 1990s, where the unanticipated pleiotrophic actions of systemically administered growth factors produced untoward effects, suggests extreme caution before large scale clinical application. With regard to using exogenous NSCs for therapy against ischemic injury, a number of interesting insights have emerged that draw on the growing field of tissue engineering. It was observed that, in conditions where the likelihood of parenchymal loss is greatest, use of a biodegradable synthetic scaffold to support exogenous NSCs transiently within the injured terrain served to prolong the reciprocal interaction between the donor and host, fix instructive molecules emanating from both, abet the inhibition of astroglial and

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inflammatory host reactions, and provide a template for donor-derived and host-fiber regrowth. Using hypoxic–ischemic injury as a prototype for insults characterized by extensive tissue loss, Park et al. (2002a) seeded NSCs onto a polymer scaffold that was subsequently implanted into the infarction cavities. Indeed, not only did this approach serve a dramatic therapeutic function, but it allowed the investigators to document for the first time the multiple reciprocal interactions that spontaneously ensue between NSCs and the extensively damaged brain: parenchymal loss was dramatically reduced, an intricate meshwork of many highly arborized neurites of both host- and donor-derived neurons emerged, and some anatomical connections appeared to be reconstituted. The NSC–scaffold complex altered the trajectory and complexity of host cortical neurites. Reciprocally, donor-derived neurons appeared to be capable of directed, target-appropriate neurite outgrowth. These “biobridges” appeared to unveil or augment a constitutive reparative response by facilitating a series of reciprocal interactions between NSC and host, including promoting neuronal differentiation, enhancing the elaboration of neural processes, fostering the re-formation of cortical tissue, promoting connectivity and prompting revascularization of new parenchyma by the host. Inflammation and scarring were also reduced. Another interesting observation is that NSCs administered intravenously in the systemic circulation may migrate into lesioned brain sites, differentiate into neurons and glia and improve functional deficits in rat models of focal ischemia or cerebral hemorrhage (Chu et al., 2003; Jeong et al., 2003; Kim, 2004). This approach to ischemic injury using an intravascular route for the delivery of NSCs intracranially extends the observations first made using animal models of intracranial brain tumors (Aboody et al., 2000) and the above-described EAE (Pluchino et al., 2003, 2005). Although results in animal models of stroke seem promising, challenges remain before attempting human therapies. For example, obtaining the requisite number of the proper cells that circumvent immunorejection yet interact effectively with host neurocircuitry and limit their impact and distribution solely to the CNS are important considerations. Some answers may be found through a better understanding of the molecular events that underlie each of the key responses of the injured adult brain to donor NSCs and vice versa.

CONCLUSION Even though studies regarding neurogenesis date from 1913, only recently has the cellular and molecular basis of this inborn plasticity begun to be understood. The stem cell appears to be the repository of much of this plasticity. The ability to identify, select, isolate, clone, culture, differentiate, genetically manipulate, and transplant NSCs has clearly advanced our understanding of this biology. Indeed, given that the NSC is the first stem cell isolated from a solid organ, insights derived from studying these cells have and will continue to help advance our understanding of development and repair of other solid organs. Although research into the use of NSCs, dating back to the late 1980s, initially focused solely on their potential for replacing injured or dysfunctional neurons, we have come to recognize that the original view was overly narrow and simplistic. The inherent biology of the NSC – the richness and complexity of which we are only now beginning to appreciate – offers many other therapeutic tools, including effecting neuroprotection and immunoregulation, induction of and inhibition of obstacles to inborn regenerative processes, axonal guidance, promotion of angiogenesis, exertion of homeostatic pressure, and presumably other actions yet to be unveiled. These multifaceted actions of NSCs make them ideally suited for anchoring the multimodal interventions that we are coming to recognize will be needed to restore function in most neurological disorders. Complex diseases require complex solutions. NSCs, as we have described in this review, have interfaced and worked synergistically with gene and growth factor therapy, anti-apoptotic and neuroprotective

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strategies, stimulation of neurogenesis, anti-inflammatory and anti-scarring approaches, material science and tissue engineering. We, therefore, propose an updated concept of the NSC. The field’s conventional view which has touched principally on the essential multipotency of lineage phenotypes (i.e., the ability of NSCs to differentiate into all neural cells) should be broadened to include the emerging recognition of the biofunctional multipotency of the NSC to mediate systemic homeostasis. Under this new conceptual context, one may begin to appreciate and seek the “logic” and teleology behind the wide range of molecular tactics the NSC appears to serve at each developmental stage as it integrates into and prepares, modifies, and guides the surrounding CNS microand macro-environment toward the formation and self-maintenance of a physiologically functioning adult nervous system. We believe that embracing this view of the NSC’s “multipotency” is pivotal for correctly, efficiently, and optimally exploiting stem cell biology for therapeutic applications including reconstituting the dysfunctional CNS.

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Meng, F., Zolova, O., Kokorina, N.A., Dobretsova, A. and Wight, P.A. (2005). Characterization of an intronic enhancer that regulates myelin proteolipid protein (Plp) gene expression in oligodendrocytes. J. Neurosci. Res. 82: 346–356. Miale, I.L. and Sidman, R.L. (1961). An autoradiographic analysis of histogenesis in the mouse cerebellum. Exp. Neurol. 4: 277–296. Miller, M.W. and Nowakowski, R.S. (1988). Use of bromodeoxyuridine-immunohistochemistry to examine the proliferation, migration and time of origin of cells in the central nervous system. Brain Res. 457: 44–52. Miller, R.H. (2002). Regulation of oligodendrocyte development in the vertebrate CNS. Prog. Neurobiol. 67: 451–467. Miller, R.H. (2005). Dorsally derived oligodendrocytes come of age. Neuron 45: 1–3. Ming, G.L. and Song, H. (2005). Adult Neurogenesis in the mammalian central nervous system. Annu. Rev. Neurosci. 28: 223–250. Mitsui, T., Shumsky, J.S., Lepore, A.C., Murray, M. and Fischer, I. (2005). Transplantation of neuronal and glial restricted precursors into contused spinal cord improves bladder and motor functions, decreases thermal hypersensitivity, and modifies intraspinal circuitry. J. Neurosci. 25: 9624–9636. Nottebohm, F. (2004). The road we travelled: discovery, choreography, and significance of brain replaceable neurons. Ann. NY Acad. Sci. 1016: 628–658. Ogle, B.M., Cascalho, M. and Platt, J.L. (2005). Biological implications of cell fusion. Nat. Rev. Mol. Cell Biol. 6: 567–575. Ortiz-Gonzalez, X.R., Keene, C.D., Verfaillie, C.M. and Low, W.C. (2004). Neural induction of adult bone marrow and umbilical cord stem cells. Curr. Neurovasc. Res. 3: 207–213. Ourednik, J., Ourednik, V., Lynch, W.P., Schachner, M. and Snyder E.Y. (2002). Neural stem cells display an inherent mechanism for rescuing dysfunctional neurons. Nat. Biotechnol. 20: 1103–1110. Park, K.I., Liu, S., Flax, J.D., Nissim, S., Stieg, P.E. and Snyder, E.Y. (1999). Transplantation of neural progenitor and stem cells: developmental insights may suggest new therapies for spinal cord and other CNS dysfunction. J. Neurotrauma 16: 675–687. Park, K.I., Teng, Y.D. and Snyder, E.Y. (2002a). The injured brain interacts reciprocally with neural stem cells supported by scaffolds to reconstitute lost tissue. Nat. Biotechnol. 20: 1111–1117. Park, K.I., Ourednik, J., Ourednik, V., Taylor, R.M., Aboody, K.A., Auguste, K.I., Lachyankar, M., Teng, Y.D., Redmond, D.E. and Snyder, E.Y. (2002b). Global gene and cell replacement strategies via stem cells. Gene Ther. 9: 613–624. Parker, M.A., Anderson, J.K., Corliss, D.A., Abraria, V.E., Sidman, R.L., Park, K.I., Teng, Y.D., Cotanche, D.A. and Snyder, E.Y. (2005). Expression profile of an operationally-defined neural stem cell clone. Exp. Neurol. 194: 320–332. Petersen, B.E., Bowen, W.C., Patrene, K.D., Mars, W.M., Sullivan, A.K., Murase, N., Boggs, S.S., Greenberger, J.S. and Goff, J.P. (1999). Bone marrow as a potential source of hepatic oval cells. Science 284: 1168–1170. Pluchino, S., Quattrini, A., Brambilla, E., Gritti, A., Salani, G., Dina, G., Galli R., Del Carro U., Amadio S., Bergami A., et al. (2003). Injection of adult neurospheres induces recovery in a chronic model of multiple sclerosis. Nature 422: 688–694. Pluchino, S., Zanotti, L., Rossi, B., Brambilla, E., Ottoboni, L., Salani, G., Martinello, M., Cattalini, A., Bergami, A., Furlan, R., et al. (2005). Neurosphere-derived multipotent precursors promote neuroprotection by an immunomodulatory mechanism. Nature 436: 266–271. Price, J., Turner, D. and Cepko, C. (1987). Lineage analysis in the vertebrate nervous system by retrovirus-mediated gene transfer. Proc. Natl Acad. Sci. USA 84: 156–60. Redmond, D.E., Jr, Bjugstad, K.B., Teng, Y.D., Ourednik, V., Ourednik. J., Wakeman, D.R., Parsons, X.H., Gonzalez, R., Blanchard, B.C., Kim, S.U., Gu, Z., Lipton, S.A., Markakis, E.A., Roth, R.H., Elsworth, J.D., Sladek, J.R. Jr, Sidman, R.L. and Snyder, E.Y. (2007). Behavioral improvement in a primate Parkinson's model is associated with multiple homeostatic effects of human neural stem cells. Proc. Natl Acad. Sci. USA 104: 12175–12180. Renfranz, P.J., Cunningham, M.G. and McKay, R.D. (1991). Region-specific differentiation of the hippocampal stem cell line HiB5 upon implantation into the developing mammalian brain. Cell 66: 713–729. Reynolds, B.A. and Weiss, S. (1992). Generation of neurons and astrocytes from isolated cells of the adult mammalian central nervous system. Science 255: 1707–1710. Reynolds, B.A., Tetzlaff, W. and Weiss, S. (1992). A multipotent EGF-responsive striatal embryonic progenitor cell produces neurons and astrocytes. J. Neurosci. 12: 4565–4574. Richards L.J., Kilpatrick T.J. and Bartlett P.F. (1992). De novo generation of neuronal cells from the adult mouse brain. Proc. Natl Acad. Sci. USA 89: 8591–8595.

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Rosario, C.M., Yandava, B.D., Kosaras, B., Zurakowski, D., Sidman, R.L. and Snyder, E.Y. (1997). Differentiation of engrafted multipotent neural progenitors towards replacement of missing granule neurons in meander tail cerebellum may help determine the locus of mutant gene action. Development 124: 4213–4224. Roy, N.S., Nakano, T., Keyoung, H.M., Windrem, M., Rashbaum, W.K., Alonso, M.L., Kang, J., Peng, W., Carpenter, M.K., Lin, J., et al. (2004). Telomerase immortalization of neuronally restricted progenitor cells derived from the human fetal spinal cord. Nat. Biotechnol. 22: 297–305. Ryder, E.F., Snyder, E.Y. and Cepko, C.L. (1990). Establishment and characterization of multipotent neural cell lines using retrovirus vector-mediated oncogene transfer. J. Neurobiol. 21: 356–375. Sanes, J.R., Rubenstein, J.L. and Nicolas, J.F. (1986). Use of a recombinant retrovirus to study postimplantation cell lineage in mouse embryos. EMBO J. 5: 3133–3142. Schaffer, D.V. and Gage, F.H. (2004). Neurogenesis and neuroadaptation. Neuromol. Med. 5: 1–9. Schwab, M.E. (2004). Nogo and axon regeneration. Curr. Opin. Neurobiol. 14: 118–124. Sidman, R.L. (1961). Histogenesis of mouse retina studied with thymidine-H3. In: Smelser, G.K. (ed.), The Structure of the Eye. New York: Academic Press, pp. 487–506. Sidman, R.L. (1970). Autoradiographic methods and principles for study of the nervous system with thymidine-H3. In: Nauta, W.J. and Ebbesson, S.O.E. (eds.), Contemporary Research Methods in Neuroanatomy. New York: SpringerVerlag, pp. 252–274. Sidman, R.L., Miale, I.L. and Feder, N. (1959). Cell proliferation and migration in the primitive ependymal zone; an autoradiographic study of histogenesis in the nervous system. Exp. Neurol. 1: 322–333. Sidman, R.L., Li, J., Stewart, G.R., Clarke. J., Yang, W., Snyder, E.Y. and Shihabuddin, L.S. (2007). Injection of mouse and human neural stem cells into neonatal Niemann-Pick A model mice. Brain Res. 1140: 195–204. Sidman, R.L., Shihabuddin, L.S., Li, J., Clarke, J., Snyder, E.Y. and Stewart, G.R. (2004). Transplantation of mouse and human neural stem cells into neonatal Niemann-Pick A mice. J. Neurochem. 90 (Suppl 1): 55. Silver, J. and Miller, J.H. (2004). Regeneration beyond the glial scar. Nat. Rev. Neurosci. 5: 146–156. Simpson, D.L., Morrison, R., de Vellis, J. and Herschman, H.R. (1982). Epidermal growth factor binding and mitogenic activity on purified populations of cells from the central nervous system. J. Neurosci. Res. 8: 453–462. Smart, I. (1961). The subependymal layer of the mouse brain and its cell production as shown by autography after [H3]thymidine injection. J. Comp. Neurol. 116: 325–327. Smith, A.G., Heath J.K., Donaldson, D.D., Wong, G.G., Moreau, J., Stahl, M. and Rogers, D. (1988). Inhibition of pluripotential embryonic stem cell differentiation by purified polypeptides. Nature 336: 688–690. Snyder, E.Y., Deitcher, D.L., Walsh, C., Arnold-Aldea, S., Hartweig, E.A. and Cepko, C.L. (1992). Multipotent neural cell lines can engraft and participate in development of mouse cerebellum. Cell 68: 33–51. Snyder, E.Y., Yoon, C.H., Flax, J.D. and Macklis, J.D. (1997). Multipotent neural progenitors can differentiate toward replacement of neurons undergoing targeted apoptotic degeneration in adult mouse neocortex. Proc. Natl Acad. Sci. USA 94: 11663–11668. Stanfield, B.B. and Trice, J.E. (1988). Evidence that granule cells generated in the dentate gyrus of adult rats extend axonal projections. Exp. Brain Res. 72: 399–406. Tator, C.H. and Fehlings, M.G. (1991). Review of the secondary injury theory of acute spinal cord trauma with emphasis on vascular mechanisms. J. Neurosurg. 75: 15–26. Temple, S. (1989). Division and differentiation of isolated CNS blast cells in microculture. Nature 340: 471–473. Temple, S. and Qian, X. (1995). bFGF, neurotrophins, and the control of cortical neurogenesis. Neuron 15: 249–252. Teng, Y.D., Lavik, E.B., Qu, X., Park, K.I., Ourednik, J., Zurakowski, D., Langer, R. and Snyder, E.Y. (2002). Functional recovery following traumatic spinal cord injury mediated by a unique polymer scaffold seeded with neural stem cells. Proc. Natl Acad. Sci. USA 99: 3024–3029. Teng, Y.D., Choi, H., Onario, R.C., Zhu, S., Desilets, F.C., Lan, S., Woodard, E.J., Snyder, E.Y., Eichler, M.E. and Friedlander, R.M. (2004). Minocycline inhibits contusion-triggered mitochondrial cytochrome c release and mitigates functional deficits after spinal cord injury. Proc. Natl Acad. Sci. USA 101: 3071–3076. Teng, Y.D., Liao, W.-L., Choi, H., Konya, D., Sabharwal, S., Langer, R., Sidman, R.L., Snyder, E.Y., and Frontera, W.R. (2006). Physical activity-mediated functional recovery after spinal cord injury: potential roles of neural stem cells. Regen. Med. 1: 763–776.

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Timsit, S.G., Martinez, S., Alliquant, B., Peyron, F., Puelles, L. and Zalc, B. (1995). Oligodendrocytes originate in a restricted zone of the embryonic ventral neural tube defined by DM-20 mRNA expression. J. Neurosci. 15: 1012–1024. van Praag, H., Schinder, A.F., Christie, B.R., Toni, N., Palmer, T.D. and Gage, F.H. (2002). Functional neurogenesis in the adult hippocampus. Nature 415: 1030–1034. Vescovi, A.L. and Snyder, E.Y. (1999). Establishment and properties of neural stem cell clones: plasticity in vitro and in vivo. Brain Pathol. 9: 569–598. Vescovi, A.L., Reynolds, B.A., Fraser, D.D. and Weiss S. (1993). bFGF regulates the proliferative fate of unipotent (neuronal) and bipotent (neuronal/astroglial) EGF-generated CNS progenitor cells. Neuron 11: 951–966. Vescovi, A.L., Eugenio, A.P., Gritti, A., Poulin, P., Ferrario, M., Wanke, E., Frolichsthal-Schoeller, P., Cova, L., ArcellanaPanlilio, M., Colombo, A., et al. (1999). Isolation and cloning of multipotential stem cells from the embryonic human CNS and establishment of transplantable human neural stem cell lines by epigenetic stimulation. Exp. Neurol. 156: 71–83. Williams, B.P., Read, J. and Price J. (1991). The generation of neurons and oligodendrocytes from a common precursor cell. Neuron 7: 685–693. Yan, J., Welsh, A.M., Bora, S.H., Snyder, E.Y. and Koliatsos, V.E. (2004). Differentiation and tropic/trophic effects of exogenous neural precursors in the adult spinal cord. J. Comp. Neurol. 480: 101–114. Yandava, B.D., Billinghurst, L. and Snyder, E.Y. (1999). “Global” cell replacement is feasible via neural stem cell transplantation: evidence from the dysmyelinated shiverer mouse brain. Proc. Natl Acad. Sci. USA 96: 7029–7034. Yu, W.-P, Collarini, E.J., Pringle, N.P. and Richardson, W.D. (1994). Embryonic expression of myelin genes: evidence for a focal source of oligodendrocyte precursors in the ventricular zone of the neural tube. Neuron 23: 1353–1362.

19 Mesenchymal Stem Cells Zulma Gazit, Hadi Aslan, Yossi Gafni, Nadav Kimelman, Gadi Pelled, and Dan Gazit

INTRODUCTION In the development of stem cell-based therapeutic platforms for tissue regeneration, the selection of which type of stem cell to use will be enormously important. Adult mesenchymal stem cells (MSCs) are considered one of the most promising tools for cell and cell-based gene therapy in bone repair (Gafni et al., 2004). Adult MSCs have been shown to possess the potential to differentiate into several lineages including bone, cartilage, fat, tendon, muscle, and marrow stroma (Haynesworth et al., 1992; Mackay et al., 1998; Yoo et al., 1998; Young et al., 1998; reviewed by Caplan and Bruder, 2001). The best known source of MSCs in adult humans is the bone marrow (BM) compartment; this region contains several types of cells, including those of the hematopoietic lineage as well as endothelial cells (ECs) and MSCs that are part of the marrow stromal system (Pittenger et al., 1999). Other sources of MSCs have also been identified, such as fat tissue (Zuk et al., 2001, 2002), cord blood (Hong et al., 2005; Jeong et al., 2005; Moon et al., 2005), and peripheral blood, although the latter finding is still controversial (Fernandez et al., 1997; Conrad et al., 2002). Several protocols were recently established to enable regeneration of large bone defects by using human MSCs (hMSCs) that have been expanded in culture. These cells differentiate into osteogenic cells and, as vehicles, deliver a therapeutic gene product such as one of the bone morphogenetic proteins (BMPs) (Turgeman et al., 2001; Peterson et al., 2005; reviewed by Gamradt and Lieberman, 2004). It has been shown that in combination with BMP-2, hMSCs are able to heal full-thickness nonunion bone defects (Turgeman et al., 2001; Dragoo et al., 2003). In addition, Lee et al. (2001) have demonstrated that, following transduction with retroviral vectors, in vivo implantation, and differentiation, hMSCs can maintain stable expression of the therapeutic gene. In these studies, MSCs were isolated from BM, expanded in culture (in some cases genetically engineered) and implanted in vivo. Reports of these studies and many others have emphasized the benefit of MSCs as vehicles for cellmediated gene therapy in the field of orthopedics (Gafni et al., 2004). In addition, MSCs have been implemented in regeneration of the heart (cardiac muscle and vascular system), skeletal muscle, nerve, liver, and pancreas, with regeneration of cardiac tissue being foremost (Burt et al., 2002; Lardon et al., 2002; Bonafe et al., 2003; Dabeva et al., 2003; Abedin et al., 2004; Kim et al., 2004; Jain et al., 2005; Sonoyama et al., 2005; Goncalves et al., 2006). In cell-based therapies, the culture expansion stage is extremely costly and time consuming, and in many cases cells may lose their multipotentiality in vivo and fail to meet the desired goal. Rubio et al. (2005) reported that cultured hMSCs can undergo spontaneous transformation as a consequence of in vitro expansion. In very few articles has the use of noncultured freshly isolated hMSCs been described. Recently, CD105 hMSCs were isolated from BM and were shown to exhibit in vivo osteogenic potential prior to in vitro expansion suggesting the utilization of these cells as freshly isolated population and avoiding the culture-expansion stage (Aslan et al., 2006b). Horwitz et al. (1999) showed that hMSCs present in unprocessed BM allografts engraft and may provide a stem cell reservoir for the differentiation and renewal of osteoblasts. The enrichment of mesenchymal

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progenitors, derived from fresh BM aspirates, in cancellous bone matrices has been found to increase bone formation and the bone union score significantly in a spinal fusion model (Muschler et al., 2003). Rombouts and Ploemacher have demonstrated that culture expansion attenuates the homing ability of MSCs after systemic infusion in irradiated mice (Rombouts et al., 2003). This indicates that MSCs may lose some of their natural stem cell characteristics following expansion in vitro. Other investigators have proposed that all known characteristics of MSCs may be an outcome of the culture stage and do not really represent the actual characteristics of MSCs residing in vivo at the BM niche (Javazon et al., 2004). The isolation of an hMSC-enriched population requires an efficient and reproducible method. Few methods have been described for the isolation of MSCs, including enhancement of the plastic-adherence property of the cells by using selected amounts of fetal calf serum (FCS) (Kadiyala et al., 1997; Pittenger et al., 1999) and immunomagnetic isolation based on the presence of the STRO-1 surface molecule (Gronthos et al., 1995, 2003). These methods have not been used in any study to show the differentiation potential of cells before culture expansion. In the study conducted by Majumdar et al. (2000), the anti-CD105 (endoglin) antibody was used to isolate cells from human BM aspirates; after expansion in culture these cells differentiated in vitro into chondrogenic cells and displayed an immunophenotype distinctive to hMSCs. We recently reported that we used the CD105-based immunoisolation method to obtain a fresh noncultured population of hMSCs and to determine these cells’ osteogenic potential both in vitro and in vivo. Our results demonstrate that this noncultured population of adult stem cells can be genetically engineered and induced to undergo osteogenic differentiation in vivo – thus showing the cells’ potential to serve as an attractive therapeutic tool for bone regeneration purposes (Aslan et al., 2006b). One striking feature of MSC therapy is the cumulative data on the tolerance shown by the host to allogeneic MSCs. The mechanisms by which this immunotolerance exist are complex and have not yet been thoroughly identified. It has been shown that there is a low expression of alloantigens by MSCs, and this might involve cell contact-dependent or -independent pathways, which are modulated by secretion of soluble factors such as interleukin (IL)-2 and IL-10, transforming growth factor-beta1 (TGFβ1), prostaglandin E2 (PGE2), and hepatocyte growth factor (HGF) among others. Immune system cells, such as dendritic cells (DCs) and T-cells, have also been shown to be affected by the presence of MSCs in mixed lymphocyte cultures (MLCs) (Beyth et al., 2005). In addition to the advantage that these cells offer the field of regenerative medicine, MSCs provide prophylaxis against graft-versus-host disease in cases of allogeneic hematopoietic stem cell (HSC) transplantation.

THE DEFINITION OF MSCS BM was the first tissue described as a source of plastic-adherent, fibroblast-like cells that develops colonyforming units (CFU-Fs) when plated in tissue culture plates (Friedenstein et al., 1982, 1987). These cells, originally designated stromal cells, elicited much attention in stem cell research during the mid-1990s and the beginning of the 21st century. The main goals of studies conducted using these cells were to find an ultimate pure cell population that could be further utilized for regenerative purposes. In these studies, cells were isolated using several methods that will be discussed later in this chapter and were given names such as MSCs, mesenchymal progenitors, stromal stem cells, among others. The precise definition of these cells remains a matter of debate. Nevertheless, to date MSCs are widely defined as a plastic-adherent cell population that can be directed to differentiate in vitro into cells of osteogenic, chondrogenic, adipogenic, myogenic, and other lineages (Pittenger et al., 1999; Javazon et al., 2004). As part of their stem cell nature, MSCs proliferate and give rise to daughter cells that have the same pattern of gene expression and phenotype and, therefore, maintain the “stemness” of the original cells. In the

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presence of specific signals such as those given by growth factors, cytokines, and components of the extracellular matrix, a certain population of daughter cells undergoes a sequential cascade of differentiation that alters the cells’ original gene expression pattern. Self-renewal and differentiation potential are two criteria that define MSCs as real stem cells; however, these characteristics have only been proved after in vitro manipulation, and there is no clear description of the characteristics displayed by unmanipulated MSCs in vivo (Javazon et al., 2004). Our limited knowledge of MSCs is due to the fact that MSCs lack a unique marker, in contrast to other stem cells such as HSCs, which are identified by the expression of the CD34 surface marker. The CD105 surface antigen (endoglin) has been recently used to isolate hMSCs from BM and such an approach enabled the characterization of freshly isolated hMSCs before culture. A distinct expression of certain surface antigens such as CD45 and CD31 was demonstrated in freshly isolated hMSCs and the expression of these molecules was lower in culture-expanded hMSCs (Aslan et al., 2006b). These data suggest, again, the alterations that hMSCs may undergo during culture. In several studies, cultured MSCs have been characterized either by using cell surface antigens or by examining the cells’ differentiation potential. The most accepted characteristics for in vitro grown MSCs are the following: (1) the ability to form CFU-Fs when plated in plastic tissue culture plates in the presence of an animal serum such as FCS in a basic medium such as Dulbecco modified Eagle medium; (2) the expandability of these cells without losing their differentiation potential; and (3) the high levels of expression of the surface antigens CD105, CD73, CD29, CD44, CD71, CD90, CD106, CD120a, and CD124, and the low levels of expression of CD14, CD34, and the leukocyte common antigen CD45 (reviewed by Deans and Moseley, 2000).

THE STEM CELL NATURE OF MSCS Stem cells are defined by their ability to self-renew and by their potential to undergo differentiation into functional cells under the right conditions. MSCs exhibit the potential to differentiate into the osteogenic, chondrogenic, adipogenic, tenogenic, myogenic, or stromal lineages (Haynesworth et al., 1992; Mackay et al., 1998; Yoo et al., 1998; Young et al., 1998; reviewed by Caplan and Bruder, 2001). In the presence of certain agents, such as a combination of ascorbic acid, β-glycerophosphate, and dexamethasone, or in the presence of BMPs, MSCs undergo a series of morphological and metabolic changes until they exhibit characteristics of osteogenic cells, which include elevated levels of alkaline phosphatase, osteopontin, and osteocalcin, and accumulation of calcium. Culturing MSCs in a three-dimensional manner (such as a pellet culture) and in the presence of TGFβ1 can induce the formation of collagen II and glucosaminoglycans within these cultures, therefore creating a cartilage-like tissue. Differentiation of MSCs into adipogenic cells has also been achieved in vitro, as demonstrated by the accumulation of fatty acid droplets within these cells (Pittenger et al., 1999). In addition to their in vitro differentiation potential, MSCs have been shown to home to and engraft into several organs and tissues when injected systemically. Human BM-derived MSCs transplanted into the peritoneum of lamb fetuses at 65 days of gestation (before the development of the immune system) engrafted and underwent site-specific differentiation into chondrocytes, adipocytes, myocytes, cardiomyocytes, BM stromal cells, and stromal cells of the thymus. Surprisingly, when this transplantation took place at 85 days of gestation (an age at which there is active hematopoiesis and a competent immune system), hMSCs also integrated in a manner similar to cells transplanted at 65 days of gestation. These results suggest that systemically administered hMSCs are widely distributed to many tissues and organs, and that within these organs, specific signals and factors induce tissue-specific differentiation of MSCs. In local models, as opposed to systemic, hMSCs can induce bone formation in vivo following transplantation in ectopic sites and in sites of segmental bone defect (Bruder et al., 1998a; Mankani et al., 2001). Direct injection of hMSCs into the brain tissue of rats resulted in the cells’ long-term engraftment and subsequent

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migration along pathways similar to those used by neural stem cells (Azizi et al., 1998). The results of these studies demonstrate the multilineage differentiation potential of BM-derived adult MSCs and aid in defining them as suitable candidates for the regeneration of several mesenchymal tissues.

SOURCE OF MSCS AND ISOLATION TECHNIQUES The embryonic origin of MSCs is still unclear, and much of our knowledge of these cells lies in the biological characteristics they display in adult tissues. Nevertheless, some findings indicate a possible origin of MSCs in a supporting layer of the dorsal aorta in the aorto-gonadal–mesonephric region in human and murine fetuses (Cortes et al., 1999; Marshall et al., 1999; Tavian et al., 1999). Consistent with these findings, the presence of circulating MSC-like cells within early human blood suggests remnants of migrating MSCs in the circulation (Campagnoli et al., 2001). In adults, MSCs are found in the BM compartments of long bones, iliac crest, sternum, and cranial bones. BM has two major compartments: the hematopoietic compartment, in which hematopoiesis occurs, and the stroma-supportive system, which is associated with the former compartment and is composed of MSCs, ECs, and adipocytes (Bianco et al., 2001). Recent reports and our unpublished data have shown the presence of a potent MSC population in the BM of the craniofacial complex (Matsubara et al., 2005; Akintoye et al., 2006). The lack of a surface marker unique to MSCs poses a further challenge for isolating them as a pure and unmanipulated population. Originally, Friendestein identified “stromal stem cells” by their ability to adhere to standard plastic in the presence of animal serum (Friedenstein et al., 1982, 1987). Pittenger et al. (1999) found that particular lots of fetal bovine serum (FBS) are highly preferable for initial cell adherence and the subsequent survival and proliferation of MSCs isolated from human iliac crest BM (Kadiyala et al., 1997; Pittenger et al., 1999) According to their report, a density gradient should first be used to separate and isolate fractions of mononuclear cells (MNCs) and red blood cells in the BM. The MNCs are then collected and seeded in medium containing 10% FBS at a density of 10–15  105 cells/cm2 growth area. Adherent spindle-shaped cells appear within 48 h after the initial seeding, and the estimated percentage of MNCs ranges from 0.001% to 0.01%. These cells continue to grow, and when they have reached 100% confluence the cells should be detached and replated in fresh culture medium at a density of 5,000–6,000 cells/cm2 growth area. This MSC isolation approach has been broadly followed by many groups. The major disadvantages of using this method are the presence of adherent cells of hematopoietic origin within the cultures during the first days and the need for a specific lot of FBS. Based on the expression of surface molecules on MSCs, some techniques have been developed to isolate MSCs at a higher yield and purity and even without the need to seed them in culture. The expression of endoglin (CD105) by MSCs was used to distinguish these cells and isolate them from other BM cells (Majumdar et al., 2000). CD105-immunoisolated MSCs exhibit the same immunophenotype and differentiation potential described for MSCs that have been isolated using the plastic-adherence method. Using antibodies directed against the CD105 molecule, MNCs can be labeled with microbeads that possess magnetic properties and attach to antiCD105 antibodies. Within the MNCs, CD105 cells become coated with a magnetic shield and can be separated from the rest of the cells by passing them through a magnetic field (Majumdar et al., 2000; Aslan et al., 2006b). Similarly, anti-Stro-1 antibodies were also used to isolate MSCs from BM (Gronthos et al., 1995; Gronthos et al., 2003). Stro-1 is an unidentified cell surface antigen expressed by a minor subpopulation of adult human BM. Anti-Stro-1 antibodies can be used to identify all clonogenic CFU-Fs within the BM, but they do not react to cells of hematopoietic origin (Simmons et al., 1991). Stro-1 cells have been shown to contain an MSC fraction with the capacity to form a supportive microenvironment for hematopoietic cells in vitro and to differentiate into stromal cell types including smooth muscle cells, adipocytes, osteoblasts, and chondrocytes (Gronthos et al., 1994; Dennis et al., 2002).

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The use of anti-CD49a antibodies to isolate hMSCs has also reported to yield a CFU-Fs-forming population that displays MSC characteristics (Deschaseaux et al., 2003).

WHICH TISSUES CONTAIN MSCS? We have already described BM as the original and main source of MSCs. However, many studies have demonstrated the presence of MSCs or MSC-like cells within other tissues such as adipose tissue (Zuk et al., 2001, 2002), cord blood (Hong et al., 2005; Jeong et al., 2005; Moon et al., 2005), BM of the craniofacial complex (Matsubara et al., 2005; Akintoye et al., 2006), and peripheral blood, although the latter finding is still controversial (Fernandez et al., 1997; Conrad et al., 2002). A plastic-adherent, CFU-F-forming cell population can be isolated from adipose tissue after treatment with enzymatic collagenase (Zuk et al., 2001, 2002; Katz et al., 2005). Following this treatment, a stromal vascular fraction is obtained that parallels the MNC fraction in BM. This fraction is collected while the adipocytes-containing fraction is removed during the first steps of centrifugation due to its high content of fatty acids. Plastic-adherent cells within the stromal vascular fraction were originally named processed lipoaspirate (PLA) cells, and were shown to have a high potential for in vitro expansion and a high potential for differentiation into several mesodermal lineages including the adipogenic, chondrogenic, myogenic, and osteogenic lineages (Zuk et al., 2001, 2002). PLA cells are quite similar to BM-derived MSCs morphologically and immunophenotypically; however, PLA cells form more CFU-Fs when plated in culture (Kern et al., 2006). Because adipose tissue is usually more available, can be collected with the use of local anesthesia, and its aspiration is associated with minimal discomfort and risks, it has been proposed as an additional or even alternative source for obtaining MSCs for regenerative medical purposes (Mizuno et al., 2003). Cord blood is a source of MSCs that has been viewed with growing interest. MSCs have been isolated from umbilical cord blood (Hong et al., 2005; Hutson et al., 2005; Jeong et al., 2005; Moon et al., 2005) following gradient centrifugation in a manner similar to that used to obtain them from BM. The success rate of MSC isolation from umbilical cord blood is less than 100% (34% Wagner et al., 2005 and 63% Kern et al., 2006) compared with the 100% rate found in using BM or adipose tissue. Other sources of MSCs include maxillofacial BM (Matsubara et al., 2005; Akintoye et al., 2006), and dermal tissue (Bartsch et al., 2005). Recent reports have shown isolation of MSCs from BM of craniofacial bones (craniofacial MSCs) and compared them to iliac crest and long bones-derived MSCs. Craniofacial MSCs were shown to have highly osteogenic potential and share the main basic characters as iliac crest-derived MSCs (Matsubara et al., 2005). Akintoye et al. (2006) compared maxillofacial- to iliac crest-derived MSCs from the same individuals and reported higher osteogenic and adipogenic potential of maxillofacial MSCs. Our unpublished data have shown that hMSCs isolated from maxillofacial BM can be genetically engineered using adenoviral vectors and utilized for inducing bone formation. MSCs appear to be “resident” stem cells in many tissues, and they function in the normal turnover of these tissues. When tissue repair is required, these cells can be stimulated to proliferate and differentiate. The use of MSCs for appropriate tissue healing may require isolation of the right stem cells and directing the differentiation of these cells into the appropriate lineage. SKELETAL TISSUE REGENERATION BY MSCS Bone Bone regeneration is required for a number of orthopedic, neurosurgical, and maxillofacial clinical indications. Spinal fusion, treatment of nonunion bone defects in long bones, and treatment of substantial bone loss due to trauma or osteoporosis are only a few examples. Currently, these conditions are treated by using synthetic

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implants that do not resemble natural bone and tend to fail in the long term. The few biological solutions that exist for restoration of bone loss include autologous bone grafts, which may cause donor-site morbidity (Quarto et al., 2001), and bone-inducing protein-based treatment (BMP-2 therapy, for example), which requires highly expensive megadoses of protein, and they do not always lead to beneficial results. Given that MSCs can differentiate into the osteogenic lineage, they are considered good candidates for tissue-engineered bone replacement. To promote bone regeneration in vivo by using cultured MSCs, it is essential to seed the cells onto a ceramic scaffold, which is usually composed of hydroxyapatite and β-tricalcium phosphate. Without the osteoinductive properties of these ceramic scaffolds, implanted MSCs tend to form a nonspecific connective tissue in bone defects, as we have shown in several studies (Moutsatsos et al., 2001; Turgeman et al., 2001). The potential for MSC-loaded ceramic scaffolds to repair bone defects has been shown in a number of animal models by using MSCs isolated from the BM of different species. Bruder et al. have shown that critically sized defects in dog femora can be filled with bone newly formed from autologous MSCs (Bruder et al., 1998b) and that a similar result can be achieved by placing hMSCs in femoral bone defects in athymic rats (Bruder et al., 1998a). Using a similar approach, Arinzeh et al. (2003) were able to regenerate femoral bone defects in adult dogs by using allogeneic MSCs, without any evidence of an immune response targeted to the tissue-engineered graft. Another animal model in which autologous MSCs have been used to repair large bone defects is sheep (Kon et al., 2000). In this instance, the same approach was used to generate a substantial amount of newly formed bone tissue to create bone fusion between adjacent vertebras, a method also known as spinal fusion. Such a fusion can eliminate the need for metal screws, which nowadays are used for spinal fusion. The validity of this approach was demonstrated in both rabbit and rhesus monkey models, in which implantation of autologous MSCs led to much greater bone formation than other experimental grafts devoid of cells (Cinotti et al., 2004). Following the solid experimental proof of principle, Quarto et al. attempted this tissue-engineering method in the treatment of three human patients who suffered a bone loss of 4–7 cm in long bones (Quarto et al., 2001). Autologous MSCs were isolated and expanded in vitro for each patient. The cells were seeded onto macroporous hydroxyapatite scaffolds, which had been molded into the shape of the missing piece of the bone, and were implanted in the defect. Two months after implantation, a good integration of implant to bone was evident. Although the patients recovered function in 6–7 months after surgery (one-half to one-third of the time needed for recovery using bone grafts (Quarto et al., 2001) and did not report any problems during a 6-year follow-up period, the ceramic scaffolds were still not absorbed after 5 years (Mastrogiacomo et al., 2005). Other approaches to MSC-aided bone repair include the use of MSCs that have been osteogenically differentiated in vitro prior to implantation in vivo. This strategy allows the seeding of cells onto nonosteoinductive scaffolds, which degrade better in vivo. However, this method requires prolonged periods of culture. Because MSCs are relatively easily isolated from BM and fat tissue, it is conceivable to use them as vehicles for the delivery of therapeutic genes in vivo, a strategy known as stem cell-based gene therapy (Gazit et al., 1999). The aim of most gene therapy studies directed at bone healing is to induce bone formation either in a model of nonunion bone fractures or as a means to achieve spinal fusion. Indeed, some studies have involved the use of primary MSCs and cell lines for expression and delivery of osteogenic genes, which induce bone formation (Engstrand et al., 2000). These studies have implemented various types of MSCs including cell lines such as C3H10T1/2 and primary marrow-derived stem cells for the delivery of BMP-2. The delivery of growth factors of the BMP family is often used in these studies, because these factors promote osteogenic differentiation and bone formation (Wozney et al., 1988). In particular, BMP-2 has commonly been used because it is a highly osteoinductive agent that has been well studied and is known to induce bone in vivo in ectopic and orthotopic sites (Wozney et al., 1988; Wang et al., 1990; Volek-Smith et al., 1996; Yamaguchi et al., 1996; Chaudhari et al., 1997; Hanada et al., 1997; Lecanda et al., 1997; Fromigue et al., 1998; Gori et al., 1999). Other members of the BMP family, such as BMP-4 and BMP-9, have also been used for stem cell-mediated

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gene therapy (Chen et al., 2002; Dumont et al., 2002; Gysin et al., 2002; Peng et al., 2002; Wright et al., 2002). The hypothesis of these studies was that healing of bone defects could be achieved by long-term production of osteoinductive agents in the vicinity of bone defects, inducing new bone formation and defect repair. BM-derived MSCs are good candidates for gene therapy directed at bone regeneration, not only because of their accessibility but also because they form the source stem cells for osteoprogenitors and osteoblasts (boneforming cells) in the bone environment (Prockop 1997). It has been hypothesized that genetically engineered MSCs may have a particular advantage (Gazit et al., 1999). When these cells are engineered to express osteogenic growth factors such as BMP-2, on transplantation in vivo the expressed transgene exerts its effect not only on host mesenchymal tissue (paracrine effect) but on the engineered MSCs as well (autocrine effect). Thus, engrafted, engineered MSCs differentiate and contribute to the bone formation process and, in parallel, recruit and induce osteogenic differentiation in other host stem cells. It has been hypothesized that the combined autocrine and paracrine effects of MSCs may promote bone formation to a larger extent than the mere paracrine effect of other cell types. The murine C3H10T1/2 MSC line, which was engineered to express BMP-2, has displayed a greater osteogenic potential than the non-MSC engineered CHO cell line, which also expresses BMP-2 (Gazit et al., 1999). Engineered MSCs have displayed the ability to heal murine nonunion radial defects better than nonosteogenic CHO cells, despite the fact that CHO cells secrete greater quantities of BMP-2 protein than engineered MSCs. Using MSCs as vehicles for gene delivery has an additional advantage over direct in vivo delivery of proteins or genes. Engineered MSCs can potentially engraft into damaged tissue in vivo and express therapeutic genes for long periods, whereas local, one-time administration of genes or protein has a limited time effect. BMP family members are known for their ability to induce bone formation in vivo and repair bone defects when applied locally in injury sites (Valentin-Opran et al., 2002; Yoon et al., 2002). To compare the efficiency of stem cell-based gene therapy with BMP-2 protein delivery, we analyzed the amount of bone tissue produced by an engineered MSC line (C3H10T1/2) expressing BMP-2 and compared it with the extent of tissue repair following a local administration of a high dose of BMP-2 in a murine model of a radial nonunion defect (Moutsatsos et al., 2001). In that study we have found that engineered MSCs produced significantly more bone tissue than that found following local administration of BMP-2 protein. In addition, we were able to show that using an inducible promoter one can exogenously regulate bone formation in vivo. The BMP-2 gene expression in this study was controlled by a tet-off system, in which the addition of tetracycline, or its analog, doxcycline, to the mice drinking water, inhibited the transgene expression. This method of gene regulation was also shown to be efficient in controlling the extent of bone formation in a posterior spinal fusion model, in vivo (Hasharoni et al., 2005). MSC- or osteoprogenitor cell-mediated gene therapy holds yet another advantage over protein delivery and other types of gene delivery. When the healing process in bone defects was analyzed following transplantation of MSCs engineered to express rhBMP-2, an interesting pattern was observed. Engineered MSCs produced bone in an organized manner by augmenting new bone formation on top of the defect edges, creating continuous regeneration between the original defect edges and the newly formed bone (Gazit et al., 1999). In comparison, BMP-2 protein delivery or the implantation of non-MSC CHO cells, which express BMP-2, resulted in the formation of diffused bone foci with no continuity to the original bone (Gazit et al., 1999). This phenomenon can be attributed to the ability of MSCs to localize and orient themselves to particular sites in the defect area following their transplantation. It was found that MSCs mainly localize to surround the defect edges rather than migrate randomly around the defect site (Gazit et al., 1999). Apparently, as stem cells, MSCs can respond to local factors and developmental signals that direct and guide their orientation within the transplantation site and affect the healing process in a manner similar to the process that takes place during bone formation in developmental stages. Liechty et al. demonstrated that hMSCs possess these characteristics by showing in sheep that these cells are able to engraft in various fetal mesenchymal tissues following

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systemic administration in utero (Liechty et al., 2000). Moreover, hMSCs are able to localize to the osteoprogenitor layers of calvarial bone in severe combined immunodeficiency (SCID) mice when transplanted subcutaneously adjacent to the calvaria (Oreffo et al., 2001). Human BM-derived MSCs are expected to have the same regenerative benefits described earlier for murine MSCs. Nevertheless, as we previously stated, these cells require the right cue to form bone in vivo. hMSCs infected with an adenoviral vector encoding hBMP-2 are able to differentiate into osteogenic cells, both in vitro and in vivo, forming cartilage and bone tissues and healing nonunion defects created in CD-1 nude mice (Turgeman et al., 2001). hMSCs infected with an adenoviral vector encoding the LacZ reporter gene have been shown to be unable to form bone or cartilage in vivo. Consequently, genetic engineering of hMSCs may elicit the osteogenic potential of MSCs, regardless of carrier type (Laurencin et al., 2001). Recently, a nonviral gene delivery approach was used to repeat these results using hMSCs. In this study, MSCs isolated from human BM were transfected with BMP-2 or BMP-9 genes by using a physical method of gene delivery known as nucleofection. In this system, the gene is introduced into the cells by applying an electric field that leads to small pores in the membrane that are to be opened. hMSCs that were transfected in this way demonstrated osteogenic differentiation both in vitro and in vivo (Aslan et al., 2006a). One can safely assume that in large bone defects, nonengineered hMSCs cannot induce repair as efficiently as genetically engineered cells. Bone tissue induced by genetically engineered MSCs has so far been analyzed using X-ray-based systems such as micro-computed tomography or by molecular analyses of gene and protein expression (Moutsatsos et al., 2001; Turgeman et al., 2001). To date there is no knowledge of the biomechanical properties of new bone tissue regenerated using this method. Recently, we have investigated the ultrastructural, chemical, and nanobiomechanical properties of ectopic bone derived from BMP-2-expressing MSCs (Pelled et al., 2006). In this study an engineered bone was analyzed using atomic force microscopy, scanning electron microscopy, and nanoindentation technologies. The engineered bone was compared with native femoral bone adjacent to the implantation site. Interestingly, the engineered bone was found to be similar in its ultrastructural and chemical composition to the native bone, but its hardness and modulus values were lower. When MSCs engineered in the same manner were implanted in a radius bone defect for a longer period of time, however, the hardness and modulus values were strikingly similar to those of the intact contralateral radius (unpublished data). Genetically engineered MSCs can also be used to find novel candidate therapeutic genes for bone repair. We have implanted MSCs expressing the BMP-2 gene under tet-off regulation in an ectopic site in vivo. RNA from the implantation site was purified at different time points during bone formation. Implants in which tetracycline inhibited the expression of BMP-2 transgene served as controls. Gene array followed by a clustering analysis generated a large database of genes playing a major role in the osteogenesis that was induced by the genetically modified MSCs. One important gene that was found was a Wnt inhibitor whose overexpression in BMP-2-expressing MSCs led to a significant reduction in osteogenesis (Aslan et al., 2003). In this manner, candidate transgenes can be found, and their overexpression in MSCs could enhance or inhibit bone formation as needed in a specific pathological condition. The aforementioned studies demonstrate the unique features of MSCs that grant them an additional advantage for the use in bone gene therapy and gene delivery. These stem cells can serve as “smart” vehicles that express the transgene in specific areas of damaged tissue and also can actively participate in the process of new tissue formation. Cartilage Regeneration of damaged cartilage presents a great challenge for orthopedic medicine, because articular cartilage has a very limited capacity for effective repair. The primary therapeutic approaches used nowadays include the surgical procedures of cartilage debridement and drilling, as well as prosthetic implants and autologous cell

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transplantation. Unfortunately, these solutions bring only short-term relief and fail in the long term. Adult MSCs have the potential to proliferate and differentiate into chondrocytes; they can therefore be considered ideal candidates for cartilage tissue repair. Several attempts have been made to implant cells in cartilage defects. The first attempt was to culture autologous chondrocytes and implant them in a cartilage defect in patients younger than 50 years of age who were believed to have healthy chondrocytes (Brittberg et al., 1994). It appeared, however, that chondrocytes can only achieve limited success in regenerating cartilage defects (Liu et al., 2002). It was also shown that chondrocytes loaded onto a polymeric carrier underwent apoptosis, which limited their therapeutic potential (Gille et al., 2002). These results prompted research into autologous pluripotent cells with chondrocyte-differentiating capacities (Caplan et al., 1997). Evidence that MSCs can produce cartilage regeneration has been controversial. Findings of some studies indicate that MSCs fail to produce full regeneration over long time periods (Tatebe et al., 2005). MSCs have also been found to have limited success in forming long-lasting cartilage tissue (Wakitani et al., 2002a, b) Other studies, in which sheep and rabbit models were used, have demonstrated the feasibility of using biodegradable scaffolds seeded with MSCs for articular cartilage repair (Im et al., 2001). In quite a few studies, researchers have investigated the use of different polymeric scaffolds for the growth of cartilage in vitro by using cultured MSCs (Wang, Y. et al., 2005). The feasibility of producing tissueengineered cartilage in this manner has been demonstrated; however, additional studies should be pursued to determine what type of scaffold is optimal for this tissue-engineering approach. Genetically engineered MSCs have also been used in an attempt at cartilage formation; however, only a few genes have been shown to induce chondrogenic differentiation in these cells. Kawamura (2005) and Palmer et al. (2005) and their associates have shown that when infected with adeno-TGFβ, but not with adeno-IGF-1, MSCs differentiated into chondrocytes in vitro. We were the first to show that the transfection of a transcription factor called Brachyury into MSCs could lead to chondrogenic differentiation in vitro and in vivo (Hoffmann et al., 2002). In this study we have utilized the MSC line, C3H10T1/2, which had been shown previously to have a similar differentiation potential to BM-derived MSCs that was stably transfected with the Brachyury transcription factor expressed the chondrogenic marker collagen II but not collagen X, a marker of hypertrophic cartilage. Moreover, the implantation of these cells in ectopic sites in vivo has led to the formation of a chondrogenic tissue composed of proliferative chondrocytes. To the best of our knowledge, this is the only study which has demonstrated an in vivo cartilage formation using genetically modified MSCs. Tendon Although they do not often occur (Hoffmann et al., 2006), tendon and ligament lesions (especially rotator cuff, Achilles tendon, and patellar tendon defects) are among the most common soft-tissue injuries (Juncosa-Melvin et al., 2005). Repairing these defects is not a simple task, and indeed the surgical treatments that are available (those in which autografts, allografts, xenografts, and/or biomaterials are used) are not satisfactory (Wang, Q.W. et al., 2005). Tissue-engineering approaches are being investigated as a means of treating this type of injury. The in vitro differentiation of MSCs into tendon or ligament cells has only been shown in a few studies and has not been induced by supplements added to growth medium, as indicated for chondrogenic, osteogenic, and adipogenic differentiation. Instead, tenogenic differentiation has been induced either by application of exogenous forces on the scaffold on which the cells are grown (Altman et al., 2002) or by the use of a specific scaffold made of hyaluronic acid, which induces ligament differentiation in hMSCs (Cristino et al., 2005). There is no evidence that MSCs that have differentiated in vitro into tendon or ligament cells can indeed repair those tissues in vivo. One possible treatment for in vivo tendon repair involves the implantation of nondifferentiated MSCs that have been seeded onto various biodegradable scaffolds. From investigations of most animal models to date, a surgically induced defect in the rabbit patellar tendon has arisen to become one of the popular models for tendon regeneration. So far it has been shown that the implantation of autologous MSCs in such defects

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improves the physical properties of the damaged tendon when compared with tendons treated only with hydrogel, scaffold, or sutures (Juncosa-Melvin et al., 2005). Dressler et al. (2005) have also observed that MSCs obtained from older animals are able to induce tendon repair in young ones. One adverse effect discovered in some of these studies, however, was the formation of ectopic bone within tendons implanted with MSCs (Harris et al., 2004). Awad et al. (1999) have also posited that there is no morphometric difference between tendons implanted with MSCs and ones implanted with collagen gel. In a recently published paper, we described the use of genetically engineered MSCs to generate tenocytelike cells in vitro and regenerate a rat Achilles tendon defect in vivo. C3H10T1/2 MSC line coexpressing BMP-2 and the Smad8 signaling molecule differentiated in vitro into tendon-like cells, as confirmed by analyzing gene expression and describing the morphological characteristics of the cells. These cells were either implanted ectopically or seeded onto a collagen sponge, creating a construct that was implanted into a 3-mm defect in a rat’s Achilles tendon defect. In both cases, tendon-like tissue was created. Moreover, double-quantum filtered magnetic resonance (MR) imaging was used to determine regeneration in the site of the tendon (Hoffmann et al., 2006). This is the only study so far that has utilized genetically modified MSCs in order to regenerate tendon tissue and it could hold great promise for the repair of cartilaginous defects in therapeutic applications like osteoarthritis and plastic surgery. Intervertebral Disk Regeneration of an intervertebral disk (IVD) poses great challenges due to the hostile environment in which implanted cells must survive. The IVD is avascular and hypoxic; in the rabbit IVD, the nearest blood vessel can be 5–8 mm away from cells at the disk center (Gan et al., 2003). The disk’s main source of nutrition lies in its end plates, which become calcified as the disk grows. As a result, disk cells (mainly nucleus pulposus (NP) cells) use anaerobic metabolism to generate energy (Gan et al., 2003; Roughley, 2004). Due to the avascular nature of this tissue, lactic acid (the main waste product of glycolysis) can accumulate, resulting in a low pH environment (Roughley, 2004). When attempts are made to regenerate an IVD, two strategies can be taken. The first, which is indicated for early disk degeneration (when only the NP is degenerated), is to regenerate only the NP. The injectable technique is very appealing, because it eliminates the need for surgical intervention; however, few experiments pursuing this route have been performed. Compared to the injection of cell suspension alone, the injection of cells suspended in hydrogel into “nucleotomized” disks provides an abundant source of cells because of improved cell survival and the location within the NP (Bertram et al., 2005). In another work, DiI-labeled rat MSCs embedded in 15% hyaluronan gel were injected into a rat-tail IVD. Good cell viability was recorded after 24 h. A decrease in the number of cells was noted after 14 days, but cell viability returned to 100% 28 days postinjection. Compared to IVDs treated with injections of blank gels, IVDs treated with injections of cellularized gel had greater heights, a finding suggestive of matrix production in the injected disk (Crevensten et al., 2004). This trend toward increased cell viability and function following transplantation was repeated when green fluorescent protein (GFP)-labeled autologous rabbit MSCs immersed in atelocollagen were injected into the rabbit lumbar NP. Forty-eight weeks after implantation, a significant increase in GFP-positive cells was noted in the NP. Moreover, some of the GFP-labeled cells expressed NP marker genes and typical NP proteins, a finding suggestive of the differentiation of the implanted MSCs. In addition, an examination of gene expression in, and a biochemical analysis of, the engineered NP tissue demonstrated that tissue function was restored to some extent (Sakai et al., 2005). An evaluation of this therapeutic avenue was performed using MR imaging and plain radiography, and the findings showed 91% disk height and 81% MR imaging signal intensity compared with untreated controls 24 weeks after injections of autologous MSCs into rabbit lumbar IVDs (Sakai et al., 2006). Those results indicate the good clinical potential of this method. Nevertheless, a comprehensive biomechanical comparison between native and engineered tissues should

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be performed to evaluate the ability of this approach to generate functional NP tissue. Genetically modified MSCs have not been utilized for IVD regeneration, yet. In a preliminary study performed in our laboratory, we have been able to generate an IVD-like tissue using two types of genetically engineered MSCs. Since an IVD is composed of a tendon-like tissue on the outer portion and a cartilage-like tissue in the center, a hybrid of cells and scaffolds should be utilized in order to regenerate it. Therefore we have seeded a biodegradable scaffold, shaped as a ring, with Smad8/BMP-2- expressing MSCs, which form tendon tissue in vivo (Hoffmann et al., 2006). The midcompartment of the ring-shaped scaffold was filled with fibrin gel containing Brachyury-expressing MSCs, which form cartilage in vivo (Hoffmann et al., 2002). Following in vivo implantation in ectopic and inter vertebral sites, an IVD-like tissue was formed demonstrating similar molecular and morphological features of a native IVD (Kimelman et al., 2006). This approach could serve as a biological solution for the replacement of degenerative IVD in the clinic. The use of MSCs for skeletal tissue repair raises several questions regarding cell survival, differentiation, and biodistribution in vivo. The use of noninvasive imaging methods is mandatory in order to answer these questions quantitatively in real time. For example, Bar et al. used the bioluminescence imaging (BLI) system (described by Honigman et al., 2001) with transgenic mice that express the luciferase gene under the human osteocalcin promoter (Iris et al., 2003). Using this system, osteogenesis, indicated by the expression of osteocalcin, is correlated with the luciferase signal. In this way, the extent of osteogenesis following the implantation of osteogenic cells, based on the intensity of the luciferase signal, could be analyzed (Hasharoni et al., 2005). The system allows to perform longitudinal studies without the need to sacrifice animals at different time points. Moreover, the BLI system can noninvasively, quantitatively, and longitudinally monitor the survival or biodistribution of luciferase-labeled MSCs in vivo. Additional imaging systems that can be applied to detect MSCs in vivo include the fibered confocal microscope (Cell Vizio™) that can detect fluorescently labeled MSCs in high resolution on a single cell level (Aslan et al., 2006b). If implanted subcutaneously, the survival of fluorescently labeled MSCs can be followed by a noninvasive imaging system (Aslan et al., 2006b) as well.

IMMUNOMODULATORY EFFECTS OF MSCS A small but increasing number of preclinical and clinical studies have been performed in which the use of MSCs resulted in alloantigen tolerance. In a pilot study, Horwitz and colleagues concluded that improvements in bone structure and function following allogeneic BM transplantation in children with severe osteogenesis imperfecta can lead to objective clinical benefits (Horwitz et al., 2001). In patients with Hurler syndrome (mucopolysaccharidosis type IH) and in those with metachromatic leukodystrophy (MLD), the clinical manifestations of the disease were partly corrected after transplantation of allogeneic HSCs. Koc et al. have postulated, however, that some of these defects may be corrected by infusion of allogeneic, multipotential, BM-derived MSCs. In their trial, MSCs, isolated and expanded from a BM aspirate, were infused and no infusion-related toxicity was observed. The overall conclusions of that study were that donor-allogeneic MSC infusion is safe and may be associated with reversal of disease in some tissues, but the role of MSCs in the management of Hurler syndrome and MLD remains unclear (Koc et al., 2002). A preclinical study was performed in baboons by Bartholomew and coworkers, aimed at elucidating whether the BM microenvironment confers on MSCs the capability of immunomodulation of lymphocytes. Results showed that MSCs failed to elicit a proliferative response from allogeneic lymphocytes when added to a mixed lymphocyte reaction or to mitogen-stimulated lymphocytes. In vivo administration of MSCs led to prolonged survival of skin grafts when compared with control animals (Bartholomew et al., 2002). MacDonald and associates have demonstrated that xenogeneic murine MSCs implanted immediately after myocardial infarction in immunocompetent adult rats survived, differentiated, and were immunologically tolerated; and that their presence led to a recovery in left ventricular function (MacDonald et al., 2005). On the

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contrary, results obtained by Eliopoulos et al. strongly suggest that MSCs are not intrinsically immunoprivileged and cannot serve as a “universal donor” in immunocompetent, major histocompatibility complex (MHC)mismatched recipients. Murine MSCs were engineered to release erythropoietin and were implanted in MHCmismatched allogeneic mice without any sign of immunosuppression. In syngeneic recipients, the hematocrit rapidly rose from baseline level and remained higher than 0.88 (88%) for longer than 200 days. However, in MHC-mismatched recipient Balb/c mice, the hematocrit rose transiently and rapidly declined to baseline values (Eliopoulos et al., 2005). Nevertheless, a remarkable clinical response was achieved in a case reported by Le Blanc et al. when haploidentical MSCs were transplanted into a patient suffering from a severe treatment-resistant Grade IV acute graft-versus-host disease of the gut and liver (Le Blanc et al., 2004). Later, this group and additional colleagues treated a female fetus with multiple intrauterine fractures (diagnosed as severe osteogenesis imperfecta) by transplantation with allogeneic human leukocyte antigen (HLA)-mismatched male fetal MSCs in the 32nd week of gestation. Coculture experiments performed in vitro after MSC injection did not show any patient lymphocyte proliferation against the donor MSCs. These investigators concluded that allogeneic fetal MSCs can engraft and differentiate into bone in a human fetus, even when the recipient is immunocompetent and HLA incompatible (Le Blanc et al., 2005). Numerous in vitro experiments have been performed in an attempt to provide an explanation for the assertion that MSCs inhibit allogeneic responses. Different approaches have included coculture of MLCs or mitogen stimulations by PHA (phytohemagglutinin) or PMA (phorbol 12-myristate 13-acetate). To date, there are probable mechanisms that may explicate why MSCs seem to escape allogeneic rejection, such as weak immunogenicity, interference in the maturation and function of DCs, abolishment of T-cell proliferation, or interaction with natural killer (NK) cells in cell-to-cell contact or through the release of soluble secreted factors. Findings of most studies have indicated that MSCs are positive for MHC class I and negative for MHC class II, although there have been discrepancies, probably due to the different experimental systems that have been implemented. However, the majority of reports have indicated no or low expression of MHC class II proteins (Majumdar et al., 2003; Gotherstrom et al., 2004). Evidence for the interference in the maturation of DCs has been provided by our collaborators, Beyth et al. (2005). These researchers demonstrated that, although hMSCs are able to promote antigen-induced activation of purified T-cells, an addition of antigen-presenting cells (APCs) – monocytes or DCs – to cultures inhibited, in a contact-dependent manner, the T-cell responses, and instead large amounts of IL-10 were secreted and the maturation of the APCs was abnormal. This inhibition could be partially overridden by the addition of factors that promote APC maturation. These data have been supported by findings of coculture experiments, in which Zhang et al. (2004) showed that both MSCs and their supernatants interfered with the endocytosis of DCs and decreased their capacity to secrete IL-12 and activate alloreactive T-cells. Similar conclusions have been reported by Aggarwal et al. (2005), who demonstrated in cocultures of hMSCs and DCs decreased tumor necrosis factor secretion in mature type I DCs and increased secretion of IL-10. Several groups support the direct interaction of MSCs and T-cells, either by cell contact or by the release of soluble factors into the culture medium. Rasmusson et al. made the distinction between T-cell stimulation in culture by mitogen and alloantigens. In a recent paper, they stated that MSCs increased the levels of IL-2 and the IL-2-soluble receptor, as well as that of IL-10 in MLCs. None of these factors are constitutively secreted by MSCs, according to Rasmusson et al. and Beyth et al. When peripheral blood lymphocytes were stimulated with PHA, decreases in levels of IL-2 and the IL-2 soluble receptor were observed, whereas IL-10 levels were not affected. Moreover, the addition of a prostaglandin inhibitor, indomethacin, restored the inhibition induced by MSCs in PHA cultures but did not influence MLCs (Rasmusson et al., 2005). Di Nicola and colleagues identified TGFβ1 and HGF as mediators of MSC effects on T-lymphocyte-suppressed proliferation by using neutralizing monoclonal antibodies. They demonstrated that cellular stimuli were effective as well as nonspecific mitogens, and that T-cell inhibition is likely due to the production of soluble factors, as shown by transwell experiments, in which cell-to-cell

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contact between MSCs and effector cells was avoided (Di Nicola et al., 2002). Using a different approach and seeking the interaction between MSCs and NK cells, Sotiropoulou et al. found that MSCs alter the phenotype of NK cells and suppress proliferation and cytokine secretion. Some of these effects were mediated by soluble factors including TGFβ1 and PGE-2 (Sotiropoulou et al., 2006). Other studies differ in findings related to TGFβ1, with investigators reporting no involvement in T-cell inhibition by MSCs (Djouad et al., 2003). The upregulation of PGE-2 in cocultures has been observed by others as well, although the role of PGE-2 in downregulation of MLCs diverged from the one mentioned above, as shown in the studies conducted by Tse et al. (2003) and Rasmusson et al. (2005). The way by which MSC avoid detection by the immune system is not thoroughly elucidated yet, it is expected that additional soluble factors or cells might result of significant impact as well as novel mechanisms might be revealed.

NONSKELETAL TISSUE REGENERATION BY MSCS During the mid-1990s, Okuyama and Wakitani and their colleagues separately presented the first two reports demonstrating the in vitro nonskeletal differentiation potential of MSCs. MSCs differentiated into endodermally, mesodermally, and ectodermally derived cell types such as ECs, adipocytes, and myocytes. This paved the way for further research to establish differentiation protocols for MSCs into nonskeletal progenitor cells and, further down the road, to create nonskeletal tissue regeneration (Okuyama et al., 1995; Wakitani et al., 1995). These first reports were validated and established within the scientific community a few years later by Liechty et al. (2000) and Fukuda et al. (2001, 2002) who stated that multipotent MSCs derived from BM can differentiate into skeletal myocytes and adipocytes after treatment with various inducers as well as following in vivo transplantation. Since then MSCs have been used as regenerators of heart (cardiac muscle and vascular system), skeletal muscle, nerve, liver, and pancreas (Burt et al., 2002; Lardon et al., 2002; Bonafe et al., 2003; Dabeva et al., 2003; Abedin et al., 2004; Kim et al., 2004; Jain et al., 2005; Sonoyama et al., 2005; Goncalves et al., 2006). The leading field in that context has been cardiac tissue regeneration. Cardiomyocytes cease cell division immediately after birth and are thought to adapt subsequently to the demands placed on the heart by undergoing hypertrophy without cell division. Recent research has revealed that, although a small number of cardiomyocytes do undergo cell division immediately after a myocardial infarction, their contribution is not sufficient to improve heart failure (Beltrami et al., 2001; Yuasa et al., 2004). Heart transplantation is traditionally performed to treat intractable severe heart failure secondary to dilated and hypertrophic cardiomyopathy, but its use is restricted by a shortage of donors. The use of pluripotent stem cells to regenerate damaged heart tissue is being advocated as the new treatment for heart failure secondary to heart disease or severe myocardial infarction. Promising results at the research stage have now led to the challenge of applying stem cell technology in the clinical setting (Fukuda 2003a, b; Itescu et al., 2003; Orlic 2003; Amado et al., 2005; Bayes-Genis et al., 2005; Fazel et al., 2005; Fukuda 2005; Jain et al., 2005; Siepe et al., 2005; Smits et al., 2005; Wojakowski et al., 2005; Yamaguchi et al., 2005; Yoon et al., 2005b; Minguell et al., 2006). Makino et al. (1999) generated cardiomyocytes from murine BM MSCs in vitro. The stromal cells were immortalized and treated with 5-azacytidine, which induced the generation of spontaneously beating cells. In addition, hMSCs from adult BM were able to differentiate into cardiomyocytes, when transplanted into the adult murine heart (Toma et al., 2002). One of the major concerns is the poor viability of the transplanted cells. It has been estimated that more than 99% of MSCs die 4 days after transplantation into the hearts of uninjured nude mice (Toma et al., 2002). Rat MSCs, genetically modified to overexpress the prosurvival gene Akt1, prevented remodeling and restored performance in an infarcted heart (Mangi et al., 2003). Nonhematopoietic MSCs (cardiomyogenic cells) expressing enhanced GFP (EGFP) were transplanted into the BM of lethally irradiated mice;

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a myocardial infarction was induced and the cells were treated with granulocyte colony-stimulating factor. The presence of EGF/actinin cells in the ischemic myocardium indicated that the cardiomyogenic cells had mobilized and differentiated into cardiomyocytes. These results suggest that most BM-derived cardiomyocytes originate from MSCs (Kawada et al., 2004). Clonally expanded novel human BM-derived multipotent stem cells (hBMSCs), a subpopulation within BM-derived MSCs, were expanded in vitro and generated cardiomyocytes and cells of all three germ layers in coculture conditions (Yoon et al., 2005a). The transplantation of hBMSCs into an infracted myocardium resulted in engraftment of the transplanted cells, which exhibited colocalization with markers of cardiomyocytes, smooth muscle cells, and ECs. Therefore, the hBMSCs differentiated into multiple lineages. Moreover, the hBMSC-transplanted hearts demonstrated upregulation of paracrine factors including angiogenic cytokines, anti-apoptotic factors, and proliferation of host ECs and cardiomyocytes (Yoon et al., 2005a). Transplantation of MSCs improved cardiac function in animal models of induced cardiac diseases, possibly through induction of myogenesis and angiogenesis as well as by inhibition of myocardial fibrosis. The beneficial effects of MSCs may be mediated not only by their differentiation into cardiomyocytes and vascular cells, but also by their ability to supply large amounts of angiogenic, anti-apoptotic, and mitogenic factors (Nagaya et al., 2005). In 2001 Reyes and associates characterized a subpopulation of MSCs that, at the single-cell level, can differentiate into cells of visceral mesoderm and can be expanded extensively by means of clinically applicable methods (Reyes et al., 2001). These cells were named multipotent adult progenitor cells (MAPCs). These cells were cultured selectively by using growth factor supplements and gave rise to clusters of small adherent cells. The MAPCs differentiated into cells of limb-bud mesoderm as well as visceral mesoderm (ECs). Continuing their research in 2002, Reyes and associates have also presented in vivo results demonstrating the contribution of human MAPC-derived ECs to neoangiogenesis in tumors and wound healing (Jiang et al., 2002; Reyes et al., 2002). Since then MSC differentiation into ECs has been further investigated and culturing and differentiation protocols have been simplified (Oswald et al., 2004). In addition MSC-based ECs have been used as neovascularization vehicles in the murine brain and heart (Davani et al., 2003; Fang et al., 2003; Gojo et al., 2003; Takizawa, 2003; Minamino et al., 2005; Silva et al., 2005). The in vivo injection of MSCs has been shown to promote neuron survival and limit the severity of neurological impairment in animal models of traumatic brain injury (Lu et al., 2001; Mahmood et al., 2003) and induced stroke (Chen et al., 2001; Zhao et al., 2002) as well as to promote recovery of motor function in mice treated with 1-methyl-4-phenyl-1,2,3,6-tetra-hydropyridine (MPTP) hydrochloride (Li et al., 2001). Direct implantation of MSCs into the spinal column has also been shown to promote functional recovery following a standardized contusion injury (Chopp et al., 2000; Hofstetter et al., 2002) and to stimulate remyelination and improve axon conduction velocities within a focal demyelinated lesion (Akiyama et al., 2002). The neuroprotective effects of MSCs are thought to result in part from their ability to replace diseased or damaged neurons via cellular differentiation (Black et al., 2001; Crigler et al., 2006). As the prevalence of diabetes increases (7% of the populations in the USA have diabetes) and with diabetes being ranked as the sixth leading cause of deaths according to US death certificates in 2002, new treatment avenues are being sought, and MSCs have been identified as prime candidates. The endocrine compartment of the pancreas consists of insulin-producing beta-cell islets and three other cell types. An inadequate mass of functional pancreatic beta cells is found in both type 1 and type 2 diabetes. Thus, beta-cell replacement therapy is thought to be a possible curative treatment for diabetes. Achieving the reconstitution of pancreatic beta cells by using BM-derived cells suggests that BM cells are a feasible source for beta-cell replacement therapy. Scientists have been able to obtain islet-like functional cells through differentiation of MSCs from BM by modifying the cell culture environment or by supplanting rat pancreatic extract (RPE) in the culture media (Chen et al., 2004; Choi et al., 2005).

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MSCs can be used for beta-cell replacement therapy by harvesting the MSCs and applying in vitro differentiation protocols before delivering the cell back to the target tissue, or by enhancing biological mechanisms of mobilization and the homing of MSCs followed by biologically induced differentiation. Sordi et al. were able to define the chemokine receptor repertoire of hMSCs derived from BM that determines their migratory activity. Using a pancreatic cell coculture, these researchers concluded that modulation of the homing capacity of MSCs may be instrumental for harnessing the therapeutic potential of MSCs derived from BM (Sordi et al., 2005). Recently it was reported that in vitro human BM stem cells are able to differentiate into insulin-expressing cells through a mechanism involving several transcription factors of the beta-cell developmental pathway when cultured in an appropriate microenvironment (Moriscot et al., 2005). Nevertheless, the insulin-producing capacity of BM-derived cells is still controversial. Recently, Choi et al. suggested that there is little evidence of transdifferentiation of BM-derived cells into pancreatic beta cells in vivo. However, their studies did not exclude the possibility that BM-derived MSCs could differentiate into beta cells in vitro by using the right inducer, for example, RPE, or as recently suggested, that the expression of the Pdx1 gene into various cells can provoke differentiation into cells similar to pancreatic beta cells (Choi et al., 2003). In vitro models of parenchymal liver cells are of great importance in toxicology and in bioartificial liver research (Azar et al., 1996; Locasciulli et al., 1997), because primary cultures of hepatocytes are hindered by a short life span and a rapid loss of hepatic function under in vitro conditions (Kim et al., 2000). Schwartz and associates reported for the first time that under in vitro conditions an adult marrow-derived stem cell, MAPC, can differentiate into functional hepatocyte-like cells (Schwartz et al., 2002) as well as into mesodermal and ectodermal cell lineages (Reyes et al., 2001; Jiang et al., 2002; Verfaillie et al., 2003). Following this study, Lee et al. used MSCs and demonstrated differentiation into cells of the endoderm as well as into those of the mesoderm (Lee et al., 2004). Rat MSCs require specific culture conditions and growth factors to differentiate into hepatocytes. Regarding this, several controversial reports have emerged: some demonstrating that differentiation was achieved only by using fibroblast growth factor-4 and HGF (Wang et al., 2004; Kang et al., 2005) and some showing that rat MSCs must be cultured in supplemented medium and the presence of freshly isolated rat liver cells (Lange et al., 2005b). Under specified culture conditions, only rat MSCs cocultured with liver cells acquired the hepatocytic phenotype. In vivo transplantation of HGF-induced differentiated rat MSCs into liver-injured rats restored serum albumin levels and significantly suppressed transaminase activity and liver fibrosis (Oyagi et al., 2005). The next generation of experiments involved hMSCs, which were examined by directly xenografting them to allylalcohol-treated rat liver. When BM-derived cells were fractioned into MSCs, CD34 cells, and non-MSC CD34- cells, and transplanted in vivo, hepatocyte-like cells were observed only in the recipient livers that contained MSC fractions (Sato et al., 2005). The ultimate goal of differentiation studies is the amendment of damaged tissue by cellular transplantation. The recovery of damaged liver may be clearly attained if one uses a syngeneic model on a larger scale of transplantation (Lange et al., 2005a, b). In summary, BM-derived hMSCs indeed possess great potential as the future treatment of choice for several nonskeletal tissue injuries and diseases.

CONCLUSIONS MSCs constitute a unique population of adult stem cells that hold great promise for various tissue-engineering applications. These cells can readily be isolated from various sites in the human body, especially from BM and adipose tissues. Established protocols exist for the induction of specific differentiation patterns of MSCs into different committed cells, most notably into osteoblasts, chondrocytes, and adipocytes. So far it has been demonstrated

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that the use of genetically modified MSCs, overexpression of various therapeutic transgenes, is a powerful tool in the induction of differentiation and in the promotion of tissue regeneration in vivo. Novel technologies, which utilize electroporation-based systems, allow for the safe and efficient gene delivery into MSCs and bypass the need for using non-safe viral vectors. It has been shown that the ultrastructural, chemical and nanobiomechanical properties of engineered bone derived from MSCs were similar to that of native origin. Bioinformatics techniques can be applied to genetically modified MSCs in order to find new candidate genes for therapeutic purposes. The conventional method of MSC isolation using plastic adherence has shown to be costly and might reduce the stemness of the cells. Therefore an attractive alternative has been developed and it includes the immediate use of immunoisolated, non-cultured MSCs for in vivo implantation. Future challenges require the identification of an optimal scaffold for MSC implantation in vivo and, finally, the development of a preservation method for future reuse of autologous cells. Noninvasive imaging will continue to play an important role in analyzing the power of MSCs to regenerate tissues in various defect models. Overcoming these hurdles will no doubt make MSCs the optimal tool for biological tissue replacement in this century.

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20 Hepatic Stem Cells: Lineage Biology and Pluripotency N. Cheng, Hsin-lei Yao, and Lola M. Reid

INTRODUCTION The development of tissues is synonymous with the evolution of metazoans, organisms having tissues. All metazoans make use of two, dynamic and interacting sets of mechanisms: (1) stem cell and maturational lineage biology and (2) the epithelial–mesenchymal relationship. The successes of current efforts toward tissue engineering are dependent upon strategies employing recognition of these mechanisms. In this review, we will first present an overview of those two fields and then will discuss them as they pertain to liver. At the end we discuss some of the many legal and ethical issues confronting all stem cell biologists. GENERAL ISSUES WITH RESPECT TO STEM CELLS AND MATURATIONAL LINEAGE BIOLOGY Stem Cells and Progenitors Stem cells are the hope for many people suffering from some form of organ or tissue dysfunction. The renowned capacity of stem cells for expansion and for reconstitution of tissue, especially damaged tissue, makes them the “magic bullets” for cell therapies, bioartificial organs, and industrial programs such as protein manufacturing. In the near term, stem cells will possibly alleviate or cure diverse conditions such as bone and cartilage disorders, some forms of liver failure, Parkinson’s disease, some genetic diseases (perhaps diabetes), and may offer plastic surgeons the tools for replacement of skin in burn patients and accident victims, or repair of neurological tissues in quadriplegics. However, the full, dramatic potential of stem cells must await continued research to establish all relevant aspects of the technology and is associated inherently with ethical and legal issues that have become and will continue to be the subject of heated debate in political, religious, and cultural circles. Stem cells are precursor cells forming the basis, the “spring,” for regeneration and renewal of tissues. The stem cells and their immediate descendents are small cells, are readily cryopreserved, and have extensive growth properties when placed into culture dishes or when injected into animals. Recent studies suggest that at least some types of stem cells tolerate ischemia (lack of oxygen) even at warm (body or room) temperatures (Smith, 2006). This tolerance for ischemia and other adverse conditions makes possible the use of tissues from asystolic donors. This means that there should be a ready supply of tissues for the harvesting of at least some types of stem cells. According to the strict definition of stem cells, they:

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are pluripotent and can give rise to multiple types of adult cells; have extensive growth potential, indeed self-replication capacity enabling them to produce daughter cells identical to themselves; this clonogenic expansion potential is determined by seeding a single cell into a dish and

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demonstrating that it yield daughter cells that can expand indefinitely and be subcultured and all of the daughter cells can be induced to differentiated to adult fates; can mature into adult cells and can reconstitute damaged tissue when injected in vivo.

The issue of self-replication is being reconsidered at present, since stem cells found in adult tissue have been found to change in subtle ways throughout the life of the host. Consequently, rigorous proof of self-replication has been achieved for only stem cells from early embryos. See the published glossary on stem cells by Smith (2006). Categories of Stem Cells Stem Cells Found Exclusively in Embryos Totipotent stem cells. Totipotent stem cells have the capacity to produce all adult cell types, can enter the germ line (i.e. contribute genetic material to succeeding generations), and have proven ability to self-replicate (i.e. produce daughter cells that are identical to the parent). The zygote or fertilized egg is, of course, a totipotent stem cell. The known and well characterized totipotent stem cells are found only in early embryonic tissues and derive usually from the first few cell divisions after fertilization. Totipotent stem cells can be derived from the fertilized eggs from in vitro fertilization (IVF), a procedure in which sperm and eggs (ova) collected by laparoscopic procedures are placed into culture dishes and permitted to undergo fertilization. The resulting fertilized eggs can be implanted into the uterus (in utero) to generate a pregnancy (Brill et al., 1994). The unused fertilized eggs can be stored in cryopreserved form in liquid nitrogen indefinitely. Couples undergoing IVF may have many cryopreserved fertilized eggs that can be used for future pregnancies or can be discarded. Totipotent stem cells are able to go through all of the stages of development in a normal way to form an animal (or human) only when implanted in utero, the site at which the countless signals and conditions occur with the correct timing and in the correct quantitation. When totipotent stem cells are implanted at ectopic sites (sites other than in utero), the cells can differentiate to many types of tissue. Differentiation occurs in a disorganized way resulting in teratomas or teratocarcinomas (“monster” tumors) that have a jumble of teeth, hair, bodily organs, etc. The frequency of tumor formation is nearly 100% (Brinster et al., 1989). Embryonic stem cells. Embryonic stem (ES) cells are pluripotent; ES cells can give rise to mature cells derived from all the germ layers but cannot give rise to amnion or placenta. Therefore, strictly speaking, they are not totipotent stem cells. Culture conditions have been identified in which ES cells maintain their undifferentiated state and can be expanded indefinitely. Yet every cell in the dish retains the capacity to produce an entire animal (or theoretically a human). The findings with the existing human ES cell cultures are ones derived from discarded fertilized eggs from IVF procedures and have been obtained with permission from the donor. A critical requirement for maintenance of ES cells in culture as undifferentiated cells are embryonic mesenchymal feeder cells that supply unidentified signals vital to the ES cells (Yin et al., 2002). This requirement involves a risk for clinical programs: particular feeder cells might harbor a virus or some other pathogen that could get into the ES cells, affecting the ability to use the ES cells clinically. Indeed, the Food and Drug Administration (FDA) is requiring that any ES cells that might one day be used in cell therapies must be grown in the absence of such feeder cells, a demand that is not possible for most ES cell cultures. There is considerable effort ongoing by many investigators to define the soluble (e.g. growth factors) and insoluble (extracellular matrix) components relevant to their ability to support stem cells with the hopes that completely defined model systems can be developed (Thorgeirsson et al., 2004). The term ES cell has been used also to mean cells capable of entering into the germ line; that is, the resulting animal demonstrates ES cell genetic material in the germ cells in the gonads of the animal derived from

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the ES cells. However, this criteria for ES cells, used routinely for non-human ES cells, cannot be tested for human ES cell cultures. Therefore, we accept the definition for human ES cells, provided by Thomson et al. (1998) that they are:

• • •

derived from early embryos (pre-implantation or peri-implantation embryos); have prolonged proliferation as undifferentiated cells; have stable developmental potential to form all known adult cell types (i.e. derivatives of all three embryonic germ layers even after culture).

When ES cells are put into culture and allowed to differentiate, the differentiation process occurs spontaneously and without the ability to be controlled in a precise way (Fraser et al., 1992). Current research has yielded some information leading to improved ability to control differentiation toward particular fates, referred to as “lineage restriction.” However, thus far, this cannot be done to yield specific fates with complete fidelity (Sicklick et al., 2006). Thus, there are conditions identified that result in lineage restriction to blood cells or neuronal cells (just two examples of those known) but always with a small percentage of the cells that do not achieve the full commitment to the designated fate. Injection of lineage-restricted ES cells into ectopic sites results in the reconstitution of damaged tissues but with a significant risk (on the order of 5%) of tumor formation. Therefore, ES cells, even those that are lineage restricted, cannot be considered for clinical use at this time. However, lineage restriction of ES cells is an area of research under intensive investigation at present and, perhaps one day, will achieve the fidelity in lineage restriction required for these cells to be used clinically. Their real potential currently is for research and for industrial programs (e.g. protein manufacturing) in which their extraordinary expansion potential and ease of cryopreservation are major assets; in such industrial uses the inability to achieve complete fidelity in lineage restriction to a given fate is not an issue. Further discussions of ES cells will not be addressed, given that numerous excellent reviews have been published in recent years (Potten and Wilson, 2004; Thorgeirsson et al., 2004; Sicklick et al., 2006). Multipotent Stem Cells

Bone marrow. Bone marrow is a well established tissue source used routinely for reconstitution of hemopoietic dysfunctions. There has been a dramatic discovery within the last few years that bone marrow transplants result in donor cells contributing to many types of tissues, including ones of ectodermal (neuronal), mesodermal (heart), or endodermal (liver) fates (Petersen et al., 1999; Theise et al., 2000). The phenomenon has been called “transdifferentiation” and has been touted as evidence of considerable plasticity in stem cells. However, analyses of transdifferentiation have demonstrated that it is due primarily to cell fusion (Wang et al., 2003; Lucas and Terada, 2004). Yet there remain findings supporting the presence of multipotent stem cells in the bone marrow and capable of giving rise to cell types of all the germ layers (Jian et al., 2002). Unfortunately, bone marrow contains such small numbers of these multipotent adult progenitor cells (MAPC), that bone marrow transplants result in exceedingly low efficacy (1% or less) with respect to reconstitution of damaged tissues (Overturf et al., 1997). Although the transdifferentiation issue remains an area of ongoing controversy and research, the general consensus is that it is a minor pathway with little hope for clinical programs. Yet, the clinical effectiveness of bone marrow transplants for damaged tissues (ones independent of hemopoietic fates) supercedes the direct evidence of transdifferentiation of cells to the desired cell type. Increasingly, there is the assumption that the bone marrow-derived cells are contributing to the restoration of the tissue by paracrine signaling mechanisms. Therefore, more research is needed to assess the requirements or limitations of bone marrow as a source of cells capable of reconstituting damaged tissue. Umbilical cord and adipocyte-derived stem cells. In vitro propagation of umbilical cord blood (UCB)derived or bone marrow-derived mesenchymal stem cells (MSCs) (Lee and Kuo, 2004), and adipose-derived stem cells (ASCs) (Seo et al., 2005) in medium containing hepatic growth factor (HGF) and oncostatin

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M yield cell populations that express a number of hepatic characteristics. The detection of mRNA transcripts for α-fetoprotein, CK18, and albumin by RT-PCR (reverse transcriptase polymerase chain reaction), as well as the detection of cells that store glycogen and secrete urea, suggest that each of these stem cell populations can differentiate into hepatocytes in vitro. However, the more stringent test for differentiation is whether they differentiate into hepatocytes when implanted into damaged livers. The data are clear that the implanted ASCs incorporate into the host liver tissue (Seo et al., 2005); this can be interpreted as evidence that the stem cells indeed have differentiated into hepatocytes in vivo. However, the mechanism of engraftment was not determined. This is significant because current data have not shown definitively that transplanted stem cells differentiate in vivo into the cell types under investigation. The mechanism of incorporation of implanted stem cells into the liver has been extensively examined only for hemopoietic stem cells. In nearly every case, repair was not mediated by hepatic differentiation of hemopoietic stem cells, but by fusion with host hepatocytes (Camargo et al., 2004; Kashofer et al., 2005; Sharma et al., 2005). It appears that hemopoietic stem cells differentiate in situ into cells of the macrophage–monocyte lineage, which exhibit a high capacity for cell fusion (Willenbring et al., 2004; Thorgeirsson and Grisham, 2006). Whether UCB- and bone marrow-derived MSCs, or ASCs also fuse with the host tissue has not yet been examined in detail. Determined Stem Cells

Determined stem cells are pluripotent cells that give rise to some (but not all) possible adult cell types, have extensive growth potential including clonogenic expansion potential, and are easily cryopreserved. They are able to reconstitute damaged tissues when injected in vivo. The lay press refers to them as “adult stem cells,” an inaccurate term, since they are found in tissues from both embryos and adults. The most well studied determined stem cells are hemopoietic stem cells, epidermal stem cells, MSCs, and neuronal stem cells. In the past, determined stem cells were assumed to self-replicate (Potten and Wilson, 2004). However, in recent years even the most well studied of the determined stem cells, the hemopoietic stem cells, are thought to change subtly and slowly over the life of the host and, therefore, are questionable in their ability to selfreplicate, in the most rigorous sense of the term. Bone marrow-derived hemopoietic stem cells from elderly donors have less renewal capacity than those from infants. This finding will be a driving force to obtain determined stem cells from as young a donor as possible. The determined stem cells are the real hope for cell therapies in the near term. They are known already to have a profound capacity to correct organ and tissue dysfunction and yet are non-tumorigenic. For example, bone marrow transplants, the original form of cell therapy with determined stem cells, have been done since the 1950s in clinical therapies and yet have no evidence for tumorigenic potential. Committed Progenitors

Committed progenitors are immature cells, precursors that are unipotent and yet have considerable expansion potential. These include the “transit amplifying cells” of the skin. They may prove just as useful as the determined stem cells except where particularly extensive growth potential is required; the more limited growth potential of committed progenitors versus their parent stem cells will put some constraints on their usefulness in clinical or commercial programs. The number of rounds of division possible for committed progenitors differs from tissue to tissue. Liver, however, is representative of quiescent tissues and has committed progenitors able to undergo 5–7 rounds of division (Antica et al., 1997). Therefore, these committed progenitors may be useful in many therapies where only a single cell type is desired and indefinite expansion potential may not be required. Isolation and Purification of Stem Cells Methods for isolation of ES cells have made use of culture technologies in which IVF-derived fertilized eggs are put into culture under very specific culture conditions. Most important is the use of particular embryonic

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stromal feeder cells supplying factors, mostly unidentified, that keep the cells from differentiating. The most commonly used embryonic stromal feeder cells are ones derived from murine embryos. Although a few of the factors are known, such as “leukemia inhibitory factor” or “LIF,” it is assumed that others are yet to be identified (Langenbach et al., 1979; Wells et al., 1980; Sato et al., 1999). There are recent reports of the development of media that permit feeder-free cultures of the ES cells (Schuldiner et al., 2000). Yet it is unknown still whether long-term maintenance under such feeder-free conditions might alter the developmental potential of the cells. The continued requirement of totipotent stem cells and ES cells for feeders has been and will be a problem for any future clinical use of these cells. The FDA has indicated that they want to try to avoid clinical trials of stem cells expanded on such feeders, since it is unknown if the feeders are contributing a virus or other pathogen that might be harmful to people. It is hoped that the newly established feeder-free conditions will prove able to sustain the cells with retention of their full developmental capacity. Methods for identification and purification of determined stem cells have made use of three approaches: 1. Selection in culture under highly restrictive conditions such as in suspension or on tissue culture plastic and in serum-free medium (Kubota and Reid, 1999; Jian et al., 2002). 2. Flow cytometric selection of cells with altered chromatin organization assessed by reduced uptake of DNAbinding dyes (e.g. Hoechst dyes), the so-called “side-pocket” cells (Goodell et al., 1997). 3. Multiparametric flow cytometric sorting of cells using forward scatter and side scatter properties and, where possible, using monoclonal antibodies to defined antigens unique to the particular type of stem cells (Brill et al., 1993; Sigal et al., 1994, 1995a, b, 1999; Brill et al., 1995; Kubota and Reid, 2000; Kubota et al., 2002; Schmelzer et al., 2007). This approach can be successful even when no antigens are known to define the stem cells of interest. One can enrich significantly for the stem cells by doing a “negative sort” using fluoroprobe-labeled antibodies to markers on contaminant cell populations to separate the population into cells that express those markers and cells that do not. Secondly, one characterizes the cell population remaining for side scatter, a flow cytometric parameter in which the more cytoplasmic particles (mitochondria, ribosomes, etc.), the greater the side scatter of the laser beam. The less mature cells are “agranular” (lower in granularity), whereas the more mature cells are more granular; this enables one to enrich for cell populations of given granularity. Antigenic profiles identified for stem cells have revealed that there are many markers in common among major classes of stem cells. For example, ES cells and most (all?) determined cell types express few antigens of the major histocompatibility (MHC) family and are, consequently, relatively non-immunogenic (Kubota and Reid, 2000; Jian et al., 2002). Similarly, most of them express pumps (e.g. multidrug resistance gene 1 or MDR1) that eliminate xenobiotics (Goodell et al., 1997). There are specific cell adhesion molecules (CAMs), such as CD34, that are present on most mesodermal stem cell types (and not just on hemopoietic stem cells as originally thought) (Goff et al., 1996; Timeus et al., 1998; Ahmed et al., 1999) and proteins critical in vascularization processes such as hedgehog proteins (Sicklick et al., 2006). This strategy of multiparametric sorting has proven the most successful and efficient in identifying and isolating stem cell populations.

Maturational Lineage Biology All tissues are organized with a compartment containing stem cells that give rise to daughter cells maturing stepwise to adult cells, transition to apoptotic cells and, finally, die and are eliminated from the tissue (Sell, 1994; Potten, 1997; Gonzalez-Reyes, 2003). The kinetics of the lineage vary with the tissue and correlate inversely with the extent of polyploidy within the tissue. The rapidly regenerating tissues have lineages with rapid kinetics, such as the intestine with a lineage that turns over within a week or skin and hemopoietic cells that turn over in 4–5 weeks. These lineages typically have only 5–10% polyploid cells, located in the tissue within the sites of the greatest differentiated functions. The newly recognized lineages are those associated with quiescent tissues, such as the liver, with turnovers estimated to be months to years. The extent of polyploidy in these tissues is from 30% to 95%.

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All fetal and neonatal tissues are entirely diploid and transition to the adult ploidy profile within a time frame varying from species to species. In mice, it occurs within 3 weeks; in rats within 4 weeks; in humans by ~20 years of age. With increasing age, the percentage of diploid cells steadily declines. Thus, tissues from geriatric donors have much lower levels of diploid cells than those from young people. The major stages of maturational lineages identified are: those in embryos only (1–3), those in both embryos and adult tissues (4–6), and those dominant in adult tissues (6–8): 1. Totipotent stem cells, able to generate extraembryonic and embryonic tissues. 2. Embryonic stem cells, able to generate all mature cells derived from all germ layers. 3. Germ layer stem cells, able to generate the fates known for each of the three germ layers: ectoderm (skin, brain), mesoderm (cartilage, bone, hemopoietic cells), and endoderm (liver, pancreas, lung, gut). 4. Determined stem cells, which have restricted their genetic potential to a subset of those known for germ layer stem cells, for example, epidermal stem cells (skin), neuronal stem cells (nervous tissue), hemopoietic stem cells (blood), hepatic stem cells (HpSCs) (liver). All are diploid, pluripotent cells capable of symmetric and asymmetric cell division enabling them to have enormous expansion potential and to produce daughter cells of more than one fate. They have a gene expression profile that comprises stem cell genes (e.g. pumps, MDR1 (Ros et al., 2003) that enables them to eliminate xenobiotics) and some genes unique to the class of stem cell (e.g. CD34). 5. Committed progenitors are diploid, unipotent, and immature cells. These precursors give rise to only one adult cell type. They no longer express some of the stem cell genes but express genes typical for cells in the fetal tissues. 6. Diploid adult cells are able to undergo complete cell division for 6–7 rounds, can form colonies in culture but have limited capacity to be subcultured. They express a subset of the adult-specific genes. 7. Polyploid adult cells are no longer able to undergo complete cell division. They can undergo DNA synthesis but with limited capacity for cytokinesis. They are much larger cells (due to the hypertrophy associated with polyploidy) and express high levels of the “late” genes. 8. Apoptotic cells express various markers of apoptosis and demonstrate DNA fragmentation.

The lineages in embryonic tissues are skewed toward the stem cell compartment with few, if any, of the polyploid cells or terminally differentiated cells. Those in young adult tissues have cells representative of all the lineage stages but without the stem cells found exclusively in embryos. The tissues of elderly people are skewed toward the older stages of the lineage even though there remains a stem cell compartment. A presumed exception is the heart, thought to mature rapidly during embryogenesis, with the last time point at which there is rigorous evidence for a stem cell compartment being ~3 months gestational age (Giroux and Charron, 1998); one is born with heart tissue having sufficient lineage intermediates to grow into the size of an adult heart but with limited regenerative capacity. Yet even heart tissue is being re-evaluated for the presence of a stem cell compartment. The data to date remain controversial, and whether or not the heart has any stem cells in the adult tissue will be defined in ongoing and future studies. Therefore, each adult tissue is comprised of a stem cell compartment (the “young” cells in the lineage) that gives rise to adult cells (“middle aged cells”) and then to apoptotic cells (“old” cells) that are sloughed off or in some way eliminated. The speed of turnover of a given lineage has a basal rate and a more rapid rate induced by injury processes. We hypothesize that maturational lineages have a feedback loop in which one or more signals from mature cells inhibit the proliferation of stem cells and/or progenitor cells. This hypothesis is based on numerous findings in studies of liver both in vitro and in vivo. In culture stem cells do not grow if they are in the presence of mature liver cells or are provided conditioned medium from the mature cells (Brill et al., 1993, 1995; Sigal et al., 1994, 1995a, b, 1999; Overturf et al., 1997; Kubota and Reid, 2000; Kubota et al., 2002; Schmelzer et al., 2007); in vivo, there is a need for loss of mature cells in vivo for expansion of the cells from the stem cell

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compartment (Gonzalez-Reyes, 2003; Ros et al., 2003). The implications of this phenomenology are that expansion of cells from the stem cell compartment occurs with injury and loss of mature cells. Moreover, one would predict that chronic loss of mature cells, as occurs in certain viral infections (e.g. hepatitis C), repeated drug exposures, or radiation would elicit chronic regenerative responses that could lead to mutational events associated with malignant transformation. The Epithelial–Mesenchymal Relationship The epithelial–mesenchymal relationship consists of a layer of epithelia bound onto a layer of mesenchymal cells (Fujita et al., 1986; Reid, 1990; Martinez-Hernandez and Amenta, 1993; Brill et al., 1994; Reid and Luntz, 1997). The most common forms are epithelia wed to stroma and epithelia wed to endothelia that are part of a blood vessel. Signaling between and within the two cell layers coordinates local cellular activity. The signaling molecules comprise soluble signals (autocrine and paracrine signals) and an insoluble complex of proteins, lipids, and carbohydrates found outside of the cells and called the extracellular matrix. These two sets of signals work synergistically to regulate the tissue at the “local” level. Soluble Signals Investigators have identified and characterized a multitude of soluble signals (growth factors, cytokines, hormones) and described cellular and molecular mechanisms associated with regulation of cells by these signals. The wealth of information on these signals is so great that the reader is directed to many recent reviews and books on this subject (Balkwill, 1995; Norman and Litwack, 1997; Matzuk et al., 2001). Extracellular Matrix

The extracellular matrix is an insoluble complex of proteins and carbohydrates found on the lateral and basal surfaces of cells (Fujita et al., 1986; Spray et al., 1987; Martinez-Hernandez et al., 1991; Reid et al., 1992; Reid, 1993; Berthiaume et al., 1996; Kim et al., 1997; Boudreau and Bissell, 1998; Pines et al., 1998). On the lateral borders, the lateral extracellular matrix couples homotypic cells (e.g. epithelia to epithelia), whereas the basal extracellular matrix glues together heterotypic cells, forming the connection between the epithelial and mesenchymal cell layers. For many years, the extracellular matrix was thought to play an entirely mechanical role, binding together cells in specific arrays. Now it is understood to be a solid state scaffold that confers persistent signaling mechanisms stabilizing cells in appropriate configurations of intracellular pathways and cell surface molecules (antigens, receptors, ion channels) and in appropriate cell shapes (flattened or three-dimensional). This enables the cells to respond rapidly to soluble signals that can derive from local or distance sources. The primary components of the lateral extracellular matrix are: (a) cell adhesion molecules or “CAMs” that are age and tissue specific (Stamatoglou and Hughes, 1994); (b) tight junction proteins that are age and tissue specific (Rahner et al., 2001); (c) proteoglycans: molecules containing a protein core to which are attached polymers of sulfated (negatively charged) sugars called glycosaminoglycans (GAGs) (e.g. heparan sulfates, heparins, chondroitin sulfates, or dermatan sulfates) (Kjellen and Lindahl, 1991; Ruoslahti and Yamaguchi, 1991; Lyon, 1993; Kim et al., 1994).

The basal extracellular matrix consists of basal adhesion molecules (e.g. the families of laminins and fibronectins) that bind the cells via matrix receptors, integrins, to one or more types of collagen scaffoldings (Brill et al., 1994; Berthiaume et al., 1996; Boudreau and Bissell, 1998). The collagens of one cell layer are cross-linked to those of the adjacent cell layer to provide stable coupling between the layers of cells. Proteoglycans are bound to the basal adhesion molecules, to the collagens, and/or to the basal cell surface. The knowledge of the collagens, basal adhesion molecules, and the integrins has grown so much over the last

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decade that again the reader is directed toward major books and reviews on these subjects (Laurent and Fraser, 1986; Kjellen and Lindahl, 1991; Ruoslahti and Yamaguchi, 1991; Zvibel et al., 1991; Lyon, 1993; Nimni, 1993; Reid, 1993; Lara et al., 2001). This discussion will note only some generalities derived from these many studies and required for strategies in dealing with stem cells. Developmental Changes in Matrix Chemistry

The chemistry of the extracellular matrix changes during development with various types of matrix molecules dominant in fetal tissues and others dominant in adult tissues. The following summary addresses findings of developmental changes occurring in two of the most critical families of matrix molecules, the collagens, and proteoglycans: (a) Collagens are the largest family of proteins known and have more than 25 subfamilies (Seyer et al., 1977; Geerts et al., 1990; Nimni, 1993; Schuppan et al., 1998). They are scaffolding for all epithelial–mesenchymal relationships. Almost all culture studies using collagens have made use of just type I collagen, a type found in vivo in tendons, bones, and in the most mature portions of tissues. However, stem cells and progenitors require collagen types present in embryonic tissues or in the stem cell compartments of adult tissues (Cortivo et al., 1990; Culty et al., 1990; Martinez-Hernandez and Amenta, 1993, 1995; Balazs et al., 1995; Prestwich et al., 1998; McClelland et al., 2006; Turner et al., 2006). Maturational lineages are associated with collagens that transition from those that turn over rapidly (e.g. type IV collagens) to those that are very stable (e.g. type I collagens). This correlates with findings of the behavior of stem cells when cultured on the different types of collagens in which the stem cells remain undifferentiated and expand on embryonic collagens or in hyaluronans and undergo differentiation on the collagens found predominantly in mature tissues (McClelland et al., 2007; Turner et al., 2007). (b) Proteoglycans are among the most complicated of the matrix components having effects through their core proteins as well as their carbohydrate moieties (Fujita et al., 1986, 1987; Ruoslahti and Yamaguchi, 1991; Zvibel et al., 1991; Bernfield et al., 1992; Kim et al., 1994; Kresse et al., 1994; Lara et al., 2001). Transmembrane proteoglycans such as the heparan sulfate proteoglycans can be found bound on the intracellular surface to cytoskeletal elements and on their extracellular domains have GAG chains that bind soluble signals and present the signals in appropriate way (conformation, stability) to their receptors. The chemistry of the GAGs can affect the signals determining the receptors to which they can bind and the turnover of both the signal and its receptors. Although not understood yet with respect to mechanisms, the poorly sulfated GAGs (e.g. heparan sulfates) bind the signals and present them in a way such that they behave as mitogens, whereas the highly sulfated GAGs (e.g. heparins) bind the same signals and present them in a way such that they act as differentiation signals. In addition, there are some transmembrane receptors that are proteoglycans; these have core proteins that are the receptors for soluble signals. Examples are the transferrin receptor (Tf-R) and colony stimulating factor receptor (CSF-R). In these cases, the receptor has its own GAG chains governing binding, stability, and conformation of the soluble signal (i.e. in these examples it would be transferrin or CSF) (Guthridge et al., 1998).

Dynamic Interactions of the Two Mechanisms The two sets of mechanisms dynamically interact with each other. In all tissues, the epithelial stem cells are partnered with MSCs and their maturation is coordinate. The size of the cells, their potential for cell division, their gene expression, and the chemistry of their lateral and basal extracellular matrix are all lineage dependent. Stem Cells and Cancer An old idea, recently rediscovered, is that cancers are actually transformed stem cells. The idea originated with the pioneering work of Van Potter in the 1960s, he proposed that cancers are cells undergoing “blocked

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ontogeny” (Potter, 1978). Later investigators, especially Barry Pierce and Stewart Sell (Stevens and Pierce, 1975; Sell et al., 1987; Sell, 1993, 1994; Sell and Pierce, 1994), characterized various cancers as mutated stem cells. Many functions, long thought to be related to cancer (e.g. α-fetoprotein expression in liver cancers) are now realized to be perfectly normal functions of an expanded stem cell population (Dabeva et al., 1998; Kubota et al., 2002). Therefore, current efforts focus on comparing cancer cells to their normal stem cell counterparts in order to identify the changes in a specific stem cell population that have given rise to the malignancy (Sigal et al., 1993; Brill et al., 1994). These themes have been discussed at length in a number of reviews (Reya et al., 2001; Fariba and Rosen, 2005; Clarke, 2006). A key idea derived from these studies is that cancer cells are blocked at a lineage stage at which cell division is a dominant feature. Indeed, investigators have found that normal stem/progenitor cells are strikingly similar to tumor cells in their appearance, their gene expression, and their growth properties, and that specific tumors, especially specific tumor cell lines, can be mapped or identified as an expanded stage of a lineage (Reid, 1990; Sigal et al., 1992; Brill et al., 1994; Sell and Pierce, 1994; Brill et al., 1995; Rosenberg et al., 1996; Fiorino et al., 1998; Zvibel et al., 1998). This includes lineage stages such as ES cells, determined stem cells, or committed progenitors; it indicates that existing tumor cell line model systems can be used to define properties of their normal stem cell counterparts, and that comparison of those tumors to those normal counterparts should be extraordinarily revealing about key aspects of the malignant transformation process. Implicit also is that clinical use of stem cells may come with an increased risk of tumors depending on the donors (e.g. if there are undiagnosed tumor cells within the stem cell compartment) and on the patient’s medical condition (e.g. severe immunosuppression). Treatment of patients also involves strategies recognizing lineage biology. If a patient’s tumor can be mapped to a specific lineage stage, then treatment of that patient (e.g. with chemotherapies, radiation therapies, etc.) must be targeted to the lineage stage(s) that has aberrant cells. If the treatment eliminates cells at a later lineage stage but not the stage with the aberrant cells, then the treatment will actually worsen the patient’s prognosis, since the feedback loop regulation will be eliminated by the treatment’s killing off of late (mature) lineage stages and subsequent disinhibition of the mutated cells.

SOURCING OF HUMAN TISSUE ES Cells The primary source of cells for ES cells is the discards from IVF procedures. Families undergoing IVF for pregnancies have many fertilized eggs produced by the procedures. Those not used immediately for launching a pregnancy are stored in cryopreserved form. Once the families have as many pregnancies as desired, they must decide whether to discard the extras, or to pay extra for the continued storage of them. Some families have voluntarily donated the remained fertilized eggs to researchers who prepared ES cell cultures from them. There are widely divergent opinions, especially by certain religious and political groups, on the morality and ethics of discarding the extra fertilized eggs (in some of the most extreme opinions, this constitutes murder) or of utilization of them to create ES cell cultures. After months of heated debate at the national and international levels on the legality and ethics of utilizing these cells, a decision was made that researchers may receive federal funding for research on the existing human ES cell cultures but cannot create new ones. Although there are more than fifty such cultures in existence, many if not most of these cultures are poorly characterized and may not yield viable, long-term model systems of human ES cells. It is unknown at the present time how many of them will prove truly useful. A number of investigators who want to pursue the establishment of novel human ES cell cultures have moved to countries where such research is still permitted.

Hepatic Stem Cells: Lineage Biology and Pluripotency 353

Sourcing Issues of Multipotent and Determined Stem Cells The sources for multipotent and determined stem cells are fetal tissues from spontaneous abortions and tissues from neonatal, pediatric, and adult donors. Fetal Tissue Most of the studies on determined stem cells are done on tissues removed from aborted fetuses. Some of the staff of certain organ and tissue procurement agencies retrieve abortuses from abortion clinics, dissect the tissue, and ship it to investigators throughout the country. The amount of tissue from abortuses is sufficient to supply the needs of many investigators. Although fetal tissues have the highest known numbers of determined stem cells/gram of tissue, this source can be used for research but not readily for clinical programs. The strong opinions on abortion held by many religious, political, and cultural groups preclude the use of this source for determined stem cells that might be used in people (see last section of the review). Brain-Dead-but-Beating-Heart Donors The current sources for determined stem cells from organs, other than fetal organs, are those from braindead-but-beating-heart donors. The numbers of determined stem cells/gram of tissue can remain similar throughout life (Schmelzer et al., 2007) but their immediate descendents, the unipotent committed progenitors, decline with age (Kabrun et al., 1997) with the highest numbers being in fetal and neonatal tissues and the lowest (if any) in the tissues of geriatric donors; the implications are that the younger the donor, the greater the yields of the determined stem cells and their committed progenitors (Gordon et al., 2000). All human organs derive from donors who have undergone massive head injury leading to brain death but not cardiac arrest, accounting for 1–2% of the deaths in the United States (www.unos.org). This is because the intact organs are exquisitely sensitive to ischemia and other biochemical changes associated with death and the cessation of heart function; they deteriorate rapidly after death and are unusable for transplantation within ~30 min of death. The organs are removed from donors, quickly chilled by flushing with and submerging in a transport buffer, and then transported to an institution where a candidate recipient is located. Empirically, it has been found that organs must be transplanted within approximately 18 h after removal from the donor to have a reasonable chance for a successful transplantation. This major time constraint has led to a nationwide program in which the country has been divided into districts; the staff of organ procurement agencies in a given district interface with families to get permission for organ donation from a brain-dead patient, and then arrange for the organs to be transferred to recipients within the same (or nearby) district. These severe time constraints limit the amount of testing that can be done on the organ. Typically, the organs are tested for diseases using serology; serological procedures are usually adequate but cannot detect newly acquired infections. The more accurate and sensitive assays (e.g. PCR assays) can detect disease regardless of when it was acquired, but they require several days to be done and so cannot be used for organ transplantation procedures. Similarly, tissue typing requires several days and so cannot be done on organs; rather the donor is checked only for blood type. As a result, the recipient must be transplanted with an organ that necessitates immunosuppression for the rest of his/her life, a major source of complications for transplant patients. Rejected organs, organs found to have a disease or aberrant vascular system, or organs that have undergone too long a period of time since removal from the donor, are made available for research. These rejected organs constitute less than 5% of those donated. Needless to say, the competition for this material by academic and industrial investigators is fierce given its scarcity. Although the federal government tries to help alleviate the competition by surgically dividing the organs and distributing the portions to more investigators, there remains an extraordinary limit to the tissues and organs available.

354 CELLS AND TISSUE DEVELOPMENT

Asystolic Donors (Also Called Tissue Donors) Given the extreme limitation of tissues and organs from brain-dead-but-beating-heart donors, investigators have begun to explore an alternate source: asystolic donors, that is, donors who have undergone heart arrest, which constitutes 98% of the deaths in the United States (Reid et al., 2000). Asystolic donors, so-called “tissue donors,” are the source for such tissues as corneas, heart valves, skin, cartilage, and bone. The nationwide network of organ donor agencies also manages the procurement of tissue from tissue donors and investigators have defined how long after death and under what conditions one can retrieve viable cells. Whereas the mature cells of organs (heart, lungs, liver, pancreas) die within an hour of cardiac arrest, the mature cells of some tissue donors can survive for a few hours. The longest-lived cells from the tissues of both organ and tissue donors are the stem cells and committed progenitors. The empirical findings are that stem cells are relatively tolerant of the ischemia that occurs after cardiac arrest making them a novel source of human cells for clinical, academic, and industrial programs (Smith, 2006). Current investigations focus on defining the restrictions, in terms of length of warm and/or cold ischemic time, and the conditions associated with the dying process that dictate the quality and the number of viable stem cells that can be obtained from the donor. Neonatal Donors A newly established source, and an ideal one for determined stem cells, is the neonate who dies at birth or on a neonatal intensive care unit (NICU). Tissues and organs from neonates have been used rarely in the past because they are too fragile for tissue or organ transplantation, and because they must be procured postmortem given that brain death cannot be defined in these donors. The first program in the world to procure neonatal organs (especially liver) for purposes of cell therapy goals was activated at the University of North Carolina, Chapel Hill, in January 2002. Although the program is in its infancy, the data from the first experiments indicate that neonatal tissues are replete with stem cells and progenitor cells that persist for hours (e.g. 6–7 h) after cardiac arrest and readily establish in culture under defined conditions (Kubota and Reid, 2000; Schmelzer et al., 2006a, b). It is hypothesized that the stem cells from neonates and from adults are relatively equally tolerant of ischemia; yet those from neonatal tissues do better when isolated (Smith, 2006). This is due, it is assumed, to the findings that most of the cells in neonatal tissues are stem cells or the committed progenitors, all being relatively tolerant of ischemia, such that the entire organ survives as an organ for greater than 6 h. By contrast, adult organs undergo massive autolysis within an hour or two of death releasing autolytic enzymes that can have an adverse affect on the surviving cells, even the stem cells. Implicit in these findings is that neonatal tissues are likely to be a primary source of stem cells and progenitors for all forms of cell therapy.

THE LIVER AS A STEM CELL AND MATURATIONAL LINEAGE SYSTEM (FIGURES 20.1–20.3) Liver is being presented as a representative quiescent tissue that has been found to be organized with a stem cell compartment giving rise to maturational lineages of daughter cells. Organization of the Liver The liver’s organizational plan is as acini that are hexagonal in shape and with six sets of portal triads (hepatic artery, hepatic vein, and bile duct) demarcating the corners of the hexagon and with a central vein in the center of the acinus (Weiss, 1983; Jungermann and Katz, 1989). Incoming blood flows from the gut and from the spleen into the liver via the portal triads. It passes across the plates of liver cells extending between the portal triads and the central vein, and then leaves the liver via the central vein which is connected then to the vena cava. The blood flow across the liver is 1,500 ml/min, constituting 25% of cardiac output. This is subdivided, with 75% being supplied by the hepatic vein (coming from the spleen), and the remaining 25% being supplied

Hepatic Stem Cells: Lineage Biology and Pluripotency 355

Interlobular vein – branch of heptic vein carries away deoxygenated blood Sinusoid Cord of hepatocytes (liver cells) Kupffer cell

Bile canaliculus Arteriolebranch of heptic artery (brings oxygenated blood ) Interlobular veinbranch of heptic portal vein (brings blood from gut )

Bile duct (takes bile to gall bladder )

Figure 20.1 Schematic drawing of liver plates showing the portal triad, central vein and plates of liver cells. From: http://www.biologymad.com/kidneys/liver lobule.

Figure 20.2 Section of liver stained with haemotoxylin and eosin to show the histology of the liver acinus and, in particular, the zonation. The image is from http://www.md.hugi.ac.il/mirror/webpath/liver.

by the hepatic artery (coming from the stomach and duodenum). Thus, there is very low shear in the blood flow across the cells. By convention, the liver is demarcated into three zones: zone 1 is peiportal; zone 2 is midacinar; and zone 3 is pericentral (Figure 20.1 is www.biologymad.com/kidneys/liverlobule and Figure 20.2 is www.md.huji.ac.il/mirror/webpath/liver.html). The properties of the cells vary, in gradient fashion, along those zones. The smallest cells, all of them diploid, are located in zone 1 and the largest cells, all of them polyploid, are located in zone 3 (Sasse et al., 1979; Weiss, 1983; Foucrier et al., 1988; Jungermann and Katz, 1989; Marti and Gebhardt, 1991; Sigal et al., 1992; Brill et al., 1994; Lindros et al., 1997). The cell division potential of the cells is maximal periportally and negligible pericentrally. Specific genes are expressed in characteristic zones and can be interpreted as “early,” “intermediate,” and

356 CELLS AND TISSUE DEVELOPMENT

Periportal area PV

Pericentral area SE CV

Zone 1

Zone 2

Zone 3

HA Key: PV - portal vein; BD - bile duct: HA - hepatic artery: SE - sinusoidal endothelium over the space of Disse: CV - Central vein. The portal triad and central vein are surrounded by a matrix which differs from the vascular basement membrane: see table below.

1

2

3

rats

2N

4N

4N & 8N

mice

2N&4N

4N&8N

up to 32N

Zones Ploidy

humans 2N Maximum Growth ECM Type IV&III collagen, laminin, HS-PG* Genes

Early

2N Limited gradient Intermediate

2N & 4N Negligible Type I & III collagens, Fibronectin, HP-PG* Late

Size(μ) 2N < 20 ; 4N = ∼20 - 35; 8N and above = >35 *HS-PG=heparan sulfate proteoglycan; HP-PG=heparin Proteoglycan

Figure 20.3 Liver lineage model. “late.” A summary of key properties is noted in Figure 20.3. Extensive reviews of the zonal properties within the liver have been published (Gebhardt and Mecke, 1983; Jungermann, 1986; Gebhardt, 1992; Eilers et al., 1993; Brill et al., 1994), and representative gene expression demonstrating such zonation includes:

• • •

Zone 1: P450A7, Ccnnexin 26, type IV collagen, laminin, heparan sulfate proteoglycans (syndecans), enzymes involved in gluconeogenesis such as PEPCK (Berthoud et al., 1992; Kojima et al., 1996; Rosenberg et al., 1996). Zone 2: Transferrin, tyrosine aminotransferase (Yeoh and Morgan, 1974; Shelly et al., 1989). Zone 3: P4503A1, type I collagen, fibronectin, heparin proteoglycan, major urinary protein (MUP), and glutamine synthetase (Gebhardt and Mecke, 1983; Liu et al., 2003).

Stem Cell Compartment of Human Livers An extensive review of the current knowledge of HpSCs has just been published (Schmelzer et al., 2006b). Below we summarize statements from that review and from recent articles on hepatic progenitors. The formation of the liver is initiated by an endodermal stem cell population in the embryonic foregut (Mobest et al., 1999; Matsumoto et al., 2001) and with processes leading to the subsequent formation of mature hepatocytes, cholangiocytes, and other hepatic cell types (Zaret, 1998, 1999). Liver development has been linked to HNF1 and HNF6b signaling in a highly localized response to cells immediately adjacent to the portal tracts (Clotman et al., 2002; Coffinier et al., 2002). These cells are referred to as the ductal plate, or limiting plate, and are the focus of intense hedgehog signaling processes (Sicklick et al., 2006) that are associated with the co-development of the liver’s vasculature and the parenchymal cells. The ductal plate has been shown now to be the reservoir of the HpSCs (Zhang et al., 2007), has characteristic intense staining with cytokeratin 19 (CK19), and with neural cell adhesion molecule, N-CAM (Ruebner et al., 1990; Fabris et al., 2000). The ductal plate transitions by unknown mechanisms to become Canals of Hering in adult livers (Theise et al., 1999). Adjacent to the ductal plates are hepatoblasts, recognizable by their intense expression of α-fetoprotein. Hepatoblasts are the dominant parenchymal cell population in fetal and neonatal livers, and have been shown to be bipotent, giving rise to the committed biliary and hepatocytic progenitors. The number of hepatoblasts declines in the livers of hosts of increasing age; they are difficult to find in adult livers except in the presence of ongoing disease such as cirrhosis or hepatitis.

Hepatic Stem Cells: Lineage Biology and Pluripotency 357

ALB, CK19, N-CAM

ALB, CK19, ICAM1,AFP

Committed Hypatocyte progenitors

Hepatocytes Zone 1 (diploid) Zone 2 (diplois) Zone 3 (tetraploid)

Hepatic stem cells

Hepatoblasts: Bipotent hepatic stem/progenitor cells

ALB, CK19

ALB, CK19

Self renewal

Transit amplifying cells? Self renewal?

Committed bile duct progenitors

PEPCK Transferrin

Aquaporins

P4503A1

MDR3, DPPIV

Bile duct epithelium

Figure 20.4 Human liver lineage working model.

Past studies have resulted in strategies for isolation of hepatic progenitors from livers (Reid et al., 1993; Sigal et al., 1994, 1995a, b, 1999) and in the development of serum-free, defined culture conditions for expansion versus differentiation of the cells (Kubota and Reid, 2000). The purified progenitors have been shown to be able to mature to adult fates after transplantation in vivo (Sigal et al., 1995b). Similar strategies have been utilized for identification of progenitors in human livers and have resulted in a startling find: that the HpSC is not a hepatoblast but its precursor, a cell type that does not express α-fetoprotein (Schmelzer et al., 2007). From studies on human livers (fetal, neonatal, pediatric, and adult), Reid and associates have defined the antigenic profiles for all known cellular components of the liver’s stem cell niche and that comprise parenchymal progenitors consisting of two pluripotent parenchymal cell populations (HpSCs and hepatoblasts) and two unipotent parenchymal progenitors (the committed biliary and hepatocytic progenitors) (Kubota et al., 2007; Schmelzer et al., 2007; Sicklick et al., 2006); hepatic stellate cell precursors (Kubota, 2007); and hepatic angioblasts (Yao et al., 2007). All four populations of parenchymal progenitors are wholly negative for hemopoietic markers (CD45, CD34, CD38, glycophorin A), making them distinct from the progenitors described from bone marrow or other sources (Petersen et al., 1999; Theise et al., 1999, 2000; Jian et al., 2002); all four subpopulations express epithelial cell adhesion molecule (EpCAM) and three of the four (hepatic progenitors and angioblasts) express prominin (CD133/1). EpCAM has been shown to be expressed by biliary cells (Ruebner et al., 1990; Blakolmerl et al., 1995; Schmelzer and Reid, 2007), but when co-expressed with albumin, is a marker for progenitors (Schmelzer et al., 2007). Prominin, a polytopic membrane protein, is found on various stem cell populations and has unknown functions (Weigmann et al., 1997; Corbeill et al., 2000). The size of the EpCAM populations (7–10 μm) is strikingly different from that of mature adult liver cells (18–25 μm). The differential antigenic profiles of the stem cells, the hepatoblasts, the unipotent progenitors, and the diploid and the tetraploid hepatocytes are given below and summarized in Figure 20.4. 1. HpSCs or ductal plate cells are multipotent, agranular, have high nucleus to cytoplasmic ratios, are 7–10 μm in diameter, and are located within the ductal plates of fetal and neonatal livers or the Canals of Hering in pediatric and adult livers. The antigenic profile of these cells is albumin, CK19, EpCAM, CD133/1, CK8/18,

358 CELLS AND TISSUE DEVELOPMENT

2.

3.

4.

5.

Indian Hedgehog, telomerase, claudin 3, and N-CAM (Schmelzer et al., 2007; Sicklick et al., 2006). The HpSCs are negative for all forms of P450s and for ICAM-1 and even, surprisingly, for α-fetoprotein (Schmelzer et al., 2006a, b; Sicklick et al., 2006). Ex vivo expansion of the cells occurs with a defined medium developed for hepatic progenitors (Kubota and Reid, 2000) and substrata of embryonic matrices (Schmelzer et al., 2007). Hepatoblasts are larger (10–12 μm) with higher amounts of cytoplasm and side scatter, are located throughout the parenchyma in fetal and neonatal livers but decline in numbers such that in postnatal livers are only greater than 0.1% of the parenchymal cells and are found tethered to the ends of the Canals of Hering. Their numbers wax and wane with injuries. Their antigenic profile overlaps with that of the HpSCs except for the following: they express ICAM-1 (not N-CAM), express α-fetoprotein intensely, and express fetal forms of P450s (e.g. P450A7) (McClelland et al., 2006; Schmelzer et al., 2006a, b; Schmelzer et al., 2007). Unipotent progenitors also wax and wane in numbers in pediatric and adult livers in a pattern similar to that in the hepatoblasts. They have low side scatter and are ~12–15 μm in diameter. There are two subpopulations: committed biliary progenitors are EpCAM, CD133, CK8/18, CK19, and negative for albumin, α-fetoprotein, and N-CAM; committed hepatocytic progenitors are EpCAM, CD133, CK8/18, albumin, α-fetoprotein, and negative for CK19 and N-CAM. Diploid adult hepatocytes (“small hepatocytes”) are present in the livers of donors at all ages and with percentages being over 85% in pediatric livers, and over 50% in adult livers. The antigenic/biochemical profile is, in part, albumin, ICAM-1, CK8/18, PEPCK, connexin 26, and with intermediate inside scatter. They are negative for EpCAM, CD133/1, and α-fetoprotein. Their size is, on average, 18–22 μm. Polyploid adult hepatocytes are present in the livers from teenagers to elderly donors, and their numbers increase with age. The polyploid cells are mostly (entirely?) tetraploid, binucleated cells. The antigenic profile is, in part, albumin, ICAM-1, CK8/18, P4503A, connexin 32 and high side scatter. They are negative for EpCAM, N-CAM, α-fetoprotein, CK19, and CD133/1. Their size is, on average, above 25 μm.

Ectopic Sources of Liver Precursors In addition to the progenitors identified in liver, multipotent precursor populations have been identified also from bone marrow and adipocytes (Brill et al., 1993; Sigal et al., 1994, 1995a; Overturf et al., 1997; Laconi et al., 1998; Shafritz, 2000). Demonstrations that bone marrow-derived cells can mature into hepatocytes both in vitro (Jian et al., 2002) and in vivo (LeGasse et al., 2000) have led to the exciting possibility that they might serve as an alternative to liver transplantation. However, the extremely low efficacy in reconstituting damaged liver tissue by bone marrow-derived cells, and the realization that most of the apparent transdifferentiation is actually fusion of donor cells with host cells, will minimize their use in clinical liver cell therapy programs (Terada et al., 2002; Vassilopoulos et al., 2003). In all studies to date, maximal reconstitution of livers occurs with liver-derived cells, especially small (12 μm) progenitor populations. It appears that HpSCs, endogenous to the liver (not bone marrow or adipocytes) remain the most promising cells for therapies involving cell transplantation or bioartificial organs. Tissue Engineering of Liver Success in tissue engineering liver, as for all solid organs, requires seeding the epithelial stem cells (HpSCs or other progenitors) into or onto a scaffold of embryonic extracellular matrix and in a medium containing the soluble signals (autocrine, paracrine, and endocrine), nutrients, and gases (e.g. oxygen) needed by the cells. The cells will differentiate into the tissue by maturing through the lineage stages found in vivo. Their ability to do this depends on the critical ability to be three-dimensional and to be able to establish gradients of signals (e.g. nutrients, hormones, oxygen) that are required to define the maturational process of the cells and leading to the heterogeneity of cell phenotypes typical of all tissues. This has been accomplished in a muted form in monolayer cultures (Xu et al., 2001) and to a greater extent in spheroid cultures (Koide et al., 1990; Ito and Chang, 1992; Lazar

Hepatic Stem Cells: Lineage Biology and Pluripotency 359

et al., 1995; Thorgeirsson et al., 2004; Cheng et al., 2007) in which cells are allowed to aggregate and form balls of cells (spheroids) that float in the culture medium. The pinnacle of three-dimensional culture systems in maintaining differentiated function have been spheroids cultured on various extracellular matrices and in serum-free, hormonally defined media (HDM) (Tong et al., 1990; Grohn et al., 1997). The alginate encapsulated hepatocytes are being sold as the Liverbeads™ (http://www.liverbeads.com) and spheroid or clusters of cells are presently being developed by DuPont. However, spheroid cultures have not been adopted by most investigators because of technical problems in handling the cultures (e.g. media changes with floating balls of cells are problematic), and because the balls of cells grow and outstrip the ability of nutrients and soluble signals to reach all parts of the tissue, resulting in pockets of necrosis or apoptosis. These problems are being resolved by the use of bioreactors that provide perfusion of media and gases in a precise way to facilitate mass transfer of nutrients. Below is discussed the microenvironmental variables and in various reviews are summarized the diverse forms of bioreactors that have been developed to facilitate mass transfer of gases and nutrients into liver cells (Gerlach, 1996; Brusse and Gerlach, 1999; Macdonald et al., 1999; McClelland and Coger, 2000; Allen et al., 2001; McClelland et al., 2003). Biological Issues for Tissue Engineering of Liver: Microenvironment Contributed by Extracellular Matrix and Soluble Signals (Tables 20.1–20.3). The matrix chemistry associated with the parenchymal cells is present in the Space of Disse, between the parenchyma and the endothelia, and undergoes a transition from that found in the periportal zone to that

Table 20.1 A serum-free, HDM for maintenance of mature liver cells Components Basal media

RPMI 1640 nicotinomide (4.4 mM)  L-glutamine (2 mM)

Trace elements

Selenium, 3  1010 M; zinc sulfate, 5  1011 M; copper sulfate, 1010 M

Lipids

High density lipoprotein (10 μg/ml), free fatty acids (see below) bound to purified human albumin (0.2% w/v)

Free fatty acids

Linoleic acid, 2.7  106 M; palmitic acid, 2.3  106 M; oleic acid, 1.0  106 M; stearic acid, 8.8  107 M, palmitoleic acid, 2.1  107 M; linolenic acid, 4.2  107 M.

Calcium

0.6 Mm

Hormones/growth factors Hydrocortisone

Insulin (5 μg/ml), epidermal growth factor (50 ng/ml), tri-iodothyronine or T3 (109 M), hydrocortisone (108 M)

Table 20.2 Kubota’s medium (KM), a serum-free HDM for HpSCs, hepatoblasts, and unipotent progenitors Components

KM for HpSCs: differentiation

KM for HpSCs: self-replication

Basal media

RPM 1640  nicotinamide (4.4 mM)  L-glutamine (2 mM)

Lipids

High density lipoprotein, HDL (10 μg/ml)  free fatty acids (see below) bound to purified human albumin (0.2% w/v)

Free fatty acids

Linoleic acid, 2.7  106 M; palmitic acid, 2.3  106 M; oleic acid, 1.0  106 M; stearic acid, 8.8  107 M, palmitoleic acid, 2.1  107 M; linolenic acid, 4.2  107 M

Shared hormone requirements

Insulin (5 μg/ml), transferring/Fe (5 μg/ml)

Trace elements

Selenium, 3  1010 M; zinc sulfate, 5  1011 M; copper sulfate, 1010 M

Selenium, 3  1010 M; zinc sulfate, 5  1011 M

Calcium

0.6 Mm

0.3 mM

Hormones/growth factors

9

EGF (50 ng/ml), T3 (10 M), hydrocortisone (108 M)

–

360 CELLS AND TISSUE DEVELOPMENT

Table 20.3 Known extracellular matrix substrata for liver cells Self-replication

Expansion

Differentiation

Stem cells hepatoblasts

Hyaluronans (Turner et al., 2006); Type IV collagen (McClelland Type III collagen (Kubota and et al., 2006) Reid, 2000; McClelland et al., 2006)

Type I collagen (McClelland et al., 2006) tissue extracts enriched in extracellular matrix components: biomatrix and matrigel

Committed progenitors

Does not occur

–

–

Diploid hepatocytes

Does not occur

Type IV collagen; limited expansion on Type I

Type I (especially if cells embedded in it) (LeCluyse et al., 2000; McClelland and Coger, 2003); Matrigel (Schuetz et al., 1988; Brown et al., 1995; LeCluyse et al., 1999); Heparin proteoglycans (Fujita et al., 1987; Spray et al., 1987; Shinji et al., 1988; Ruoslahti and Yamaguchi, 1991; Zvibel et al., 1991; Bernfield et al., 1992; Brill et al., 1994; Shinji et al., 1988)

Polyploid hepatocytes

Does not occur

Does not occur

–

found pericentrally. The matrix chemistry periportally (zone 1) is similar to that found in fetal livers and consists, in part, of type IV and type III collagens, hyaluronans, laminin, and forms of heparan sulfate proteoglycans and chondroitin sulfate proteoglycans. It transitions to one in the pericentral zone (zone 3) with stable forms of fibrillar collagens (e.g. type I), forms of fibronectin, and heparin proteoglycans (Martinez-Hernandez et al., 1991; Reid et al., 1992; Sigal et al., 1992; Martinez-Hernandez and Amenta, 1993a, b, c, 1995). Extracellular matrix is known to regulate the cell’s morphology, growth, and cellular gene expression (Fujita et al., 1986; Spray et al., 1987; Reid, 1990; Mooney et al., 1992, 1994; Reid, 1993; Singhvi et al., 1993; Brill et al., 1994; Ingber et al., 1995; Griffith and Lopina, 1998). Achieving liver histological structure and possibly organotypic functions is important for expression of tissue-specific functions of cells cultured under ex vivo conditions (Xu et al., 2000; Macdonald et al., 2002). These requirements can be achieved ex vivo by using purified extracellular matrix components, now available commercially, and using surfaces that are porous and flexible to permit critical cell shape changes (Xu et al., 2000; Macdonald et al., 2002). This is especially important for HpSCs/progenitor cells with a high nucleus to cytoplasm ratio, resulting in intolerance for attachment to impervious and rigid surfaces (Xu, 2001). Optimal survival, expansion, and differentiation of the cells depend also on use of serum-free medium conditions, since serum drives the cells toward responses appropriate for wound formation (fibrosis, i.e. scar formation) and, in parallel, loss of tissue-specific functions. Serum-free, basal media supplemented with defined mixtures of supplements can be tailored to elicit an appropriate response, either growth or differentiation, of the cells (Brill et al., 1994; Kubota and Reid, 2000; Xu et al., 2000; Macdonald et al., 2002; Schmelzer et al., 2007; Wanthier et al., 2007). Thus, there are HDM for expansion and others for differentiation of a given maturational stage of parenchymal cell (see Tables 20.1–20.3). The details of the isolation of liver cells, of the development and use of HDM, and the use of extracellular matrix substrata are given in a lengthy methods

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review published in 2002 (Macdonald et al., 2002) and in several recent articles (McClelland et al., 2007; Schmelzer et al., 2007). The Need for Feeders The epithelial–mesenchymal relationship is mediated by known soluble signals and extracellular matrix substrata, as discussed above. However, there remain paracrine signals not yet identified or not yet fully characterized. These are provided by the use of tissue- and age-specific feeder cells. In the liver’s stem cell compartment, the paracrine signals are derived from angioblasts, hepatic stellate cell precursors, and MSCs (Kubota and Reid, 2006; Sicklick et al., 2006; Yao et al., 2006). Liver development is induced in a step-wise process with the signals from cardiac mesoderm and then from the angioblasts, the precursors of endothelium and of stroma (Matsumoto et al., 2001). Initial stages of hepatogenesis require fibroblast growth factors (FGFs) secreted from pre-cardiac mesoderm and bone morphogenetic proteins (BMPs) from the septum transversum mesenchyme (STM) (Hebrok et al., 1998; Rossi et al., 2001). Newly specified hepatic cells delaminate and migrate into the surrounding STM and intermingle with precursors to endothelia and to stroma. The mesenchymal cells remain in contact with hepatic cells throughout development (Lammert et al., 2003; Bautch and Ambler, 2004). Mutant animals that lack endothelial cells have the initial indications of hepatic induction but no proliferation of cells into the surrounding STM (Shalaby et al., 1995, 1997). Vascular endothelial growth factor-A (VEGF-A) is a factor critical in angiogenesis (Hogan, 2004; Kearney et al., 2004; Roberts et al., 2004); it increases proliferation of liver endothelial cells (LSECs) by activating VEGF receptor-2 (VEGFR-2/flk-1/KDR), and promotes growth of hepatocytes through paracrine signaling from the endothelia via VEGF receptor-1 (VEGFR-1/flt-1) (Bautch et al., 2000; Ambler et al., 2003; Bautch and Ambler, 2004). Once the paracrine signals from the angioblasts and hepatic stellate cell precursors are defined, it will be feasible to maintain the stem cells ex vivo under wholly defined conditions. This is clearly a major focus for future research. Liver Regeneration Two forms of liver regeneration have long been known, and the stem cell compartment plays roles, albeit distinct ones, in both. 1. Liver regeneration following toxic injuries (due to chemicals, viruses, radiation) involves selective loss of the mature parenchymal cells in zones 2 and 3. This creates a “cellular vacuum” followed by proliferation of cells in zone 1. The zone 1 cells comprise progenitors and diploid adult cells, all of which differentiate to the mature parenchymal cells typically found in the pericentral zone. This phenomenon is the classic “oval cell” response in which small cells with an oval shaped nucleus are induced to proliferate following toxic injury to the liver (Strain et al., 2004; Thorgeirsson et al., 2004). This in vivo phenomenon is paralleled by the findings, noted earlier, that stem cells in culture are inhibited by soluble signals released from mature hepatocyte, the “feedback loop.” The feedback loop explains why purification of diploid subpopulations away from polyploid ones is required to observe clonal growth of diploid cells in culture and why significant expansion of transplanted liver cells occurs only in hosts in which there is a “cellular vacuum” in the pericentral zone. 2. Liver regeneration after partial hepatectomy (surgical removal of a portion of the liver) has long been thought to be mediated only by mature liver cells (Michalopoulos et al., 1987). However, it has now been shown to involve the stem cell compartment (Sigal et al., 1999). In the first 24 h after partial hepatectomy, there is a wave of DNA synthesis across the liver plates, but with limited cytokinesis, resulting in elevated polyploidy and a sharp decline in the diploid subpopulations (Liu et al., 2003). The ploidy profile of the parenchymal cells is restored slowly and gradually over several weeks by contributions from the stem cells.

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3. Implications of liver regeneration and lineage phenomena for cell transplantation with stem cells/progenitors. HpSCs and hepatoblasts are the lineage stage of cells most likely able to provide the maximum reconstitution of livers after transplantation given their maximum potential for proliferation (Dabeva et al., 2000; Sandhu et al., 2001). Yet these progenitor subpopulations will behave distinctly in animals, or people, depending on the mechanisms of regenerative stimuli. The maximum regenerative responses of HpSCs and hepatoblasts are predicted to occur in animals (or in humans) when there is a cellular vacuum created by drugs, radiation, or genetics, and the resulting loss of the feedback loop signals from the old cells in the pericentral zone (Dabeva and Shafritz, 1993; Overturf et al., 1997; Dabeva et al., 1998; Grompe et al., 1999; Gupta et al., 1999, 2000) (reviewed in Susick et al., 2002 and in Schmelzer et al., 2006a). Humans suffering from acute liver failure would be representative of those in whom the feedback loop signals are lost and in whom the maximum proliferation of progenitor subpopulations would be predicted.

By contrast, transplantation of progenitors into patients with inborn errors of metabolism should undergo some limited growth in parallel with the growth of the patient’s liver cells, but there will be no selection of the transplanted cells over the host cells. Thus, these patients will have intact feedback loop signals and should require much higher dosages of transplanted cells than those with liver failure. Transplantation of stem/progenitor populations into patients suffering from a disease, such as cancer, and in which the treatment involves a partial hepatectomy, should result in some expansion of the transplanted cells but with the growth potential intermediate between that found in patients with inborn errors of metabolism and that in patients with liver failure. The most difficult patient population of all is likely to be patients with cirrhosis who have an aberrant liver infrastructure containing excessive scar tissue. Engraftment of transplanted cells might be inhibited, and those cells that do engraft will be in a microenvironment that could inhibit their growth and cause them to terminally differentiate (e.g. excessive amounts of type I collagen). Therefore, strategies for transplantation of stem/progenitor cell populations will be different depending on the disease state in the patient.

CLINICAL, COMMERCIAL, AND RESEARCH APPLICATIONS OF STEM CELLS (WITH A FOCUS ON LIVER) The known properties of the different classes of progenitors help to define their potential in academic, clinical, or industrial programs. These will be summarized both generally and with specific examples for current or future uses of the cells. Properties of Stem Cells Several features of stem cells make them ideal as “off-the shelf ” products for clinical, academic, and industrial programs: ability to be cryopreserved, expansion potential, and behavior after transplantation. Cryopreservation of Stem Cells versus Mature Cells The ability to be cryopreserved is a property of all known stem cells (Chen, 1992; Resnick et al., 1992; Maltsev et al., 1993; Thomson et al., 1998; Schuldiner et al., 2000). The cells are suspended in a cryopreservative buffer, aliquoted into small aliquots (e.g. 3 ml cryovials) or larger volumes (e.g. 50–1,000 ml cryocyte bags), and frozen to liquid nitrogen temperatures (–160°C) using a computerized control rate freezer. The samples are stored in the vapor phase of liquid nitrogen (–160°C) to minimize cross contamination. The ability to cryopreserve the cells enables the establishment of a cell bank in which stem cells can be rigorously screened for diseases, genotyped, and tissue typed and then stored indefinitely, greatly facilitating the logistics of getting well characterized cells from donor to recipients.

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By contrast, cryopreservation of adult cells (e.g. mature cells from lung, liver, pancreas, etc.) has met with failure or with only limited success, and even that limited success is achieved only by embedding the cells in alginate or in a form of extracellular matrix (Koebe et al., 1990; Lin et al., 1994; Guyomard et al., 1996; Swales et al., 1996).The freshly thawed mature cells rarely survive for more than a few days. It is unknown why the mature cells are so difficult to cryopreserve; changes in chromatin (e.g. increased ploidy) is one possibility. Expansion Potential of Stem Cells versus Mature Cells The ability of stem cells to expand even at very low cell densities (Chen, 1992; Resnick et al., 1992; Maltsev et al., 1993; Thomson et al., 1998; Schuldiner et al., 2000) is one of their most important features. It permits the generation of large numbers of cellular offspring that can be used clinically, industrially, or for research. Indeed, the most rigorous way to define a stem cell is to seed it as a single cell into a dish under specified conditions, allow it to expand into a population of daughter cells, and then demonstrate that the daughter cells are capable of differentiating to multiple adult fates (Kubota and Ried, 2000). The expansion potential of stem cells should make them ideal for gene therapies, for establishment of bioartificial organs, and for protein manufacturing. ES cells are especially renowned for their ability to expand ex vivo without differentiating if maintained under precise culture conditions. They can then be lineage-restricted to at least some defined fates with use of particular soluble factors and/or components of the extracellular matrix. Adult (mature) cells of all tissues are very different in their growth potential from that of stem cells and consist of two subpopulations differing qualitatively in their ability to divide: 1. Some diploid subpopulations are able to expand in culture, are able to be subcultured through a few rounds, and are able to form colonies of cells when seeded at very low cell densities under precise culture conditions (Kubota and Reid, 2000). The number of divisions possible for the diploid, adult subpopulations is limited; cell division numbers of 5–7 divisions are typical. 2. Many mature cells, especially all polyploid subpopulations, cannot undergo complete cell division; rather they undergo DNA synthesis without cytokinesis resulting in hypertrophy (increase in cytoplasmic mass) and increased polyploidy (Mitaka et al., 1995; Tateno and Yoshizato, 1999). These mature cells survive for a matter of days in culture, or with appropriate extracellular matrix and medium conditions, will survive for a few weeks (Enat et al., 1984; Reid and Jefferson, 1984; Reid and Luntz, 1997; LeCluyse, 1999; LeCluyse et al., 2000a).

Reconstitution of Tissues by Transplantation of Stem Cells The major feature of stem cells in their potential for clinical programs is their ability to reconstitute damaged tissues in vivo. The ability of ES cells to give rise to all or almost all possible adult fates makes them appealing as a “one serves all” approach for cell therapies and makes them the most exploitable of the classes of stem/progenitor cells. However, their use in cell transplantation for patients is precluded by their well known tumorigenic potential; in animal studies, they are virtually 100% tumorigenic (Martin, 1981; Chen, 1992; Thomson et al., 1998; Mendiola et al., 1999). The tumorigenicity of ES cells when injected at ectopic sites is being investigated extensively, especially by biotechnology companies, in hope that it can be controlled to enable ES cells to reach their full potential both industrially and clinically (Mendiola et al., 1999). Investigators are experimenting with lineage-restricted ES cells that are transfected with genes that might control the tumorigenicity, such as ones to control telomerase, an enzyme present in stem cells and in tumor cells. While determined stem cells are more restricted in their adult cell fates, they have not been found to be tumorigenic, enabling them to be the first choice for clinical programs in cell transplantation (Susick et al., 2001). Comparative studies of the ability of adult cells (Rhim et al., 1994, 1995; Overturf et al., 1997, 1999; Gupta et al., 1999, 2000) versus stem/progenitor cells (Sigal et al., 1995; Petersen et al., 1999; LeGasse et al., 2000; Theise

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et al., 2000; Dabeva et al., 2000) to treat patients with tissue/organ dysfunctions have indicated that stem/ progenitor cells are preferred. Adult (mature) cells will integrate into tissues and can survive but usually demonstrate little if any ability to expand. By contrast, stem/progenitor cells show remarkable abilities to divide in vivo when under conditions of a “cellular vacuum” in the recipient in which a significant percentage of the mature cells are lost due to (1) drugs (Laconi et al., 1998; Gagandeep et al., 2000); (2) viral infections (Bilir et al., 1998, 2000); (3) oncogenic insults (Grisham and Thorgeirsson, 1997; Shafritz, 2000); or (4) aberrant genetics (whether occurring naturally or artificially induced) (Sandgren et al., 1991; Overturf et al., 1996). Thus, inoculation of progenitor cells into normal tissues is associated with integration and rapid maturation into adult cells (Sigal et al., 1995). Inoculation of cells into organs or tissues associated with injury-induced cellular vacuum results in extensive hyperplasia followed by maturation (Sandgren et al., 1991; Gupta et al., 1999; Braun et al., 2000; LeGasse et al., 2000). The implications for strategies in the use of stem cells have relevance for all future applications of stem cells: expansion ex vivo (e.g. in bioreactors or for protein manufacturing) will require relatively purified progenitor cells under defined culture conditions. Clinical cell therapy strategies will differ between recipients with organ/tissue failure and those with inborn errors of metabolism. Patients having a tissue associated with massive cell loss (preferentially of the mature cells) would be predicted to require fewer donor cells, since the donor cells should expand under those in vivo conditions. By contrast, recipients with inborn errors of metabolism and with normal cell numbers but with an aberrant function(s) should require high numbers of donor cells that should demonstrate limited growth and rapid differentiation. Immunological Issues Although research into the immunological issues with respect to stem cells is still in its infancy, it is predicted that immunological rejection is likely to be alleviated or eliminated by using stem/progenitor cell populations. Stem cells have been found to have minimal immunogenicity due to the complete absence or low levels of MHC antigens (Wang et al., 2003). Although the stem cells’ descendants should certainly become immunogenic, their extraordinary expansion potential and cryopreservability enable the samples to be tissue typed, facilitating the matching of donor to recipients. Alternatively, the recent studies from Nelson Chao and associates and others suggest that cell therapy coupled with bone marrow transplantation could result in only a transient need for immunosuppressive drugs (Benedetti et al., 1997). Tissue-Specific Gene Expression Is or Can Be Lineage-Position Dependent Tissues have long been known to be heterogeneous in expression in their specialized functions and each tissue demonstrates discrete patterns (Traber et al., 1988; Gumucio, 1989; Gebhardt, 1992; Sigal et al., 1992). Numerous recent studies suggest that whereas some genes may be expressed throughout the lineage of a tissue, especially its common genes, other genes are expressed at discrete stages of the lineage, for example, only in the cells of the stem cell compartment (Wang et al., 2003). Based on the maturational lineage models, the heterogeneity can be interpreted as a combination of distinct microenvironments and maturational changes within the cells. The net sum of the two results in lineage-position dependence of gene expression, resulting in “early,” “intermediate,” and “late” gene expression. These findings are relevant for those circumstances in which specific gene expression is desired or needed: one must isolate and utilize the lineage stages that optimally express those genes. Representative Future Uses of Stem Cells For the sake of brevity, this section will focus on liver only and present discussions of applications utilizing ES cells, multipotent stem cells, and HpSCs. The themes and strategies discussed here with respect to liver are applicable to strategies with most tissue types.

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Cell Therapies Cell therapies are ones in which suspensions of cells are injected into the blood stream or into a tissue or organ with expectations that the cells will repair any injury process that has occurred or that is ongoing. The expectations are that cell therapies are likely to replace or supplement much of organ transplantation within 5–10 years. This makes the organ transplant surgeons central to the development of the clinical programs and bodes well for an expansion of the number of patients who can be treated. The cell therapy protocols, even the demanding ones, are far easier on patients; they involve fewer side effects and offer the likelihood of lower or even minimal immunological complications relative to those experienced by organ transplantation. In addition, the procedures are faster, can be done on infants or frail patients, and can be performed for a small fraction of the costs of routine organ transplants. Very importantly, the procedures do not have to be done at tertiary care centers; they can be done at primary care clinics almost anywhere in the world, opening the door to treatment for patients in developing countries. These facts are driving forces in the extraordinary interest in cell therapies by countries like China, India, and Korea. Determined stem cells, but not ES cells, are the first forms of stem cells to enter into clinical trials of cell therapies, due to their restorative potential and absence of evidence for tumorigenicity (Gluckman et al., 1989; Mayani et al., 1992; Lian et al., 1999). The early data from these trials are very encouraging for the future for progenitor cell therapies. The problems with determined stem cells include: (1) identification of tissue sources, a special problem for organs that until now have derived only from fetal tissues or from brain-dead-but-beating-heart donors; (2) the need for the development of purification schemes for isolation of the cells; and (3) defining the ex vivo expansion and differentiation conditions. The use of ES cells for cell therapies is years away. Estimates of over 10 years have been made for how long it will take to overcome the difficulties with tumorigenicity of ES cells. Yet if that risk can be eliminated, then the potential for ES cells is enormous. Bioartificial Livers Bioartificial livers are emerging as potential therapeutic approaches for acute or chronic liver failure or inborn metabolic disorders (Macdonald and Wolfe, 1999; Xu et al., 2000). Clinical trials are ongoing for bioartificial livers (Mullon and Solomon, 2000). These assist devices are projected to be used transiently to rescue patients from acute organ failure with the hope that the patients organ(s) can recover from an acute crisis such as a drug overdose. At present, these patients must be rescued by routine medical intervention or with organ transplants; the former is often unsuccessful, and the latter has severe, life-long consequences such as chronic immunosuppression. All of the various forms of bioartificial organs being developed today consist of: cells inoculated into a bioreactor, most commonly one of the hollow fiber designs, and providing a three-dimensional space for cells with adequate mass transfer of essential nutrients, factors, and oxygen, and removal of metabolic wastes; and a microenvironment comprising a nutrient medium with serum and/or purified hormones and growth factors, and pumped through the cell compartment either directly or via the hollow fibers in direct contact with the cells within the bioreactor’s cell compartment (Knazek et al., 1972; Wolf and Munkelt, 1975; Jauregui, 2000). The extant bioreactors work well for cells that float, such as the hemopoietic cells, in which Starling flow can adequately provide gases and nutrients but are limited in their utility for adherent cell types derived from solid tissues or organs. Adherent cell types bind to the hollow fibers and deposit their extracellular matrix and cellular proteins, resulting in occlusion of the pores and, therefore, blocking transfer of nutrients across the hollow fibers. The limited capacity of mature cells to grow is equally important as a variable for the potential of bioreactors, since the bioartificial organs must be seeded with sufficient cells to provide the requisite cellular mass needed for the patient. The sourcing of some types of cells is extremely difficult.

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The tremendous ability of stem cells to expand and differentiate makes them candidates as the seed material for bioartificial organs. Indeed, the ease of obtaining some of the human ES cells, their ability to proliferate, and the ability to lineage restrict them into some adult fates makes them the ideal seed material for bioartificial organs. Clearly, in a bioreactor device, the inability to achieve 100% fidelity in commitment to a given adult fate is irrelevant and is greatly outweighed by the extraordinary ability to generate vast numbers of a specific adult cell type very rapidly. It is likely, therefore, that ES cells will be favored as a cell source for bioartificial livers, since they demonstrate far greater ease of sourcing, and their tumorigenic potential is eliminated as a concern by the fact that the cells are in a bioreactor device. Drug Testing and Protein Manufacturing with Liver Cells Pharmaceutical companies have long desired readily available human liver cells for drug testing or protein manufacturing given the often human-specific responses to some drugs and the modifications of proteins that can occur uniquely in some human proteins. Although there is constant use of human hepatic cell lines, these model systems are flawed in being often tumorigenic and/or aberrant in their responses due to genetic mutations, or in response to inadequate cell culture conditions. Therefore, the goal is to have human cells behaving as normally as possible. Stem cells are projected to ease the limitation in supply of normal human cells through their enormous expansion potential. Transition to this strategy awaits the development of optimal methods for differentiating the stem cells to their mature fates under ex vivo conditions. Some projected uses for stem cells by industry, such as protein manufacturing of human-specific proteins, will most easily be done with ES cells used in unrestricted form for proteins produced by all cell types and in lineage-restricted forms for proteins generated by specific cell types. The reasons for ES cells being preferred are in their relative ease of expansion and differentiation that translate to ease of manufacturing constraints. By contrast, determined stem cells from diverse donors are likely to be favored in drug testing that is increasingly being tailored for specific genotypes. Gene Therapies The hope of gene therapies is to use the wealth of molecular biological techniques to correct gene defects such as diabetes. Although there have been some dramatic improvements for some diseases (e.g. severe combined immunodeficiency disorders or SCIDs), the ability to use gene therapies has proven problematic. The focus for some years has been to use “target injectable vectors” in which the gene is introduced to the tissue and is able to get into the relevant cells by means of molecularly tailored vectors. To date this approach has been very limited in its success, both because the ability to target given vectors has been difficult and because the expression in the tissues has proven transient. The only reproducibly successful forms of gene therapies have been those in which stem/progenitor cells have been isolated from the patient, modified (usually ex vivo), and the modified cells given back to the patient (Anderson et al., 1989). It is hypothesized that in the future, vectors targeted to the stem cell compartment of a tissue may prove as successful. Thus, the availability of stem/progenitor cells, ideally determined stem cells, from diverse tissues should translate into great potential for gene therapies that utilize those cells. Vaccine Production Lineage-dependent gene expression has ramifications for viral infections and vaccine production. A number of viruses (e.g. papilloma virus, hepatitis C) replicate in one lineage stage (e.g. the stem cell compartment) and then mature along with the cells so that they express mature viral proteins in later lineage stages (Kwong et al., 2001).

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Some viruses that have proven difficult to grow in culture may prove amenable to expansion in cultures of stem cells that are then differentiated. If so, the stem cell cultures should facilitate the development of novel vaccines and drugs for control of the pathogens. Research Stem cells will remain a topic of great interest to investigators for diverse fields including developmental biology, cell biology, molecular biology, and biochemistry. The ability to expand an undifferentiated cell ex vivo and then to differentiate it to some adult cell type offer unprecedented opportunities to dissect molecular controls on commitment and differentiation, on facets of cancer, on analyses of infections such as viral infections that may be lineage-position dependent, etc. Indeed, the extraordinary excitement and intense discussions now ongoing in the lay press are paralleled by excitement of scientists awaiting opportunities for discovery. Scientists engaging in ES cell research believe that the benefit to humanity from this endeavor may be so great that the research should be allowed to proceed in order to learn how to control the tumorigenicity issues of these stem cells. The assumption is that all forms of cell therapy that initially will be mediated by determined stem cells will eventually be doable with ES cells once the ability to control the tumorigenic potential of these stem cells is accomplished.

ETHICS AND LEGAL ISSUES IN THE USE OF STEM CELLS This section is unusual for a typical scientific review of a cell biological field. However, we feel compelled to include it, since stem cell biologists are facing legal and ethical hurdles everywhere, and there is variation on those hurdles depending on the country, the societal and cultural values, and the religions. General Issues Future clinical and commercial programs using stem cells face hotly debated controversies and legal issues. These center around two issues (1) sourcing of human tissue and (2) use of embryonic cells. 1. Sourcing of human cells and tissues involves ethics and laws that apply to dealing with people who have donated tissue or dealing with family members of a donor who has died. Sourcing of tissues or organs must be done within the rules and policies established for tissue and organ donation, rules that have been codified by laws and are described in detail in documents from organ procurement organizations (OPOs). For example, a donor or a donor’s family must agree to the donation and probable future usage of the cells (e.g. transplantation, research, industry). Also, if a disease is identified in the donated tissue, the families (or donors) are provided the information through established clinical channels only if the disease is one in which state and federal laws require registration with appropriate health divisions (e.g. HIV). There is ongoing discussion about whether the families should be alerted for various genetic diseases if those diseases might be expressed in other members of the family. A major legal issue is a financial one: in most countries, human tissues and organs cannot be bought or sold. Therefore, they must be obtained voluntarily from donors or with permission of the donor’s family and without a profit motive. The rules are identical to those for organ donation in which the costs of procurement and processing can be billed to recipients (or to the research or commercial group receiving the tissue, but the bill cannot include a line item for the organ/tissue/cells). 2. ES Cells: Fully accepted forms of stem cells are: multipotent stem cells and determined stem cells from adult tissues and ES cells from spontaneous abortuses. The focus of all controversies is for use of ES or determined stem

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cells from tissues of induced abortuses or from IVF-derived fertilized eggs. The extent of opposition to use of ES cells or of tissues from induced abortuses depends, in part, on religious or cultural beliefs with respect to opinions about when life begins. Somatic nuclear transfer (SNT) to derive ES cells that can be used in industry or for bioartificial organs are variably accepted. However, use of SNT to generate ES cells that are used clinically for cell therapies is forbidden for now, since so many aberrations are being found in SNT-derived cells, but use of SNT might be possible in the future if this problem is solved. There is universal opposition to cloning of humans by any method or use of non-lineage-restricted ES cells for clinical programs. As noted in the introductory overview, the non-lineage-restricted ES cells are known to be essentially 100% tumorigenic when injected in vivo and efforts to lineage restrict the cells to a safe lineage stage have not yet succeeded. Many of the debates are fueled by specific religious or cultural beliefs. These are discussed in more detail below.

Survey by Philosophy and Religion Modern international human rights principles attach significance to human beings and respect for their dignity. There is a long-standing debate in embryonic research because of the potential for the embryo to develop into a human being. The debate is ongoing in a religious as well as secular context. The opinions expressed by major religions influence the debates, because bioethics concerns itself with the fundamental issues of human life. Clearly there are disagreements as to what extent embryo research is compatible with religious beliefs and the sanctity of human life. The ethical legitimacy of performing research on stem cells derived from embryos rests largely on the status we attribute to embryos. If the embryo is viewed as a human being, then the manipulations it can be subjected to are limited to those that are allowed to a person. But if the embryo is viewed as a mere collection of cells, then the restrains are much fewer. Because of its capacity to develop into a functioning organism, an embryo has a unique status in biological terms; in this capacity it is distinct from other clusters of cells. The difference can be described as the embryo’s potential: the potential to become a human being. Does this potential imply that our ethical notion of valuing a human being should also apply to the embryo? If the embryo has the full membership of our human community, then it cannot be treated and used as a means to an end. It should be protected from destruction. A major subject of the debates is the potential of the embryo. The proponents of embryo protection argue that the embryo has the potential to be a person; therefore, it is wrong to prevent it from fulfilling its potential. Opponents of this view argue that the potential to become a human does not warrant it automatically the status of a human. They argue that ova and sperm are the components of a zygote which then becomes an embryo and yet we do not give to ova and sperm the same sanctity that we do the fertilized egg. If we do not assign fetal status to sperm and ova, why do we assign fetal and even human status to a cluster of cells newly derived from the sperm and ova? Moral philosophers have challenged the assumption that the embryo has the full status of a human being. They consider that a membership in the human community requires an ability to experience those features in life that defines value and meaning to life. From their biological perspective, individuality can be attributed to an embryo only after day 14 of fertilization; before that day, the embryo can be split into normal identical twins. Day 14 is also the time when the primitive streak appears (a band of cell along the caudal of the embryo); 40% of fertilized eggs do not reach this stage in development. The appearance of a nervous system or any feelings of sensation comes much later. These arguments form the core of the ongoing debate with little compromise possible on either side. These core issues are central also to various religions. Among the major religions there is a wide range of positions on the status of an embryo and the permissibility of using embryos in research. Some allow the use

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of an embryo for therapeutic purpose and for research. Others absolutely prohibit the use of embryo in such a fashion. Even within one religious tradition, the view of embryo research varies greatly. Both Islam and Judaism believe that the full status of a human being does not occur with fertilization. Another relevant feature of Judaism is that, the embryo outside of the womb, not unlike gametes, has no legal status unless it is to be implanted, which then gives it the potential for life. An embryo derived from IVF can be donated for therapeutic and research purpose, especially if the therapy is life saving. The sanctity and value of life is a central tenet of Judaism. Interestingly, the Muslim and Judaic religions do not object to the use of embryonic tissue, since they recognize a spiritual entity as a neonate surviving for at least 40 days at which time the “spirit enters the embryo.” Protestants are similar to Muslims and Jews in their belief that a fertilized egg cannot be equated with a human being (or with a far more advanced embryo). Some branches of protestant tradition consider that the human being is formed by a slow and gradual process. The essence or soul of the human may occur at a very late stage of embryonic development. Yet Protestant theology is very diverse, and there are those who are strong in their opposition to research on embryos and others who are very accepting of it. Therefore, it is difficult to find a single authority who can represent the entire religion. Roman Catholics are among the most fierce opponents to the use of embryos for therapy and research. Their view is that a human being comes into existence both physically and spiritually at the precise moment of fertilization. Therefore, an embryo is considered to be a unique human individual having all the rights to its own life. An embryo must be given the opportunity to develop into a mature organism. Because of this view of human life, Catholic tradition considers it of the utmost importance to strictly control the fertilization of ova in vitro. Consequently, it is impermissible to utilize supernumerary embryos in therapeutic and research purpose. The “right to life” groups object to any use of the cells for any purpose (especially when a profit motive may be involved, even if indirect) because they consider it immoral to use any fertilized egg or cells descendent from that egg (life begins at conception). These beliefs affect whether there will be acceptance of the use of IVFderived fertilized eggs. Those opposing their use worry that the descendent cells might be aberrant and/or that one should not “manipulate life.” Shintoism involves beliefs of the sanctity of ancestry. These beliefs result in attitudes resulting in an avoidance of cell therapies, bone marrow transplantation, or organ transplantation, since it would involve transfer of cells or an organ from someone of different ancestry into a recipient. They believe that this would “pollute” their own ancestry both in this life and in the next. Therefore, cell therapies (e.g. bone marrow transplantation) or organ transplantation are permissible among family members but not for people outside of the family. By contrast, there is wide-spread acceptance of bioartificial organs that might transiently supply organ function to a family member with severe organ dysfunction. These beliefs are facets of the enormous focus on the development of bioartificial organs in Japan and other countries of the Far East. Survey of Opinions by Country (Table 20.4) United Kingdom For the last 10 years, British researchers have been allowed to use spare embryos from IVF clinics and to create embryos for research. A 1990 law limits such research to infertility, contraception, and congenital diseases, but in 2001, the Parliament relaxed the limitations on human embryo research to include understanding human disease and development of cell-based therapy. Since 1990, 53,000 human embryos have been used in research in the UK. Under the Human Fertilization and Embryology Act (HFA Act) a license is needed for creating embryos in vitro or performing research on embryos.

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Table 20.4 A list of Permitted and Banned sources, by country Country

Permitted source

Banned source

Canada

Left-over embryos. Human–animal Pre-existing ES hybrids. Creating cell lines embryos for research

Brazil/Peru

UK

France

Germany/ Austria Sweden

Denmark/ Finland

Norway

Comments

Goal

Protestant majority and catholic minority No storage and manipulation of embryos Formation of Advancing the nervous system treatment of on 14-day infertility and embryos congenial disease. Increase knowledge of embryonic development and of disease

License is needed. Embryos older Creation of than 14 days embryos for the derivation of stem cells. Creation of embryos by SNT. All embryo research must terminate on day 14 Left-over embryos Cannot create embryo for stem cell research ES cells produced No supernumerary Nazi era eugenics in other embryos and Catholicism countries Left-over embryos. Has the most ES No explicit ban cell lines on creating meeting Bush’s embryos for funding research criteria (24) Left-over embryos. Embryos created Embryo research for research must terminate on day 14 ES cell lines produced in other countries

Fertilized eggs

Belgium, Greece, and Luxembourg Italy Health minister Production of supports the embryos for cloning of research human embryos for derivation of stem cells

Religion

Anglican

Improve IVF

Catholic

Cells can be examined as part of IVF Improve infertility treatments

Protestant and Catholic

Improve IVF techniques and a broad range of medical research Law resembles Improve infertility that in Germany treatments but it is currently under revision No legislation concerning human embryo research Disease therapy

Catholic

Hepatic Stem Cells: Lineage Biology and Pluripotency 371

Table 20.4 (Continued) Country

Permitted source

Ireland

All human embryo research ES derived from Nuclear transfer Improve IVF spare embryos (3-year techniques and moratorium). disease therapy Human–animal hybrid embryos ES cells derived Placing cloned New guidelines will Disease therapy from spare embryos in allow labs to and infertility embryos. uterus. Research start studies on treatments Cloning embryo on cloning building tissues in vitro humans, from ES cells creating sperms/ova Stem cells from Embryonic tissue Building a state run Treatment and umbilical cord is generally stem cell prevention of and afterbirth banned complex. diseases Including stem cell bank, transplant center and engineer center. Regulation is lax Stem cell research Government spend is on going 500 million dollar last year to promote research in private sector Embryos less than Cloning of somatic 14 days, cells (nuclear left-over from transfer). Hybrid IVF. Adult stem of human and cells animal No law regulating 1999 prohibits Talmudic law Israel is in the stem cell reproductive places distinct forefront of stem research and cloning for value on embryo cell research embryo 5 years only after destruction for implantation and stem cell considers it to research is achieve “formed” allowed human only after 40 days Aborted fetuses Human embryos Researches have that come from to be compatible miscarriages or with Islamic therapeutic sensitivities; abortions Jeddah BioCity is under construction

Australia

Japan

China

Singapore

South Korea

Israel

Saudi Arabia

Banned source

Comments

Goal

Religion

Catholic Protest majority

Shinto

Communist

Mixed

Christianity, Buddhism

Judaism

Islam

372 CELLS AND TISSUE DEVELOPMENT

Continental Europe Legislation for stem cell research is most restrictive in Germany and Austria. The Nazi era has haunted German consciousness for many years; any hints of eugenic research are strongly repugnant to the German people. Therefore, research on human embryos is forbidden. In Austria, egg and embryo donation is prohibited; cells can only be used to test the viability of embryos during IVF procedures. No supernumerary embryos are available because the number of eggs fertilized in IVF is limited. In both countries, the sole purpose of embryo production is to start pregnancy. German embryo protection regulations forbid any research that harms the embryo. However, in 2002, Germany passed a law that allows the importation of ES cell line produced in other countries for research. Laws in Scandinavia and France are more permissive. Researchers can derive stem cells from “spare” embryos. In Finland, licensed agencies are allowed to carry out medical research using embryos up to 14 days after conception. All human embryo research is prohibited in Norway. Currently, Belgium, Greece, Italy, and Luxembourg have no guidelines and legislation regarding human embryo research. The Italian National Committee on Bioethics has opposed human cloning, but has no consensus on the use of supernumerary embryos and therapeutic cloning. The United States and Canada The United States controls the ES cell research via the purse. It prohibits federal funding of embryo research; unless it is confined to the approved existing ES cell lines. However, there is no federal level or state level of control over private research. In 2000, the Bioethics Advisory Committee recommended that federal regulation should permit research into ES cells using supernumerary embryos. However, it remains opposed to therapeutic cloning and nuclear transfer. National Institutes of Health (NIH) issued guidelines for where federally funded researcher can engage in these investigations. One condition was that no embryo could be destroyed for the purpose of generating stem cells. In the wake of federal opposition to the stem cells, there has been growing support at the state levels. California has passed a major bill offering funding to stem cell research; the funding from this bill has yet to reach researchers only because of opponents putting forth legal challenges. Once the challenges are overcome, California will likely become a major mecca for stem cell investigators. Similar bills have been passed in other states or are under consideration in state legislatures. Similar to the Scandinavians, Canadian law permits the derivation of stem cells from supernumerary embryos. However, research on embryos after the 14th day of existence is prohibited, and consent of the couple who supply the embryo is required. Canadian investigators are also organizing toward major efforts in stem cell research. Stem cell centers have been formed in multiple cities, most notably in Toronto, and diverse funding sources are being established to help further investigations. South America Brazil has a law prohibiting storage and manipulation of human embryos. Peru specifically prohibits the fertilization of ova for purpose other than reproduction. The Peruvian policy is based on the concept that life begins at the moment of conception. Asia and Pacific Rim Countries

Australia’s guideline for ES cell research permits the research of human embryos that results in their destruction under exceptional circumstances and with the approval of an Institutional Ethics Committee. The exceptional circumstance includes significant advances in therapeutic technology. Japan’s Parliament enacted the Human Cloning Regulation Act in 2001 that recommends allowing ES cell research but prohibits human cloning. Many Japanese people reject forms of cell therapy and even of organ

Hepatic Stem Cells: Lineage Biology and Pluripotency 373

transplantation except among family members due to tenets of Shintoism; there is a reverence for ancestry and revulsion at the thought of acquiring cells or an organ from someone of a different ancestry. Thus, the focus on the use of stem cells in Japan will involve cell therapies among family members, bioartificial organs, and industrial uses of the cells. Some of the most active research groups doing studies on bioartificial organs are in Japan supported by large government grants. China has some of the world’s most liberal policies toward tissues from abortuses and with respect to stem cells. Tissues from induced abortions are used routinely for forms of cell therapy that, unlike in the United States, have been ongoing for several years. In 2001 the Chinese Ministry of Health announced that it would allow closely monitored ES cell research for treatment and prevention of disease. The Chinese government is building its first state run stem cell complex. Stem cell research is being conducted in Singapore. In late 2001, the government appointed a panel of experts on philosophy, science, and law to study the ethical issues regarding ES research. Singapore is the base for huge funding for biotechnology companies in the Asia Pacific, and a major focus of some of those companies and academic research groups is on stem cell research. Thus, it is likely that it will become a center for stem cell research and stem cell policy in the world. South Korea Ministry of Health put forward guidelines for ES cell research in December 2001. The guideline bans the cloning of somatic cells, but embryos less than 14 days old (left-over from IVF) will be permitted for research. The Middle East Countries Israel has extremely liberal laws with respect to stem cells, whereas most of the Muslim countries are far more restrictive. Policies are largely a reflection of the dominant religions of the countries, that is, Judaism versus Islam. Strictly speaking, Talmudic (Judaic) and Islamic law condemn human cloning because they believe it does not show respect for the sanctity of life, especially given the risks in giving birth to a human with severe flaws. By contrast, use of IVF-derived fertilized eggs and ES cells for therapeutic and research purposes is acceptable. Both Talmudic and Islamic law believe that embryos before the 40th day of fertilization are acceptable for therapeutic and research purposes, because embryos less than 40 days are not “ensouled.” There is divergent opinion on ethics with respect to the use of tissues from abortuses. United States Federal and State Law The federal government has taken some action without new legislation. Federal funding for research on human cloning has been barred with policies issued during the Clinton administration. As a response to Dr. Richard Seed’s declaration to clone himself, the FDA announced that it had regulatory jurisdiction over human cloning under existing federal statutes. It is debatable whether the FDA has jurisdiction, and it would be a question for the courts. However, at this point, no lawsuits challenging its authority are known. Given that the creation of an embryo is required for the establishment of ES cells, non-reproductive cloning is also affected by federal rules on embryo research. This issue has been extremely controversial at the federal level with regard to federal funding for such research. A 1994 National Institutes of Health Human Embryo Research Panel would have allowed the use of human embryos for federally funded research including, with specific limitations, the production of embryos for this purpose. The report was not adopted as policy by NIH. Congress banned “the creation of a human embryo and embryos for research purposes.” The National Bioethics Advisory Commission issued its report, “Ethical Issues in Human Stem Cell Research,” in January 2000. This was followed by release in August 2000 by NIH of its Guidelines for Research Involving Human Pluripotent Stem Cells. The guidelines allowed NIH funded investigators to conduct research on ES

374 CELLS AND TISSUE DEVELOPMENT

cells obtained from private sectors, provided the source is supernumerary embryos produced to treat infertility and that are donated without compensation. Federal funding for the creation of stem cells from abortions, their derivation from embryos, and the production of embryos to serve as sources of stem cells, either by sexual combination or by nuclear transfer for research, were prohibited. The NIH guidelines were in turn limited by President Bush’s 2001 decision to allow federal funding for ES cell research only for cell lines established before the date of his announcement. This would prohibit federal funding for research with ES cells produced through cloning. State Law It is important to note that the rules above apply only to research that involves federal funds; privately funded research on non-reproductive cloning is not affected by these policies although it would, at some point, be regulated by the FDA. This limitation was highlighted by the work by Advanced Cell Technologies (California) in using nuclear transfer technology and human eggs to produce what it called early embryos. Other than restrict federal funding, the national government has left the authorization of non-reproductive embryo research to the discretion of each state. Only five states have, so far, passed statutes prohibiting human cloning: California in 1997, Michigan, and Rhode Island in 1998, Louisiana in 1999, and, most recently, Virginia in 2001. The California statute, the first one adopted, bans reproductive cloning for a period of 5 years. It does not deal with non-reproductive cloning, but is restricted to situations where a cloned embryo is implanted in a woman’s uterus. The Rhode Island and Louisiana statutes were modeled generally on California’s. The Michigan statute is much different. It bans reproductive and non-reproductive cloning and contains no “sunset” date. Virginia’s statute is similarly broad, banning completely the transfer of any human cell nucleus into oocytes. Several other states have passed legislation barring state funding for human cloning research or prohibiting such research at state institutions. More than twenty states have laws banning or restricting research with human embryos. These laws were passed many years ago in response to concerns expressed largely by “pro-life” groups. These statutes could prohibit certain forms of non-reproductive cloning. They could also be construed to prohibit reproductive cloning at least at its early, experimental, and research stages, in an effort to avoid regulating IVF and other forms of assisted reproduction. However, many of these statutes expressly state that they do not govern research that aims to result in the birth of a living child.

SUMMARY With continued basic research on stem cells, the prospects for growing opportunities for clinical and commercial programs based on this field are seemingly limitless. It is this realization that is proving the driving force and motivation for increasing numbers of investigators pursuing research into stem cell biology. Many of the most hotly debated issues are likely to become more subdued with improvements in isolation and utilization of forms of stem cells that can be obtained from postnatal tissues and as forms of therapy derived from stem cells make their way into clinical programs. The driving force for acceptance of stem cells is already in governments at the state and local levels with the growing number of citizens realizing the possibilities of therapies for diseases and conditions that have long been problems for mankind. ACKNOWLEDGMENTS Funding derived from NIH grants (DK52851, AA014243, IP30-DK065933), a Department of Energy Grant (DE-FG02-02ER-63477), by a Whitaker grant, and by a sponsored research grant from Vesta Therapeutics (Durham, North Carolina). The paragraph on stem cells from umbilical cords and from adipocytes was written by Dr. Joseph Ruiz, Director of Research at Vesta Therapeutics (Durham, North Carolina). We thank Mara Gabriel, a professional writer for the pharmaceutical industry, for her extensive editing of the manuscript.

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21 Skeletal Muscle Stem Cells Jason H. Pomerantz and Helen M. Blau

INTRODUCTION Somatic stem cells have been described most prominently in tissues that exhibit high rates of turnover, such as the hematopoietic and gonadal tissues. Skeletal muscle, by contrast, is a relatively stable tissue under normal conditions. However, damage in the form of exercise, trauma, or disease elicits a remarkable regenerative response in muscle that is characterized by augmentation or replacement of significant degrees of muscle mass. It is in such situations that the need for stem cell function in skeletal muscle becomes apparent. Skeletal muscle cells are syncytial structures comprised of organelles including the specialized endoplasmic reticulum and t-tubule system that are involved in transmitting excitatory signals to coordinate muscle contraction, as well as bundles of fibrils and contractile proteins in organized units called sarcomeres. All of these structures are generated via instruction from hundreds or thousands of nuclei spanning the length of each muscle cell. During development or in response to injury, nuclei are incorporated into muscle fibers by the fusion of mononuclear myoblasts with each other or with existing larger syncytia. In adult animals, myoblasts are derivatives of muscle stem cells and may be considered analogous to the transit-amplifying cells described in other somatic tissues such as the skin or blood. This chapter focuses on the cells responsible for maintaining, repairing, and regenerating skeletal muscle. Research over decades addressing the nature of skeletal muscle stem cells has generated a solid body of knowledge regarding their origin, morphological and physiological characteristics, and relation to disease. Recently, technological advances including genetic labeling, confocal microscopy, and the fluorescence activated cell sorter (FACS) have enabled a new wave of investigation that has brought us closer to understanding the more intricate molecular mechanisms of muscle stem cell function. In addition, elegant techniques for extracting cells from muscle tissue have allowed direct observations of key stem cell functions. The definition of a stem cell continues to evolve. Indeed, identification of a cell as a stem cell remains difficult due to a need for rigorous experimental verification coupled with problems of semantics and long-held dogma. First, physical criteria are required that endow stem cells with tangible qualities that can be reliably used to detect, isolate, and manipulate them. Second, the assignment of necessary functional characteristics to stem cells is crucial. Characterization of specific cells as stem cells requires comparison with the classic stem cell definition: stem cells must self-renew, yield progeny that are multipotent, and regenerate significant amounts of tissue. Notably, muscle stem cells diverge from this definition as they are only known to give rise to one specialized cell type and are therefore unipotent. Skeletal muscle stem cells have been identified based on physical properties and function. Muscle stem cells have a characteristic anatomical location and morphology, express a specified set of proteins, and are capable of both self-renewal and the production of mature, functional muscle tissue. We suggest that designation of a cell as a “muscle stem cell” must meet the criteria of clear identification of the cell of interest, followed and a

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demonstration of the ability of the cell to both generate functional stem cell while at the same time generating “effector cells” that build and repair muscle in significant quantities.

SKELETAL MUSCLE REQUIRES STEM CELL FUNCTION AFTER INJURY OR IN DISEASE For muscle to respond to acute needs and repair itself after injury, stem cells are required (Studitsky, 1964; Carlson, 1986). This is in contrast to blood, for example, that requires stem cells to be active continuously in order to meet the need for extensive cellular turnover even in the absence of stress. Supported by studies of telomere length, muscle is thought to undergo relatively little turnover, except when stressed by exercise or other types of tissue damage (Decary et al., 1996). In such cases, muscle is repaired or augmented by the addition of new myonuclei to existing fibers or by the generation of new fibers (Mastaglia et al., 1975; Sloper and Partridge, 1980; Irintchev and Wernig, 1987; Allen et al., 1995; Kadi and Thornell, 2000; Charge and Rudnicki, 2004). In the classical regenerative response, skeletal muscle damage and disruption of muscle fiber integrity lead to infiltration of inflammatory cells, followed by satellite cell activation, proliferation, and integration into damaged or necrotic fibers by fusion (Grounds, 1998). The activation of satellite cells and the regenerative response is inhibited by irradiation that damages satellite cells, but not muscle fibers (Dmitrieva, 1960; Popova et al., 1968; Rosenblatt and Parry, 1992; Rosenblatt and Parry, 1993; Adams et al., 2002). Finally, it has been suggested that myofibers may be generated de novo during normal growth or regeneration (Mazanet and Franzini-Armstrong, 1986; Grounds, 1991). Thus, there is a need in adults, for stem cells in skeletal muscle that are capable of regenerating tissue that has suffered injury. Certain disease states result in extensive muscle degeneration and regeneration, presumably requiring stem cell-mediated contribution of nuclei (Lipton and Schultz, 1979). A clear histological feature of dystrophic muscle is fiber regeneration as evidenced by centrally located myonuclei. Satellite cells are increased in number and activation state as evidenced by nuclear euchromatin content (Mauro, 1979). This histological evidence for regeneration is gradually lost in patients who suffer from Duchenne muscular dystrophy (DMD), as fibers are increasingly replaced by fibrotic tissue. Muscle cells from dystrophic patients have defects in growth and division in culture that are not primary effects of the disease-causing dystrophin mutation, but rather result from a reduction in replicative capacity gradually acquired after excessive demand (Blau et al., 1983; Webster and Blau, 1990; Heslop et al., 2000). Exhaustion of stem cell reserve may cause the onset of clinical symptoms in DMD. MUSCLE STEM CELL CRITERIA In order to classify any particular cell as a muscle stem cell, certain criteria must be met that fit the accepted definitions of other types of stem cells. While definitions necessarily place constraints that may make it difficult to expand our concepts of stem cells, there are criteria that are essentially universally agreed upon. First, there must be clear, unambiguous characterization of the stem cell entity. Such a characterization may be in the form of clear morphological attributes or anatomic location. Anatomic location was the first criterion used to define muscle stem cells: mononuclear cells juxtaposed to the sarcolemma and beneath the basal lamina (Mauro, 1961). More recently muscle stem cells have begun to be characterized by expression of patterns of surface markers such as c-met, M-cadherin, syndecans 3 and 4, CD34 and the absence of CD45 and Sca1 (Irintchev et al., 1994; Allen et al., 1995; Cornelison and Wold, 1997; Beauchamp et al., 2000; Cornelison et al., 2001; Montarras et al., 2005). Surface marker characterization is an active area of research that continues to evolve and inform our understanding of the nature of a stem cell. A case in point is the hematopoietic stem cell (HSC) that has been characterized and designated to be a stem cell based on the expression of a certain set of surface markers. However, these markers are not conclusive. When used to isolate HSCs prospectively, only a fraction of the cells exhibit stem cell function, the ability to reconstitute the blood when transplanted as single cells into

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lethally irradiated animals (an assay not yet available for muscle stem cells). This may reflect technical problems with transplantation of single cells, but it is also possible that the definition of HSCs by surface markers is incomplete and yields a population that is enriched for functional stem cells. In addition to surface markers that are useful for isolation, muscle stem cells express a number of genes that are expressed intracellularly and are therefore not generally suitable for isolation of stem cells, but are notable for their developmental and or functional relevance. Examples include the myogenic basic helix-loop-helix proteins Myf-5, MyoD, and the paired box transcription factors Pax-3, and Pax-7 (Tajbakhsh et al., 1996; Cornelison and Wold, 1997; Seale et al., 2004). The pattern of expression of skeletal muscle stem cell genes differs among tissues, for example head, diaphragm, and limb, suggesting that satellite cells have heterogeneous phenotypes and functions. A correlation of phenotype with function is an area of intense investigation and the relationships of cells of different phenotypes to one another remain to be determined (Ordahl and Le Douarin, 1992; Rantanen et al., 1995; Rosenblatt et al., 1996; Pavlath et al., 1998; Yoshida et al., 1998). Once a candidate muscle stem cell is characterized based on markers, it must be shown to function as a stem cell. First and foremost, it must be capable of extensive regeneration of the tissue that it serves. A prerequisite of each individual stem cell is the potential to give rise to a significant portion of the tissue. For example, a single HSC is capable of replacing the entire cellular blood compartment of the body. Similarly, a muscle stem cell must be capable of contributing a significant number of nuclei to replace or replenish physically associated skeletal muscle tissue after stress or injury, a criterion unmet by the majority of muscle stem cell candidates. Muscle stem cells must also be capable of self-renewal. Stem cells, by all definitions, not only give rise to differentiated progeny that provide tissue function, but for each division they maintain the stem cell pool. By contrast to some other types of stem cells the issue of self-renewal of muscle stem cells is somewhat simpler. In tissues in which stem cells are oligopotent (i.e. give rise to a few different cell types) self-renewal is required to maintain oligopotency, or multipotency in the case of embryonic stem cells (ES). For oligopotent stem cells, the stem cell gives rise to progenitors that are still capable of extensive division, but become less potent in terms of the ability to give rise to multiple types of cells. By contrast, under normal conditions muscle stem cells are only thought to give rise to skeletal muscle. Therefore they are unipotent stem cells. In muscle, division of the stem cell gives rise to myoblasts that are capable of both differentiating into mature muscle and extensive division to produce more myoblasts. The concept of quiescence is crucial to defining stem cells and clarifies the difference between muscle stem cells and myoblasts. Stem cells in most organs divide rarely and when they divide, give rise to a proportion of progeny capable of returning to quiescence, presumably through asymmetric division. As a result, one hallmark of stem cells is the generation of progeny that are labeled with BrdU long term, designated as “label retaining cells.” The maintenance, release and re-acquisition of quiescence are crucial characteristics of muscle stem cells that underly self-renewal.

HISTORICAL PERSPECTIVE ON MUSCLE STEM CELL BIOLOGY In the late 1950s and early 1960s, investigators began to actively question the source of mononucleated myoblasts required for muscle regeneration. Various hypotheses were entertained such as the possibility that myoblasts arise from peripheral fiber nuclei becoming re-cellularized after injury, still thought to be a major mode of muscle regeneration in urodeles. As an alternative, myoblasts were postulated to arise from “satellite” cells found adjacent to muscle fibers, by Alexander Mauro in 1961. In mammalian skeletal muscle, the satellite cell was quickly recognized as a stem cell candidate based on light and electron microscopic imaging of tissue explants and cultured muscle fibers (Mauro, 1961; Moss and Leblond, 1971). As described by Mauro, electron

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microscopic imaging of satellite cells in the tibialis anticus muscle of the frog showed them to be mononucleated cells situated on the periphery of myofibers, “ ‘wedged’ between the plasma membrane of the muscle fiber and the basement membrane”…“On the inner surface, the plasma membrane of the satellite cell is in apposition with the plasma membrane of the muscle cell.” Mauro noted the very high nucleus to cytoplasm ratio, such that the cell takes on “the shape of the nucleus.” The existence of satellite cells was confirmed by electron microscopists Swann, Peachey, and Palade in other skeletal muscles of the frog and rat. Intriguingly Moore and Palade recognized their absence in electron micrographs of cardiac muscle, a tissue not shown to regenerate significantly in mammals (Mauro, 1961) and for which a definitive stem cell has yet to be identified. It was recognized that after injury or during culture of muscle, the satellite cells enlarge, nuclei exhibit chromatin decondensation, and the cells divide to form additional mononucleated cells (Mauro, 1979; Bischoff, 1986). Furthermore, the suspected progeny of satellite cells were observed by light microscopy to fuse to form myotubes in culture. These early studies, like those used today, relied on the use of enzymes to digest the basal lamina of muscle fibers in order to release the satellite cells from their anatomical compartment (Bischoff, 1974; Yablonka-Reuveni et al., 1987). Reportedly, myogenic cells were not obtained without disruption of the basal lamina. Muscle regeneration was also studied using minced muscle samples. After mincing, investigators using electron microscopy observed viable mononuclear cells beneath the basal lamina adjacent to degenerating muscle fibrils and non-viable appearing, pyknotic myonuclei. In order to distinguish whether these viable mononuclear cells, presumed to be the source of regenerative muscle cells, were derived from satellite cells or myonuclei, tritiated thymidine was used to label nuclei undergoing DNA replication. When the nucleotide was delivered continuously to muscle during development and the muscle tissue analyzed at later times after animals matured, only myonuclei and not satellite cells were labeled (Snow, 1978; Mauro, 1979). After muscle injury, the label appeared only in the pyknotic, non-viable myonuclei, whereas the viable mononuclear cells were not labeled, suggesting that cells responsible for muscle regeneration are not derived from myonuclei in muscle fibers. A converse experiment, in which the satellite cells were labeled by a pulse dose, revealed that viable mononucleated cells were labeled after injury, whereas myonuclei were not. These experiments provided support for the satellite cell as the mediator of skeletal muscle regeneration after injury. Notably, a study using electron microscopic imaging characterized additional cells located in muscle tissue that “invaded” fibers, gaining access to regions beneath the basal lamina. These invasive cells were hypothesized to be monocytic cells, morphologically distinct from satellite cells at the electron microscopic level and were not thought to be involved in regeneration (Mauro, 1979; Mazanet and Franzini-Armstrong, 1986). Remarkably, these findings of 20–40 years ago have stood the test of time and remain true even in the advent of more sophisticated analytic methods. A recent revival of muscle stem cell biology was spawned by technological advances. At the time of Mauro’s original studies of satellite cells, it was “virtually impossible to discern the cellular nature of this entity in the light microscope, as it appears to be indistinguishable from a peripheral muscle nucleus proper” (Mauro, 1961). Now, the increased resolution afforded by laser scanning confocal microscopy, in conjunction with immunostaining, greatly facilitates efforts to distinguish adjacent cells without the need for electron microscopy. Second, the FACS permits the isolation of particular cells from tissues based on the phenotypic markers they express. Third, genetic labeling using transgenes encoding beta-galactosidase or fluorescent proteins, or the use of chromosomal markers detectable by in situ hybridization has enabled the tracking of single cells over time. As a result, in transplantation experiments it became possible to follow putative muscle stem cells to assess their contribution to muscle regeneration (Ferrari et al., 1998; Blaveri et al., 1999; De Angelis et al., 1999; Heslop et al., 2001; LaBarge and Blau, 2002; Corbel et al., 2003; Doyonnas et al., 2004; Palermo et al., 2005). The results of these experiments have refueled the debate over which cells have the capacity to contribute to muscle regeneration. Another important question that labeling experiments have helped to answer relates to the developmental origin and potential of satellite cells. Viral infection to express beta-galactosidase in muscle stem cells in vivo

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demonstrated that individual muscle stem cells could give rise to progeny that participated in both fast and slow skeletal muscle fiber development (Hughes and Blau, 1992). Other studies have begun to define the transcription factors that control the development of satellite cells, facilitating studies to determine whether satellite cells arise postnatally and from non-muscle sources, such as the bone marrow. During development, skeletal muscle and presumably satellite cells are derived from mesodermal tissue comprising segmental, paired somites (Buckingham et al., 2003). Among the earliest genes thought to be involved in the specification of muscle tissue are the paired box transcription factors Pax-3 and Pax-7 (subgroup III). In certain muscles (diaphragm) these two genes are co-expressed, whereas in other muscles (limb), Pax-3 is absent, suggesting that they have distinct functions (Relaix et al., 2004, 2005). In mice null for the Pax-7 gene, the absence of satellite cells correlates with the failure to develop post-natal muscle (Seale et al., 2000). Subsequent studies have suggested a role for Pax-7 in muscle stem cell self-renewal (Olguin and Olwin, 2004; Zammit et al., 2004). In addition, Pax-3 appears to be capable of mediating some degree of muscle development in Pax-7 null animals. Nonetheless, whether in the absence of Pax-7, muscle is formed by satellite cells or interstitial cells is still a subject of debate (Kuang et al., 2006; Relaix et al., 2006). Until recently, the evidence that satellite cells are responsible for muscle regeneration was entirely circumstantial, as their location and observed proliferation in response to injury suggested their involvement. Transplantation of satellite cell derivatives resulted in a relatively low contribution to myofibers. Recently, however, the transplantation of relatively undisturbed satellite cells, still juxtaposed to their parent fiber and with an intact basal lamina into skeletal muscles of mdx mice, the mouse model of DMD, has afforded a potent experimental system for monitoring satellite cells. After transplantation, using this paradigm, satellite cells proliferate to give rise to substantial numbers of progeny that generate large clusters of myofibers expressing dystrophin, providing the most direct evidence to date that satellite cells are indeed muscle stem cells (Collins et al., 2005). In addition, in the same study, satellite cells were isolated from single fibers by mechanical trituration, reportedly preserving their ability to regenerate muscle significantly, by contrast to satellite cells isolated by enzymatic methods. Notably, if plated in culture for even a few days, satellite cells exhibit a markedly reduced potential to engraft in vivo (Montarras et al., 2005). Thus, significant refinements in isolation of muscle stem cells have set a high standard for assays of muscle stem cell function.

PUTATIVE MUSCLE STEM CELLS AND THEIR DEFINED CHARACTERISTICS Since the initial characterization of the satellite cell, a number of different entities have been investigated for their ability to function as muscle stem cells. However, none has reached the bar set by studies of the satellite cell. The issue of which cellular entity harbors muscle stem cell function is clouded somewhat by the fact that different “muscle stem cells” may give rise to one another. For example, while studies to date generally implicate the satellite cell as the key, if not sole, mediator of muscle repair and regeneration, other putative muscle stem cells may represent intermediates in the pathway to becoming a satellite cell. It is not known whether all satellite cells are formed and localized in their niches during embryogenesis, or whether new satellite cells may be formed postnatally. Studies of a satellite cell at a given point in time do not provide information about its spatial or temporal origin. Therefore, studies that demonstrate that the vast majority of muscle stem cell function resides within the satellite cell do not exclude significant contributions by other cells, but if these cells play a role it is mediated by the satellite cell, an important if not essential intermediate step. Non-satellite cells with intriguing putative muscle stem cell properties include whole bone marrow derived cells (BMDC), HSCs, muscle interstitial cells, mesangioblasts, and mesenchymal stem cells. These will not be discussed individually in detail here, because each example while demonstrating some of the properties of a muscle stem cell lags significantly behind the satellite cell in terms of meeting all the criteria. None except the satellite

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cell has been rigorously demonstrated to replenish a large amount of muscle tissue. Mesangioblasts and mesenchymal stem cells remain difficult to define by any prospective criteria. Both of these cell types are produced in vitro from tissue explants and cannot be predictably isolated based on morphological characteristics or surface markers (Schubert, 2003; Vaananen, 2005). Until this can be accomplished, it will be difficult for different investigators to study these populations with confidence that they are investigating the same cells, and their in vivo relevance will remain questionable. Muscle interstitial cells, on the other hand, have been relatively extensively characterized based on surface markers and Hoechst dye exclusion, exhibit some in vitro myogenic potential, but have not been shown to contribute significantly to in vivo muscle regeneration (Sherwood et al., 2004; Kuang et al., 2006). Whole bone marrow derivatives, including HSCs, can be characterized based on surface marker expression, and have been shown to contribute nuclei to skeletal muscle fibers (Ferrari et al., 1998; LaBarge and Blau, 2002; Doyonnas et al., 2004). However, the level of contribution of BMDC in terms of number is extremely low, and the extent of muscle gene expression by the nuclei of these cells after fusion remains uncertain. BMDC have also been shown to take up residence beneath the basal lamina in the satellite cell position and to express satellite cell markers (LaBarge and Blau, 2002; Sherwood et al., 2004). Thus BMDC fulfill many of the criteria for satellite cells. However, they have thus far fallen short of displaying muscle stem cell function, because they have not been shown to replenish a significant portion of muscle mass, or to self-renew as satellite cells. Definitive experiments to test these attributes remain to be published, but a gold standard has now been set by the elegant experiments performed using single fiber isolation (Collins et al., 2005). In principle, similar experiments could be performed using BMDC, since these cells become satellite cells in muscle.

MUSCLE STEM CELLS IN THEIR NATURAL ENVIRONMENTS IN VIVO The importance of efforts to study stem cells in their natural environments in vivo must be emphasized. Culture artifacts likely form the basis for much of what may be envisioned as “stem cellness,” such as extensive proliferation without differentiation. The ES cells are a case in point. ES constitute a population of cells that can be expanded in culture indefinitely while remaining multipotent. Yet, in vivo ES cells do not exhibit significant expansion, since when they comprise the inner cell mass of the blastocyst their division must be very tightly controlled. Another example is the quintessential somatic stem cell, the HSC, which has not been reproducibly demonstrated to proliferate in vitro without differentiating, unless exposed to artificially high intracellular levels of Hox B4 (Antonchuk et al., 2002). Indeed, the best evidence that the HSC is a stem cell is the demonstration that a single cell transplanted into a lethally irradiated mouse can reconstitute all of the cells of the blood, a feat which requires both expansion and differentiation. Thus, recapitulation of the in vivo microenvironment in vitro remains a challenge in the study of stem cells behavior. Accordingly, in the case of muscle stem cells, interpretation of in vitro experiments must be tempered with the understanding that relevance to in vivo function is questionable. Often, publications purport to investigate muscle stem cells in vitro. It must be remembered that in all studies published thus far, any attempt to culture satellite cells in vitro results in activation and division to yield myoblasts that proliferate or differentiate depending on the culture conditions. Return of a satellite cell or muscle stem cell to quiescence in culture has not been definitively demonstrated. Studies of muscle stem cells in culture are in truth studies of myoblasts. The importance of these distinctions lies in the need to attribute biological significance to findings in vitro. Cultured myoblasts do not contribute to skeletal muscle when transplanted into muscle tissue to the degree expected of stem cells. Recently, isolation of muscle stem cells with minimal activation has been achieved to some extent, permitting their successful transplantation with concomitant satellite cell compartment replenishment as well as tissue repair (Collins et al., 2005; Montarras et al., 2005). However, in an ideal biological experimental system, muscle stem cells would be observed in their natural environment, in vivo, without external manipulations.

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The above paragraphs are not meant to dissuade investigators from performing in vitro experimentation with muscle stem cells. The point is simply to increase awareness of the distinction between in vitro observations showing what are believed to be important stem cell attributes and in vivo relevance, or stem cell function in the body. In vitro studies of stem cell properties should be interpreted with caution. That being said, attempts to further the use of muscle stem cells, and stem cells in general, for therapeutic purposes will necessarily require extensive understanding of how these cells may be manipulated and expanded in culture while preserving their “stemness.” Perhaps one of the greatest goals in stem cell biology is re-creating the microenvironment or niche in which stem cells reside, preserving “stemness,” and, in particular, retaining and cultivating self-renewal by asymmetric division outside of the body.

THE MUSCLE STEM CELL NICHE In contrast to many stem cells in other tissues, the muscle stem cell is defined as a cell occupying a highly circumscribed locale, or niche. Specifically, the satellite cell niche is a compartment, classically identified by electron and more recently by confocal light microscopy, beneath the basal lamina and external to the sarcolemma of mature muscle fibers. Although the vast majority of satellite cells are quiescent, disruption of the physical structure of the niche results in their activation. Niches have been most extensively characterized in Drosophila germ cells and in some mammalian tissues including the germ line, hematopoietic system, the epidermis, and the intestinal epithelium (reviewed in Watt and Hogan, 2000; Spradling et al., 2001; Yamashita et al., 2005). In these tissues shared niche characteristics include the presence of secreted signaling molecules as well as adhesion molecules tethering the stem cell to its immediate surroundings. The molecular components of the satellite cell niche and their role in maintaining quiescence or in stimulating activation remain ill-defined. The extent to which the niche confers active signals to the satellite cell to maintain quiescence or protects it from signals that could cause activation is a relatively unexplored area, but likely involves both types of influences. Examples of candidate molecules in the niche microenvironment that could play a role in the activation of satellite cells are the insulin-like growth factor, hepatocyte growth factor, fibroblast growth factor, and the notch ligand, delta. The importance of soluble factors in the microenvironment in regulating muscle stem cell function was recently highlighted in a series of remarkable experiments investigating the effects of aging on the ability of skeletal muscle to mount a regenerative response to damage (Conboy et al., 2005). These studies showed that the failure to regenerate muscle in aged animals is in large part due to the absence of secreted factors in the niche, which if replaced artificially, could elicit a robust regenerative response from the aged satellite cells. Many niche components have been identified that are in contact with quiescent satellite cells but have yet to be characterized in terms of their ability to maintain quiescence (Dhawan and Rando, 2005). Possibly, factors present in various somatic stem cell niches share common features that control stem cell function. Undoubtedly, the location of a stem cell within its microenvironment, including its physical orientation with respect to the surrounding matrix, provides a complex yet precise network of signals that together instruct the stem cell to remain quiescent, but poised to become activated, divide, and self-renew in times of need, as signals change. CHALLENGES FACING THE POTENTIAL THERAPEUTIC USE OF MUSCLE STEM CELLS The study of muscle stem cells has significant implications for therapy on a number of levels. In muscle disease such as DMD or other muscular dystrophies, the role of the stem cell remains unclear. Is the ultimate failure to regenerate muscle due to the exhaustion of muscle stem cells that relentlessly attempt to repair degenerating myofibers? Or is dystrophy a result of defects intrinsic to the muscle stem cells? Possibly, in some disease states, aberrant environmental influences extrinsic to muscle stem cells dictate their function (Oexle and Kohlschutter, 2001). Regardless, efforts toward treating disease could focus on the stem cell. For example,

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a loss of stem cell function would suggest the need to either enhance the function of the native cells or to locate and tap into other potential sources of muscle stem cells, either derived from transplant donors or from the affected individual. Accordingly, a missing factor could be delivered to diseased muscle by muscle stem cells that express the wild-type gene or by defective genetically engineered muscle stem cells (Huard et al., 1998; Partridge et al., 1989; Blau and Springer, 1995). A difficult problem common to all potential ES and somatic stem cell therapies is finding a means of delivering the transplanted cells into the recipient tissue that allows for tissue integration and contribution to function. Transplantation of cultured muscle cells into skeletal muscle has led to disappointing results. Myoblasts fused with nearby myofibers, produced dystrophin, but remained highly localized near the site of injection (Gussoni et al., 1992, 1997; Morgan et al., 1993; Rando et al., 1995). As described above for mice, it may now be possible to isolate human satellite cells with minimal activation, either using surface markers or by dissecting single fibers with intact basal laminae. In contrast to the myoblasts used in clinical trials thus far, transplanted satellite cells may exhibit an enhanced ability to disperse from the site of injection and a more robust contribution to muscle regeneration, including the formation of new satellite cells, that is, self-renewal (Collins et al., 2005). However, these properties of satellite cells are lost following any growth or expansion of the cells in vitro, limiting their supply. Perhaps a greater understanding of niche biology will permit the development of extracorporeal incubators that will support the manipulation and cultivation of satellite cells without compromising their ability to function as stem cells when re-implanted into muscle. Finally, it should be remembered that all functions of muscle stem cells are not necessarily beneficial. With respect to cancer development, the cell of origin of skeletal muscle tumors, in particular rhabdomyosarcomas, is unclear at present. Rhabdomyosarcomas arise within skeletal muscles as well as other tissues, and are characterized by cells expressing skeletal muscle proteins. There is debate about whether these tumors arise from muscle stem cells. Electron microscopic studies suggest that the mitotic cells in rhabdomyosarcomas resemble dividing satellite cells and lack components of mature muscle cells. In addition, overexpression of c-met, a marker expressed by satellite cells but not mature muscle cells, leads to rhabdomyosarcoma development (Sharp et al., 2002). However, conditional mouse models of Pax3:Fkhr translocation-driven rhabdomyosarcoma suggest that these tumors may be initiated in differentiating myofibers as opposed to satellite cells (Keller et al., 2004a; Keller et al., 2004b). Although the cellular etiology of skeletal muscle tumors remains to be determined, muscle biologists should be mindful of the satellite cell and experimental approaches designed to enhance muscle stem cell function or proliferative capacity must consider potential for malignant transformation.

CONCLUSION For the past four decades the regenerative capacity of skeletal muscle has been attributed to muscle stem cells, mononucleate cells residing in an anatomical niche as “satellites,” juxtaposed to myofibers. The attribution of muscle stem cell activity to the satellite cell was largely supposition based on location and appearance buttressed by circumstantial evidence of contribution to muscle regeneration. Evidence that satellite cells could be awakened from a dormant quiescent state to exhibit changes in nuclear and sub-cellular structure, and to divide in the vicinity of tissue damage suggested a central role in muscle repair. Additional experiments using labeling methods for DNA synthesis as well as functional perturbations such as irradiation strengthened the case. More recently, the role of satellite cells has been confirmed by the molecular characterization of satellite cells in terms of the markers they express, allowing prospective isolation and designation of the role of particular genes necessary for their formation and function. Ectopic genetic markers such as GFP or LacZ, along with progress in transplantation with minimal perturbation have now permitted a direct demonstration of the potential for satellite cell self-renewal and contribution to muscle regeneration to a functionally significant extent.

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Experiments centered on satellite cells have set the standard for searches for additional non-satellite muscle stem cells. To date, none has come close, although several candidates have partially fulfilled the criteria. One possible reconciliatory view is that satellite cells may represent a focal point in muscle stem cell biology. The ontogeny of satellite cells is not fully elucidated and may involve multiple routes, including somite derived embryonic precursors, bone marrow derived adult precursors, muscle interstitial cells, and perhaps others. More distant is the possibility that significant muscle regeneration may occur via avenues bypassing the satellite cell entirely. Convincing data for the latter is lacking and any prospect will require nothing short of rigorous demonstration of robust functional muscle repair. The current challenges in muscle stem cell biology include understanding the basis of satellite cell depletion in disease and in designing rational therapeutic interventions. Toward these ends, current and future insight into normal satellite cell biology may provide the clues toward re-creating intrinsic satellite cell properties by nuclear reprogramming of other cells or utilizing extrinsic environmental signals for cultivation ex vivo with a view to replacement and the correction of defective cells. Finally, recognizing that satellite cell malfunction may result in some cases from altered environmental signals should suggest approaches for re-vitalizing otherwise functional cells.

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22 Islet Cell Therapy and Pancreatic Stem Cells Juan Domínguez-Bendala, Antonello Pileggi, and Camillo Ricordi

INTRODUCTION Replacement of insulin-producing cell function represents an appealing approach for the treatment of diabetes mellitus, a condition characterized by loss of β-cell mass and/or function (Ricordi, 2003; Ricordi and Strom, 2004) consequent to autoimmunity (type 1 diabetes mellitus, T1DM), metabolic disorders (i.e. cystic fibrosis, hemochromatosis, and liver cirrhosis), surgery (i.e. iatrogenic diabetes following pancreatectomy for relapsing, chronic pancreatitis) (Ricordi, 2003), or β-cell dysfunction secondary to insulin resistance and hyperinsulinism (type 2 diabetes mellitus, T2DM). Exogenous insulin injections have represented a life-saving treatment in T1DM, changing the natural history of diabetes and remarkable progress has been achieved in recent years in the management of glycemic control by combining diet, exercise, and improved exogenous insulin treatment options. However, this approach requires continuous adjustments in insulin administrations with significant challenges in attaining tight glycemic control in the absence of severe hypoglycemic episodes. Tight metabolic control with avoidance of wide glycemic excursions is necessary to decrease the risk of development and/or progression of the chronic complications that can negatively impact the quality of life and life expectancy of patients with diabetes. Hundreds of thousands endocrine cell clusters, from 50 to 500 μm in diameter (islets of Langerhans) are scattered into the pancreatic tissue, representing approximately 1–2% of the entire organ. The islets are “micro-organs” with a unique cytoarchitecture, composed of heterogeneous cell subsets specialized in the production and secretion of endocrine hormones (α-cells for glucagon; β-cells for insulin; δ-cells for somatostatin; PP-cells for pancreatic polypeptide) that are essential for the regulation of glucose homeostasis in the blood (Brissova et al., 2005; Cabrera et al., 2006). Complex interactions between the cell subsets composing the islets, their innervation and the rich vascular bed result in “real-time” secretion of endocrine hormones that maintain glycemic values within physiologic ranges. Better understanding of pancreatic islet cell ontogeny, biology, and physiology will be of assistance in developing efficient protocols for cellular therapies for the restoration of metabolic control in patients with diabetes. Considerable progress has been achieved in the last two decades in the field of β-cell replacement therapy, either by transplantation of the pancreas as a vascularized organ or by infusion of islet cell products. The encouraging results of recent clinical trials support the value of this approach, which has been shown to improve both quality of life and metabolic control in patients with T1DM following intrahepatic islet transplantation (Ryan et al., 2001, 2002, 2005; Froud et al., 2005a, b; Pileggi et al., 2005, 2006; Poggioli et al., 2006). Current challenges to the widespread application of β-cell replacement therapies include the shortage of transplantable tissue and the need for more effective and safer immune interventions that favor longterm graft function. Ultimately, successful strategies for immunoisolation, tissue engineering with local

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immunosuppression, or the development of successful strategies for tolerance induction to avoid the need for life-long immunosuppression of the recipients will be necessary for the widespread applicability of islet cell therapy. In fact, the current requirements for life-long immunosuppression of the recipients severely limit the current indications for islet transplantation to the most severe cases of T1DM or in patients already undergoing organ transplantation and therefore already undergoing immunosuppressive treatment (Pileggi et al., 2001; Ricordi and Strom, 2004). When islet transplantation will be possible without chronic recipient immunosuppression, current sources of donor pancreata will clearly be insufficient to meet the demand. That is why it is so critical to continue to work toward the identification of alternative sources of insulin-producing cells. Encouraging data are emerging in the field of islet cell neogenesis and stem cell research that justify a cautious optimism for the years to come (Ricordi et al., 2005; Pileggi et al., 2006). This chapter will review the current status, challenges, and perspectives in clinical islet transplantation for treatment of diabetes and the progress of selected areas of stem cell and β-cell regeneration.

BENEFITS OF β-CELL REPLACEMENT THERAPY Transplantation of β-cells is currently performed as vascularized pancreas or isolated islet cell grafts. Both procedures can result in improved glycemic control in patients with diabetes (Pileggi et al., 2006). Transplantation of pancreatic islets offers substantial advantage over whole pancreas transplantation because of the lower risks for procedure-related complications and the possibility of preconditioning the graft in vitro prior to implantation (Pileggi et al., 2006). Islets are isolated from the donor pancreas by a mechanically enhanced, enzymatic digestion process that allows for the physical dissociation of pancreatic tissue into small fragments and liberation of the endocrine cell clusters with preserved integrity (Ricordi et al., 1988). The dissociation phase is followed by purification on density gradients that enriches for fractions with higher endocrine cell clusters (2% of the whole pancreatic tissue) while minimizing contamination with non-endocrine tissue (Alejandro et al., 1990; Ichii et al., 2005b. After isolation and culture, fractions with different degrees of purity are pooled for transplantation. Islet transplantation is performed using minimally invasive interventional radiology techniques consisting of percutaneous, transhepatic cannulation of the portal vein, and infusion of the islets by gravity. After intraportal infusion, the islet cell clusters remain trapped at the presinusoidal level (Alejandro and Mintz, 1988; Baidal et al., 2003; Froud et al., 2004; Pileggi et al., 2005). The purification procedure allows to substantially reduce the volume of tissue to be infused, therefore minimizing the previously reported risk of portal thrombosis and portal hypertension consequent to the intrahepatic embolization of the islet grafts (Froud et al., 2004, 2005), which has been described when unpurified islet preparations or inadequate islet infusion techniques were used. Islet transplantation is indicated in patients who have lost insulin-producing cell function. Recent clinical trials have shown the importance of intensive insulin treatment to obtain tight glycemic control and its ability to prevent or delay the dreadful complications of unstable glycemic control, including neuropathy, vasculopathy, and nephropathy (DCCT, 1993). Unfortunately, intensive insulin treatment cannot maintain glycemic values within normal ranges throughout the day and is associated with an increased risk of severe hypoglycemia, at times fatal. Restoration of islet β-cell function is a highly desirable goal for patients with T1DM as it can provide a more physiological glycemic metabolic control than exogenous insulin. Transplantation of autologous islets (autotransplantation) is generally performed to prevent iatrogenic diabetes in patients undergoing pancreatectomy due to severe pain for chronic, relapsing pancreatitis, or for

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non-enucleable benign neoplasm of the pancreas (Robertson et al., 2001; Oberholzer et al., 2003). The islets are isolated from the recipient’s pancreas after total pancreatectomy and then transplanted into his/her own liver. Transplantation of allogeneic islets (obtained from the pancreas of deceased multiorgan donors) is generally performed in patients with T1DM for whom loss of pancreatic β-cells in the pancreatic islets is due to an autoimmune process (Ricordi, 2003). The transplant is indicated in non-uremic, c-peptide negative patients with unstable diabetes complicated by severe hypoglycemia and performed as solitary islet transplantation (islet transplantation alone, ITA) in patients with end-stage renal disease receiving a kidney graft before (islet after kidney, IAK) or at the time of islet transplantation (simultaneous islet kidney; SIK) (Shapiro et al., 2000; Ricordi, 2003; Ricordi and Strom, 2004). Allogeneic islet transplantation has also been performed in patients with diabetes associated with metabolic diseases (i.e. cystic fibrosis, hemochromatosis, and liver cirrhosis) (Tzakis et al., 1990, 1996; Brunicardi et al., 1995; Ricordi et al., 1997; Tschopp et al., 1997; Angelico et al., 1999) and surgical removal of the pancreas (for trauma or benign abdominal diseases) in combination with liver, lung, or clustered abdominal organs (Johnson et al., 2004). After transplantation of pancreatic islets, dramatic improvement of metabolic control is generally observed with reduction of mean amplitude of glycemic excursions and of insulin requirements, normalization of glycated hemoglobin, and absence of severe hypoglycemia (Alejandro et al., 1997; Ryan et al., 2002, 2004, 2005a; Geiger et al., 2003; Froud et al., 2005). Insulin independence is achieved when a sufficient islet mass is implanted, a goal generally obtained using islets obtained from one or more donor pancreata (Shapiro et al., 2000; Markmann et al., 2003; Froud et al., 2005; Hering et al., 2005; Pileggi et al., 2005). Long-term graft function has been reported after islet autotransplantation (Robertson et al., 2001) and allogeneic islet transplantation (Carroll et al., 1995; Alejandro et al., 1997; Froud et al., 2005; Pileggi et al., 2005; Ryan et al., 2005a), with improved metabolic control and absence of severe hypoglycemia even in patients under exogenous insulin treatment. Recent clinical trials of allogeneic islet transplantation have shown that insulin independence can be obtained in approximately 80% of the patients at 1year, but progressive graft dysfunction has been observed over time, with approximately 10% of patients insulin free by 5 years, despite sustained c-peptide production and good metabolic control with reintroduction of exogenous insulin (CITR, 2004; Froud et al., 2005; Ryan et al., 2005a; Pileggi et al., 2006). The benefits of replacing β-cell function by islet transplantation include a dramatic improvement of the quality of life associated with the enhanced glycemic control and reduced fear of severe hypoglycemia (Barshes et al., 2005; Poggioli et al., 2006). The positive effects of islet transplantation are maintained even in patients experiencing partial graft dysfunction and requiring reintroduction of exogenous insulin, until measurable c-peptide persists (Alejandro et al., 1997; Pileggi et al., 2005). Additionally, as previously reported for pancreas transplantation, islet transplantation is associated with improved survival and function of renal allografts (Fiorina et al., 2003, 2005), improvement of vasculopathy (Fiorina et al., 2003), better cardiovascular function (Fiorina et al., 2005) in IAK recipients, stabilization of diabetic retinopathy, and neuropathy in recipients of ITA (Lee et al., 2005). The transplantation procedure has been associated with a relatively low incidence of side effects to date (Goss et al., 2003; Markmann et al., 2003; Owen et al., 2003; Frank et al., 2004; Froud et al., 2004, 2005; Hafiz et al., 2005; Ryan et al., 2005a; Venturini et al., 2005). Expected untoward complications of the immunosuppressive drugs utilized to prevent graft rejection have been described in recent clinical trials (Hirshberg et al., 2003; Cure et al., 2004; Frank et al., 2004, 2005; Andres et al., 2005; Froud et al., 2005; Hafiz et al., 2005; Molinari et al., 2005; Ryan et al., 2005a; Senior et al., 2005), which are similar to those observed for other organs and tissues.

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CURRENT LIMITATIONS TO β-CELL REPLACEMENT THERAPY Hurdles to the widespread application of β-cell replacement therapy based on the transplantation of allogeneic islets include the relatively high numbers of islets required to achieve insulin independence, due to the shortage of deceased donor organs available for transplantation. While improved donor management after brain death, refined organ procurement (Lee et al., 2004), and preservation (Kuroda et al., 1988; Matsumoto et al., 1996; Fraker et al., 2002) techniques have allowed for better results in recent years, expansion of the donor pool to marginal donors (Ricordi et al., 2003; Tsujimura et al., 2004a, b) and donation after cardiac arrest (Goto et al., 2005; Matsumoto and Tanaka, 2005) appear promising avenues to increase organ utilization and obtain adequate (both qualitatively and quantitatively) islet cells from a single donor pancreas for transplantation. Unfortunately, a large number of organs suitable for transplantation are underutilized (Krieger et al., 2003), indicating the need for improved management of potential pancreas donors and organ recovery policies to increase organ availability. An appealing alternative to overcome donor organ shortage is the use of living donor organs (namely distal pancreatectomy) as source of islets (Matsumoto et al., 2005), although for a largescale application of this approach a thorough evaluation of risks/benefits for both donors and recipients should be undertaken to avoid onset of T2DM in the pancreas segment donor later in life (Robertson, 2004) and prevent loss of transplanted islets in the recipients due to the lack of safe and non-diabetogenic immunosuppressive/tolerogenic protocols at the present time. Steady improvements in islet cell processing, purification, and culture have been implemented in recent years (Ricordi et al., 1988; Alejandro et al., 1990; Lakey et al., 1999; Ichii et al., 2005a) that have allowed for the recovery of better islet yields from a single donor pancreas and therefore maximizing organ utilization for islet transplantation. Additionally, active research is ongoing toward the definition of sensitive predictive tests of islet potency (Street et al., 2004; Ichii et al., 2005a) that could discriminate preparations yielding adequate islets for transplant from those that are not as they could contribute to improve islet transplantation outcomes. Islet transplantation is regulated by the Food and Drug Administration as Investigational New Drug (IND) (Wonnacott, 2005). Implementation of current Good Manufacture Practice (cGMP), availability of specific infrastructures and of dedicated personnel is required to warrant high-quality standards and consistency in islet cell quality for transplantation (Weber, 2004). These requirements impose a remarkable economic burden on clinical islet transplantation programs (Markmann et al., 2003; Guignard et al., 2004). The creation of “regional” human islet cell processing facilities that can provide cGMP quality islet cell products for research and clinical transplant applications may represent a viable option to improve the consistency and quality of the final islet cell products, while containing the costs (Oberholzer et al., 2000; Goss et al., 2002, 2003, 2004; Lee et al., 2004; Kempf et al., 2005). The relatively high islet numbers required for successful post-transplant outcome also depend on the quality of the islet cell product infused into the recipients and the impaired engraftment of a relatively large proportion of islets due to the generation of inflammation in the liver microenvironment (Pileggi et al., 2001). Inflammation and hypoxia (due to lack of vascularization in the early period of post-implantation) could contribute to functional impairment and/or islet cell death early after islet transplantation. Engraftment of a suboptimal islet mass may also result in graft dysfunction due to metabolic exhaustion (Froud et al., 2005) that could be further worsened by the relatively hyperglycemic liver environment and toxic levels of immunosuppressive drugs in the hepatic vascular district (Desai et al., 2003; Shapiro et al., 2005; Pileggi et al., 2006). The steady improvement in islet cell processing will be of assistance in optimizing both yields and quality of islets from donor pancreata therefore contributing to increase the number of transplants in the years to

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come. It is conceivable that this approach will not suffice overcoming the increasing demand of islet grafts due to the disproportioned pancreas donor-to-recipient ratio: there will be a large number of patients with diabetes who would benefit of restoration of β-cell function and not sufficient pancreata for processing. Alternative sources of insulin-producing β-cells (from either allogeneic or xenogeneic donors) or induction of self-regeneration of the patient’s own β-cells (in combination with adequate immunomodulation to prevent recurrence of autoimmunity) (Ogawa et al., 2004) may help achieving the desired metabolic control in the near future (Ricordi et al., 2005). For β-cell replacement therapies to become the treatment of choice for patients with diabetes, successful restoration of metabolic function should be achieved long term. For this reason, implementing a sequential, integrated approach that combines strategies aiming at improving β-cell mass together with those focusing on the modulation of the immune response (i.e. preventing rejection and recurrence of autoimmunity) could represent an essential element toward definition of successful therapeutic strategies (Ricordi and Strom, 2004; Ricordi et al., 2005). Promising data on the induction of donor-specific unresponsiveness and tolerance to transplanted tissues in experimental models justify optimism for the near future, and may allow achieving long-term function of transplanted insulin-producing cells in the absence of rejection and recurrence of autoimmunity without the need for chronic immunosuppression in the clinical setting (Inverardi and Ricordi, 2001; Inverardi et al., 2004; Ricordi and Strom, 2004).

ALTERNATIVE SOURCES OF INSULIN-PRODUCING CELLS: STEM CELLS AND β-CELL REGENERATION Stem cells could be defined as undifferentiated cells that have the ability to proliferate while retaining the potential to fully mature into other cell types. The extent to which stem cells can be induced to proliferate or differentiate depends on their origin and stage of development. Arguably, the most powerful stem cells available are embryonic stem (ES) cells. These cells, which are obtained from the inner cell mass (ICM) of the blastocyst, can be maintained indefinitely in an undifferentiated, proliferative stage in vitro (Odorico et al., 2001; Thomson et al., 2005; Thomson et al., 1998). When transplanted into immunodeficient animals or otherwise induced to spontaneously differentiate, they can give rise to cells of all three embryonal layers (endoderm, ectoderm, and mesoderm). The prospects of turning human ES (huES) cells into islet cell types are therefore substantiated, but not exempt of safety and ethical concerns. Stem cells of fetal origin may still retain some degree of multilineage differentiation, as well as the potential to proliferate in vitro. Despite their embryonic origin, these cells should not be confused with the blastocyst-derived ES cells. In fact, in many respects, these cells are more akin to adult cell types than to ES cells. This, together with the controversy surrounding their procurement, makes them unlikely candidates to become an alternative source of islets. Expansion of fully differentiated, adult β-cells has been reported in vitro. However, the induction of β-cell proliferation has been generally associated with loss of mature cell phenotype and of functional competence that is only partially recovered after re-differentiation. Many adult tissues have also stem cells involved in their physiologic maintenance and repair mechanisms. Whether the adult pancreas contains endocrine stem cells or not is still the subject of heated debate. In general, tissue-specific stem cells are elusive and difficult to culture in vitro, and their differentiation potential is much more restricted than that of ES cells. One possible exception to this rule is the bone marrow (BM)-derived multipotent adult progenitor cells (MAPCs) described by Verfaillie and colleagues. These cells have been shown to proliferate extensively in vitro and are able to give rise to the three embryonal layers when injected into mouse blastocysts (Jiang et al., 2002). However, the routine isolation and culture of these cells is still far from mainstream, as it has proven more challenging than working with ES cells.

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An additional potential approach to obtain insulin-producing cells is transdifferentiation of adult cells (i.e. hepatocytes) under selected conditions both in vivo or ex vivo.

PANCREATIC DEVELOPMENT Research conducted over the last decade has outlined a basic “roadmap” of the major molecular events that shape islet development (Figure 22.1) (Edlund, 2001; Edlund, 2002). Critical developmental milestones are: (a) generation of endoderm/gut endothelium; (b) pancreatic differentiation; (c) endocrine specification; and (d) β-cell differentiation. Transition between each of these stages of development appears to be catalyzed by a surprisingly manageable number of transcription factors, which are highly conserved between mouse and man. Generation of Endoderm/Gut Endothelium ES cells are an artificially frozen snapshot of the ICMcells found at the blastocyst stage (embryonic day e3.5). Expression of genes such as telomerase, Oct3/4, and Nanog make these cells immortal and pluripotent under defined conditions in vitro. Subsequent differentiation will be marked by the permanent down-regulation of these genes. Visceral endoderm and epiblast, respectively, constitute the outer and inner layers of the ICM immediately before gastrulation. The visceral endoderm will become part of the yolk sac, without contribution to the embryo proper. In contrast, the definitive endoderm is formed during gastrulation when epiblast cells leave the ICM through the primitive streak. There is an intermediate stage in definitive endoderm formation, called mesendoderm. Although visceral and definitive endoderm are similar, mesendoderm-specific genes such as Gsc and Bry do not appear during visceral endoderm differentiation (Kispert and Herrmann, 1994; Tam et al., 2003; Kubo et al., 2004; Yasunaga et al., 2005), and therefore can be used to identify true definitive endoderm (Yasunaga et al., 2005). The anterior part of the definitive endoderm will evolve into the foregut, from which pancreas, liver, and lungs will eventually bud out. The posterior definitive endoderm, on the other hand, becomes the midgut and hindgut, which will differentiate into large and small intestine. Graded Nodal signaling is responsible for the initial patterning of the primitive gut endothelium. Many genes

Figure 22.1 Schematic representation of the differentiation pathway from ES cells to β-cells. Genes whose expression is necessary for the transition between each step are indicated in italics.

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have been associated with the formation of true endoderm, including Foxa2, Mixl1, GATA-4, and several members of the Sox family, chiefly Sox17 (de Santa Barbara et al., 2003). Pancreatic Differentiation There is a cross-communication between the gut endoderm and the surrounding mesoderm, mediated by Shh signaling. Shh is highly expressed throughout the gut endothelium, but is down-regulated in a Pdx1-positive region that will later become the pancreas at e8. Both Shh repression and Pdx1 activation are defining events of pancreatic specification. Pdx1 knockouts are born without pancreas (Jonsson et al., 1994). Chemical inhibition of Shh enhances pancreatic differentiation at the expense of intestine (Kim et al., 1997). Conversely, ectopic expression of Shh under the control of the Pdx1 promoter induces intestinal fates at the expense of the pancreas (Apelqvist et al., 1997). Endocrine Specification Endocrine differentiation occurs through a lateral inhibition process, mediated by Notch signaling. Cells in which the Notch receptor is activated by the ligands delta or serrate express high levels of HES-1, which in turn represses the pro-endocrine gene Ngn3. Lower levels of Notch signaling may randomly occur in individual cells, where HES-1 expression will not be up-regulated. In the absence of its repressor, Ngn3 will be expressed robustly, and the cell will adopt a pro-endocrine fate (Apelqvist et al., 1999; Gradwohl et al., 2000; Jensen et al., 2000). The differentiation into each of the five endocrine cell types within the islet (α-, β-, δ-, PP- and ε-cells) is preferentially observed at specific time points during embryogenesis, suggesting that Ngn3-positive cells adapt their responses to an evolving milieu of signals. β-Cell Differentiation Little is known about the extracellular signals that drive β-cell specification from Ngn3-positive progenitors. Animals lacking Nkx6.1 (Sander et al., 2000) and Nkx2.2 (Sussel et al., 1998) have defects in β-cell formation. However, several observations point to Pax4 as the main hallmark of β-cell differentiation: (i) the knockout of this gene results in the total absence of β-cells (Sosa-Pineda, 1997), but not α-cells; (ii) its expression peaks between e13.5 and e15.5, which coincides with the period of maximal differentiation of β-cell precursors (Sosa-Pineda et al., 2004); and (iii) shortly after endocrine specification, Ngn3 co-localizes with Pax4 (Wang et al., 2004), which suggests that the latter may be one of the targets of the former. Recent evidence indicates that Pax4 and Arx are mutually repressed, and that the balance between the two determines α- (Arx) or β-cell (Pax4) specification from Ngn3 progenitors (Collombat et al., 2003, 2005).

ISLET NEOGENESIS FROM ES CELLS Ideally, the “education” of human ES (huES) cells along the β-cell lineage would require the exact recapitulation of the differentiation steps described earlier. If we could identify the “instructive” extracellular signals that naturally drive this process, such signals could then be added in the proper sequence to the culture medium in the hope that the cells would respond accordingly (Figure 22.1). However, our understanding of the fine regulation of extracellular signaling is still somewhat limited at the present time. In fact, the combined action of signals such as FGF, Nodal, Hedgehog, Notch, BMP/TGF-β, or Wnt is responsible for the patterning and development of most organs (Edlund, 2002). During development, cells respond differentially to environmental cues depending on their exact location, their interaction with other developing tissues and time. Fine gradients of Nodal (for endoderm/gut endothelium specification), FGF, and Shh (for pancreatic

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differentiation), as well as random cell-to-cell interactions in the Notch pathway, are examples of the complex differentiation mechanisms that we are just starting to unravel. Given these limitations, it is not surprising that all attempts at generating β-cells from ES cells have been unsuccessful so far. The observation by Assady and colleagues that spontaneous in vitro differentiation of huES cells resulted in the scattered appearance of insulin-producing cells (Assady et al., 2001) merely confirmed the well-known fact that these cells have unlimited differentiation potential. Protocols for the efficient differentiation of β-cells were still necessary, and several groups set up to develop them. Lumelsky and colleagues, for instance, described a five-step method to generate islet-like cells from murine ES cells (Lumelsky et al., 2001) based on the derivation of cells positive for the intermediate filament protein Nestin, a known marker of neuroectodermal and mesodermal fates. Further analyses on these cells demonstrated that they were not true pancreatic endocrine cells, but rather neuroectodermal derivatives that absorbed insulin from the culture medium. Further refinements on this method have led to somewhat improved results, although the amount of insulin expressed by these cells is still quite reduced compared to that of mature β-cells (Fujikawa et al., 2005). Using a genetic engineering approach, Soria and colleagues (Soria et al., 2000) generated murine ES cell lines where a selectable marker (neomycin, which confers resistance to the drug G418) was placed under the control of the insulin promoter. Thus, when allowed to spontaneously differentiate, G418 selection yielded insulin-producing clones. Although elegant, this method requires the introduction of foreign genes. Also, it must be taken into account that insulin expression is not a very stringent criterion for the selection of β-cells, as many other tissues do express it. Indeed, the same authors later confirmed the ectodermal identity of some of the selected clones (Roche et al., 2005). The most exciting developments in the field of ES cell differentiation have been the result of a seemingly less ambitious approach. Instead of attempting the direct differentiation of ES cells into insulin-producing cells, several groups have focused on the key first step of the process, namely endoderm specification. The difficulty of this enterprise is highlighted by the fact that standard culture conditions strongly favor ectoderm and mesoderm over endoderm specification (hence the proliferation of ectoderm-based differentiation protocols). Also, early attempts to generate endoderm could not direct ES cells specifically toward definitive, rather than visceral, endoderm. Kubo and colleagues were the first to report the generation of definitive endoderm from murine ES cells, albeit at a low frequency (Kubo et al., 2004). Far more striking results were successively described by Tada and colleagues (Tada et al., 2005), also in mouse ES cells, and D’Amour and collaborators in huES cells (D’Amour et al., 2005). The latter was based on the addition of high concentrations of Activin A, a TGF-β-related agonist of Nodal, in low-serum conditions. Since endoderm specification had been widely regarded as the main obstacle toward pancreatic differentiation, we now expect a steady, accelerated progress of these lines of research. Of particular interest are the new differentiation strategies that make use of protein transduction technology (Wadia and Dowdy, 2002, 2003) for the delivery of key transcription factors (Pdx1, Ngn3, Pax4, etc.) to stem cells in vitro. This approach would be particularly useful to bypass the enormous challenge of mimicking in vitro the complex signaling pattern that is responsible for the sequential activation of such transcription factors in vivo (DomínguezBendala et al., 2005).

ISLET NEOGENESIS FROM ADULT STEM CELLS The ability of adult pancreatic islet cells to retain regenerative potential during adulthood has been recognized. Several experimental models such as partial pancreatectomy (Bonner-Weir et al., 1993), cellophane wrapping of the pancreas (Rafaeloff et al., 1992), duct ligation (Wang et al., 1995), or treatment with streptozotocin (Guz et al., 2001), as well as physiological conditions such as pregnancy (Nielsen et al., 1999;

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Figure 22.2 Islet regeneration during adulthood may occur through several mechanisms. Yet unidentified endocrine stem cells within the islet may be responsible for beta cell turnover, although data obtained in a mouse model suggest that islet self-maintenance is preferentially due to replication of existing beta cells. Other investigators hypothesize that islets can be regenerated from ductal or acinar tissue, although it is not clear yet whether this phenomenon would be mediated by putative stem cells or by transdifferentiation. Finally, the BM has also been proposed as a reservoir of β-cell progenitors. Recent evidence, however, suggest that the regenerative capacity of migrating BM cells might rather be due to their in situ differentiation into supporting endothelial cell types.

Johansson et al., 2006; Sorenson and Brelje, 1997) and perhaps long-standing T1DM (Meier et al., 2005) confirm that insulin-producing cells can regenerate in adult life. However, the quest for endocrine pancreatic stem cells has been an elusive one (Figure 22.2). Numerous observations suggest that these cells may reside in the ductal epithelium. Aside from countless microphotographic snapshots showing insulin-positive cells that appear to sprout from the pancreatic ducts (Bonner-Weir et al., 1993; Wang et al., 1995; Meier et al., 2005; Sarvetnick and Gu, 1992), cultured ductal cells respond to various stimuli in vitro by expressing several β-cell markers and even secreting low levels of insulin (Bonner-Weir, 2000). Other groups have identified Nestinpositive cells within the pancreas with a remarkable ability to expand, although their ability to emulate βcells upon differentiation was less impressive (Zulewski et al., 2001). More recently, Gershengorn and colleagues demonstrated that adult islets can undergo a reversible epithelial-to-mesenchymal transition in vitro (Gershengorn et al., 2004). Upon “de-differentiation,” these cells could be expanded by a factor of 1012, which is well within the realm of clinical applicability. However, when “re-differentiated,” these putative βcells expressed a mere 0.02% of the amount of insulin found in mature, primary islets. A variation on this protocol resulted in enhanced insulin production (Ouziel-Yahalom et al., 2006), but the ability of these cells to proliferate was much more modest. Finally, it has been proposed that acinar tissue may also contain endocrine stem cells (Hao et al., 2006). In this case, transdifferentiation was almost negligible, and no effort was made at characterizing either the neogenic insulin-positive cells or their putative progenitors. In short, thus far nobody has been able to present conclusive evidence that adult stem cells can generate genuine β-cells in vitro. Most of these cellular byproducts are, at best, oddities that co-express markers found in many diverse cell types. Concerns that these cells may just be culture artifacts are justified, and were further fueled when Dor and collaborators (Dor et al., 2004) suggested that adult β-cells regenerate by replication rather than differentiation. Lineage-tracing experiments conducted in rodents convincingly demonstrated that the regeneration and normal turnover of islets occur preferentially by division of existing β-cells.

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This report did not rule out the possibility that stem cells may still exist in the pancreas, but their importance was suddenly – and dramatically – reduced. Although the burden of proof is now on those who defend the existence and biological significance of pancreatic stem cells, the jury is still out. For instance, it has been argued that human β-cells replicate at a much lower rate than their mouse counterparts, which would be inconsistent with the “replication only” hypothesis (Meier et al., 2005; Hao et al., 2006; Tyrberg et al., 2001; Finegood et al., 1995). Also, it is conceivable that the adult human pancreas may have evolved different mechanisms for normal β-cell turnover and damage-induced regeneration. The impossibility of conducting lineage-tracing experiments in humans will keep this controversy alive for years to come.

TRANSDIFFERENTIATION Adult, differentiated cells from specific tissues can turn into completely different cell types in certain conditions. This phenomenon has been termed transdifferentiation. We will examine here two cell substrates (namely bone marrow and liver) that have shown some promise at transdifferentiating into pancreatic cell types. Stem cells derived from the bone marrow (BM) have been associated with numerous examples of tissue repair and regeneration in vivo. It has been documented that transplanted BM cells can migrate from their niche to various tissues in response to injury (Goodell, 2001). In some cases, this migration was accompanied by a significant regeneration of the damaged tissue, which led to the hypothesis that some cell types within the BM may have either ES cell-like properties or the ability to transdifferentiate. However, as it was confirmed in a model of liver disease (Grompe, 2003), the regenerative effect can also be due to the fusion of the BM cells with cells of the target tissue. Regarding the pancreas, an early study showed that up to 3% of islet β-cells were of donor origin 1month after BM cell transplantation, without evidence of cell fusion (Ianus et al., 2003). The conclusions of this report, however, were recently contested by Lechner and colleagues (Lechner et al., 2004), who could not find any significant contribution of the BM to islets either in healthy mice or in models of pancreatic injury. Furthermore, Kang and colleagues reported that while BM cell transplantation was enough to prevent diabetes onset in nonobese diabetic (NOD) mice, there was little or no involvement of the BM cells in islet cell regeneration once the disease was overt (Kang et al., 2005). Still, yet another report presented evidence that donor BM cells do promote islet regeneration in a mouse model of diabetes (Hess et al., 2003). Interestingly, the authors of this study did not find any evidence of transdifferentiation of BM cells into β-cells: the beneficial effect was rather due to the recruitment of donor-derived endothelial cells to the injured islets, where they induced their regeneration. There is also a wealth of observations indicating that liver and pancreas are especially susceptible of interconversion. Many invertebrates have a single organ that comprises both hepatic and pancreatic functions, which suggests that the separation of these two organs is a relatively late evolutionary event. Indeed, both originate from common endodermal progenitors in the early foregut of vertebrate embryos (Deutsch et al., 2001; Jung et al., 1999). In general, hepatocytes and β-cells share not only many developmental features, but also similar molecular machinery for glucose sensing and secretion (Nordlie et al., 1999; Kim and Ahn, 2004). Many studies confirm that interconversion of liver and pancreas occurs under a variety of experimental conditions (Rao et al., 1988; Rao and Reddy, 1995; Rao et al., 1986; Shen et al., 2000), as well as in certain diseases (Lee et al., 1989; Wolf et al., 1990). Based on the above evidence, Ferber and colleagues (Ferber et al., 2000) set up to demonstrate that ectopic expression of Pdx1 in liver cells could induce transdifferentiation into pancreatic cell types. Using an adenovirus vector, a Pdx1 cassette was delivered to the livers of recipient mice, where normally silent, β-cell-specific

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genes were activated. However, the putative β-cells obtained (which seemed to share a dual hepatocyte/β-cell nature) were not properly characterized. Furthermore, the very low frequency at which this “transdifferentiation” event occurred led the authors to hypothesize that the cells that switched fates might have been resident stem cells rather than fully differentiated hepatocytes. More conclusive results were more recently reported by Slack and collaborators, who showed that large portions of the liver could be completely transdifferentiated into pancreas in transgenic frogs where a Pdx1VP16 fusion cassette is expressed under the control of the liver-specific promoter TTR (Horb et al., 2003). The rationale for the use of VP16, a potent transcriptional transactivator from the herpes simplex virus (Sadowski et al., 1988; Triezenberg et al., 1988), is that non-pancreatic cells may lack the appropriate molecular partners for Pdx1 to exert its biological function. Indeed, no transdifferentiation was observed when Pdx1, without VP16, was used. This observation suggests that Pdx1 is necessary, but not sufficient to promote true pancreatic differentiation from the liver. Additional progress in this direction may open very exciting avenues, as hepatocytes can be easily obtained in large numbers either from adult livers or from ES cells (Rambhatla et al., 2003; Shirahashi et al., 2004).

WHAT THE FUTURE MAY HOLD Steady progress in the field of β-cell replacement has made of islet cell transplantation a therapeutic reality for patients with the most severe forms of diabetes. The benefits of this approach both in terms of metabolic control and quality of life after islet transplantation support the advantage of restoring β-cell function, compared to exogenous insulin treatment. The pace of stem cell research over the last decade has also been significant. Diseases thus far considered incurable now seem within the reach of our ever increasing therapeutic arsenal. Stem cells, be it of embryonic or adult origin, may provide in the future an unlimited supply of insulin-producing cells for treatment of diabetes. It is important, however, not to lose perspective of the many challenges ahead. First, no protocol for the efficient derivation of fully competent β-cells from stem cells has been described as yet. In our opinion, the most promising approaches are based on the generation of true endoderm from ES cells, but this would be just the first of several steps. Terminal differentiation of β-cells may require further advances in our ability to mimic their unique biological niche, which is known to be highly oxygenated through extensive vascularization. Another important consideration is safety. While islet transplantation is generally considered a safe procedure, ES cell-based approaches may require additional precautions to prevent the formation of tumors by carryover undifferentiated cells. The same considerations may apply to protocols aiming at in vivo β-cell regeneration in the native pancreas, since stimulation of β-cell proliferation may be associated with increased risk of hyperplasia or neoplastic transformation (e.g. nesidioblastosis, insulinoma, or other tumors). Solving the problem of supply is just one component of the puzzle. T1DM will not be cured unless we can protect the β-cells from the host’s immune system (Ricordi et al., 2005). In this direction, ES cells may have the edge over adult cell types (which could be potentially obtained from the patient himself) because there is no advantage in transplanting autologous cells in the context of an autoimmune process. In addition, recent reports suggest that ES cells, as well as their differentiated derivatives, may require less intensive immunosuppressive regimes compared to adult cell types (Li et al., 2004; Drukker et al., 2006). Stem cell research bears an invaluable potential for the treatment of T1DM and many other disease conditions. The enormous potential impact of stem cell-derived therapies in future medical practices warrants a renewed investment of resources in this field of investigation in the context of academic institutions, under strict ethical and regulatory oversight. Support of stem cell research by government agencies would allow for a faster, regulated, and safer advancement of a field that is currently limited by political restrictions.

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Notwithstanding the challenges, it appears that the prospect of defeating T1DM is within reach and that successful therapeutic strategies can be developed as a result of a multidisciplinary, integrated approach.

ACKNOWLEDGMENTS Supported by: National Institutes of Health/National Center for Research Resources, Islet Cell Resources (ICR; U42 RR016603, M01RR16587); Juvenile Diabetes Research Foundation International (#4-2000-946); National Institutes of Health/National Institute of Diabetes and Digestive and Kidney Diseases (5 R01 DK55347, 5 R01 DK056953, R01 DK025802); American Diabetes Association; State of Florida; a contract for support of this research, sponsored by Congressman Bill Young and funded by a special congressional out of the Navy Bureau of Medicine and Surgery, is currently managed by the Naval Health Research Center, San Diego, CA; and the Diabetes Research Institute Foundation (www.diabetesresearch.org).

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23 Regenerative Medicine for Diseases of the Retina Deepak Lamba and Thomas A. Reh INTRODUCTION The vertebrate retina is subject to a variety of degenerative conditions. Glaucoma, diabetes, macular degeneration, and retinitis pigmentosa are among the more common conditions that lead to loss of one or more of the retinal cell types and frequently result in partial or complete blindness. While some of these disorders have treatments that can slow the progression of visual loss, in most cases, there will be an untreatable visual impairment. The development of effective cell therapies is thus a goal of many individuals working in the visual sciences, and there have been steady advances using a variety of approaches toward this end. Some of these approaches have relied on the interesting fact that in many non-mammalian vertebrates, the retina can spontaneously repair itself to a truly remarkable degree. In the early days of regeneration research, investigators used the eye as one of the key model systems to study the phenomenon of regeneration. In this chapter, we will review (1) the basic developmental biology of the eye, describing the relationships between retinal stem cells and progenitors during development, (2) the sources of retinal stem cells and progenitors in mature animals that mediate retinal regeneration, and (3) the potential for derivation of retinal stem cells or progenitors from embryonic stem (ES) cells for transplantation. This review is not meant to be exhaustive, but we hope it will illustrate the main currents of research in this field. The vertebrate retina arises from the ventral diencephalon of the neural tube (Figure 23.1). Paired evaginations, known as the optic vesicles, emerge from the anterior region of the neural plate. The optic vesicle cells express a unique complement of transcription factors, termed eye-field transcription factors (EFTFs), which set them apart from the surrounding regions of the neural plate (see later). The optic vesicle cells undergo extensive proliferation over the next phases of retinal development and will ultimately generate all the various cell types of the neural retina, as well as several non-retinal ocular structures, such as the ciliary epithelium, the pigmented epithelium, and the iris. In this chapter, we will first briefly outline the current understanding of the molecular biology of eye development, describe the intrinsic potential for regeneration in the retina of non-mammalian vertebrates, and finish with research into ES cells and their use in retinal repair. EFTFs: SPECIFICATION OF THE EYE The presumptive eye-forming region of the embryo, or eye field, was first defined by transplantation experiments of Hans Spemann (Figure 23.2). More recently it has been possible to identify the same region by monitoring expression of a group of transcription factors called EFTFs (Figure 23.2). The EFTFs that are expressed early in eye field specification include Rx, Pax6, Six3, Lhx2, and Optx2 (Six6). The eye field forms late in gastrulation at the anterior end of the neural plate in the diencephalic region of the forebrain. The eye field initially extends across the midline as a single domain. This single field is subsequently split into two lateral

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Figure 23.1 The development of the various parts of the eye that are derived from the neural tube. In the first stage of eye development, the optic vesicle appears as an evagination from the diencephalons of the neural tube. Even at this early stage of development the vesicle is patterned into a presumptive RPE domain (gray) and a presumptive neural retina domain (red). The vesicle then becomes the optic cup as it folds in on itself, and the lens pinches off from the overlying ectoderm. At the optic cup stage, the first neurons emerge in the central retinal domain (blue), but not in the more peripheral regions. In the next stage, the embryonic eye begins to show distinct gene expression patterns in the more anterior (peripheral) regions, which will become the epithelial part of the ciliary body (ciliary epithelium – yellow) and the iris (green). Note that the iris and ciliary epithelium each have two layers: a pigmented and a non-pigmented layer. The pigmented layer of each region is continuous with the RPE, whereas the non-pigmented epithelial layer of both the iris and the ciliary epithelium is continuous with the retina. The CMZ, which contains the persistent progenitors/retinal stem cells in non-mammalian vertebrates, arises at the junction between the ciliary epithelium and the retina (red) and may be similar to the very early optic vesicle cells. A small part of the anterior eye is shown at the right of the figure to show the eventual relationships among the various domains in the mature eye.

domains due to the repression of EFTFs by sonic hedgehog (Shh), which is released from the prechordal mesoderm at the midline (Li et al., 1997). The EFTFs are essential for eye development; mutations in each of these genes are associated with either anophthalmia (no eye) or microphthalmia (small eyes) (Hill et al., 1991; Mathers et al., 1997; Porter et al., 1997; Carl et al., 2002; Zuber et al., 2003). Prior to the development of the eye field, the anterior of the nervous system becomes distinct from the posterior, and the Otx2 transcription factor (a member of the orthodenticle family) is critical in the control of this distinction (Simeone et al., 1993). Otx2 is downregulated in the eye field as a related transcription factor, Rx, is expressed (Andreazzoli et al., 1999). Otx2 expression persists in the periphery of the eye field and becomes restricted to the pigment epithelium and to some post-mitotic retinal cells (Bovolenta et al., 1997). Although there is evidence that Otx2 is important in eye-field development, it is difficult to precisely define its role in early ocular development because Otx2/ mutants do not form any structures anterior to rhombomere 3 (Matsuo et al., 1995). Among the first, if not the first, transcription factors to define the eye field is Rx/Rax, a paired-like homeobox transcription factor. Rx expression begins in areas that will give rise to the ventral forebrain and optic vesicles. Once the optic vesicles form, Rx expression is restricted to the ventral diencephalon and the optic vesicles and is eventually restricted to the developing retina (Furukawa et al., 1997). Homozygous null mutations of the Rx gene in mouse result in anophthalmia, with no eye development after the optic vesicle stage (Mathers et al., 1997). The region also lacks other EFTFs like Pax6 and Six3, indicating that Rx may have a role in inducing these genes

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Figure 23.2 The origin of the eye. (a) The eye field of the neural plate can be visualized in frog embryos using an in situ hybridization for Pax6 (from Zuber et al., 2003). (b) Drawing of the presumptive eye forming region on the left side of the embryo, as deduced by transplantation experiments of Hans Spemann, with the EFTF expression pattern superimposed. (c) Pax6 in situ in a chick embryo showing the early optic vesicle (arrow).

(Zhang et al., 2000). A similar anophthalmia phenotype was observed in loss of function experiments in Xenopus embryos using morpholino oligonucleotides against the Xenopus homolog to Rx (Andreazzoli et al., 2003). A mutation in the Rx gene in humans has been identified in a patient suffering from anophthalmia and sclerocornea (Voronina et al., 2004). Overexpression of Rx in Xenopus embryos results in hyperproliferation of the neural retina and retinal pigment epithelium (RPE), as well as formation of ectopic retinal tissue (Mathers et al., 1997). Similar results were obtained in misexpression studies carried out in zebrafish (Chuang and Raymond, 2001). The most studied EFTF is Pax6. It has been proposed that Pax6 is the master regulatory gene in eye development. It belongs to the family of paired box homeodomain genes and has been highly conserved across species. Pax6 is expressed in the anterior neural plate at the end of gastrulation and then becomes restricted to the region of the optic vesicle as well as lens ectoderm. Its expression persists throughout optic development and ultimately into adult animals in ganglion, horizontal, and amacrine cells (Grindley et al., 1995; de Melo et al., 2003). Mutations in the Pax6 gene result in a variety of phenotypes, depending on the gene dosage. Homozygous mutations that cause a loss of all Pax6 expression, result in anophthalmia in mice and rats (Matsuo et al., 1995; Grindley et al., 1995). Pax6 mutant mouse embryos have normal Rx expression, suggesting that Pax6 is downstream of Rx (Zhang et al., 2000). Misexpression studies with Pax6 have been carried out in Drosophila (Halder et al., 1995) and Xenopus (Chow et al., 1999), and in both species this induces ectopic eye tissue. Overexpression of Pax6 in Xenopus results in multiple ectopic eyes all along the dorsal central nervous system (CNS) along with ectopic expression of other EFTF including Rx in these areas. This suggests that Pax6 also has a role in the induction of Rx. The ectopic eyes display a similar morphology to the normal eye having both a neural retina and a lens. Thus, loss of function studies, as well as misexpression studies, lend support to the idea that Pax6 is a master regulatory gene during eye development. In addition to Rx and Pax6, there are several other members of the EFTFs. Lhx2 is an EFTF belonging to the family of Lim-homeodomain genes. It is expressed in the optic vesicles just before the completion of gastrulation

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(Xu et al., 1993; Porter et al., 1997). Lhx2 null mutants fail to form eyes (Porter et al., 1997). Developmentally, eye formation gets stalled at the optic vesicle stage and the optic cup and lens do not form. Analysis of Pax6 expression in these mice shows a normal pattern of Pax6 in the optic vesicle, and so Lhx2 may lie downstream of Pax6. Overexpression of Lhx2 in Xenopus embryo results in both large eyes as well as ectopic retinal tissue (Zuber et al., 2003). Six3 belongs to the Six-homeodomain family of genes. Six3 appears in the region of the presumptive eye field around the same time as Pax6 (Oliver et al., 1995; Bovolenta et al., 1998; Loosli et al., 1998). Six3 inactivation in medaka fish has been shown to result in anophthalmia and forebrain agenesis (Carl et al., 2002). Misexpression of Six3 in medaka fish results in multiple eye-like structures that express other EFTFs (Loosli et al., 1999), while in zebrafish it results in enlargement of the optic stalk (Kobayashi et al., 1998). Optx2 (Six6, Six9) also belongs to the Six-homeodomain family of genes and is expressed from the optic vesicle stage (Toy et al., 1998; Jean et al., 1999; Lopez-Rios et al., 1999; Toy and Sundin, 1999). Misexpression studies carried out with Optx2 in Xenopus embryos result in a large expansion of the retinal domain as well as hyperproliferation of cultured retinal progenitors transfected with XOptx2 (Zuber et al., 1999; Bernier et al., 2000). Although a considerable amount has been learned about the role of EFTFs in ocular development, little is known about the factors that control their expression. Recently, a few investigators have looked into the role of Wnt signaling in the initiation and regulation of the eye fields (Rasmussen et al., 2001; Cavodeassi et al., 2005). Wnts and their receptors belonging to both the canonical β-catenin pathway and the non-canonical pathways are expressed at the site of the prospective eye field. Wnt1 or Wnt8b, both of which are known to activate the canonical Wnt-β-catenin, can cause reduction in the eye fields and suppression of Rx and Six3 expression, when overexpressed in Xenopus embryos. On the other hand, Wnt11, which works through the non-canonical pathway, results in larger eyes in Xenopus when overexpressed (Cavodeassi et al., 2005). Misexpression of Wnt receptor Frizzled-3 (Fz3) in Xenopus results in the formation of multiple ectopic eyes. Fz3 is believed to preferentially activate the non-canonical Wnt pathway (Rasmussen et al., 2001). Wnt4, which probably acts through Fz3 receptor, is required for Xenopus eye formation (Maurus et al., 2005), by activating EAF2, which in turn regulates Rx expression in Xenopus. Loss of EAF2 function results in loss of eyes, while loss of Wnt-4 function can be rescued by EAF2 misexpression in frogs. Cell–cell signaling is also critical for the movement of eye-field precursors to the correct location prior to the activation of EFTFs. In Xenopus, all cells destined to form the eye field accumulate together through ephrin B1 signaling. This can be inhibited by fibroblast growth factors (FGFs), and activated FGF receptors modulate the activity of ephrin B by phosphorylating their intracellular domain (Moore et al., 2004). Thus, activating FGF signaling prior to gastrulation prevents cell movement and eye-field formation, whereas inhibiting FGF results in expansion of the eye primordial size.

RETINAL PROGENITORS: FROM OPTIC CUP TO RETINA The next phase of retinal development involves a massive proliferation of a group of cells that occupy the structure known as the optic cup (Figure 23.1). The optic cup forms from an involution of the optic vesicle, and as noted above, mutations in the EFTFs prevent the eye from progressing to this stage, or much beyond it. The cells of the optic cup resemble neural progenitors from other regions of the CNS (Figure 23.3). They have a simple bipolar morphology, span the width of the neuroepithelium, undergo mitosis at the scleral (ventricular) surface, and progress through stages of interkinetic nuclear migration during S-phase. These cells were once thought to be homogeneous, although more recently they have been shown to have distinct patterns of gene expression. Ever since the first birthdating studies of Sidman (1961), it has been consistently found that the different types of retinal neurons are generated in a sequence, with ganglion cells, cone photoreceptors, amacrine cells, and horizontal cells generated during early stages of development, and most rod photoreceptors, bipolar

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ROD BIP

Muller glial cell

Figure 23.3 Optic cup to neural retina. (a) The central regions of the optic cup contain a mix of mitotically active progenitor cells (d and e) at various stages of the cell cycle and differentiating ganglion cells (a, b, and c) and cones (f) (from Cajal, 1893). (b) Labeling the progeny of the mitotic cells at early stages of retinal development gives clones that contain all the cell types of the differentiated retina [(g, ganglion cell; a, amacrine cell; m, Müller glia; b, bipolar; p, photoreceptor (rod or cone) (from Holt et al., 1988)]. (c) Proposed model of retinal development. Birthdating studies in a variety of vertebrates indicate that there is a sequence to the production of the different retinal cell types, and that the progenitors have qualitatively different properties when isolated from early stages of development versus late stages of development. Müller glia are likely to be the last cell type generated in most species. cells, and Müller glia generated in the latter half of the period of retinogenesis. Clonal analysis of the progeny of these cells shows that they can give rise to all the different types of retinal neurons, and that the clones have mixed neuronal and glial lineages (Figure 23.3; for a review, see Cepko, 1993). Clonal analysis of retinal progenitor cells has also demonstrated a wide variety of clone size (Fekete et al., 1994), with some clones containing thousands of progeny.

Regenerative Medicine for Diseases of the Retina 423

The mechanisms that direct progenitor cells to different fates have been the subject of much investigation in the retina. While a thorough discussion of these findings is beyond the scope of this chapter, there have been two basic hypotheses proposed. First, it has been proposed that retinal progenitor cells undergo a progressive change during development that constrains them to a smaller range of fates (Reh and Kljavin, 1989; Figure 23.3). This implies that there is some sort of “molecular clock” keeping track of the developmental stage. The conserved birth order of the different classes of neurons can then be explained in that those cells that become post-mitotic at a particular stage of development are constrained to a specific cell fate. An alternative model is that a changing environment directs the cells to progressively later fates, but the progenitor cells themselves remain competent to generate all retinal cell types throughout the period of retinogenesis (James et al., 2003). There is an experimental support for both the models, and both the environment and intrinsic state of the cell are likely to be important factors in determining its ultimate fate (for a review, see Reh and Cagan, 1994; Livesey and Cepko, 2001). Are all retinal progenitor cells the same? Are there differences between the retinal progenitors and a more primitive retinal stem cell or founder cell? Several lines of evidence suggest that progenitors are not all identical. Proneural bHLH gene expression profiles appear to differ among the progenitors. For example, the bHLH transcription factor Ascl1, also known as Mash1 or Cash1, is expressed in only a subset of retinal progenitors (Jasoni et al., 1994; Jasoni and Reh, 1996). Two other proneural genes, Ngn2 and NeuroD1, appear to be expressed in subsets of progenitors as well (Nelson et al., unpublished observations). Another transcription factor, Foxn4, is also expressed in only a subset of retinal progenitors; this gene is thought to specifically bias progenitors to generate either amacrine or horizontal cells (Li et al., 2004). Retinal progenitors can also be distinguished by their response to growth factors and intracellular signaling. Progenitor cells isolated from late stages of embryonic development, or from neonatal retina are induced to differentiate by treatments that raise cAMP; progenitors isolated from early stages of embryogenesis have the opposite response – their proliferation is stimulated by increasing intracellular cAMP (Taylor and Reh, 1990). Progenitors isolated from the early embryonic retina are stimulated to proliferate by FGF, but are only minimally responsive to epidermal growth factor (EGF) or transforming growth factor-α (Anchan et al., 1991; Lillien and Cepko, 1992; Anchan and Reh, 1995). At later embryonic stages, and in the postnatal retina, the progenitors acquire a robust response to EGF (Anchan et al., 1991; Lillien and Cepko, 1992; Anchan and Reh, 1995). The evidence that all retinal progenitors are not identical is somewhat at variance with the fact that lineage studies have not demonstrated distinct subpopulations of progenitors that generate specific cell classes. One possibility is that the different types of progenitor cells can interconvert among themselves. There is some evidence for this type of inter-conversion; deletion of Ascl1 in mice results in an expansion of the number of progenitors (Akagi et al., 2004). More generally in the developing CNS it is thought that neural stem cells can convert from being FGF-responsive to being EGF-responsive (Ciccolini and Svendsen, 1998). In summary, the multipotent progenitors make up the majority of mitotically active cells in the embryonic retina. At early stages of retinal development, these cells are competent to generate the entire complement of retinal neurons and glia; however, at later stages of development, their progeny become restricted to rod photoreceptors, bipolar cells, and Müller glia. Although these cells are typically referred to as multipotent progenitors, those isolated from the early stages of retinogenesis could also be considered as retinal stem cells because (1) they generate all retinal cell types, (2) they can generate very large clones, and (3) many of their divisions are symmetric. In addition, several groups have shown that the early progenitors can be cultured as “neurospheres,” a capacity that neural stem cells are known to possess (see for example Klassen et al., 2004). As will be described later, the adult retina of some vertebrates continues to add new neurons and glia at the peripheral margin, and thus true retina stem cells exist. Presumably these cells were derived from a population of similar cells in the developing retina, but at this point there is no definitive way to distinguish the stem cells from the progenitors during retinogenesis.

424 CELLS AND TISSUE DEVELOPMENT

(a) Frogs and Fish CMZ

CMZ-derived retina

(d)

CB

Embryonic retina

(b) Birds

(e)

CMZ CB

(c) Mammals

CMZ-derived retina

Embryonic retina

Embryonic retina

(f)

CB

Figure 23.4 The CMZ of non-mammalian vertebrates contains retinal stem cells. (a–c) Diagramatic representations of the regions of the retina generated in embryonic development (blue) or by the CMZ (yellow). In frogs and fish, most of the retina of the adult animal is generated by the CMZ, whereas in chicks, only a small region is generated by the CMZ, and in mammals, this zone is absent. (d) The frog CMZ cells are labeled with H3-thymidine. Arrow points to the anterior-most point of labeling, where the CMZ joins with the ciliary epithelium (from Reh and Constantine-Paton, 1983). (e) Chick CMZ are labeled with BrdU and Islet1 to show new neurogenesis (small arrows – double labeled ganglion cells), as well as the point where the CMZ joins with ciliary epithelium (large arrow) (from Fischer and Reh, 1999). (f) Nestin-BrdU double labeling shows a CMZ-like zone (arrow) in a mouse that is haplo-insufficient for the patched Shh receptor (from Moshiri and Reh, 2004).

RETINAL STEM CELLS AND PERSISTENT PROGENITORS IN ADULT VERTEBRATES: THE CILIARY MARGINAL ZONE The development of the amphibian, fish, or avian retina is not complete after the embryonic or neonatal period. In these animals, the retina continues to add new neurons into adulthood. This process is most apparent in teleost fish, which shows a dramatic growth of the eye during their lifetime, of up to 100-fold. New retinal neurons are generated from a zone of cells at the peripheral margin of the retina, where it joins with the ciliary epithelium. These cells form a ring around the ciliary margin of the retina called the ciliary marginal zone (CMZ) (Hollyfield, 1968; Figure 23.4). The CMZ cells of non-mammalian vertebrates resemble the early

Regenerative Medicine for Diseases of the Retina 425

progenitor cells of the eye, and possibly even the “founder” cells of the optic vesicle. In fact, most of the retina of the mature frog (Reh and Constantine-Paton, 1983) and fish are generated by the CMZ cells. Wetts and Fraser (1988) carried out lineage-tracing studies of these cells, similar to those done in embryos. They found that these cells can give rise to clones that contain all types of retinal neurons, like those of the embryonic retina. Therefore, it is likely that the CMZ contains a population of true retinal stem cells. Recent molecular analysis of this region in frogs and chicks has shown that CMZ cells express most, if not all, of the EFTFs (Perron et al., 1998; Fischer and Reh, 2000; Wehman et al., 2005). The CMZ cells also express bHLH transcription factors, like Ngn2 and Ascl1 (Perron et al., 1998), and at least some of the CMZ cells respond to the same mitogenic growth factors as the embryonic progenitors (Mack and Fernald, 1993; Fischer and Reh, 2000; Moshiri et al., 2005). The CMZ is highly productive in fish and in some amphibians, but in birds it is greatly reduced, and is absent in mammals. Although the CMZ is robust in fish and amphibians, it is not known what percentage of the cells in this zone represents true retinal stem cells and what proportion of them are progenitors. In birds, most of the retina is generated during embryonic development and only a small number of retinal neurons are generated by the CMZ (Prada et al., 1991). It is not known whether this zone persists throughout the lifetime of a bird, but new retinal neurons are generated at peripheral edge of the retina in chickens up to 1 month of age (Fischer and Reh, 2000), and in the quail eye for up to a year (Kubota et al., 2002). In addition to their potential to generate new retinal neurons, the chicken CMZ cells express many of the EFTFs, including Pax6 and Chx10 (Fischer and Reh, 2000). The CMZ is greatly reduced or absent in the mammalian eye. Several groups have analyzed various species for evidence of ongoing proliferation in the margin of the retina, near its junction with the ciliary epithelium, but no mitotic cells are present in normal mice, rats, or macaques (Ahmad et al., 2000; Kubota et al., 2002; Moshiri and Reh, 2004). However, there is evidence that this zone may be repressed in the mammalian retina. Moshiri and Reh (2004) analyzed mice with a single functional allele of the patched gene, a negative regulator of Shh signaling. They found a small number of proliferating cells at the retinal margin of these mice, into adulthood (Figure 23.4F). Moreover, when the patched / mice were bred onto a background with photoreceptors degeneration, the proliferation in this zone was increased. This is reminiscent of the response to retinal damage observed in the CMZ cells of lower vertebrates, and suggests that the CMZ-like zone in patched mice has much in common with the CMZ of frogs and fish. In addition, recent studies have found that proliferation can be stimulated after the progenitor cells have normally withdrawn from the cell cycle in the neonatal mammalian retina by the injection of specific growth factors (Zhao et al., 2005; Close et al., 2005), suggesting that the proliferation at the retinal margin may be suppressed in the mammalian retina by factors in their microenvironment.

TRANSDIFFERENTIATION AND RETINAL REGENERATION The Pigmented Epithelium One of the most striking examples of regeneration in vertebrates is the regeneration of the newt eye. These animals are capable of remarkable regeneration of a variety of tissues, and the eye is no exception. The neural retina can be completely removed in these animals, and within 5 weeks it is restored and the animal can respond to a visual stimulus. Regeneration of the retina in newts, and in many other amphibians, occurs through a highly stereotypic process (Figure 23.5). Shortly after the retina is removed, the adjacent pigmented epithelial tissue re-enters the cell cycle (Stone, 1950; Reyer, 1971; Stroeva and Mitashov, 1983). The proliferating pigmented epithelial cells lose their pigmentation and begin to express markers of retinal progenitors; this process was one of the first examples of transdifferentiation (Okada, 1980). The de-differentiated pigment epithelial cells go on to generate new retinal neurons in a manner that resembles normal retinal histogenesis (Reyer, 1971; Reh et al., 1987; Sakaguchi et al., 1997). Over a period of just a few weeks, the developmental process is recapitulated and the new retinal ganglion cells re-grow connections with the brain.

426 CELLS AND TISSUE DEVELOPMENT

(a)

Neural retina

Pigmented epithelium

(d) Dissociate and culture (b)

(e)

(f)

(c)

Figure 23.5 Regeneration in amphibian retina from the RPE. (a–c) Sections through the regenerating newt retina from 2, 3, and 5 weeks after removal, showing the progressive restoration of the retina (from Sanae Sakami). (d) Schematic of technique for studying retinal regeneration in vitro. The pigmented epithelium can be dissected free from the neural retina, dissociated, and cultured. (e–f) In the presence of laminin or FGF, the pigmented cells lose their pigmentation and develop into spheres containing neurons and retinal progenitors (From Reh et al., 1987).

Most of the details of the cellular transformations that occur during retinal regeneration in amphibians have been well established for many years; however, only recently have there been studies into the molecular mechanisms underlying this process. The use of molecular markers has established that the de-differentiating pigment cells progress through a stage in which they resemble retinal progenitors (Reh et al., 1987; Sakami et al., 2005); however, it is possible that these cells go through a stage where they resemble stem or “founder” cells, because the RPE cells can regenerate the entire retina in some species, up to four complete times (Stone, 1950; Stone and Steinitz, 1957). A similar process of de-differentiation of the pigmented epithelial cells occurs in embryonic chick and mammals, and this also leads to retinal regeneration. However, the ability of RPE cells to transdifferentiate into retinal stem or progenitor cells is present only in the early stages of eye development (Pittack et al., 1997; Park and Hollenberg, 1993; Zhao et al., 2005; Coulombre and Coulombre, 1965). A key stimulus for retinal regeneration from the RPE in both amphibians and chick embryos is FGF. When added to cultures of pigment cells or in vivo, this factor stimulates the RPE cells to adopt a retinal progenitor identity (Park and Hollenberg, 1989, 1991; Pittack et al., 1997; Sakaguchi et al., 1997), and new, laminated retina is generated. Recent evidence also indicates that Shh is also playing a critical role in the process of RPE transdifferentiation (Spence et al., 2004).

Regenerative Medicine for Diseases of the Retina 427

The Ciliary Epithelium The ciliary body has also been proposed to harbor retinal stem cells or progenitors. This region of the eye is made up of derivatives of both the neural tube (the pigmented and non-pigmented ciliary epithelia; Figure 23.1) and the neural crest. The ciliary epithelia are developmentally analogous to the choroid plexus in the rest of the CNS. The non-pigmented ciliary epithelium is the anterior extension of the neural retina, whereas the pigmented ciliary epithelium is the anterior-most extension of the pigmented epithelium. It is likely that both regions have potential to generate neurons, at least in some animals. A few years ago, we found that intraocular injections of growth factors (insulin, FGF2, and EGF) stimulated the proliferation and ultimate neuronal differentiation of cells within the ciliary epithelium (Fischer and Reh, 2003). Like the CMZ, these cells also express the EFTFs Chx10 and Pax6. The neurons that develop in this region following growth factor treatments resemble amacrine cells, ganglion cells, and Müller glia, but they do not express markers of bipolar cells or photoreceptors. The mammalian ciliary epithelium also has some ability to generate neurons. Fischer and Reh (2001) described cells with neuronal and proliferation markers in the non-pigmented ciliary epithelium of the mature Macaque eye. Several groups have found that cells expressing neuronal markers can be generated from dissociated cell cultures of either the pigmented or non-pigmented cells of the ciliary epithelium (Ahmad et al., 2000; Tropepe et al., 2000; Das et al., 2005; Englehardt et al., 2005; Inoue et al., 2005). Tropepe et al. (2000) found that a small subpopulation of the pigmented cells form neurospheres and can be passaged to form new spheres. As a result of these characteristics, the cells have been termed “retinal stem cells” (Ahmad et al., 2000; Tropepe et al., 2000). Human eyes also contain these cells (Coles et al., 2004) and they can be grown in vitro for extended periods of time, expanded, and transplanted. The non-pigmented epithelial cells from mammalian eyes can also be maintained in vitro, and these cells are also capable of expressing neuronal markers. In these cases, however, the cells frequently express these markers without taking on the morphological characteristics of retinal neurons. Therefore, at this time, it is not known whether these cells will be useful for reconstructing functional retinal circuits, and more work needs to be done to assess the potential of these cells. The relationship between the sphere-forming pigmented cells and the true retinal stem cells present in the CMZ of fish and frogs is also not clear, since the latter are not thought to be pigmented. In addition, cells within the iris, the most anterior derivative of the primitive ocular neuroepithelium, are able to express photoreceptor genes when transfected with Crx (Haruta et al., 2001). Nearly all of these studies have been carried out in vitro, and ultimately it will be necessary to determine to what extent the cells that proliferate in these assays are truly acting as retinal stem cells, or whether they only activate a part of the neural gene expression profile. Intrinsic Stem Cells, Rod Precursors, and Müller Glia Müller glia are the primary glial cell intrinsic to the retina and the only glia generated by the multipotent retinal progenitors. They are among the last cell type generated during development, and genetic profiling studies have shown a great degree of similarity between the Müller glial cell and the retinal progenitor (Blackshaw et al., 2004). Despite their similarity, some key progenitor-specific genes are not expressed in Müller glia; for example, the Müller glial cell does not normally express proneural genes, like Ngn2 and Ascl-1. However, in post-hatch chickens and rodents, damage to the retina by neurotoxins causes some of the Müller glia to re-enter the cell cycle and re-express the proneural gene, Cash1 (Jasoni et al., 1994). Some of the proliferating Müller glia go on to generate cells that express markers and morphology of neurons (Fischer and Reh, 2002; Fischer et al., 2002a, b; Ooto et al., 2004), indicating that Müller glial re-entry into the cell cycle may initiate a regenerative process. Curiously, this regenerative response is largely abortive, since the majority of the Müller glial progeny remain as un-differentiated cells. At this point, it is unclear why so many of the cells do not replace the neurons destroyed by the neurotoxin. However, this is not the case in fish. Yurco and Cameron

428 CELLS AND TISSUE DEVELOPMENT

(2005) have found that lesioning the retina in mature zebrafish leads to Müller glial proliferation, much like that observed in the chick, but in fish, the repair of the retina is almost perfect (for a review, see Otteson and Hitchcock, 2003). Although it is not known what percentage of the new neurons are derived from Müller glia, as opposed to intrinsic stem cells or rod precursors, the regenerative process is very coordinated in fish. Retinal Neurons from ES Cells Although the retina of non-mammalian vertebrates has a variety of different strategies for repair, and these typically involve the de-differentiation of existing retinal cells into new retinal stem or progenitor cells, the sources for repair in mammalian retinas are more limited. A number of investigators have therefore attempted to transplant fetal retinal progenitors into the retinas of animals with retinal degenerations, with some success (Lund et al., 2003). However, it is difficult to imagine that fetal human retinal progenitors will ever be readily accessible, as fetal tissue has been limiting in other cell-based strategies elsewhere in the nervous system. Thus, several investigators are developing methods to direct human ES cells to a retinal progenitor and retinal neuron identity. In the next section, we will review the progress in this area. ES cells, derived from the inner cell mass of the blastocyst, can self-renew indefinitely under appropriate culture conditions (Thomson et al., 1998), and their ability to differentiate into most, if not all, cells in the body makes them an attractive alternative to endogenous retinal stem/progenitor cells for tissue engineering. Table 23.1 details the studies to date that have attempted to direct mouse ES cells into a retinal differentiation pathway (Zhao et al., 2005; Hirano et al., 2003; Meyer et al., 2004; Tabata et al., 2004; Ikeda et al., 2005; Sugie et al., 2005; Aoki et al., 2006). Some of the early work on neural induction involved the use of retinoic acid (RA) (Bain et al., 1995; Fraichard et al., 1995; Bain et al., 1996). RA has generalized neural fate-inducing properties, though it does bias cells to a more posterior neural identity (i.e. spinal cord). An alternate protocol has been to treat the mouse ES cell aggregates (embryoid bodies) with basic FGF (FGF-2) and a combination of insulin, transferrin, selenium, and fibronectin (ITSFn) (Okabe et al., 1996; Lee et al., 2000). This approach has yielded high proportion of neuroepithelial cells, which can then be induced to differentiate into neurons and glia. Some groups have tested a two multi-step protocol, with the rationale that once ES cells have been neuralized using one of the above two methods, subsequent placement of these cells in either a neurogenic environment in vitro, like dissociated newborn rat retinal cells (Zhao et al., 2005) and dissociated embryonic chicken retinal tissue (Sugie et al., 2005), or a degenerative retinal environment in vivo (Meyer et al., 2004) would further direct the cells to a retinal identity. These groups found that either RA or ITSFn/FGF-2 results in cells that express neural precursor markers like Pax6 and nestin. Upon co-culture with neurogenic retinal tissue, some cells even expressed markers of photoreceptor precursors like Crx and Nrl, but rarely expressed differentiated photoreceptor markers like rhodopsin and interstitial retinol-binding protein. Using a similar approach, coupled with transplantation into the posterior chamber of the eye in a mouse model of neuronal and photoreceptor degeneration, Meyer et al. (2004) found that the cells penetrated into the retinal layers and acquired neuronal-like morphology. However, the cells did not express any photoreceptor markers, though they did seem to promote survival of the remaining host photoreceptors (Meyer et al., 2004). An alternative approach to generating retinal progenitors from ES cell lines has employed stromal cell lines like PA6. The PA6 cell line has been shown to effectively induce neural differentiation in mouse ES cell lines (Kawasaki et al., 2000; Hirano et al., 2003; Yoshizaki et al., 2004; Aoki et al., 2006). The signaling molecule causing this effect is yet to be determined, but has been called SDIA (stromal cell-derived inducing activity). Eye-like structures from mouse ES cell lines can be formed by using the PA6 stromal cell line as a feeder layer (Hirano et al., 2003; Aoki et al., 2006). Researchers showed that culturing these ES cells in the presence of FGF-2 and dexamethasone along with cholera toxin for first 3 days resulted in differentiation into eye-like structures resembling the lens, the RPE, and the neural retina. The mechanism by which dexamethasone or

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Table 23.1 Papers published on induction of retinal fate in mouse, monkey and human embryonic stem cell. Cell Line

Method

Analysis

Reference

D3 mouse ES cell line

RA or ITSFn  FGF-2, followed by co-culture with P1 rat retina Co-culture with PA6 stromal cell line  FGF-2  dexamethasone  cholera toxin Transfection by electroporation with Rx gene, followed by RA or PA6 cells and retinal explant co-culture RA or ITSFn then FGF-2  laminin, followed by co-culture with E6 chicken retina Dkk1  lefty-A with addition of FCS  activin after 3 days for 3 days, followed by retinal explant co-culture RA followed by transplantation into retinal degeneration mouse Co-culture with PA6 stromal cell line  FGF-2  dexamethasone  cholera toxin, followed by Wnt2b and transplant into E2 chicken embryos Co-culture with PA6 stromal cells

ICC and RT-PCR

Zhao et al. (2005) Hirano et al. (2003) Tabata et al. (2004)

CCE mouse ES cell line

EB3 mouse ES cell line EB5 mouse ES cell line

CCE and D3 mouse ES cell line

CMK6 and CMK9 monkey ES cell line Cynomolgus monkey ES cell line Cynomolgus monkey ES cell line H1, H7, and H9 human ES cell line HES-1 human ES cell line H1 and H5&6 human ES cell line

FGF-2 followed by co-culture with PA6 stromal cells Co-culture with PA6 stromal cells, followed by transplantation in vivo Overgrowth under adherent conditions Noggin followed by FGF-2  EGF and transplant into adult and newborn rat retinas Dkk1, Noggin, IGFI for 3 days followed by Dkk1, Noggin, IGFI4 and FGF for 3 wks

ICC and RT-PCR ICC, RT-PCR, IHC, and electrophysiology ICC, IHC, and RT-PCR ICC and IHC

IHC IHC and RT-PCR

Sugie et al. (2005) Ikeda et al. (2005) Meyer et al. (2004) Aoki et al. (2006)

ICC and RT-PCR

Kawasaki et al. (2002) ICC and Western blot Ooto et al. (2003) ICC, IHC, RT-PCR, and Haruta et al. Western blot (2004) ICC, Western blot, RT-PCR, Klimanskaya and chip analysis et al. (2004) ICC, IHC, and RT-PCR Banin et al. (2006) ICC, IHC & RT-PCR Lamba et al. (2006)

RA, retinoic acid; ITSFn, combination of insulin  transferrin  selenium  fibronectin; ICC, immunocytochemistry; IHC, immunohistochemistry; RT-PCR, reverse transcriptase polymerase chain reaction; FCS, fetal calf serum.

cholera toxin cause this eye induction is not known. This combination results in cells expressing lens markers like crystalline and photoreceptor markers like rhodopsin and recoverin as well as pigmented cells. This effect was recently shown to be further enhanced by the addition of Wnt2b (Wnt13), a Wnt expressed in the CMZ and believed to play a role in the maintenance of retinal progenitor state (Kubo et al., 2003). Upon transplantation of these cells into embryonic chicken eyes, the cells integrated, but did not contribute to retinal neurons in the host. Similar experiments were carried out using primate ES cells with this same PA6 feeder layer. Cynomolgus monkey ES cells also differentiated into RPE-like cells (Kawasaki et al., 2002; Ooto et al., 2003; Haruta et al., 2004); transplantation of these cells into the retinas of a rat model of RPE dystrophy resulted in improved survival of the photoreceptor layer and some improvement in visual function. When the PA6 coculture experiments were carried out in serum-depleted media in the presence of FGF-2, a number of differentiated cells formed transparent bodies expressing α-crystallin and Pax6 characteristic of the lens.

430 CELLS AND TISSUE DEVELOPMENT

Another study combined the use of the PA6 stromal cell treatment with overexpression of EFTFs like Rx (Tabata et al., 2004). Cells that overexpressed EFTFs and were then subject to either RA or PA6 treatment were then co-cultured with retinal explants. These cells migrated into the host retina and some cells express neuronal and glial markers. One of the most efficient protocols for producing neural retina from mouse ES cells has relied on the same factors that are normally involved in neural and retinal induction during embryogenesis (see EFTF above) (Ikeda et al., 2005). Ikeda et al. (2005) used lefty-A, which is known to have a neural induction effect in animal experiments (Meno et al., 1997), dkk1, which induces anterior neural fates (del Barco Barrantes et al., 2003), and activin A, which has a role in inducing retinal genes as well as photoreceptor differentiation (Davis et al., 2000; Fuhrmann et al., 2000). This protocol resulted in almost 30% of all cells expressing Pax6 and Rx. Upon co-culture of these cells with re-aggregated adult retinal neurons, a large proportion of these cells expressed rhodopsin and recoverin, markers of photoreceptors. Transplantation of these cells onto retinal explants resulted in their integration into host retina in vitro. Recently, a group was able to produce Pax6 expressing cells more efficiently from human ES cells (Banin et al., 2006). They did this by culturing the cells on mitotically inactivated mouse fibroblasts in the presence of noggin for 8 days. The cells were then passaged and cultured in the presence of FGF-2 and EGF. Almost 30% cells expressed Pax6, though very few (1%) of all cells expressed other retinal markers like Chx10 or Crx. Upon transplantation into adult and newborn rats, few of the transplanted cells expressed rhodopsin and Nrl. Although this study is encouraging, the absence of large number of cells expressing any of the other EFTF markers suggests that most of the cells may not have been retinal progenitors; Pax6 is also expressed in spinal cord, olfactory system, and the forebrain. Recently, our lab has developed a protocol using a combination of dkki, Naggin, IGF-1 and bFGF to efficiently induce hES cells to take up retinal progenitor fate (~80% of the cells and express various EFTFs (Lamba et al., 2006). While the aforementioned studies have concentrated on the production of retinal progenitors and photoreceptors from ES cells, several groups have also developed protocols for the production of pigmented epithelial cells from ES cells. Klimanskaya et al. (2004) found that overgrowing human ES cells in the presence of mouse fibroblast feeder layer resulted in spontaneous differentiation of pigmented epithelial cells in the absence of any factors or signaling molecules. Colonies of pigmented cells could be manually picked and analyzed for markers of RPE proteins. The cells could also be expanded to generate large numbers of pigmented epithelial cells. These cells could potentially find a use in repair of the RPE layer in individuals with age-related macular degeneration.

CONCLUSION Retinal diseases that cause blindness through the loss of one or more retinal neuron type are becoming increasingly common in the population. The ability to treat blindness by cell-replacement therapy would therefore be a useful addition to the research efforts in the prevention and treatment of blindness. The retina has been a classic model for regeneration studies, particularly in lower vertebrates, and focused efforts to uncover similar mechanisms in mammals are meeting with some success. While an effective cell-based therapy is still many years away, there are several promising approaches such as (1) stimulating endogenous repair, (2) harvest, in vitro expansion, and transplantation of adult stem cells from the eye, and (3) directing human ES cells to a retinal identity for transplantation. Toward these goals, it is clear that a number of key pieces of biology need to be better understood. For example, at the present time we cannot distinguish between stem cells and progenitors at any stage of development or in adult animals. Although there are some assays that claim to distinguish between these two potential types of cells in other regions of the nervous system, these typically rely on the fact that stem cells are multipotent and self-renewing. In the retina, lineage analysis during development has shown that the

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majority of dividing cells are multipotent and can make a variety of different clone sizes, some reaching many thousands of cells. Moreover, both symmetric (e.g. two dividing cells) and asymmetric (dividing cell and neuron) divisions are equally common. Thus, at present, in vitro assays that show multipotentiality or selfrenewal based on sub-cloning procedures cannot discriminate among the different types of dividing cells in the retina, as they have been used to do in other areas of the CNS. We also do not understand the process of de-differentiation or cellular plasticity. RPE can de-differentiate into stem or progenitor cells that generate an entire new retina in some species. In mammals, pigmented cells can lose their pigmentation in vitro and go on to express a variety of proteins normally present only in the neural retina; however, the cells do not appear to recapitulate the entire program of regeneration, making a new layer retina, and most of the cells generated in these cultures do not resemble neurons morphologically or functionally. Are there intrinsic limitations to the potential of these cells to generate true functional neurons in mammals, or are necessary factors in the local microenvironment not present in the damaged mammalian retina? After either neurotoxic or surgical damage, both fish and birds can generate new neurons from intrinsic sources. In fish, the intrinsic retinal source of regeneration may include rod progenitors, an intrinsic stem cell, and/or Müller glia. In the case of the chick retina, the Müller glia appear to be the only source of the new neurons. However, there is a large difference in the regenerative responses between these two animals. In fish, the regeneration is nearly perfect, whereas in birds, most of the proliferating glia do not go on to make new neurons, but rather remain in an undifferentiated state. What is the block to efficient regeneration in the bird? Further studies of the factors that regulate Müller glial proliferation after damage in both chick and mammals may lead to clues for stimulating the process of regeneration. Moreover, gene expression profiles between Müller glia and retinal progenitors may also lead to a better understanding of the process of de-differentiation in Müller glia that precedes neuronal regeneration. Lastly, the future of retinal repair may well require the transplantation of retinal cells that have been generated from progenitors or stem cells in vitro. While the work on human ES cells is proceeding at a rapid pace, there are still some fundamental questions that will need to be resolved before a cell-based transplantation therapy can become a reality. The transplantation studies using fetal cells that have been carried out over the past two decades suggest that survival and integration of the transplanted cells may be two key barriers to functional restoration of degenerated retina. Moreover, the in vitro expansion and appropriate cell-type differentiation of retinal progenitors, either derived from ES cells or an adult retinal cell, will require a better understanding of the factors that normally control retinal cell fate during development. We have come a long way in our understanding of the phenomenon of retinal regeneration, far enough to appreciate the enormity of the task ahead for the translation of this knowledge to clinical practice.

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24 Peripheral Blood Stem Cells Shay Soker, Gunter Schuch, and J. Koudy Williams INTRODUCTION Adult stem and progenitor cells have been isolated from a wide variety of sources. Lineagecommitted progenitors are found in a number of tissues including skin, fat, muscle, heart, brain, liver, pancreas, bladder, and so on. Adults have another source of stem and progenitor cells that are not restricted to a specific tissue. These “universal” stem and progenitor cells are found circulating in peripheral blood, allowing them to reach and integrate into all tissues. The bone marrow is, most likely, the source of peripheral blood stem and progenitor cells. Hemangioblasts are the embryonic precursors of hematopoietic stem cells (HSC), giving rise to committed hematopoietic progenitors such as lymphoids, thymocytes, myeloids, granulocytes–monocytes, megakaryocytes–erythrocytes, and mast cells. These progenitor cells complete their differentiation in the bone marrow, peripheral blood, and thymus and in the target tissues. Extensive research in hematology/oncology has resulted in the identification of a wide variety of cell surface markers that allow the characterization and isolation of HSC at different stages of their differentiation. Initially, adult bone marrow mesenchymal cells (MSC) were isolated, expanded in vitro, and examined for their multilineage differentiation potentials. These early studies were followed by extensive research on bone marrow-derived multipotent adult progenitor cells (MAPC). This special cell population can proliferate long-term without senescence and can differentiate to multiple lineages in vitro and contribute to the regeneration of several tissues in vivo (Verfaillie, 2005). Like HSC, MSC may leave the bone marrow environment and be found in peripheral blood. Identification and isolation of MSC is based on differential expression of cell surface markers that distinguish them from circulating HSC. Among the most studied circulating MSC are the endothelial progenitor cells (EPC). This population is probably derived from the same hemangioblasts precursors of HSC, but they take a separate path of differentiation in the bone marrow. The identification of circulating EPC suggested that the process of vasculogenesis, previously believed to be restricted to the embryonic stages, continues into adulthood. The circulating EPC have specific cell surface markers that are not found on mature endothelial cells (EC) and lose them when they differentiate to EC. This chapter will briefly review the types and source of stem cells in peripheral blood, their specific cell surface markers, and factors that change their abundance in peripheral blood. We will focus on the isolation and in vitro expansion of peripheral blood-derived MSC and EPC and describe their therapeutic applications for regenerative medicine. We will further describe the role of peripheral blood-derived stem cells in normal and pathological processes. Although much information was gathered in the past on the identification of different populations of peripheral blood stem cells, their clinical potential for therapy is just now being explored. Since peripheral blood is readily obtainable, it can be as a viable source of cells for regenerative medicine deserves special attention. TYPES AND SOURCE OF STEM CELLS IN THE PERIPHERAL BLOOD It is well documented that the bone marrow is the major source of cells in peripheral blood. HSC are characteristically quiescent, multipotent cells, with the capacity for both self-renewal and differentiation. After

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development in the fetus, HSC reside in adult bone marrow and serve to replenish lymphoid, megakaryocytic, erythroid, and myeloid hematopoietic lineages throughout adulthood. Observations that systemically administrated MSC could come back to the bone marrow suggested that MSC may also reside in the bone marrow. Results of our recent studies indicate that mature cells, such as EC, may enter the circulation (Beaudry et al., 2005). A different study tested the fate of muscle progenitor cells introduced into the circulation of lethally irradiated recipient mice together with distinguishable bone marrow cells. All recipients showed high-level engraftment of muscle-derived cells representing all major adult blood lineages (Goodell et al., 2001). Collectively these results indicate that there is a constant exchange of cells from the bone marrow to peripheral blood. On the other hand, bone marrow transplantation studies have indicated that this process may be reversed and cells from peripheral blood may repopulate the bone marrow. Mobilization of Bone Marrow Cells Stem cell numbers in peripheral blood are very low compared to those in the bone marrow. Although stem cells can be collected by apharesis, this requires the processing of large volumes of blood. Amplification of peripheral blood stem cells can facilitate collection and allows for rescuing autologous stem cell from the bone marrow. Mobilization of HSC from bone marrow into peripheral blood can be achieved by hematopoietic growth factors. Recombinant human granulocyte (G)- or granulocyte-macrophage (GM) colony-stimulating factor (CSF) have been used as stimulators of hematopoiesis. Results of studies indicate higher numbers of circulating progenitor cells in patients receiving G-CSF or GM-CSF (Gianni et al., 1990; Baumann et al., 1993; Kawano et al., 1993). In fact, transplantation of G-CSF mobilized stem cells harvested from peripheral blood is replacing bone marrow biopsy, the method of choice for collection of stem cells for autologous bone marrow transplantation. However, it is important to find better mobilizing techniques to provide more efficient harvesting and faster hematopoietic recovery. Recently, elegant studies were designed to prove the role of angiogenic factors in EPC mobilization. Rafii and colleagues reported that mobilization of HSC and EPC from bone marrow is mediated through the activation of metalloproteinases and adhesion molecules (Eriksson and Alitalo, 2002; Hattori et al., 2002; Heissig et al., 2002; Rafii and Lyden, 2003). In the bone marrow, vascular endothelial growth factor (VEGF) and placental growth factor (PlGF) induce MMP-9 expression. Activation of MMP-9 results in the release of stem cell-active soluble kit ligand, which mobilizes quiescent HSC and EPC to the vascular zone where they are released to the circulation. The results of these studies indicate that co-mobilization of EPC and HSC contribute to the revascularization processes.

EPC Initial evidence that EPC can be detected in peripheral blood came from research conducted mainly by the groups of Isner and Asahara in Boston and Rafii in New York (Rafii et al., 1995b; Asahara et al., 1997). They showed that cells with EC characteristics can be isolated from peripheral blood and expanded in vitro. They and others have shown that the numbers of EPC in peripheral blood are significantly increased as a result of acute vascular injuries, angiogenic stimuli, and estrogen and nitric oxide (NO) synthase, but reduced by certain chronic disease states (e.g. coronary artery disease) (Gill et al., 2001). Circulating EPC originate primarily from the bone marrow and can be identified by differential expression of hematopoietic and EC markers. This is important because hematopoietic and EPC probably share a common precursor, the hemangioblasts (Hirschi and Goodell, 2001). Hemangioblasts reside mainly in the bone marrow and differentiate into HSC and angioblasts. This process occurs mainly during early embryogenesis but was shown to exist in adults (Gill et al., 2001; Hattori et al., 2001, 2002). Angioblasts will give rise to EPC that upon stimulation with angiogenic factors such as VEGF and PlGF are mobilized from bone marrow to peripheral blood (Gill et al., 2001;

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Shintani et al., 2001; Heissig et al., 2002; Rafii et al., 2002a). Once in peripheral blood, EPC can be recruited to sites of active neovascularization as seen in wounds, diabetic retinopathy, and tumors (review in Rafii and Lyden, 2003). The role of EPC in physiological and pathological neovascularization and their therapeutic applications are described below. Identification and Isolation of EPC Marrow and peripheral blood cells expressing CD34 can give rise to EPC (Asahara et al., 1997; Shi et al., 1998; Bhattacharya et al., 2000; Peichev et al., 2000; Dimmeler et al., 2001). Although CD34 is commonly used to isolate EPC, CD34 expression is also shared by HSC and MSC and cannot be used to distinguish between these populations. Likewise, VEGF receptor 2 (human KDR and mouse Flk-1), which is used to identify EPC, is expressed also on HSC (Asahara et al., 1997; Isner and Asahara, 1999). In humans, CD133 (AC133) is used to distinguish EPC from mature EC, since CD133 is not expressed by mature EC (Peichev et al., 2000; Rafii et al., 2002b). CD133 is a stem cell marker with as yet unrecognized functions (Rafii, 2000). Additionally, Hebbel and colleagues have used P1H12 antibodies that recognize CD146 (MUC18) on circulating EC (CEC) in peripheral blood but not on monocytes, granulocytes, platelets, megakaryocytes, or T- or B-lymphocytes (Solovey et al., 1997, 2001; Sodian et al., 2000). Other markers common to progenitor and mature EC are the cell surface receptors KDR and Tie2 (Rafii and Lyden, 2003; Asahara and Kawamoto, 2004; Ishikawa and Asahara, 2004). Purified populations of CD133/KDR EPC proliferate in vitro in an anchorage-independent manner and can be induced to differentiate into mature adherent EC (Rafii and Lyden, 2003). It is thought that CD133/KDR EPC are a population of immature EC that are mobilized from the bone marrow to participate in neovascularization. As myelomonocytic cells have lost surface expression of CD133, this marker also provides an effective means to distinguish true EPC from cells of myelomonocytic origin. Yet, recent studies showed that cells expressing CD14, considered as a typical monocytic lineage marker, can give rise to EC (Kim et al., 2005; Romagnani et al., 2005). Collectively, these studies suggest that identification of circulating EPC may be achieved using different markers that may define subpopulations of EPC based on their differentiation stage and origin. The number of EPC in bone marrow is very low, 10 per 10  105 mononuclear cells, and the reported numbers vary a great deal, based on which identifying markers are used among the different studies. For practical applications, EPC fraction may be enriched using cell surface markers such as CD34, CD133, and KDR (Asahara et al., 1997; Shi et al., 1998; Ishikawa and Asahara, 2004). One functional assay capitalizes on in vitro growth kinetics to discriminate bone marrow-derived EPC and CEC from vessel wall-derived mature EC (Rafii and Lyden, 2003). In this assay, the isolated cells are incubated with VEGF, basic fibroblast growth factor (bFGF), insulin-like growth factor (IGF), and fibronectin or collagen. EC colonies that appear early are derived from the recipient vessel wall CEC, whereas late-outgrowth cells or colonies originate mainly from bone marrow-derived EPC. Therefore, late-outgrowth endothelial colonies (CFU-EC) may be considered as angioblast-like EPC. Results of recent studies indicate that VEGF acts at different levels in the bone marrow to increase the number of EPC found in peripheral blood. Besides its known activities to induce EC proliferation and migration, VEGF induces secretion of EC MMP-9. These factors stimulate the release of soluble kit ligand which promotes proliferation and migration of EPC into the vascular zone of the bone marrow (Heissig et al., 2002; Rafii et al., 2002a). In Vitro Expansion of EPC Most studies derive EPC from the mononuclear fraction of bone marrow and peripheral blood. The mononuclear fraction is placed in fibronectin-coated plates containing endothelial basal medium which contain angiogenic growth factors such as VEGF and bFGF. Other growth factors such as epidermal growth factor and IGF contribute to cells’ growth but not differentiation. In one of their earlier studies, Asahara et al. (1999b) showed

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Table 24.1 Cell surface markers expressed on progenitor and mature endothelial cells

Proliferative capacity Proposed source and mechanism for mobilization Markers VEGFR-2 (KDR) CD34 CD31 (PECAM) AC133 (CD133) MUC18 (P1H12)

Putative EPC

Vessel wall derived CEC

High Bone marrow release and proliferation to VEGF and other stimuli

Limited EC damage. VEGF decreases apoptosis

  /  

    

that VEGF and not bFGF is important for EPC differentiation, and bFGF may be used by the differentiated EC for subsequent proliferation. Inclusion of angiogenic factors in the media helps to prevent “contamination” by other cell types, including lymphocytes, macrophages, and dendritic cells. VEGF appears to inhibit dendritic cell maturation from CD34 MNC fraction (Gabrilovich et al., 1996, 1998, 1999). Within 7–10 days of culture in fibronectin or collagen-coated dishes, colonies with spindle-shape cells appear in the dish. These are “slow growing” cells defined as “late-outgrowth” EPC. They differ from the mature CEC that are readily proliferate in vitro (Gill et al., 2001). EC cultures from EPC may be obtained after 2–3 weeks. The cells assume a typical flat EC morphology and present mature EC markers such as CD31, VE-cadherin, and CD146 (P1H12). They metabolize acetylated low-density lipoprotein (acLDL), bind Ulex Europaeus agglutinin 1 (UEA-1), and produce NO, consistent with EC properties. Proper characterization of EPC-derived EC requires the analysis of a combination of cell surface markers that can be measured by fluorescent antibody flow cytometry (Table 24.1). The Role of EPC in Physiological and Pathological Neovascularization Blood vessels form by two processes: (1) angiogenesis, the sprouting of capillaries from preexisting blood vessels, and (2) vasculogenesis, the in situ assembly of capillaries from undifferentiated EC. Vasculogenesis takes place mostly during the early stages of embryogenesis (Folkman and D’Amore, 1996; Yancopoulos et al., 1998). Vascular channels in the yolk sac originate from the mesoderm by differentiation of angioblasts, which subsequently generate primitive blood vessels (Breier et al., 1997). The early findings that EPC can participate in angiogenic processes indicate that postnatal neovascularization does not rely only on sprouting from preexisting blood vessels (angiogenesis), but may be assisted by EPC via postnatal vasculogenesis (Asahara et al., 1999a, b; Takahashi et al., 1999; Young et al., 1999). VEGF has an important role in angiogenesis, but new studies suggest that it also has a role in promoting adult vasculogenesis. Administration of VEGF in vivo by protein injection, DNA transfection, or adenovirus (Ad) infection results in a rapid and transient elevation of CEC numbers (Asahara et al., 1999b; Schuch et al., 2002; Beaudry et al., 2005). In burn and coronary artery bypass grafting patients, plasma VEGF upregulation was correlated with transient increase in the number of CEC (Gill et al., 2001). We observed that implantation of encapsulated cells secreting high levels of VEGF significantly induced EPC mobilization in mice, as measured by the number of CEC and β-galactozidase (LacZ) expressing MNC from Tie2/LacZ mice (Schuch et al., 2003). Continuous release of VEGF resulted in the formation of a large number of EPC colonies, which expressed specific EC markers such as KDR when cultured in the presence of VEGF. Bone marrow-derived EPC contribute to adult tissue neovascularization in several models including wound healing, cornea, and tumor angiogenesis (Asahara et al., 1999a; Rafii et al., 2002c). Bone marrow-derived EPC

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could be detected in normal organs, including spleen, lung, liver, intestine, skin, hind limb muscle, ovary, and uterus, indicating their participation in the maintenance of physiological neovascularization (Asahara et al., 1999a). Hormonally induced ovulation cycles were also associated with localization of bone marrow-derived EPC in corpus lutea and in the uterus endometrium and stroma. These findings indicate that EPC contribute to physiological neovascularization associated with postnatal regenerative processes. A recent study examined the presence of endothelial, smooth muscle, and Schwann cell chimerism in patients with sex-mismatched (female-to-male) heart transplants (Minami et al., 2005). The Y chromosome was used to determine chimerism. Biopsy specimens taken at increasing times after heart transplantation showed that EC had the highest degree of chimerism (24.3%), Schwann cells showed the next highest chimerism (11.2%), and vascular smooth muscle cells (SMC) the lowest (3.4%). Results of this study indicate that circulating progenitor cells are capable of repopulating most major cell types in the heart, but they do so with varying frequency. The signals for endothelial progenitor recruitment occur early and could relate to the injury during the surgery. In parallel, EPC were found incorporated into the vasculature of pathological lesions such as atherosclerotic plaques, tumors, the retina, and ischemic brain tissue. Vascular SMC proliferation results in neointimal hyperplasia and the development of restenosis. Bone marrow-derived SMC can integrate into the hyperplastic neointima and atherosclerotic plaques (Luttun et al., 2002; Sata et al., 2002). Evidence for the contribution of bone marrow MSC to human atherosclerotic plaques originated from a study showing donor-derived neointimal cells within the plaques (Caplice et al., 2003). Also, decreased EPC in the circulation have been correlated with a higher risk of cardiovascular complications (Hill et al., 2003). It was hypothesized that lower levels of peripheral blood EPC were associated with an impaired capacity to repair the damaged vessels, but the pathophysiological role of bone marrow-derived EPC remains unclear. Recruitment of peripheral blood EPC to damaged or diseased tissues is dependent on the underlying pathology and is probably due to the release of specific growth factors and chemokines by these tissues (Hillebrands et al., 2001, 2002). Abnormal retinal neovascularization contributes to the pathogenesis of proliferative retinopathy in diabetes and age-related prematurity and macular degeneration. Bone marrow-derived hemangioblasts were shown to contribute to retinal neovascularization in models of proliferative retinopathy (Grant et al., 2002; Otani et al., 2002). This study documented the incorporation of EPC into mature endothelium of the retinal blood vessels. Cerebral infraction is associated with neovascularization of the ischemic zone and new vessel growth. Bone marrow transplantation studies showed that EPC could be detected in the neovessels at the repair sites after 3 days (Hess et al., 2002; Zhang et al., 2002). Taken together, the results of these studies indicate that EPC’s contribution to neovascularization is not restricted to normal healing processes and they contribute significantly to several pathological processes. One of the most intensively studied models of EPC and neovascularization is tumor angiogenesis, as described below. The Role of EPC in Tumor Growth Compelling evidence for the role of EPC in tumor vascularization comes from a study by Lyden and colleagues using an angiogenesis-defective mouse model. Mice lacking both alleles of Id1 (id1/) and Id3 (id3/) died by embryonic day 13.5 and exhibited massive vascular malformation (Lyden et al., 1999). The Id3//id1/ mice survived but could not support the growth of several tumor types due to insufficient tumor vascularization. However, transplantation of id3//id1/ mutant mice with bone marrow from wild-type mice gave rise to tumors that were indistinguishable from tumors grown on wild-type mice (Lyden et al., 2001). Furthermore, 90% of the tumor vessels contained bone marrow-derived EC, indicating the contribution of EPC to tumor neovascularization. VEGF treatment failed to elevate the number of EPC in id3//id1/ mutant mice but not in id3//id1/ transplanted with wild-type bone marrow. Further evidence is provided by a model in which transplantation of human bone marrow-derived MAPC into tumor xenograft-bearing mice resulted in the

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incorporation of human cells as 40% of the tumor vessel endothelium, indicating the importance of CEC for tumor neovascularization (Reyes et al., 2002). Different tumors secrete different types and concentrations of angiogenic factors that may have a different capability to induce mobilization of EPC. Although a formal correlation between tumor type/stage/size and number of EPC has not been established in human cancer, some tumor types may be more dependent than others on CEC as a source of endothelium (Rafii et al., 2002b). EPC As a Surrogate Marker Vascularization is a crucial factor for tumors to grow and metastasize. The recent observation that the tumor vascular network is a combination of both angiogenesis and vasculogenesis requires a more complex understanding of this process. The notion that EC are present in the circulation and can contribute to neovascularization has implications for the development of therapeutic agents for cancer and implementation of these agents into clinical trials. Results of studies have shown that cancer patients have a higher number of EPC in their blood compared to healthy volunteers and suggest that these cells may play a role in tumor neovascularization in human cancers (Mancuso et al., 2001). We have recently tested the effects of angiogenic and antiangiogenic factors on EPC in mice (Schuch et al., 2003; Beaudry et al., 2005). Unlike cytotoxic agents used for chemotherapy, antiangiogenic treatment is intended to specifically target the tumor vasculature (Figure 24.1). Antiangiogenic treatment does not reduce tumor volumes significantly in a short period of time. This presents a difficulty in the assessment of the efficiency of the antiangiogenic treatment. Thus, there is an urgent need for surrogate markers to assess the potential benefit of antiangiogenic therapy. Measurement of EPC numbers may represent such a marker. We observed that mice treated with VEGF had elevated numbers of CEC, whereas co-administration of VEGF and endostatin significantly reduced these numbers. We observed a significant change in CEC numbers as early as 5 days after initiation of endostatin injections. In order to validate these results we have analyzed EPC in a Tie2/LacZ transgenic mouse model (Schlaeger et al., 1997), where EC can be stained blue, and found similar effects. In another study we investigated changes in circulating

Figure 24.1 Tumors secrete angiogenic factors and cytokines that induce mobilization, differentiation, and integration of bone marrow-derived EPC (white circles) into tumor blood vessels. In contrast, mature EC can be released from the tumor blood vessels into the circulation (gray circles), where they undergo apoptosis (black circles). Antiangiogenic therapy may enhance shedding and apoptosis of CEC (Beaudry et al., 2005).

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mature EC and EPC after treatment with ZD6474 (Beaudry et al., 2005). This is a VEGFR-2 tyrosine kinase inhibitor which has been shown to inhibit angiogenesis and slow tumor growth in a broad range of mouse models and is currently undergoing clinical testing for patients with lung cancer and other types of solid tumors (Holden et al., 2005). Our results indicated that ZD6474 may have differential effects on circulating mature EC and EPC. In Lewis lung carcinoma-bearing mice, ZD6474 treatment inhibited of tumor angiogenesis and was accompanied by early increase in mature EC and reduced EPC. Taken together, these results indicate that EPC are a target for antiangiogenic drugs and that their relative numbers may serve as a surrogate marker for the bioactivity of antiangiogenic drugs.

MSC MSC are multipotent cells that can differentiate into mesenchymal lineages including bone, cartilage, fat, and muscle. MSC were initially found in adult bone marrow (Friedenstein et al., 1987; Caplan, 1991), and were first identified as osteogenic progenitors capable of forming bone-like structures in vitro (Friedenstein, 1976; Owen, 1988). These early studies suggested that bone marrow MSC are also adipogenic progenitors (Caplan, 1994). Further studies report that MSC may be found in every mesenchymal tissue that has regeneration capacity. In addition to bone marrow, MSC were isolated from muscle, fat, skin, cartilage, bone, and blood vessels (Peng and Huard, 2003; Bartsch et al., 2005). MSC have some of the basic properties of stem cells including self-renewal, multilineage differentiation capacity, clonality, and the ability to regenerate tissues in vivo (Verfaillie, 2002a, b; Roufosse et al., 2004). In addition, Verfaillie and colleagues have shown that adult bone marrow MSC proliferate for many passages without senescence. They analyzed telomere length in these cells and showed that it was longer than in neutrophils and lymphocytes and was not different among young or old donors (Reyes and Verfaillie, 2001). Their results indicated that bone marrow MSC have high telomerase activity in vivo and came from a population of quiescent cells. Identification and Isolation Because of the multiple sources and methods of isolation of MSC, their identifying markers vary between researches. Some of the “classical” markers of bone marrow-derived MSC include CD34, CD44, CD45, c-kit, Sca-1 (murine), CD133 (human) and CD105 (Thy-1), and higher concentrations of CD13 and stage-specific antigen I (SSEA-I) (Jiang et al., 2002). As stated above, MSC were isolated from multiple sources but only a few studies have analyzed their presence in peripheral blood. Systemic infusion of MSC showed that they may be engrafted in various mesenchymal tissues. These results suggest that MSC may be present in peripheral blood. In fact, MSC were isolated from peripheral blood of cancer patients who were given G-CSF and GM-CSF. The cells were grown in vitro and had a fibroblast-like phenotype (Fernandez et al., 1997). The cells were negative for hematopoietic markers and CD34, but expressed CD105, SH3, I-CAM, and V-CAM. MSC were also isolated from normal human peripheral blood without “mobilization” (Zvaifler et al., 2000). The cells were isolated by gradient centrifugation and plated in growth media. After 2 weeks, adherent fibroblast-like cells appear in the culture. These cells were positive for CD105, Stro-1, vimentin, and BMP receptors, but were negative for CD34. Taken together, these results indicate that a small population of MSC exists in peripheral blood. These cells are difficult to isolate, but may be identified by their morphology and the expression of a subset of MSC markers. In Vitro Expansion Peripheral blood-derived MSC are obtained through density centrifugation using Histopaque™ or Ficoll™. There are several factors that are important for successful maintenance of MSC, including cell density, pH of the medium, source of sera, and the type of culture dishes. Human MSC require densities of 1,500–3,000 cells/cm2 in

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order to prevent spontaneous differentiation at higher cell densities. The basal media may be DMEM or α-MEM with 10% fetal serum (Kuznetsov et al., 2001). Collectively, the methods used for MSC expansion in vitro do not differ from those used to expand bone marrow-derived MSC. Reviews from Catherine Verfaillie and Arnold Caplan describe these methods in detail (Caplan and Bruder, 2001; Verfaillie et al., 2003). Following expansion, MSC can be differentiated in vitro into the mesenchymal lineages and tested in vivo. Interestingly, marrow-derived MSC were induced to differentiate into cells with functional properties of EC (Reyes et al., 2002), hepatocytes (Schwartz et al., 2002), and neuroectodermal cells (Jiang et al., 2003). In vitro differentiated cells may be used for future therapeutic applications. However, we need to define the appropriate phenotype and functional properties of the differentiated cells before they can be used clinically.

HSC HSC constitute a very small pool of undifferentiated cells that divide, and have the capability to differentiate into committed progenitor cells for most of all lymphoid and myeloid cell lineages (Lu et al., 1996). The frequency of HSC among bone marrow cells has been variously estimated at between 1 per 10,000–100,000 cells. Four tissue sources of HSC are bone marrow, umbilical cord blood, fetal liver, and adult peripheral blood. Importantly, they are capable of reconstituting the hematopoietic system of a lethally irradiated recipient (Suda et al., 1983; Sutherland et al., 1989). Identification and Isolation In humans, there have been numerous attempts to purify or enrich HSC using density gradient centrifugation and cell sorting based on cell surface marker expression. CD34 is a marker for human stem and progenitor cells. However, it is not specific for HSC (Brandt et al., 1988; Bernstein et al., 1991; Verfaillie, 1992; Lu et al., 1996). Other markers used to enrich HSC are CD38, CD33, CD133, and CD117 (c-kit). Ex Vivo Expansion The hematopoietic microenvironment is dependent on non-hematopoietic cells in the bone marrow to support and regulate hematopoiesis. The marrow stroma is composed of fibroblasts, EC, macrophages, and other cells that are responsible for the production of an extracellular matrix and hematopoietic growth factors (Dexter and Fairbairn, 1993). Dexter-type long-term bone marrow cultures, which are stromal cell-dependent long-term cultures, are thought to mimic the marrow microenvironment closely. Primitive progenitors can differentiate and be maintained when these cells are non-contact co-cultured with stromal layers (Verfaillie, 1992). Long-term HSC cultures can be established with CD34 cells in a stroma-free system when defined cytokines are repeatedly added. Cytokines thought to be important in the induction of differentiation and/or proliferation of these primitive hematopoietic progenitors include G-CSF, GM-CSF, interleukin (IL)-l, IL-3, IL6, IL-11, stem cell factor (SCF), and steel factor (Dexter and Heyworth, 1994). Although there is great interest in the ex vivo expansion of HSC for a variety of applications, ex vivo maintenance and generation of functional hematopoietic cells are complex processes and are poorly understood. For example, it is not clear whether the earliest progenitors are the cells being expanded using currently established protocols.

THERAPEUTIC APPLICATIONS OF PERIPHERAL BLOOD STEM CELLS The physiological role of MSC in tissue regeneration prompted researchers to evaluate their use in therapeutic applications. The ethical discussions regarding embryonic stem cells underscore the need to explore the clinical applications of adult stem cells, including MSC. MSC were first tested in several animal models and have recently been used in clinical studies. Although the results of the animal experiments are promising, the

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mechanisms behind the regenerative potential of peripheral blood MSC are not fully understood. The therapeutic applications can be divided into three groups: (1) tissue engineering, (2) cell delivery applications, and (3) MSC as a vehicle for gene therapy (review in Rafii and Lyden, 2003). The main advantage of MSC for therapeutic use is their presence in peripheral blood. However, as discussed above, further work is needed to evaluate their culture and expansion properties. EPC In many cases, organ and tissue regeneration require reestablishment of the vascular network. There are two possible sources of endothelialization: (1) mature EC that migrate from preexisting vessels (Hanahan and Folkman, 1996) and (2) circulating EPC from peripheral blood (Shi et al., 1998; Peichev et al., 2000; Rafii, 2000). Cultured EPC offer a robust cell source for tissue engineering and cell delivery applications. EPC can be obtained from the same patient to avoid immune rejection. Although EPC were shown to contribute to tissue revascularization, their function in a clinical setting has not been established. The use of EPC for tissue engineering requires ex vivo expansion that is not optimal for clinical use because of animal products and inadequate tissue culture environment. Tissue Engineering Vascular diseases are the leading causes of morbidity and mortality in the United States each year (Ross, 1993). Over 500,000 coronary bypass grafts and 50,000 peripheral bypass grafts are performed annually in the United States (www.americanheart.org) (Sowton, E., 1991). However, up to 30% of the patients who require arterial bypass surgery lack suitable or sufficient amounts of suitable autologous conduits such as small caliber arteries or saphenous veins (Edwards et al., 1966; Motwani and Topol, 1998; Pomposelli et al., 1998). Synthetic grafts, such as polytetrafluoroethylene or Dacron (polyethylene terephthalate fiber), have been used successfully to bypass large caliber, high-flow blood vessels. However, these grafts invariably fail when used to bypass small-caliber, lowflow blood vessels due to increased thrombogenicity and accelerated intimal thickening leading to early graft stenosis and occlusion (Stephen et al., 1977; O’Donnell et al., 1984; Sayers et al., 1998; Ao et al., 2000). It has been shown that a confluent EC monolayer on small-caliber prosthetic grafts may provide immediate protection from thrombus formation following implantation (Furchgott and Zawadzki, 1980; Cybulsky and Gimbrone, 1991; Seifalian et al., 2002). However, the use of allogeneic EC is limited by rejection, whereas the use of autologous human EC for the construction of vascular grafts has not been widely explored. The idea to use EPC to seed the lumen of engineered blood vessels came from the observations that MSC contributed to the lining of vascular grafts in vivo (Shi et al., 1998; Bhattacharya et al., 2000). We have shown that EPC might be an ideal source of autologous EC for seeding small diameter grafts, eliminating the need to remove native vessel from which to culture EC. By seeding EPC-derived EC onto a scaffold, a non-thrombogenic barrier between blood and vessel wall is created, thereby promoting patency in vivo. EPC-seeded collagen matrices derived from decellularized porcine arteries were used for carotid artery reconstruction in sheep (Kaushal et al., 2001). These bioengineered arteries remained patent for more than 4 months, whereas control grafts without autologous EC occluded within 15 days. Thus, functional vessels can be engineered using decellularized arteries and EPC. Moreover, we have shown that these bioengineered blood vessels, after a brief period of healing in vivo, develop a fully cellularized wall of three distinct layers analogous to normal adventitia, media, and intima. Although these are exciting results, bioengineered grafts will need to be constructed in a mechanically relevant environment. In vitro engineering of blood vessels should mimic the flow conditions that exist in vivo in order to enhance tissue formation. Neram et al. have shown that local blood flow properties induce changes in EC morphology and orientation (Nerem et al., 1981; Nerem, 1984). Further studies showed that the levels of shear stress and the duration of exposure induced changes in EC morphology, proliferation, and differentiation (Sprague et al.,

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H&E SMC

EPC

Figure 24.2 Acellular porcine arterial segment (stained with hematoxylin and eosin, H&E) and seeded with peripheral blood-derived SMC (dyed red with PKH26) and EPC (dyed green with PKH27).

1987; Levesque et al., 1990). EPC cultured under biologic-like shear conditions expressed higher levels of VEcadherin than those cultured under static conditions (Yamamoto et al., 2003). A recent study assessed the use of EPC for bioengineered heart valves. Two EC types, valve-derived mature EC and EPC, from peripheral blood were used (Dvorin et al., 2003). The study showed that both sources of EC, when seeded on PGA/P4HB scaffolds, proliferate in response to VEGF. The EPC could be induced to transdifferentiate to a mesenchymal phenotype on PGA/P4HB in response to transforming growth factor beta-1. These results indicate that EPC can respond to soluble signals that induce events that occur during valvulogenesis (Figure 24.2). One common problem of these studies is that heterogeneous cell populations are being expanded for seeding onto vascular scaffolds. As mentioned previously, one solution is to isolate MSC and to differentiate them to EPC. Another general problem with these bioengineered vascular grafts is immediate availability. For instance, when an emergency bypass needs to be performed, growth of an artificial vessel and preparation for implantation would take too much time if autologous cells are to be implemented. Alternatively, these bioengineered grafts could be seeded with stem cells that were differentiated into EC. Tissue Regeneration Several studies have suggested that EPC participate in the vascular healing process, in part by recruitment of EPC to the regenerated site (Asahara et al., 1997; Takahashi et al., 1999). Genetically labeled EPC were detected in ischemic limbs of mice and were shown to accelerate the revascularization process. Administration of cytokines such as G-CSF and GM-SCF appear to enhance mobilization of EPC and revascularization. In humans, EPC contributed to wound healing of patients implanted with left ventricular assisted device (Rafii et al., 1995a). The EPC adhered to the device and formed a non-thrombogenic surface. These studies suggested that EPC may be recruited to assist endothelialization and served the basis for preclinical and clinical studies as described later. Given the morbidity associated with limb ischemia, EPC may be used for vascular therapy as an alternative to bypass approaches. In preclinical studies, introduction of bone marrow-derived EPC significantly improved collateral vessel formation and minimized limb ischemia (Asahara et al., 1999b; Takahashi et al., 1999; Kalka et al., 2000b). In patients suffering from peripheral arterial disease, injection of autologous whole bone marrow mononuclear cells into ischemic gastrocnemius muscle resulted in restoration of limb function (TateishiYuyama et al., 2002). The improvement in muscle perfusion suggested that it was due to the presence of EPC

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in the cell preparation. However, it remains to be determined whether the improvement was due in part to the introduction of myelomonocytic cells. Bone marrow-derived MSC were recently shown to contribute to myocardial regeneration and revascularization. In nude rats that underwent myocardial infarction, cytokine-mobilized EPC homed to the infarcted tissue and contributed to neoangiogenesis (Orlic et al., 2001a). In similar studies, bone marrowderived MSC were injected into the infarcted border and were shown to differentiate into myocardial cells and EC (Jackson et al., 2001; Kocher et al., 2001; Orlic et al., 2001b). In most studies, direct introduction of these cells into an active angiogenic site, such as infarcted or ischemic myocardium, was essential for successful incorporation of the cells and improvement of cardiac function. Acute myocardial infarction, or chronic ischemic heart disease, results in the loss of cardiomyocytes and vasculature. Several animal studies have shown that introduction of autologous bone marrow MSC contributes to neoangiogenesis in the ischemic myocardium (Rafii and Lyden, 2003). In patients, whole autologous bone marrow mononuclear cells were delivered into the coronary arteries feeding the infarcted and ischemic tissue (Rafii and Lyden, 2003). In all of these studies, there was improved cardiac perfusion and left ventricular function, suggesting that delivery of autologous progenitor cells is feasible, safe, and may have a short-term therapeutic benefit. However, follow-up studies in animals and humans detected only a few bone marrow-derived cells in the regenerated vascular network, suggesting that only a small portion of the cells may contribute to revascularization. Despite the excitement for these initial observational clinical trials, it remains to be determined in double-blind placebo-controlled randomized clinical trial whether this cellular therapy approach will result in any long-standing cardiac benefits. Importantly, it remains unclear if any long-term toxicity exists with this therapy. Such toxicity may result if myeloid cells are incorporated into regenerating myocardium and generate noncardiac or fibrotic tissues. Therefore, progenitor cells that have been pre-differentiated into EPC should be used with caution and long-term monitoring. MSC In the case of MSC, the lineage-committed cells can generate a variety of specialized mesenchymal tissues including bone, cartilage, muscle, marrow stroma, tendon, ligament, fat, and a variety of other connective tissues (Caplan, 1994). As such, MSC may have a dramatic impact on the overall health status of individuals by controlling the body’s capacity to naturally remodel, repair, and upon demand, rejuvenate various tissues. In human clinical research, initial efforts are focused on applications of MSC-based tissue repair using cell delivery approaches. An example of such application is the use of MSC is to regenerate nonunion bone defects. A number of studies showed that MSC from animals and humans, delivered in a porous, calcium phosphate vehicle, were able to regenerate bone tissue (Bruder et al., 1994, 1998; Jaiswal et al., 2000). Additionally, these cells may be beneficial for cartilage repair. The cartilage is a tissue that cannot repair itself in adults. MSC have been applied in hyaluronan scaffolds for cartilage tissue repair with good results and are now in clinical trials (Solchaga et al., 1999, 2000). Bone marrow-derived MSC have also been used for muscle repair and fuse with the host myotubes and formed functional muscle fibers (Shake et al., 2002; Toma et al., 2002). Systemic delivery of bone marrow-derived MSC showed that they can home back to the bone marrow. This observation prompted clinical studies to use MSC to restore the bone marrow in patients undergoing radiation and chemotherapy-mediated myeloablation (Lazarus et al., 1995; Koc et al., 2000). The Use of Peripheral Blood Stem Cells for Gene Therapy Gene and cell therapies have been proposed for regenerative medicine and tested in a number of clinical trials. Genetically modified MSC offer a unique approach as cells with growth potential may represent a useful tool

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for tissue engineering and cell therapy. A detailed knowledge of vector delivery systems is critical for practical applications. One of the most popular vectors used for gene delivery to progenitor cells is replication-deficient Ad. Ad vectors offer two important advantages that make them ideal for gene therapy. First, they can efficiently infect non-dividing cells, which is important for MSC that live primarily in the G0/G1 phase of the cell cycle (Hawley, 2001; Alessandri et al., 2004). Second, the Ad vector can offer transient expression of the recombinant gene for a time period of approximately 3 weeks (Iwaguro et al., 2002). However, Ad vectors have shown to elicit an unwanted inflammatory response. Genetically modified stem cells have been explored in a number of studies to regenerate bone and cartilage or for neovascularization (Grande et al., 2003; Kondoh et al., 2004; Shen et al., 2004). The most common genes used in these studies are growth factors such as VEGF. VEGF, as mentioned earlier, is a potent angiogenic factor that supports the differentiation of MSC along endothelial lineages. In order to enhance vascularization of engineered muscle tissue, we have transfected primary cultures of rat myoblasts with a plasmid encoding VEGF and green fluorescence protein (GFP). Cells expressing GFP were selected by fluorescent activated cells sorter and injected mixed with gelatin, into the subcutaneous space of immune-deficient mice (De Coppi et al., 2005). Tissue volumes of VEGF-transfected cells increased during 21 days and tripled their size. In contrast, the volume of tissues containing cells, which were transfected with control plasmid, gradually decreased and the tissues were minimally visible after 21 days. Immunohistochemical analysis of VEGF-expressing tissue with anti-von Willebrand factor revealed typical muscle formation and a developed vascular network. VEGF gene transfer to stem cells has been used by in situ neovascularization and angiogenesis in order to salvage ischemic limbs (Kalka et al., 2000a, c; Iwaguro et al., 2002). Other studies looked at the combinations of growth factors to mimic the environment of vascular development. Both bFGF and angiopoietin-1 have been transfected with VEGF into progenitor cells to induce the development of mature blood vessels including the medial and outer adventitial layers (Kondoh et al., 2004). This approach also succeeded in reducing the VEGF-mediated permeability and fluid leakage of the new vessels. The future of stem cell-mediated gene therapy is dependent on the resolution of some key questions. The efficiency of gene transfer need to be close to 100% to ensure that unmodified cells do not interfere with the regenerative process. The most feasible stem cell source needs to be used for successful clinical applications. Finally, the mode of cell delivery, systemic or local injection, needs to be adjusted for each application. Regardless of the solution to each of these questions, stem cells-based therapies will benefit enormously from gene modification.

CONCLUSIONS AND FUTURE DIRECTIONS The bone marrow is probably the source of peripheral blood stem and progenitor cells. Hemangioblasts are the embryonic precursors of HSC, giving rise to committed hematopoietic progenitors. The bone marrow is also a source for other progenitor and stem cells, the MSC, which can be expanded in vitro, and have multilineage differentiation potentials. Numerous studies, described here, have shown that there is a constant exchange of cells from the bone marrow to peripheral blood. On the other hand, bone marrow transplantation studies have indicated that this process may be reversed and cells from peripheral blood may repopulate the bone marrow. Future success in applying adult peripheral blood-derived stem cells for clinical applications will depend on the development of strategies to mobilize, isolate, expand, differentiate, and to deliver these cells. For example, EPC may be isolated from peripheral blood and used for therapeutic angiogenesis directly or after a period of ex vivo expansion. Understanding the signals involved in the recruitment of these cells to the regenerating tissues will play a crucial role in optimizing this technology for clinical use. The studies summarized here provide evidence to the presence of stem cells in peripheral blood and mechanisms by which they can be mobilized from bone marrow in order to increase their numbers in blood. Although various attempts have been made to use peripheral blood-derived stem cells in humans, and some encouraging

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results were obtained, standard clinical use of these techniques must await further validation and long-term toxicity evaluations.

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25 Prospects of Somatic Cell Nuclear Transfer-derived Embryonic Stem Cells in Regenerative Medicine Z. Beyhan and J.B. Cibelli

INTRODUCTION “Since their first description 25 years ago (Evans and Kaufman, 1981) in mice, embryonic stem cells (ESCs) have provided invaluable tools for addressing many biological questions related to cell differentiation, gene function, transgenesis, genetic, and degenerative diseases.” Various methods have been established to induce differentiation of these cells into somatic cell types, including, but not limited to, oligodendrocytes, glial cells, neurons, cardiomyocytes, insulin-producing B-like cells, and hematopoietic cells. The derivation of human ESCs (hESCs) (Thomson et al., 1998) has set the stage for realizing the long sought tools to design cell-based therapies for many human genetic and degenerative diseases, such as Parkinson’s disease, Duchene’s disease, diabetes, spinal cord injuries, and cardiomyopathies, among others. Even though no clinical treatment schedule based on hESCs is available at present, considerable progress in this area has been made, and one private company has announced plans to file for an Investigational New Drug application with the US Food and Drug Administration and to start clinical trials in 2007 using hESC-derived oligodendrocytes for the treatment of acute spinal cord injuries. These developments are encouraging, considering the limitations imposed by a variety of factors, such as the availability of human oocytes and embryos and ethical and legal impediments. One of the major physiological (technical) concerns regarding the use of ESC-derived somatic cells and tissues in transplantation treatments is the immune rejection of the grafted tissue, which is a common complication of allogeneic transplantations (Prentice, 2006). Another scientific breakthrough in 1997 has brought the possibility of addressing this problem by evading the surveillance of recipient’s immune system. This breakthrough was the birth of the first mammal, a lamb, produced by using somatic cell nuclear transfer (SCNT) (Wilmut et al., 1997). The efficiency of SCNT is low; however, this drawback did not deter investigators from cloning a large number of other species. A combination of SCNT and ESC technologies would provide tools to produce isogeneic (patient-specific) ESC cell lines to treat a number of human conditions that originate due to aging, trauma, or degenerative diseases. In this chapter, we will summarize recent progress in this area of research and the state of the art; then we will discuss the prospects and limitations of SCNT in obtaining ESCs for the purpose of cell therapies.

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BRIEF HISTORY The idea of nuclear transfer (NT) was first proposed by Hans Spemann during the late 1930s – he called it “the fantastical experiment” – to address the question of cell totipotency and differentiation of cell lineages during embryo development (Foote, 2002). However, testing Spemann’s proposed model had to wait until 1952, when the first successful use of a nuclear transplantation method resulted in development of feeding tadpole stage Xenopus embryos when blastula stage nuclei were used as nuclear donors (Briggs and King, 1952; Gurdon, 1962). Earlier NT studies in amphibia led to the concept of loss of totipotency during the cell differentiation process, since development of embryos failed as more differentiated nuclei were used as donors (Gurdon, 1986). It is fair to say that NT studies in amphibians have paved the way to development of successful techniques for mammalian cloning. Technically, the first cloned mammals were obtained by splitting early preimplantation embryos of sheep (Willadsen, 1979, 1981), cattle (Ozil et al., 1983; Williams et al., 1984), and rats (Foote, 2002). However, the first successful cloning of a mammal by transferring nuclei of embryonic blastomeres was achieved in sheep (Willadsen, 1986) and cattle (Prather et al., 1987) almost simultaneously. The fame of Dolly, the sheep, was due to the fact that she was the first mammalian clone originated from a differentiated adult somatic cell. This achievement shattered the well-established dogma that a differentiated somatic cell nucleus could not be reprogrammed to an embryonic state (Wilmut et al., 1997). Soon thereafter, several laboratories independently confirmed this study by producing live offspring in a number of species using fetal and adult somatic cells in mice (Wakayama et al., 1998), cattle (Cibelli et al., 1998a; Kato et al., 1998), pigs (Betthauser et al., 2000; Polejaeva et al., 2000), goats (Keefer et al., 2001, 2002; Reggio et al., 2001), rabbits (Chesne et al., 2002), zebrafish (Lee et al., 2002), cats (Shin et al., 2002), mules (Holden, 2003), horses (Galli et al., 2003), dogs (Lee et al., 2005), and ferrets (Li et al., 2006). The ability to produce offspring using cultured somatic cells opened up a number of interesting possibilities in science and technology, such as cloned animals producing pharmaceuticals and nutraceutical proteins or organs for xenotransplantation, engineered animal models for human disease research, derivation of patient-specific and genetically modified ESCs for cell therapies, preservation of endangered species, and genetic improvement of domestic species for important production traits (Foote, 2002). The most important limitation of this technology, in its current state, is the inefficiency of methodology in producing healthy live offspring in all species studied so far. The frequency of development to term is well below the rate that is observed during in vivo and in vitro development. Overall efficiencies (number of live births/number of reconstructed embryos) of NT experiments in published studies have ranged between 0% and 10% (Wilmut et al., 1997; Kato et al., 1998; Wakayama et al., 1999; Wells et al., 1999; Kubota et al., 2000; Polejaeva et al., 2000; Reggio et al., 2001; Forsberg et al., 2002; Keefer et al., 2002; Campbell et al., 2005). Although the majority of cloned embryos complete preimplantation development and reach blastocyst stage, more than half of them are lost during the first trimester, while the rest proceed through pregnancy, gradually failing at different stages and reaching term in substantially reduced numbers (Pace et al., 2002). The high level of prenatal mortality in cloned fetuses is related to a number of developmental abnormalities, including retarded development, placental abnormalities (less numerous and enlarged placentomes, less vascularization, epithelial abnormalities, and hydroallantois), cardiovascular abnormalities, endocrine deficiencies, increased fetal weight, and failure in parturition (Thibault, 2003). Apparently, a great proportion of cloned embryos are not able to reprogram donor nuclei into an embryonic state where a developmental program is initiated and executed appropriately to produce live offspring. However, the fact that a substantial number of cloned embryos reach the blastocyst stage at reasonable rates suggests the possibility of using these embryos for ESC isolation. The first demonstration that embryonic cells can be produced using somatic cells was performed in the bovine model, albeit these cells had limited differentiation 457

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Therapeutic cloning

Reproductive cloning

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SCNT embryo Embryonic stem cells In vitro culture

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Figure 25.1 Schematic representation of therapeutic and reproductive cloning.

capability (Cibelli et al., 1998b). Munsie et al. have reported establishment of the first bona fide ESC line from an SCNT blastocyst in the mouse model (Munsie et al., 2000), and the first differentiation of mouse NT ESCs (mnt-ESCs) into several tissues, including neurons and gametes, was reported thereafter (Wakayama et al., 2001). These and several other achievements (Kawase et al., 2000; Rideout et al., 2002; Barberi et al., 2003) in the field have given rise to a distinction between two types of applications for the cloning technology, bestowing us with two commonly used terminologies, “reproductive cloning” and “therapeutic cloning.” (Figure 25.1) The major distinction between these two terminologies relies on the end points they refer to, even though the procedures that created the preimplantation embryos are the same. Reproductive cloning refers to “producing a cloned embryo and transferring it to a surrogate mother with the aim of obtaining live offspring,” while the aim of therapeutic cloning is to create a preimplantation stage embryo/blastocyst and to use this embryo to isolate isogeneic ESCs (Rideout et al., 2000). Since these ESCs and the donor cell have identical genomic content, any cell or tissue type engineered from these cells will be immunocompatible with the original donor organism, leading to a way to tackle one of the major problems of transplantation technology.

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STATE OF THE ART For therapeutic cloning to be a viable approach to treating human degenerative and genetic diseases, three major milestones need to be reached: 1. Development of efficient human SCNT methods to readily produce patient-derived cloned blastocysts. 2. Establishment of protocols for efficient derivation of ESCs from these blastocysts. 3. Creation of tools for robust differentiation of these ESCs into certain cell types in reasonable quantities and genetic manipulations needed for therapies.

Isolation and culture of mouse ESCs (mESCs) are well established and have been a driving force in stem cell biology by allowing genetic manipulations and providing a model system for stem cell differentiation and transplantation studies (Hall et al., 2006). Since the establishment of the first hESCs by Thomson et al. (1998), more than 200 hESC lines have been isolated and are available for use by the scientific community (Loring and Rao, 2006). Several groups have reported the differentiation potential of hESCs into various cell types, such as skin cells (Schuldiner et al., 2000), neurons (Reubinoff et al., 2001; Schuldiner et al. 2001; Zhang et al., 2001), blood (Kaufman et al., 2001), endothelial cells (Levenberg et al., 2002) cardiac muscle (Kehat et al., 2001), cartilage (Tanaka et al., 2004; Olivier et al., 2006), and pancreatic cells (Assady et al., 2001; Segev et al., 2004). Transplantation experiments have already begun. Mouse, monkey, and hESCs were partially differentiated and transferred into mouse, rat, or other animal models to treat Parkinson’s disease, spinal cord injury, and cardiac muscle degeneration (Street et al., 2003; Faulkner and Keirstead, 2005; Keirstead et al., 2005; Kimura et al., 2005; Liew et al., 2005; Takagi et al., 2005; Lensch and Daley, 2006). In most cases, transplanted cells survived for certain periods of time and partially restored the impaired functions of the model animals despite the considerable amount of cell death after transplantation. These results are promising despite the fact that several questions and complications regarding the treatment schemes need to be addressed to optimize the methodologies. At present, the only model for therapeutic cloning that has worked is in mice (Kawase et al., 2000; Munsie et al., 2000; Wakayama et al., 2001). While their ability to differentiate into a wide array of cell lineages and to serve as plausible sources of donor cells for regenerative treatments were not investigated extensively, earlier reports indicate that mnt-ESCs potential may not be compromised by such factors observed in cloned fetuses and embryos as chromosomal, genetic, and epigenetic abnormalities. Live, healthy offspring were obtained by using ntESCs as nuclear donors in a second round of NT, indicating at least some of the ntESCs are competent enough to support full-term development (Wakayama et al., 2005a). In addition, mouse chimeras generated with ntESCs resulted in germ-line transmission of the injected cells, strongly supporting their functional similarity to conventional mESCs (Wakayama et al., 2005a, b, c). A recent study by Wakayama et al., employing 150 mnt-ESC lines, has reported that these ntESCs are comparable to their in vivo-derived counterparts in terms of their differentiation capacity, pluripotency marker expression profile, global gene expression profile, and methylation characteristics on certain selected regions (Wakayama et al., 2006). Considering all the data available, it is reasonable to assume that differentiation protocols developed for hESCs could be employed for the yet-to-be-described human ntESCs as well. The elegant experiments performed by Rideout et al. have proven that the concept of therapeutic cloning can be coupled with ex vivo gene therapy (Rideout et al., 2002). In this study, a recombination-activating gene 2 (Rag-2) mutant mouse which is characterized by severe combined immunodeficiency syndrome with the lack of mature T and B cells was used to produce SCNT blastocysyts from tail tip fibroblasts. The resulting blastocysts were used to isolate ntESCs and to restore one functional Rag-2 allele by homologous recombination. Genetically modified ntESCs were induced to differentiate into hematopoietic precursor cells and transplanted into irradiated Rag-2 mutants to treat their immunodeficiency. Treated mice had their myeloid and lymphoid cells repopulated with functional B and T cells, clearly showing that gene and isogeneic cell

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Expand HSC and transplant

Rag2-/-

Infect with HoxB4/GFP

Culture tail tip cells

Nuclear transfer Differentiate into EBs

NT blastocyst Repair Rag2 gene in ES cells Isogenic Rag2-/- ES cells

Figure 25.2 Scheme for therapeutic cloning combined with gene and cell therapy. A piece of tail from a mouse homozygous for the Rag-2 mutation was removed and cultured. After fibroblast-like cells grew out, they were used as donors for nuclear transfer by direct injection into enucleated MII oocytes using a piezoelectric-driven micromanipulator. ESCs isolated from the NT-derived blastocysts were genetically repaired by homologous recombination. After repair, the ntESCs were differentiated in vitro into EBs, infected with the HoxB4iGFP retrovirus, expanded, and injected into the tail vein of irradiated, Rag-2-deficient mice. Adapted from Rideout et al. (2002). therapy could be facilitated using ntESCs (Figure 25.2). A recent study, even though it did not employ ntESCs, has shown that concomitantly knocking down a mutant gene and introducing a wild-type allele were possible and could correct the sickle cell anemia phenotype (Samakoglu et al., 2006), underscoring the enormous possibilities that could be explored by using ntESCs (Figure 25.2). In diseases like Parkinson’s, Duchene’s, spinal cord injury, and diabetes, where administering ample amounts of non-modified cells is needed for treatment, differentiating ntESCs into the desired cell type in reasonable quantities would be enough to alleviate the associated phenotypes without risking the immune rejection of the transplanted cells. Furthermore, the unlimited proliferation capacity of ntESCs could make possible the administration of multiple doses of isogeneic cells at different intervals as needed. Although the last two milestones – i.e. efficient derivation of ESCs and establishing robust differentiation protocols – have been at least partially achieved in humans, the first milestone, generating cloned human blastocysts efficiently, has yet to be met. Two papers by Hwang et al. reported the development of methods to clone human blastocysts and, subsequently, to establish patient-specific ESCs with reasonable efficiency (Hwang et al., 2004, 2005). These studies have, befittingly, created great excitement in the field of stem cell research by providing the very possibility of turning the therapeutic cloning concept into a practical reality of human medicine. However, an investigation by the South Korean National Bioethics Committee on the ethical concerns related to this work concluded that there have not been any patient-specific ntESCs created by the Hwang team

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(Chong, 2006). These revelations have been recorded as one of the most damaging frauds the scientific community has experienced. At present, only two studies are available where human cells were fused with enucleated human oocytes. In the first one, only a six-cell cloned embryo was obtained (Cibelli et al., 2001); in the most recent one, one cloned blastocyst was obtained – albeit using an hESC as a nuclear donor (Stojkovic et al., 2005). Non-human primate models were also unable to create live offspring or ntESCs using somatic cells as donors, even though embryonic blastomeres supported development to term in an earlier study (Meng et al., 1997). It seems that the major limitation on producing SCNT embryos and offspring in human and non-human primates is the scarcity of the oocytes and the heterogeneity of the species compared to other animal models used for cloning studies. With the limited information we currently have, it is impossible to predict the number of human oocytes that will be needed to make one human ntESC line. The investigative committee from Seoul National University concluded that 2,236 oocytes were used by Hwang’s team and that no ntESC line was ever produced. In addition, 122 women underwent superovulation solely for the purpose of generating ntESCs. Fifteen of them developed different degrees of ovarian hyperstimulation syndrome. Knowing that the Korean team was trained for SCNT and that they had the expertise to derive fertilized hESCs and yet were unsuccessful, the authors of this manuscript would like to appeal to the scientific community to reconsider the recruitment of women solely for the isolation of oocytes in an attempt to generate ntESCs. At the expense of slowing scientific progress, we humbly suggest that for the foreseeable future, and until the efficiency of SCNT in animal models dramatically improves, we should use spare oocytes from in vitro fertilization (IVF) clinics. In the murky atmosphere created by the false allegations by the Hwang team, it will take more than creating a few blastocysts to recapture the public’s interest and to restore the credibility of the solid science behind the concept of therapeutic cloning.

PROSPECTS AND CHALLENGES As an alternative source of differentiated cells for regenerative therapies, ntESCs have the potential of overcoming some of the limitations posed by allogeneic ESCs. As opposed to earlier presumptions about their reduced immunoreactivity, allogeneic ESCs have been shown to express major histocompatibility complex class I (MHC-1) molecules when induced to differentiate, thus complicating their use in clinical applications (Drukker et al., 2002; Drukker and Benvenisty, 2004). In such a situation, all ramifications of conventional allogeneic tissue transplantation need to be dealt with, namely, various methods of immunosuppression have to be administered to the recipient for prolonged periods of time. Therefore, the major limiting factor for current tissue transplantation therapies would be handed down to any allogeneic ESC-based treatment strategy. Strategies to consider are (1) building a generic ESC collection that could cover the majority of human leukocyte antigen (HLA) types, (2) implementing bone marrow chimerism before transplanting the cells, and (3) genetically engineering the MHC molecules to match those of the patient. All three strategies have their shortcomings. An hESC bank, no matter how large, will only cover approximately 30% of the population (Taylor et al., 2005). Bone marrow chimerism has shown promise not only in animals but also in humans; however, efficient protocols to generate long-term grafting ESC-derived hematopoietic cells are not yet available. And, finally, genetic modifications of the MHC molecule, while a scientifically exciting proposition, are likely to be impossible to implement on a large scale. Undoubtedly, the best way to obtain matched HLA-type ESCs from a patient is to dedifferentiate a cell from that patient, and the only method known to work is SCNT. Although SCNT procedures are not very efficient and development to term of cloned embryos is compromised, they could readily develop to the blastocyst stage at which ESCs are traditionally established. These blastocysts could provide the source for patient-specific

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ESCs, eliminating the tissue incompatibility problems and the deleterious effects of immunosuppressive treatments. However, widely reported developmental abnormalities, epigenetic and genetic aberrations, and ploidy problems in cloned embryos have created genuine concern about the use of ntESCs for regenerative treatments. Recent studies in mice have brought about promising results by demonstrating that mouse ESCs could be readily established from NT blastocysts, and these cells carry all characteristics of conventional ESCs, such as unlimited ability to replicate, forming teratocarcinomas, and contributing to chimeras when injected into immunodeficient mice and blastocysts, respectively (Hochedlinger and Jaenisch, 2003). They are also able to differentiate into various tissues in vitro and form embryoid bodies (EBs), indicating that the abnormalities observed in cloned animals do not extend to ntESCs or, at least, do not interfere in the function of these cells (Wakayama et al., 2001, 2003, 2006). Both in the context of SCNT animals and ntESCs, heteroplasmy has created legitimate concern, since the compatibility between mitochondrial and genomic DNA is critical for the function of the organelle (Barrientos et al., 1998; Dawson and Dawson, 2004), and the mitochondrion itself has some antigens that could contribute minor histocompatibility antigen (miHA) complex (Simpson, 1998). Furthermore, involvement of mitochondria and mitochondrial DNA mutations in several degenerative diseases is well established (Kang and Hamasaki, 2005; Simmons, 2006), and numerous mutations in mitochondrial DNA (mtDNA) (DiMauro and Davidzon, 2005) have been detected, underlying the reasonable concerns in regard to ESC and other reproductive technologies (Hawes et al., 2002). Several lines of evidence, however, argue against this notion, The first line of evidence comes from cloned animals, where mitochondrial heteroplasmy is the likely outcome of SCNT procedures (Hiendleder et al., 1999; Steinborn et al., 2000; Do et al., 2002) despite the presence of a few homoplasmic cloned animals (Evans et al., 1999). In either outcome, the cloned animals were healthy, and no adverse effects of the heteroplasmy were reported. It is worthwhile to mention the application of cytoplasmic transfer to treat certain types of infertility in humans. When oocyte quality in some women was compromised, possibly due to aging, transfer of ooplasm from younger women into compromised oocytes improved the outcome of IVF treatments, and several babies were born after such assistance (Brenner et al., 2000; Barritt et al., 2001; Dale et al., 2001; Fulka et al., 2005). Although many of these babies were heteroplasmic, no significant adverse effect of this situation has been reported so far. These data, however, do not exclude the possibility of long-term effects of such a condition, and suggest the need to address the concerns carefully. The second line of evidence, regarding the concerns about the role of mitochondria in immunorejection of heteroplasmic tissue and cells, comes from tissue graft experiments conducted among cloned animals. A recent study in cloned pigs has shown that skin grafts between cloned animals were not rejected, as opposed to those between unrelated animals, indicating that the ooplasmic origin of mitochondria in SCNT-derived animals does not constitute a problem in the context of immunocompatibility (Martin et al., 2003; Shimada et al., 2006). This notion was further supported by observations in bovine and mouse SCNT models. Skin grafts between cloned cows that were genetically identical but carrying different mtDNA haplotypes were not rejected, while genetically different skin grafts induced a strong immune reaction and rejection (Theoret et al., 2006). Injection of mouse SCNT-derived fetal liver stem cells to induce myocardial regeneration did not result in the rejection of the heteroplasmic cells, yet resulted in significant improvement in the regeneration process (Lanza et al., 2004). Like all other human-embryo-based technologies, ESCs, SCNT, transgenesis, and any combination thereof understandably create serious legal and ethical concerns. At the basis of the issue lay the questions of whether a human embryo is the equivalent of an individual that should therefore be protected by individual rights (de Wert and Mummery, 2003). The answers to these questions naturally vary, based on one’s social, religious, and political background, and manifest themselves in the context of the legal framework to regulate research and research funding for embryo-based technologies. To that end, reaching a consensus among different countries, states, and/or regions (like the United States or Spain) is of utmost importance. Unfortunately, the current legal

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landscape acts to deter many talented scientists. They would rather avoid working in this area, when indeed their engagement in this field is imperative. Ethical and legal issues are discussed in detail elsewhere in this book. Taken together, these data suggest that human ntESCs can certainly offer a solution to tissue compatibility problems; however, legal and feasibility issues still remain. These problems are solvable. However, due to their complex and multifaceted nature, a strong will and a multidisciplinary approach will be required to tackle them.

CONCLUSIONS AND FUTURE DIRECTIONS At present, no hESC line exists that has been created from NT embryos. The lack of these cells is mainly due to logistical, technical, ethical, and legal restrictions imposed on experimentation with human embryos. The authors believe that there is no sound scientific evidence to suggest that there are physiological limitations to attaining SCNT embryos/blastocysts in humans compared to other species. The practical use of therapeutic cloning as a source of patient-specific ESCs, however, is complicated by a number of other factors, such as availability of recipient oocytes, efficiency of NT and ESC establishment methods, as well as the time frame needed to develop usable differentiated cells for treatment of diseases. The use of patient-specific ESCs may not be a routine medical treatment option for most of the degenerative and genetic diseases until significant improvements are made in the processes mentioned above. Therefore, this area of research in stem cell biology needs to be conceived as a complementary approach to understanding the hallmarks of stem cell physiology, along with such other means as engineering generic stem cell lines, use of adult cells, and direct dedifferentiation of somatic cells into pluripotent cell types. Despite the nocent effect of ill-fated patient-specific ESCs reported by Hwang’s team, this area of research must be pursued, given the promise of scientific and clinical rewards associated with therapeutic cloning. Despite the astounding progress accomplished in developmental biology decades after Spemann’s suggestion of “the fantastical experiment,” the question he was trying to address – the question of totipotency and cell lineage commitment and its regulation – is still a point of convergence for developmental biology, cell biology, and the medical sciences. As stem cell biology pursues these intricate challenges and pieces together the puzzle of ontogeny, by using any tool available, mankind will undoubtedly benefit from the knowledge accumulated.

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26 Somatic Cells: Growth and Expansion Potential of T Lymphocytes Rita B. Effros INTRODUCTION One of the fundamental characteristics of all human somatic cells that are mitotically competent is an innately programmed barrier to unlimited proliferation (Hayflick and Moorhead, 1961). This property, known as replicative (or cellular) senescence, may serve as one of many safeguards to maintain cellular integrity necessitated by the extended longevity of humans. It is thus possible that a restriction in the number of cell divisions serves as a protection against the potential for multiple mutations that are required for the development of a cancer cell from a cell that is normal (Effros et.al., 2005). Nevertheless, for some cell types, the replicative senescence cellular program can lead to deleterious consequences, particularly by old age. Immune responses are characterized by an extraordinary expansion of lymphocytes due to the low frequency of cells that can respond to each single foreign pathogen. Under most circumstances, the limited proliferative potential of T lymphocytes, the cells that are key to controlling infections and cancer, does not hamper primary or even secondary immune responses (Effros and Pawelec, 1997). However, by old age and/or during certain chronic viral infections, there is an accumulation of clones of T cells that show signs of having reached their maximum replicative limit. This chapter will discuss the nature and underlying mechanism of replicative senescence in human T cells, a specific facet of immune system activity that seems particularly well suited to regenerative medicine approaches. One of the signature changes associated with aging is the significant decline in immune function. Immune system failure is believed to underlie the increased risk of morbidity and mortality from influenza and other infections. Even the response to vaccines intended to prevent infection is reduced in the elderly, providing further evidence of the diminished immune function. There is also increasing evidence that many of the pathologies and diseases associated with aging have an immune component, or, in some cases, even an immune-based etiology (Effros et al., 2003). In fact, a cluster of immune parameters (including high proportions of T cells with characteristics of replicative senescence) that correlates with early all-cause mortality has been identified in longitudinal studies on persons aged 80 years and older. These and other studies suggest that improved health and quality of life may be possible if the immune system of the elderly is either prevented from “aging” or can be rejuvenated in some way. The immune exhaustion associated with chronic HIV disease is another situation that might also be amenable to similar therapeutic strategies. In the sections below, we will provide an overview of the immune system, a summary of the features of T cell replicative senescence, evidence demonstrating the presence and consequences of high proportions of senescent T cells in vivo, and finally, our ongoing approaches to reverse or retard this process.

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T CELLS ARE KEY TO IMMUNITY TO INFECTIONS AND CANCER The immune system is a complex and highly integrated network of cells and lymphoid organs that functions to protect the body from foreign pathogens. Immunity is generated by two interacting components, namely, the innate and the adaptive immune systems. The innate immune response is capable of dealing with certain pathogens in a rapid, albeit, somewhat non-specific manner. By contrast, the adaptive immune response takes longer to develop, but has the advantage of exquisite specificity and memory. Indeed, this anamnestic response is the basis for the efficacy of vaccines. All the cellular components of both the innate and adaptive immune systems, including B lymphocytes, T lymphocytes, monocytes, and dendritic cells, are derived from primitive stem cells in the bone marrow. One of the noteworthy aspects of T and B lymphocytes, the main players in adaptive immunity, is the presence of antigen receptors on the surface of each cell that confer the ability to recognize a specific region of a particular pathogen. These antigen receptors are generated during the complex transition from hematopoietic stem cells to mature lymphocytes by an intricate process of cutting and splicing that leads to random joining of DNA segments from several different gene families (Janeway et al., 2001). The end result of this process is that each lymphocyte expresses a unique antigen receptor, and if that lymphocyte becomes activated as a result of encounter with the appropriate antigen, the identical receptor is expressed on all the resulting daughter cells. The generation of antigen receptors by this stochastic process leads to an extremely large repertoire of antigen specificities, thereby enabling the immune system to have broad coverage over multiple and varied types of pathogens. However, precisely because of the huge spectrum of antigen specificities within each individual, the number of lymphocytes that can respond to any single pathogen is extremely small, leading to the requirement for massive cell division and clonal expansion of the few cells whose receptors recognize the invading pathogen. Whereas both B and T cells generate their antigen receptors by similar processes, they function in quite distinct ways when that antigen is encountered (Janeway et al., 2001). B cells produce soluble proteins called antibodies, which can neutralize or otherwise inactivate pathogens that are present within the blood. T cells, on the other hand, only recognize pathogens that have already infected other cells. In the case of a viral infection, for example, the infected cells become decorated with a portion of the virus, indicating to the immune system that the cell is no longer normal and must be eliminated. Those cytotoxic T cells whose receptors recognize the specific viral antigens on the surface of the infected cell become activated and then undergo massive cell division, migrate into the tissues, where they actually kill infected or otherwise abnormal cells, thereby controlling the infection. Once the antigen-specific T cells complete their function, most of the expanded cell population dies by apoptosis, leaving only a few memory cells to handle possible future encounters with the same antigen. Thus, proliferation and the ability to undergo repeated rounds of clonal expansion is a critical feature of effective T cell function.

GROWTH AND EXPANSION POTENTIAL OF T CELLS The protagonists in this chapter are cytotoxic T cells (also known as CD8 T cells), which are the immune cells responsible for control of infections and cancer. Consistent with the increased severity of infections and steep rise in cancer incidence in the elderly, the major defects in immune function associated with aging are, in fact, within the T cell compartment (Miller, 1996; Effros et al., 2003; Swain et al., 2005). Multiple approaches, in both human and animal models, have been utilized in an effort to analyze the underlying mechanisms for the age-related decline in T cell function. Our own research strategy has been to address the immune decline of aging from the perspective of replicative senescence, a process first identified by Hayflick and Moorhead (1961). As it happens, this basic property of somatic cells is highly relevant to the field

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of regenerative medicine, since the strict barrier to uncontrolled cell division has a significant impact on the potential utility of normal cells to generate large numbers of cells for replacement therapy. Indeed, the limited proliferative potential of somatic cells is the reason that most current approaches to regenerative medicine rely on stem cells, which have unlimited expansion potential. However, if it becomes possible to manipulate the process of replicative senescence in normal somatic cells to allow increased proliferation, this might greatly expand the field of regenerative medicine to include specific types of differentiated cells with known functions. In the case of T cells, the added advantage would be the ability to focus on cells directed at a specific viral or tumor antigen. As noted above, due to the random nature of DNA regions utilized in creating the T cell antigen receptor, there is a huge repertoire of different T cell specificities. This broad spectrum of specificities is necessary to counter the huge universe of potential pathogens. The corollary to this is that T cells of any given specificity are low in frequency, requiring extensive and rapid clonal expansion in order to reach the numbers needed for an effective response to pathogens. Although T cells can divide faster than any other vertebrate cell type, the extensive cell division is not without consequences. Indeed, some T cells can actually reach the end stage of replicative senescence, particularly by old age, but also in younger people during certain chronic infections. Thus, our laboratory has focused on analyzing the process of replicative senescence in cytotoxic T cells, and on developing methods to retard or prevent this process.

CELL CULTURE MODELING OF T CELL REPLICATIVE SENESCENCE Extensive research, beginning in the 1960s, has characterized the process of replicative senescence in a variety of human somatic cell types (Hayflick and Moorhead, 1961). The major focus of these in vitro studies was the fibroblast, a cell involved in maintaining intracellular matrix integrity, and in enhancing wound healing (Harley et al., 1990; Smith and Pereira-Smith, 1996; Campisi, 1997). Other cell types, such as keratinocytes, epithelial cells, hepatocytes, and endothelial cells, have also been characterized, albeit less extensively (Le Guilly et al., 1973; Johnson and Longenecker, 1982; Saunders et al., 1993). Ironically, T cells, whose function is critically dependent on extensive proliferative activity, were late-comers to the field of replicative senescence research. Nonetheless, it is now clear that T cells do, in fact, have a limited proliferative potential in culture and probably in vivo (Effros, 1998, 2001, 2004; Pawelec et al., 2000), an observation which has broad implications for immune function during aging and chronic HIV disease, as well as cancer immunotherapy and regenerative medicine. The in vitro model system used for our studies was developed in an effort to mimic, albeit imperfectly, the in vivo immune response of human cytotoxic (CD8) T cells. Although the system is isolated from the normal in vivo environment, it has the unique advantage of allowing longitudinal analysis of the same population of T cells over time, which would be impossible to do in vivo. The basic protocol of our in vitro model is to isolate peripheral blood mononuclear cells from venous blood samples, and stimulate the cells with irradiated foreign (allogeneic) tumor cells in the presence of IL-2. This procedure leads to the expansion of those cells which have receptors that recognize the tumor cells. After a period of 2–3 weeks, the vigorous cell proliferation induced by antigenic stimulation subsides, and the cells became quiescent. The cycle of antigenic stimulation–proliferation– quiescence is repeated multiple times until the culture reaches an irreversible final stage of quiescence that cannot be overcome by antigen stimulation or growth factors (Perillo et al., 1989, 1993). This terminal state is known as replicative senescence. A similar scenario occurs in cultures of virus-specific CD8 T cells established from blood samples of HIV-infected individuals (Dagarag et al., 2004). By using donors with a particular human leukocyte antigen (HLA) type (A2), antigen stimulation of the HIV-specific CD8 T cells can be accomplished with autologous irradiated cells pulsed with appropriate HIV peptides. Such cultures also follow the pattern of eventual cessation of proliferation. The overall finding from numerous different laboratories using a variety of

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antigens for stimulation is that human T cells are able to undergo a limited number of replications, after which they cease dividing. It is important to note that this end stage of replicative senescence does not imply loss of viability. Indeed, with appropriate feeding, senescent cells remain viable and metabolically active for several months (Wang et al., 1994; Spaulding et al., 1999).

REPLICATIVE SENESCENCE ALSO AFFECTS GENE EXPRESSION AND FUNCTION The state of irreversible growth arrest is the most easily discernable characteristic of replicative senescence. Thus, senescent cultures are initially identified by the inability of the cells to enter cell cycle. However, the functional, genetic, and phenotypic alterations associated with senescence may be at least as important to the biology of cells as the inability to proliferate (Campisi, 1997). In the case of fibroblasts, for example, cells that have reached senescence in culture cease producing matrix-enhancing proteins, and start secreting such substances as collagenase, which can destroy the intracellular matrix. In addition, senescent, but not early passage, fibroblasts enhance the growth of pre-malignant tumor cells both in cell culture and when injected into mice (Krtolica et al., 2001). For CD8 T cells, one of the major changes observed in cultures that have reached replicative senescence is resistance to apoptosis, a property they share with senescent fibroblasts (Wang et al., 1994). Whereas cells from early passage cultures undergo brisk apoptosis in response to a variety of stimuli (e.g. mild heat shock, antibodies to Fas or to the T cell receptor), cultures of senescent CD8 T cells show significantly reduced ability to undergo apoptosis, and increased expression of the anti-apoptotic protein, Bcl2 (Spaulding et al., 1999). This change in the ability to initiate timely and efficient programmed cell death is highly relevant to effective immune function in vivo, since elimination of the massive numbers of activated virus-specific CD8 T cells is an essential event once the infection has been resolved (Effros and Pawelec, 1997). Another notable characteristic of CD8 T cell replicative senescence in cell culture is an alteration in the pattern of cytokine production (Effros et al., 2005). Cytokine secretion by T cells is essential for cell–cell communication and efficient immune function. Our cell culture studies show that as the cultures progress to senescence, there is an increasing concentration of two pro-inflammatory cytokines in the culture medium. Specifically, the levels of both tumor necrosis factor-alpha (TNFα) and IL-6 increase progressively as the cells reach senescence. These two cytokines are often associated with frailty in the elderly, and TNFα serum levels in HIV-infected persons are closely linked to adverse disease outcomes. A second important change in cytokine secretion is the anti-viral cytokine, interferon-gamma (IFNγ), which CD8 T cells secrete in conjunction with their cytotoxic function. With progressive cell divisions in culture, HIV-specific CD8 T cells show significantly reduced production and secretion of IFNγ, along with reduced lytic capacity and diminished production of perforin, a protein involved in cytotoxicity (Dagarag et al., 2003, 2004; Yang et al., 2005). A second important alteration in gene expression relates to the enzyme telomerase, which has the capacity to counteract the normal telomere shortening that accompanies cell division. High levels of telomerase are present in the developing embryo, but after birth, telomerase activity is retained only in stem cells and germline cells. Although most normal somatic cells lack telomerase, during activation, cells of the immune system are able to upregulate telomerase activity (Hiyama et al., 1995; Bodnar et al., 1996; Weng et al., 1996). Therefore, it seemed somewhat paradoxical that T cells, which produced high levels of telomerase activity in concert with antigen recognition, should ever undergo replicative senescence. To carefully analyze telomerase dynamics in T cells, long-term cultures were followed over time and tested for telomerase activity and telomere length at various points along the trajectory to senescence. Our studies showed that the overall loss of telomere sequences occurs at a rate of 50–100 bp/cell division, as had been shown for a variety of cell types (Harley et al., 1990; Vaziri et al., 1993). During the period following activation,

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telomerase activity is as high as that present in tumor cells and telomere length is actually maintained (Bodnar et al., 1996). Nonetheless, this high telomerase activity induced in response to the first and second encounters with antigen is not sustained during subsequent stimulations. In fact, by the fourth antigenic stimulation, CD8 T cells show no detectible telomerase activity. Interestingly, CD4 T cells from the same donors and subjected to identical rounds of antigenic stimulation retain high levels of telomerase activity even as late as the seventh antigenic stimulation (Valenzuela and Effros, 2002). At the point when telomerase was undetectable in the CD8 T cell cultures, the cells had undergone the same number of population doublings as the CD4 T cell cultures, suggestive of an intrinsic difference in telomerase dynamics between the two T cell subsets. In any case, telomere loss and a critically short telomere length seem to be intimately involved in the ultimate signaling of the cell cycle arrest associated with replicative senescence. In comparison to early passage cells, senescent T cells also show a significantly blunted upregulation of the hsp 70 protein in response to a mild, brief heat stress (Effros et al., 1994b). Finally, as cells age in culture, they show increased microsatellite instability, an indicator of reduced DNA mismatch repair capacity, which is capable of rectifying errors in DNA replication (Krichevsky et al., 2004). Thus, as T cells progress to the end stage of replicative senescence in cell culture, they are altered in a variety of processes reflecting cellular integrity and defense. Arguably, one of the most significant changes associated with T cell replicative senescence in cell culture is the complete and irreversible loss of expression of the major signaling molecule, CD28 (Effros et al., 1994a; Vallejo et al., 1998). Signaling through this so-called co-stimulatory molecule is involved in a variety of T cell functions, including activation, proliferation, stabilization of cytokine messenger RNA levels, and glucose metabolism (Shimizu et al., 1992; Holdorf et al., 2000; Sansom, 2000; Frauwirth et al., 2002). Importantly, the absence of CD28 expression is in marked contrast to the sustained expression of a variety of T cell-specific cell surface markers reflecting lineage, memory, and cell–cell adhesion. Thus, the permanent loss of CD28 expression in senescent T cell cultures constituted a biomarker that provided the unique opportunity to address the critical issue of whether CD8 T cell replicative senescence might be occurring in vivo.

SENESCENT CELLS ARE PRESENT IN VIVO Flow cytometry analysis of peripheral blood samples has clearly demonstrated that persons aged 70–90 have high proportions of CD8 T cells that lack CD28 expression. Indeed, in some elderly persons, more than 50% of the CD8 T cells within the total peripheral blood T cell pool do not express the CD28 molecule (Effros et al., 1994a). Cells in this category have shorter telomeres than CD28-expressing CD8 T cells from the same donor, and they also show minimal proliferative activity (Effros et al., 1996). Thus, by several criteria, they resemble CD8 T cells that have reached replicative senescence in culture. Importantly, the presence of these putatively senescent T cells in vivo is not restricted to aging. Rather, the proportion of these cells increases progressively over the lifespan, starting at 1% at birth (Azuma et al., 1993; Effros et al., 1996; Boucher et al., 1998). Moreover, these putatively senescent CD8 T cells are significantly increased in situations of chronic infection, such as HIV (Borthwick et al., 1994; Brinchmann et al., 1994; Jennings et al., 1994), so that 40 year olds who are HIV-positive show proportions that are as high as uninfected 90 year olds. CD8 T cells with the same phenotype have been reported in the context of certain forms of cancer as well. For example, in advanced renal carcinoma, the proportion of CD8 T cells that are CD57-positive (a marker present on the majority of CD28-negative T cells) has predictive value with respect to patient survival (Characiejus et al., 2002). Also, in patients with head and neck tumors, it has been shown that the CD8CD28 T cell subset undergoes expansion during the period of tumor growth, consistent with the notion that the increased antigenic burden may drive the tumor-reactive cells to senescence (Tsukishiro et al., 2003).

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The premature cessation of telomerase induction in the CD8 versus CD4 subset seen under the controlled conditions of cell culture may explain the in vivo preponderance of CD8 versus CD4 T cells with characteristics suggestive of senescence during aging (Boucher et al., 1998) and chronic infection (Effros et al., 1996). Indeed, studies on Epstein–Barr virus (EBV) infection have shown that telomere length of antigen-specific CD8 T cells is maintained during the acute infection stage, but once the virus establishes latency, these cells do undergo telomere shortening (Hathcock et al., 1998; Maini et al., 1999), presumably due to the same phenomenon observed in telomerase downregulation in repeatedly stimulated cultured CD8 T cells. What is the driving force for the generation of senescent CD8 T cells in vivo? The most likely cause of the excessive division of certain CD8 T cells in the intact organism is chronic antigenic stimulation, which could be the result of long-term exposure to antigens associated with latent viral infections as well as certain tumor antigens. It has been suggested that latent infection with several herpes viruses, which are endemic and persist throughout life in infected individuals, is the main culprit (Pawelec et al., 2004). Clinical data on bone marrow and organ transplant recipients indicate that under conditions of immunosuppression, cytomegalovirus (CMV), and other latent herpes viruses are often reactivated. Moreover, these patients show increased incidence of EBV lymphomas. In the elderly, many of whom are also immunocompromised, another herpes virus, varicella zoster, is often reactivated, manifesting itself as shingles. Reactivation rarely occurs in healthy individuals with normal immune systems, suggesting that maintaining latency requires active participation by the immune system, and that the constant and prolonged CD8 T cell activity to help maintain the latent state may drive certain virus-specific T cells to senescence.

SENESCENT CD8 T CELLS ARE ASSOCIATED WITH A VARIETY OF NEGATIVE HEALTH OUTCOMES The presence of senescent CD8 T cells in vivo may have a variety of effects on the proper function of both the immune system as well as other organ systems. In terms of immune function, senescent CD8 T cells undoubtedly influence the quality and composition of the memory T cell pool. Due to the property of apoptosis resistance, once senescent CD8 T cells are generated, they persist, leading to their progressive accumulation over time. Since homeostatic mechanisms are believed to independently regulate the memory and naive T cell pools (Freitas and Rocha, 2000), a high proportion of senescent cells will result in the reduced proportion of proliferation-competent, non-senescent memory cells. The fact that CD28-negative T cells are usually part of oligoclonal expansions (Posnett et al., 1994; Schwab et al., 1997) would presumably also lead to a reduction in the overall spectrum of antigenic specificities within the T cell pool. The repertoire of antigenic specificities is, in fact, reduced in elderly persons who have high proportions of CD8 T cells lacking CD28 (Ouyang et al., 2003). A more direct effect of senescent CD8 T cells is their putative suppressive activity on other cell types. For example, a population of CD8CD28 T cells generated in the course of in vitro and in vivo immunizations has been shown to suppress immune reactivity by affecting the process of antigen presentation (Cortesini et al., 2001). In the context of organ transplantation, the suppression may actually work to the benefit of the patient, by suppressing immune-mediated organ rejection. Indeed, donor-specific CD8CD28 T cells are detectable in the peripheral blood of those patients with stable function of heart, liver, and kidney transplants, whereas no such cells are found in patients undergoing acute rejection (Cortesini et al., 2001). By contrast, in other situations, the suppression of immune reactivity by CD8CD28 T cells may not be beneficial. An illustration of the possible negative outcome of this suppression emerges from the observed correlation between poor antibody responses to influenza vaccines in the elderly and the presence of high proportions of senescent CD8 T cells (Goronzy et al., 2001; Saurwein-Teissl et al., 2002).

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Additional effects of CD8 T cells with a senescent phenotype have also been reported. CD8 T cells that express a marker known as CD57 (the expression of which is associated with loss of CD28) exert suppressive influences on effector functions of HIV-specific CTL (Sadat-Sowti et al., 1994) and CD8CD28 T cells also accumulate and mediate liver damage in hepatitis C infection (Kurokohchi et al., 2003). A novel cellular interaction between CD8CD28 T cells and endothelial cells has recently suggested by in vitro experiments, which if confirmed in vivo, would have major implications on HIV pathogenesis. It has been reported that primary human endothelial cells that are exposed to culture supernatants from CD28-negative, but not CD28-positive, T cells show increased expression of a series of cell surface molecules that are specific markers of Kaposi’s sarcoma (KS) (Alessandri et al., 2003). The endothelial cells also acquire proliferative and morphological features of KS cells. The effect is mediated by soluble factors, such as TNFα, which, as noted above, are significantly increased in cultures of senescent CD8 T cells. A variety of pathological conditions have been correlated with the presence of senescent CD8 T cells. For example, a population of TNFα-producing CD8CD28 T cells has been identified in patients with cervical cancer (Pilch et al., 2002). Expanded populations of CD8CD28 T cells are present in anklylosing spondylitis patients, and, in fact, correlate with a more severe course of this autoimmune disease (Schirmer et al., 2002). Cells with the same phenotype accumulate in persons with coronary artery disease, suggesting some chronic antigenic exposure related to atherosclerosis (Jonasson et al., 2003). Finally, as noted above, there is a progressive expansion of CD8CD28 T cells in patients with head and neck tumors. Interestingly, the proportion of these cells is reduced upon tumor resection (Tsukishiro et al., 2003). The common thread in many of these reported accumulations of CD8CD28 T cells is chronic antigenic stimulation, be it by virus, alloantigen, autoantigen, or tumor-associated antigen. The regulatory effects of senescent CD8 T cells are not restricted to the immune system. For example, there is accumulating evidence indicating that bone biology is directly linked to immune system activity, and that chronic immune activation is associated with bone loss (Arron and Choi, 2000). The CD8 T cell subset, in particular, has been implicated in both bone resorption activity (Buchinsky et al., 1996; John et al., 1996) and osteoporotic fractures in the elderly (Pietschmann et al., 2001). One of the central regulators of bone resorption is expressed on, and also secreted by, activated T cells. This molecule, known as “RANKL” (receptor activator of NFkB ligand), binds to RANK on osteoclasts, inducing these bone-resorbing cells to mature and become activated (Kong et al., 2000). Under normal circumstances, the bone-resorbing activity induced by RANKL is kept in check by IFNγ, a cytokine also produced by the activated T cells (Takayanagi et al., 2000). However, CD8 T cell replicative senescence is associated with reduced ability to produce IFNγ (Dagarag and Effros, 2003). A second type of defect relates to the fact that activated T cells affect not only osteoclasts, but can also modulate bone mass by producing cytokines that inhibit the bone-forming activity of osteoblasts. Notably, IL-1 and TNFα inhibit osteoblast bone-forming activity, and also affect bone mass by inducing formation of certain cytokines by osteoblasts that increase bone resorption (Lorenzo, 2000). Senescent CD8 T cell cultures contain high levels of TNFα (Cenci et al., 2000), further implicating this class of T cells in the modulation of bone metabolism.

RETARDING OR PREVENTING REPLICATIVE SENESCENCE IN AGING AND HIV DISEASE Based on the emerging picture of pleiotropic negative effects exerted by senescent CD8 T cells in vivo, efforts have been directed at developing strategies to prevent or retard the process of replicative senescence. It has been suggested that, since a large proportion of senescent CD8 T cells are directed at CMV, childhood vaccination against this latent virus might offer a practical preventive approach (Pawelec et al., 2004). However, CMV vaccines are apparently quite technically challenging to develop. Moreover, there is evidence that even

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natural immunity that occurs as a result of actual infection does not prevent reinfection with CMV (Grundy et al., 1988). A second possible approach might be to physically remove senescent cells from the circulation, thereby stimulating the expansion of more functional memory cells. It is unclear, however, whether this expansion might ultimately lead to the generation of additional senescent cells. Given the central role of telomere shortening in the replicative senescence “program” in T cells, our own approach at modulating senescence has focused on strategies to enhance telomerase activity in CD8 T cells. An excellent model system for these studies is the virus-specific CD8 T cell response, which is known to decline with age and chronic infections, such as HIV. Therefore, using the long-term culture system described above, we followed HIV-specific CD8 T cells that had been isolated from persons infected with HIV, and tested the effect of gene transduction with the catalytic component of telomerase (hTERT). Comparisons were made between the hTERT-transduced cultures and the empty-vector-transduced cultures (Dagarag and Effros, 2003; Dagarag et al., 2003). Results of these experiments showed significant effects of hTERT on proliferative and functional aspects of the T cells. Briefly, we observed that hTERT transduction led to telomere length stabilization and reduced expression of the p16INK4A and p21WAF1 cell cycle inhibitors, implicating both of these proteins in the senescence program (Dagarag et al., 2004). Indeed, the transduced cultures showed indefinite proliferation, with no signs of change in growth characteristics or karyotypic abnormalities. In terms of protective immune function, the “telomerized” HIV-specific CD8 T cells were able to maintain the production of IFNγ for extended periods, and showed significantly enhanced capacity to inhibit HIV replication. The loss of CD28 expression was delayed considerably, although ultimately not prevented, suggesting that additional genetic manipulation of the CD28 gene itself may be required for full correction of this important senescence-associated alteration. Similarly, virus-specific cytolytic function was not restored by hTERT transduction (except in selected clones). Thus, hTERT corrects most, but not all, the alterations associated with replicative senescence in CD8 T cells isolated from HIV-infected persons. Ongoing studies are addressing whether transduction at earlier time points along the trajectory to senescence will enhance the telomerase effects. Telomerase enhancement may also be achieved using non-genetic strategies, which would offer more practical approaches to therapeutic interventions in the elderly. Pharmacologic enhancement of telomerase has the important advantage over gene therapy approaches of allowing control over the dose and timing. Several categories of non-genetic telomerase-enhancing treatments show preliminary promise in cell culture studies. Estrogen or modified “designer” versions of the hormone may have the desired effect. It is well established that estrogen is able to enhance telomerase activity in reproductive tissues. The complex formed when estrogen binds to its receptors migrates to the nucleus and functions as a transcription factor. In normal ovarian epithelial cells, this complex actually binds to the hTERT promoter region (Misiti et al., 2000). It has been known for some time that T cells can bind to estrogen via specific estrogen receptors. Thus, we tested whether pre-incubation of T cells to 17β-estradiol prior to activation might augment telomerase activity. Our preliminary data suggest that estrogen does, in fact, enhance T cell telomerase activity (Effros et al., 2005). The enhancement is observed in both CD4 and CD8 subsets, and can also be seen when estrogen is conjugated to bovine serum albumin (BSA), indicating that surface estrogen receptor interaction may be sufficient to mediate the telomerase effect. In another set of preliminary experiments with small molecule telomerase activators, we have shown a significant enhancement of telomerase activity in T cells from both healthy and HIV-infected persons. The increased telomerase activity is accompanied by enhanced proliferation and anti-viral functions (Fauce et al., 2006). Thus, therapeutic approaches that are based on telomerase modulation would seem to be promising candidates for clinical interventions that are aimed at reversing or retarding the process of replicative senescence in T cells.

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CONCLUDING REMARKS Normal somatic cells have limited expansion potential, a feature that has a dramatic effect on certain cells within the immune system. Indeed, the presence of senescent T cells in vivo has been documented in a variety of contexts, including aging, HIV infection, and cancer. Moreover, certain forms of cancer immunotherapy that are dependant on continued expansion of functional anti-tumor CD8 T cells will also be severely limited by the innately restricted expansion potential of immune cells. Therefore, manipulation of the process of replicative senescence in CD8 T cells constitutes a novel approach to regenerating functional immune cells. This strategy is relevant to a wide spectrum of clinical scenarios and broadens the spectrum of approaches to regenerative medicine. ACKNOWLEDGMENTS The research described in this chapter was made possible by the following sources of support: National Institutes of Health, University-wide AIDS Research Program, UC Discovery grant, Geron Corporation, and Telomerase Activator Therapeutics (TAT), Ltd. The author holds the Elizabeth and Thomas Plott Endowed Chair in Gerontology.

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27 Mechanical Determinants of Tissue Development Jonathan A. Kluge, Gary G. Leisk, and David L. Kaplan

INTRODUCTION The field of tissue engineering offers promising new solutions for replacement or repair of damaged tissues and organs. The ultimate goal of these strategies is to fully restore normal tissue function. The most common tissue engineering approach toward this goal is to develop viable constructs in vitro that can be implanted in the human body. Ideally, the implanted tissue continues to develop, providing the structure, composition, cell signaling, and functions that the native tissue exhibits (Vunjak-Novakovic et al., 2005). To develop viable constructs in vitro, it is believed that the in vivo conditions which promote growth and differentiation of target cell types should be replicated as closely as possible. Unfortunately, in vivo environmental conditions, native tissue mechanical loading, and the complex signaling critical to cell function and tissue development are all difficult to quantify, let alone replicate. Toward this end, a wide array of bioreactors has been developed by researchers to provide an in vitro environment that recapitulates the in vivo environment as faithfully as possible. Providing a limited set of loading and environmental conditions, often specific to the type of tissue being produced, modern bioreactor designs are increasingly sophisticated and produce ever-improving tissue quality. In this chapter, we focus on the effect that a specific epigenetic factor, mechanical stimulation, has on tissue development. To understand the complex relationship between mechanical loading and tissue development, we draw on research from biomechanics, cell biology, and biochemistry, three fields rapidly discovering overlapping themes and unsolved challenges. These challenges will be met through collaborative interdisciplinary efforts, sparked by recent initiatives (Kaplan et al., 2005). By the end of this chapter, the reader will have an understanding for the role of mechanics on two levels of cell and tissue function: the macro, full tissue level, and the microscopic cell level, including the extracellular matrix (ECM) and intracellular signaling mechanisms. By reviewing principles in both mechanics and biology, and then proceeding from a macroscopic to a microscopic perspective, the interplay between levels, and how they influence tissue engineering, will become clearer. The challenges presented to the field due to the complexity found in biology require a confluence of modeling, bioreactor design, and biomaterials engineering that best replicate in vivo tissue development in vitro. This complexity derives in part from the structural hierarchy found in biological materials, which creates difficulty in measuring and applying mechanical forces in a developmentally relevant temporal and spatial manner in vitro. Compressed time frames are needed to satisfy potential therapeutic benefits in vivo. Further complicating the situation are the diversity of cell types and states of cell function that exist cooperatively in any given tissue type, the presence of gradients of structure and function, and the effects of water and related environmental variables on tissue structure and thus mechanical properties and overall function. The challenges ahead

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are immense, however, scientific progress achieved over the last 5 years suggest future options remain bright to bridge the needs of biomechanics and functional tissue engineering.

MECHANICAL FORCES EXPERIENCED BY TISSUES Proceeding with a top-down approach to the role of mechanical forces in tissue development, we begin with a brief introduction to mechanics of materials, a review of the gross loading conditions on the body’s major connective tissues, and the measurement of tissue response. The reader seeking additional information on mechanics of materials, specifically biomechanics, is referred to full texts on the subject (Fung, 1993; Mow and Huiskes, 2005). Mechanics of Materials An understanding of how biological materials respond under mechanical loading requires knowledge of the types of external loads that may be applied, the internal forces and stresses that are generated, and the properties that govern the material response. This area of biomechanics, known as mechanics of materials, is briefly introduced here. This information will be important when studying gross mechanical forces on tissues and mechanical forces that act directly on cells and their local environment. Force and Stress An object that is externally loaded may move if it is unconstrained and deform if made from a deformable material. For an object that is physically constrained, such as a girder in a building frame or a tendon of the human body, forces and moments (i.e. bending or twisting action) at the constraints counteract the applied loads and may restrict object motion or deformation. This presence of simultaneous external loads and reaction forces and moments can generate a variety of internal forces within an object. At an arbitrary section through the object, one could isolate: normal forces that act perpendicular to the section, pushing or pulling on the object; shear forces that act along the plane of the section; torsional moments (torques) that twist the object about an axis perpendicular to the section; and bending moments that bend the object about an axis within the plane of the section (Hibbeler, 2000). Stress, which is a quantity representing the intensity of force, is separated into two types: normal and shear stress. Normal stress acts to cause local expansion or contraction within a material, while shear stress causes distortion. The deformation due to normal stress is simply called strain, while the distortion caused by shear stress is called shear strain. In the case of simple normal and shear force application, resulting normal stress and shear stress will cause the material to undergo deformation and distortion, respectively. Torsional moments tend to cause distortion only, while bending moments generate both deformation and distortion (Hibbeler, 2000). The magnitude of stress experienced by an object is affected by its geometric shape and dimensions and the nature and magnitude of external loading. The nature of loading on human connective tissues, for example, can range from biomechanical loads, imposed on the body through the actions of sitting or engaging in athletic activities, to physiological loads, such as pressure and flow effects of blood and other bodily fluids. Stress is a derived quantity and, therefore, not directly measurable. To quantify the stress in an object of known geometry, indirect measurement techniques are often used, such as measuring strain directly and correlating it to the stress from which it was produced (Hibbeler, 2000). Material Properties How a material responds to stress is dictated by its mechanical properties. These properties are derived from destructive or non-destructive testing, usually employing standardized equipment and test procedures.

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For example, the uniaxial tension test is a standard destructive test that involves pulling a carefully prepared material specimen with uniform tension, while recording the applied load and resulting material deformation. Material properties, such as Young’s modulus, which characterizes material stiffness, as well as strength properties, such as yield and failure strength, can be calculated from a tension test. Many other tests can be applied, such as compression, shear, bending, fatigue, and torsion tests. In material property testing, the closer the test conditions, such as the loading magnitude, rate, and specimen geometry of a material, resemble the actual parameters, the greater the confidence in the derived property. Forces in Biological Tissues The types of loading conditions experienced by biological tissues are varied, depending on the specific tissue. Some tissues experience continuous loading and unloading cycles, often in response to the body’s movement (e.g. bone and cartilage response to walking); others experience a state of prestress, in which a low load level is constantly applied (e.g. ligament tension and bone compression). In contrast to most commonly used engineering materials, many biological tissues respond to loading regimes with nonlinear, time-dependent deformation (i.e. a viscoelastic response). Such nonlinear response is more challenging to characterize and model. Mechanical Properties of Tissues Like most commonly used engineering materials, the tissues of the body do not last forever. Just as every component in an automobile has a finite lifetime, individual structures in a human body eventually fail, whether due to catastrophic events, disease, or normal wear and tear. One approach that automobile manufacturers employ to ensure component longevity is to design relevant assemblies such that stress levels and the number of stress cycles experienced by the component are minimized. Manufacturers can then select constituent materials whose mechanical properties (e.g. breaking strength, cycles to failure) comfortably exceed anticipated stress levels. We can view tissues of the human body in a similar light: to function properly over time and through many cycles, tissue strength properties should comfortably exceed stress levels generated by anticipated loading conditions. The reader should note that in tissue engineering practice, one does not necessarily exercise growing tissue constructs to the perceived structural potential of functioning tissue in vivo. Typically, low-level, continual loading regimes are applied to the constructs in a bioreactor environment. Regardless of the loading magnitudes used, one must ensure that the tissues have sufficient mechanical integrity through the regenerative process. The mechanical properties of specific (engineered) tissues are discussed in other chapters. Measurement of Force in Tissues In determining the mechanical properties of tissues, special equipment is often needed to isolate the tissue in question, measure its response to a predetermined loading criterion, and capture and record the data (Fung, 1993). This equipment ranges from implantable strain transducers and data acquisition devices to standard material testing equipment and non-standard video imaging, depending on whether measurements are to be made in vivo or ex vivo. Since such techniques typically provide individual measurement parameters, such as a load level or an amount of deformation, they are often combined with analytical or empirical information on a tissue’s constitutive behavior to derive additional measurement parameters. Especially for more complicated tissue structures, such as intervertebral disks, the in vitro or in vivo measurements are used to corroborate mathematical or computer-based models, sometimes referred to as in silico modeling (Prendergast et al., 2005). With the understanding that researchers in the field of biomechanics are continually discovering new ways to measure the mechanical properties of tissues, only a brief review of those techniques will follow.

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Ex Vivo Measurements Whole sample tissues can be tested ex vivo. For example, the distention forces caused by applied pressure can be measured on isolated vascular tissue explant samples. Measurement data generated in this fashion can be valuable for the design of viable tissue constructs (McCulloch et al., 2004). In vitro testing involves subjecting samples to a limited set of environmental and loading conditions that mimic in vivo conditions. Specific responses can be isolated in this way. Despite the prevalence of in vitro testing, some disadvantages exist. For example, in vitro testing tends to be short term and cannot exactly mimic natural conditions. In addition, the loading conditions of tissues in the body, such as biaxial loading in pressurized vessels, are difficult to replicate in bench-top testing. In Vivo Measurements Another common practice for deriving mechanical properties of tissues is through in vivo measurement of native tissue function. Various invasive and non-invasive techniques have been pursued by researchers to acquire data regarding internal tissues. Invasive techniques include the use of a shaped indenter or an aspirator to create a measurable deformation of the tissue, from which the material’s elastic response can be deduced. Surgical instruments that have been modified to incorporate force and position sensors have also been used. Each of these techniques is intended to quantify the resistance of the material to deformation (Ottensmeyer and Salisbury, 2001). Several non-invasive techniques, all categorized as “elastography,” are also based on tissue deformation. Strain fields (deformations) produced using this technique are measured using magnetic resonance imaging (MRI), optics, ultrasound, or another technology. Another new non-invasive ultrasonic technique involves the use of an ultrasonic pulser to send an ultrasonic wave through a tissue and the use of a second ultrasonic sensor to measure the displacement (Doyley et al., 2005). Unfortunately, there are limitations to many invasive techniques used for in vivo measurement, including an inability to isolate tissue response from a single variable, a dearth of appropriate internal force sensors, and ethical concerns. In vivo measurements on animal subjects are sometimes pursued as an ethical and practical alternative to invasive human procedures. In these cases, attempts should be made to select animal models whose morphologies and relative sizes closely match the human tissue of interest. Strain gauge-based force transducers have been implanted to monitor tissue response in certain animals. In the case of research on rabbit tendons, force response was monitored during various activity levels, such as “in-cage” and vigorous activities (Juncosa et al., 2003). Similarly, strain gauges used to measure microlevel deformations in various animal bone tissues revealed the prevalence of different strain regimes throughout regular daily loading (Fritton et al., 2000). This type of in vivo monitoring may be used to develop specific design parameters for tissue engineering.

THE CELL AS A SIGNAL RECEIVER AND PROCESSOR Shifting the discussion from the macroscopic tissue level to the microscopic cell level, we now focus on the underpinnings of tissues, cells, and their molecular constituents within the framework of mechanics. The cells are the workforce behind tissue-engineered constructs, as they serve to generate the ECM, or the material which gives tissue its mechanical integrity. In addition, the cells contain their own internal structural hierarchies and means of adaptation to external mechanical forces, which may lead to the formation of new tissue (proliferation and metabolism) or establishment of terminal phenotypes (differentiation). The following sections will review the mechanosensing components of the cell, the overall process by which cells integrate mechanical signals to direct tissue-specific growth, and the role that mechanosensation (MS) has on cell proliferation and differentiation. Cell Receptors and Sensors Cells within living tissues transduce mechanical force by using a variety of mechanisms. Although the signaling processes of MS are complex, involving many different molecules and pathways, they may all be activated

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by similar mechanisms in response to a variety of incoming signals (Huang et al., 2004). Among possible mechanical determinants previously discussed, shear forces due to fluid flow, strains imparted through cell/ECM constraints, and high-frequency vibrations are among the most prevalent to which a cell will respond (Hamill and Martinac, 2001). Surface mechanisms, which allow cells to transmit these forces throughout the cell, and the inner supportive cell structures (the cytoskeleton) are the vehicles by which signals can be integrated from the host tissue. In this section, transmembrane matrix molecules will be reviewed in the context of their suspected ability to convert external physical forces to intracellular biochemical signals, followed by a review of the tensegrity model, which offers a correlation between surface-level alterations and widespread cell changes in a global network. Cell–Matrix Adhesions The cell makes contact with its surrounding ECM through “adhesions,” a term used to describe a wide array of protein-mediated molecular links. The membrane portion of a cell’s ECM adhesions contains specific integrins, which are heterodimers of α and β subunits that bind to specific sequences on ECM molecules through a large extracellular domain (Geiger et al., 2001). Intracellularly, these integrins will interact with plaque proteins, which could be bridging proteins that connect the integrins to the cytoskeleton or signaling molecules. Several intracellular multi-molecular proteins serve to link the actin portion of the cell cytoskeleton to membrane integrins; these linkers include α-actinin and talin among others. Signaling molecules, another widely classified group of intracellular plaque proteins, are often activated by integrins or their bridging proteins, and include focal adhesion kinase (FAK) and mitogen-activated protein kinase (MAPK). Extracellularly, ligands can act as part of the ECM adhesion receptors such as fibronectin (α5 and β1), vitronectin (αv and β3), and various collagens (α1 and β1), which are all supplemented by membrane-bound non-integrin proteoglycan components such as syndecan-4 and CD44 (Geiger et al., 2001) (Figure 27.1). Following the occupation of integrins by their ligands, the initial step in reinforcing adhesions involves the clustering of integrin molecules. Focal complexes, small dot-like structures associated with the cell lamellipodium (thin, flat extensions at the cell periphery), are typically the first structures formed at cell/ECM junctions, and are either transient or evolve into more stable focal adhesions. The creation of a more stable adhesion is thought to be generated internally by responses of the cytoskeleton to applied forces (Geiger et al., 2001). Intracellular activators (such as talin) are thought to interact with either the α or β subunit tail of integrins and induce their separation, thereby causing further conformational changes that open the binding site on the headpiece and allow the integrin to create this enhanced focal adhesion (Giancotti, 2003). Since most intracellular adhesion components are multi-domain molecules, having the ability to partner with several different molecules, there are innumerable combinations of molecular interactions that could be occurring to create a signal pathway or generate a stimulus response (Geiger et al., 2001). The complexity of such relationships does not diminish the impact that individual integrins have on the development of tissue. Specifically, studies using β1 integrin-deficient knock-out mice showed basement membrane defects (Stephens et al., 1995). Additionally, the early organization of collagen fibrils in vitro depends on fibronectin, whose attachment is determined by integrins (Geiger et al., 2001). Integrins exist in two major allosteric conformations that are determined by their activity: an inactive (low-affinity state) and active (high-affinity) states. When ECM ligands bind to the molecular headpiece, they induce conformational changes in the integrin that are propagated along its length. Release from the inactive state causes the intracellular tails to move away from each other so that the β subunit is free to engage the underlying cytoskeleton (Giancotti, 2003). Since the cytoplasmic segment undergoes conformational changes via intracellular linkers, and the extracellular segment is controlled by ECM interactions, integrin molecules appear to have two functions: to regulate the extracellular binding activity from inside the cell (inside-out

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Figure 27.1 Diagrammatic view of mechanical force propagation (seen as dark arrows) throughout tissue from macroscopic to nanoscopic levels. Human connective tissue, such as bone, is affected by repeated external loads (left). The underlying ECM (mostly collagen, represented by random fibrils) will be subjected to these forces, which are then transferred to cells through insoluble ligands (fibronectin, circles at cell periphery). Surrounded by ECM, the cell contains its own sub-structure: lines, both continuous and dotted, represent the cytoskeletal microtubules and actin microfilaments, respectively (middle). The result of external forces is intracellular signaling: recruitment of bridging proteins, paxillin (Pax), α-actinin (α-Act), and talin (Tal), and signaling molecules, such as FAK, to the site of developing integrin-mediated focal adhesions. The result is a force balance which presumably affects actin-bound signaling proteins (actin helices with connecting myosin) (right). signaling) and to elicit intracellular changes through ECM binding (outside-in signaling) (Giancotti and Ruoslahti, 1999). The signaling molecules are associated with enzymes that can trigger pathways, which ultimately lead to changes in protein production and cell fates. By following the response of a signaling molecule, FAK, one can appreciate the complexity of signaling pathways and the importance of focal adhesion in initiating cell responses. For several years, FAK has been associated with both the growth and the disassembly of integrinbased focal adhesion sites (Geiger et al., 2001). Recently, the relationship of FAK with various GTPase proteins (Rho, Rac, and Cdc42) and indirect association with integrins through bridging proteins have been elucidated, and place FAK at the forefront of several intracellular signaling pathways (Mitra et al., 2005). The molecular signaling pathways are too numerous and complex to provide a full review in the context of this discussion; instead, the possible role of FAK activation in various signaling channels will be outlined. To briefly illustrate these channels, a simple schematic is provided which links FAK to intracellular activity (Figure 27.2). Formation of new integrin-mediated focal complexes or the transduction of forces through integrins may lead to an activation of FAK signaling mechanisms, mainly recruitment of other focal adhesion proteins. One such mechanism, responsible for the assembly and disassembly of focal complexes, is FAK’s ability to control phosphorylation of the bridging protein α-actinin, which can cross-link and tether actin/myosin stress fibers (Mitra et al., 2005). The second mechanism, although not completely understood, may be the activation of the Rho effector diaphanous (mDia) which leads to the stabilization of the cytoskeleton (i.e. microtubules) at the leading edge of migrating cells (Geiger et al., 2001). The third mechanism is the activation and/or inhibition of the various GTPase proteins that lead to regulation of cytoskeletal extensions, such as stress fibers,

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Cadherin contacts

Focal contacts FAK activators and/or inhibitors (Rho, Rac, Cdc42, mDia)

alterations in polymerization

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Figure 27.2 Cell signaling mediated by integrin responses at membrane interfaces, such as due to changes in external mechanical states, leads to intracellular cascades, as shown. FAK plays a central role in these responses. Mediation of cell interactions with the external environment are summarized, including responses such as cell adhesion, spreading, and movement based on changes in focal contacts related to cell–matrix adhesion and cadherins related to cell–cell-mediated interactions. FAK functions to recruit other focal contact proteins or their regulators, leading to changes in the internal structure through polymerization and stabilization of cytoskeletal elements. All of the events illustrated occur in a complex symphony of orchestrated events to modulate internal and external changes in response to changes in external mechanical signaling (Figure is in part patterned after Figure 1 from Mitra et al. (2005)).

lamellipodia, and filopodia. The final mechanism is the formation and disassembly of cell–cell (cadherin-based) connections, providing an added route for solute exchange and signaling (Mitra et al., 2005). The importance of relationships between the cytoskeleton and molecular signaling pathways will gain further emphasis in the following discussion. Tensegrity Model of Cell An understanding of intracellular microstructure and hierarchy is critical for grasping the interactions between incoming signals and their propagation throughout the cell. The cytoskeleton is not merely a randomly configured collection of molecules; instead, it is believed that each different cytoskeletal molecule is integrated in a unique way to maintain the mechanical signaling pathways. Furthermore, the mechanical behavior of the whole cell is driven by both the cytoskeletal elements found just below the surface of the plasma membrane and also the internal cytoskeletal lattice, a component often overlooked because of its misunderstood complexity. The role of the cytoskeleton as a support and shape-retaining structure has long been recognized, but it is now known that it can also provide directed signals to the intracellular elements, and is thus capable of inducing endogenous changes to occur (Ingber, 2003a). Structural molecules that make up the cytoskeleton can be broken down into three major groups: actinbased microfilaments associated with the cell’s cortical cytoskeleton (adhesion complexes), stiff hollow tubulinbased microtubules that radiate from an organizational center (centrosome), and thick intermediate filaments, such as lamin, vimentin, and keratin, that integrate with the cell’s nucleus and attachment sites (desmosomes). Formation of larger and stronger cytoskeletal structures is possible when these molecules are supplemented by

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other proteins, such as actin-bound myosin chains, which combine to form stress fibers around the cell periphery. Altogether, these structural elements organize throughout the cell interior to form a complex network (Ingber, 2003a). A “Tensional Integrity” (or tensegrity) model, one that assumes compressive-bearing struts (microtubules and ECM adhesions) is resisted by the pull of surrounding tensile elements (microfilaments, intermediate filaments), has been espoused by Donald Ingber and his colleagues, and serves as the most prominent cytoskeletal structure-function model to date. This model adapts the principles used in the design of a ship mast and riggings, as linear reinforcing elements can be linked together to form tension-resistant scaffolding around a hull (cell membrane). Following this model of tensegrity, the majority of structural elements need only to have good tension resistance to maintain shape and stability, while avoiding the need for buckling or compression resistance of large rigid struts by creating a network of triangulated shorter members (Ingber, 1997; Boal, 2002). These intracellular structures, not unlike the tissues they inhabit, have been regarded as soft materials because of their shrinking and stiffening response to temperature and their relative ease of deformation. Cytoskeletal elements can exchange energy with their surroundings, permitting their shapes to fluctuate as they bend and twist in response to transverse loading. Whatever the deformation mode, energy is required to distort the filament from its “natural” shape (Boal, 2002). According to the tensegrity model, many cytoskeletal molecules in their natural state have a certain level of prestress or isometric tension, generated by the contractile function of actin and myosin sliding, osmotic forces, and/or forming new ECM adhesions. This prestress can be visually confirmed when the plating of cells on a flexible substrate does not lead to distortion, or when the cutting of a cell leads to spontaneous retraction of its intracellular cytoskeleton (Ingber, 2003a). The structural assembly of these elements, in addition to being designed for optimal structural stability, has also been designed for effective transport of molecules throughout the cell; therefore, their configuration must resemble that of a spider web or city plan. This two-dimensional network, common to all three main groups of molecules, can be found attached to the plasma or nuclear membrane, and exhibits many deformation modes in response to an applied force. Actin filaments and microtubules require linking proteins, such as Actin-binding Proteins (ABPs), in order to form these cross-link networks and composite structures (Boal, 2002). Permanent cross-links and a high intracellular density of filaments will add a compressive and shear resistance to cells. Mechanochemical Transduction Mechanochemical transduction (mechanotransduction) is the process whereby cells sense and respond to external stimuli. One widely held belief is that ECM proteins and integrins will undergo conformational changes in response to mechanical stimuli. Another belief is that intracellular perturbations of the preexisting cytoskeletal tension will initialize a response. It would seem as though the two are not mutually exclusive events, but, instead, coincide to facilitate signal transduction and to produce intracellular change. Ligands and Cryptic Binding Sites As an immature tissue develops and grows, the ECM must have a role in regulating cell interactions. For instance, it is believed that local changes in ECM structure and mechanics will alter the adhesion characteristics of epithelial cells, as they encounter new sections of the emergent basement membrane. In this instance, high turnover in growing sections of new ECM may lead to relatively thinner and more flexible (compliant) regions of the basement membrane (Ingber, 2003b). In response, fibronectin attached to these flexible regions will undergo large strains, leading to a possible exposure of cryptic self-association sites necessary for fibronectin polymerization (Geiger et al., 2001). It has been shown that forces as low as 3–5 pN are sufficient to unfold these cryptic subdomains in fibronectin, which can then lead to fibronectin fibril formation. In summary, once extracellular proteins are altered, integrin binding and cytoskeletal signaling will presumably occur (Huang et al., 2004).

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Mechanosensitive Ion Channels The lipid bilayer is believed to be one of the major mechanosensory (MS) components of a cell. When a membrane, such as the plasma membrane in eukaryotic cells, is deformed by a force, two changes may occur. Disturbance of the lateral force balance around a lipid bilayer may first lead to conformational changes of transmembrane proteins, with or without necessarily activating a second receptor (Janmey and Weitz, 2004). A second change may occur as membrane forces trigger opposing local curvatures that could reorganize the membrane chemically. It is believed that these changes lead to activated ion channels, which, in turn, respond with changes in their permeability (Hamill and Martinac, 2001; Kung, 2005). Whether the forces affect membrane proteins or lipids, and whether the magnitude of these forces can be correlated to channel properties is not clearly understood (Janmey and Weitz, 2004). The original study of channel gating (regulatory open-and-close mechanisms) in eukaryotic cells was heavily focused on mechanosensory neurons, in which gated channels of Xenopus oocytes could be activated for a latent response with pressure-clamp techniques (McBride and Hamill, 1999). After more than 20 years of research, patch-clamp studies have illustrated not only the prevalence of these channels across many eukaryotic species but also their key influences on cell volume regulation and the possibility that tight seal formation could lead to mechanosensitivity in focused K channels (Hamill and Martinac, 2001). Many of these experiments controlled membrane tension by suction pressure in a micropipette attached to a small region of the cell membrane. In these experiments, increased pressure to just below that which would cause the membrane to rupture was shown to increase pore dimension on the order of 0.5 nm in MS channels of large conductance. This “pressure relief valve” mechanism, in addition to its role in MS, can be seen as a cell’s natural defense against large osmotic gradients. Similar mechanisms have been linked to calcium ion (Ca2) channel activation in the stereocilia of hair cells in the inner ear and fluctuations in intracellular ion concentration of endothelial cells (ECs) (Hamill and Martinac, 2001; Huang et al., 2004). Although the existence of stretch-activated ion channels has been well documented for “specialized” cells (i.e. human cells associated with auditory function, visual function, etc.) and also for non-specialized cells, its mechanisms are not clear, nor are its connection with cytoskeleton-related mechanisms (Hamill and Martinac, 2001). Furthermore, a study in which a Triton buffer was used to remove the cytoplasm and apical cellular membrane showed a binding of paxillin, pp125FAK, and p130CAS to the Triton-insoluble cytoskeleton following a 10% stretch of collagen substrate (Sawada and Sheetz, 2002). These and other results indicate that, in addition to ion channels, there are other mechanisms that enable cells to sense physical forces. These mechanisms, in conjunction with transmembrane proteins discussed earlier, are part of the underlying cell cytoskeletal structure. Altering Intracellular Mechanics The major intracellular change that seems to occur as a result of external forces is in cytoskeletal molecular mechanics, as the internal lattice endures global change. Immediately after mechanical signals are sensed by surface integrins, the cytoskeleton will realign in the direction of the applied tensional stimulus, through deformations of the cytoskeletal lattice and nuclear scaffolds (Ingber, 1997). In keeping with the tensegrity model, forces transmitted by integrins to microfilaments in focal adhesions can be passed to microtubules at distant sites via intermediate filament connections (Ingber, 2003a). Actin and tubulin are dynamic polymers; their fundamental protein building blocks can both polymerize and depolymerize, depending on the conditions, changing the length of the filament in the process. Rapid depolymerization releases the contents of the microtubule to the cytoplasm and permits it, or nearby microtubules, to start reconstruction elsewhere (Boal, 2002). These mechanisms allow a cell to constantly adjust its internal prestress, and thus alter the tightness with which the cytoskeletal lattice is held together, if tensegrity is indeed at play. To account for this change, one must first recognize that many of the enzymes and substrates

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that mediate protein synthesis, glycolysis, and signal transduction appear to be immobilized on the insoluble cytoskeleton. It is believed that if the cytoskeletal molecules and their immobilized proteins distort without breaking following focal adhesion stimulation, then those attached molecules must similarly change shape. Altered biophysical properties may result in altered local thermodynamic properties, or altered kinetic behavior, just as a spring would change its vibration frequency following distortion (Ingber, 2003b). Similarly, the altered cytoskeleton has been shown to influence protein synthesis by destabilizing cytoskeleton-associated mRNAs at the intersections of actin filaments, and through polymerization at vertices within highly triangulated microfilament networks (Bassell et al., 1994; Ingber, 1997). The Hard-Wired Nucleus One major principle of the tensegrity model is that structural hierarchies exist on many levels between muscles and bones of connective tissue, between cells and the ECM, connecting surface receptors to the cytoskeletal elements, and sub-structures within the cytoskeleton, including a nuclear scaffold. The tensed intermediate filaments that connect to the nucleus and its proximal network may be a route by which the signals transduced through surface-level integrin complexes are delivered (Ingber, 2003a). Nuclear scaffolds may be “hard wired” to the integrins, such that distortions of adhesion complexes result in synchronized realignment of structural elements (mainly intermediate filaments) to the nuclear envelope, via the underlying laminin network (Huang et al., 2004). To prove this hypothesis, ligand-coated beads were used to pull focal adhesions of cultured ECs at very high rates, and then the nucleoli were shown to deform and elongate in the direction of applied force. This same phenomenon was also observed in cultures with extracted membranes and intracellular components, suggesting that this signal was transduced directly through the cytoskeletal lattice, and not through a signaling cascade (Maniotis et al., 1997). Cell Fates: Growth, Differentiation, and Apoptosis Cell proliferation (multiplication through mitosis), differentiation (changes in phenotype and matrix production), and apoptosis (programmed cell death) are all heavily reliant on the signaling mechanisms which were previously discussed. Most importantly, experimental observations combined with the tensegrity model account for how cellular interactions with their local environment, whether from other cells or ECM, can cause these different modes to be triggered. As previously explained, damaged or reconstructed ECM exhibit degraded mechanical properties. Following injury of normal tissue and subsequent loss of cellular elements, residual ECM will remain intact and will promote organized cellular tissue regrowth, ensuring the correct cellular placement and alignment. The intracellular changes that result from local alteration in ECM permit the cells to respond to soluble growth factors and other mitogens, thereby driving changes in tissue phenotype (Ingber, 2003b). Qualitative data on these intracellular changes serve to corroborate the mechanotransduction models. In a study by Chen et al., cell shape seemed to determine whether individual cells proliferate or undergo apoptosis, independent of the growth factor used to stabilize cell adhesions. The study first confirmed that attached capillary ECs display a flattened nucleus with an extended morphology when cultured on flat beads, as opposed to suspended cells which remain small and spherical. More importantly, the study also confirmed that, while most cells survived when spread on larger beads coated in fibronectin, cells seeded on the same beads with decreased diameter became more rounded, and matched the apoptotic pattern of non-adherent cells. The molecular trigger for this programming may be linked to the “hard-wired” nuclear mechanisms, affected by cytoskeleton rearrangement (Chen et al., 1997). The results of a more recent study would also indicate that the formation of ECM, via cooperative interactions between integrins, the cytoskeleton, and threedimensional tissue organization, confers prevention of apoptosis. Furthermore, this study showed that

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laminin-induced integrin ligation directs tissue polarity and promotes resistance to apoptosis, regardless of growth status (Weaver et al., 2002). It has been shown that transitions between growth and differentiation stimulated by these mitogens can also be influenced by the elasticity and geometry of the growth substrate, like the developing ECM. For example, a study by Mochitate et al. (1991) showed that collagen gels seeded with human fibroblasts which underwent stress relaxation (transient hypercontraction followed by dissipation) led to differences in cell morphology and biosynthetic activity, including disruption of their actin filament bundles, loss of cell surface fibronectin, and marked decrease in both cellular DNA and protein synthesis. Although mechanical stimulation appears to trigger the transition between growth and differentiation, what mechanisms are involved are not firmly established.

OVERVIEW OF BIOREACTORS Bioreactors provide an in vitro environment for tissue generation and growth. Ideally they mimic the mechanochemical regulation that tissues experience in vivo in their native environment. The key functions of the bioreactor are to: (1) allow the seeding of uniform concentrations of cells to a scaffold; (2) control physiological conditions in the cell culture medium (e.g. temperature, pH, oxygen levels, nutrients); (3) supply sufficient metabolites; and (4) provide physiologically relevant signals in the form of mechanical loads (Altman et al., 2002; Freed et al., 2000). Since there are many types of bioreactors in current use, one must choose a tissue-appropriate design that incorporates the unique set of in vitro environmental and mechanical loading conditions that can produce a tissue that is as similar as possible to the native tissue. The following is a survey of some bioreactor designs in current use. For additional details on bioreactor design and specific comparisons, the reader is referred to additional sources (Barron et al., 2003; Vunjak-Novakovic et al., 2004). Types The simplest type of bioreactor is the static flask; a tissue construct is fixed in place in a culture medium. Gas aeration is provided by surface aeration of the culture medium. Mass transfer, therefore, occurs by molecular diffusion since there is no fluid flow at the surface of the tissue construct (Barron et al., 2003). The structures formed under static conditions tend to exhibit limited cellular ingrowth, resulting in two-dimensional tissue structures. To better produce clinically relevant three-dimensional tissues, recent advances toward more biomimetic bioreactor designs with complex environments have been implemented. A stirred-flask bioreactor uses a magnetic stirrer to mix a dilute cell suspension around a stationary scaffold, aiding in cell distribution through the scaffold. Stirring of the culture medium produces mass transfer through turbulent convection, generates shear stresses that enhance cell and tissue growth in comparison to static incubation conditions, and improves nutrient supply through the scaffold (Barron et al., 2003; Nasseri et al., 2003; Martin et al., 2004). A wavy walled bioreactor is similar to a spinner-flask bioreactor, with the exception that the flask wall contains wavy contours that mimic baffles. This design provides a range of hydrodynamic forces, enhancing mixing of the culture medium (Bilgen et al., 2005). A rotating-wall vessel bioreactor provides a dynamic environment in which two concentric cylinders are horizontally rotated. Cells are grown on tissue constructs which are freely suspended in the annular volume, in essentially a microgravity environment, which is filled with culture medium (Nasseri et al., 2003). The tissue constructs benefit from the laminar flow, low shear stress fluid environment, and improved supply of nutrients and outflow of wastes (Barron et al., 2003; Martin et al., 2004). Direct perfusion bioreactors involve the perfusion of culture medium through tissue constructs. The perfusion can produce higher density and more uniform distribution of cells than with stirred-flask bioreactors

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and provide mechanical stress in the form of flow-induced shear (Martin et al., 2004). More advanced bioreactor designs include strain actuation that can apply static and dynamic loads, as found in the Flexcell line of bioreactors (Shukla et al., 2004). Modeling Given the need for future bioreactors to more faithfully represent the in vivo environment, which is very complex, the use of analytical and computational modeling will become more important. Computational fluid dynamic (CFD) software is a powerful tool to calculate flow fields, shear stresses, and mass transport within and around three-dimensional tissue constructs. CFD models have been use to study oxygen transport in a rotating-wall bioreactor, to model direct perfusion in various scaffold designs, and to evaluation the effect of pore structure and interconnectivity on tissue development (Martin et al., 2004). CFD models can help optimize bioreactor design and flow conditions (Bilgen et al., 2005). To aid in modeling efforts, additional modern tools have been employed. For example, computed tomography scanning can be used to construct computerbased models of tissue engineering scaffolds (Cioffi et al., 2005). Additional technology is showing promise for tissue engineering, including rapid prototyping, the introduction of smart materials in scaffolds, and advanced manufacturing techniques like electrospinning.

PRACTICAL EXAMPLES OF MECHANICAL DETERMINANTS The previous reviews of mechanotransduction and general cell responses outline cellular behavior in their mechanical environment. Several examples of these phenomena are offered next, as they occur in specific tissue lineages: vasculature, bone, and cartilage. The goal is to provide a sense of the physiological loading regimes, signaling that is transduced to matrix and cells, and their application in bioreactor design. Vasculature – Endothelial and Smooth Muscle Cell Loading Conditions Vasculature is made up of smooth muscle cells (SMCs), ECM (collagen and elastin fibrils), and ECs. ECs form a monolayer that covers the innermost aspects of vasculature, providing a barrier between flowing blood and the tissue wall. The SMCs and ECM provide the proper shape and size for blood flow, constrain the ECs, and provide structural integrity to withstand internal and external stresses (Davies, 1995). The three main components of vasculature, SMCs, ECM, and ECs are all subjected to stretching (strain) as a result of pulsatile blood flow. The amount of strain that is generated is directly related to blood pressure. Shear stress, due to fluid flow-generated frictional forces, is experienced principally by the ECs. These shear stress levels vary with the blood velocity profile generated during the cardiac cycle and with the shape and size characteristics of the vasculature. For example, curves in arterial walls can lead to flow separation and the development of vortices, which affects the shear stress near the vessel wall. It should be noted that pulsatile blood flow and blood elements also influence EC responses by varying the luminal concentrations of growth factors and other soluble mitogens that interact with apical surface integrin receptors (Davies, 1995). Only effects of fluid-induced shear stress on vasculature cell responses will be the focus herein. Cell Response and Transduction Mechanisms In response to fluid-induced shear stress, monolayers of ECs change morphology and become torpedo shaped, aligned in the fluid flow direction. It would follow that the cytoskeletal backbones of EC undergo major alterations, in which stress fibers reinforce the EC membrane (Satcher and Dewey, 1996). Many improvements in the estimation of cytoskeletal structure and strength of EC have aided in the understanding

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of how these and other human cells respond to external stimuli through cytoskeletal remodeling (Fung, 1993; Satcher and Dewey, 1996; Helmke and Davies, 2002). In the same way that mathematical models of whole tissue structures, based on their estimates of substructure and geometry, are used to aid in mechanical analysis of load distribution, the same tools are widely used in EC cytoskeletal analysis. Because of cytoskeletal responses, shear stresses acting on the luminal cell membrane of ECs in vivo are transmitted to the basal attachment sites. It is believed that the collection of plaque proteins ABP and spectrin, used to reinforce integrins at focal complexes on the basal side of EC, work in concert with reassembly effects (Satcher and Dewey, 1996). It is unclear whether focal complex enhancement is solely driven by basal side integrin activation, or if further support is also provided by the translocation of inactive apical side integrins to the basal membrane following shear stress (Shyy and Chien, 2002). In either case, the development of focal adhesions will lead to recruitment of cytoplasmic signaling molecules and MAPK signaling pathways. Focal adhesion sites, like the cytoskeleton, align their shape parallel to the flow direction without changing their overall contact area (Helmke and Davies, 2002). Activated luminal cell surface mechanisms (stretch-activated or potassium ion channels) have been linked to EC shear strain response. Similarly, G-protein activation due to distortions of the plasma membrane from shear has also been documented (Helmke and Davies, 2002). These and the integrin-dependent mechanisms are part of either the inside-out or outside-in signaling routes that develop from an EC’s complex response to shear. Bioreactor Design In the engineering of cardiovascular tissue, it is believed that bioreactor design should involve laminar fluid flows that induce a uniform distribution of shear stress and laminar convective mass transfer. Rotating-wall bioreactors have been used to create engineered cardiac tissues that are structurally and functionally superior to those grown in static or mixed flasks (Barron et al., 2003; Martin et al., 2004). Other bioreactors have been used which include strain actuation, mimicking the dynamic mechanical stimuli present in vivo. For example, it is thought that since arteries experience axial strains through connective tissue, tubular scaffolds that represent a cardiovascular vessel should experience the same strain. In addition, circumferential strains can be provided by a pulsatile force through the tissue scaffold, mimicking pulsatile blood flow in actual arteries (McCulloch et al., 2004). Bone – Osteocytes, Osteoblasts, and Osteoclasts Loading Conditions The enduring principles established by Wolff (i.e. Wolff ’s Law) state that the rate and degree of new bone tissue deposition is dependent on the tissue’s stress levels, and that the pattern of bone architecture coincide with stress trajectories (Wolff, 1870). Because of the rigidity (overall structural stiffness) of bone tissue, its deformation resulting from gross loads is small, typically on the order of microstrain (where 10,000 microstrain is the same as a 1% change in length). One study, reporting the use of high-resolution ( 0.08 microstrain) strain gauges to measure in vivo bone strains, showed that during the course of a day, few high-magnitude (1,000 microstrain) events occur. Furthermore, daily strains which fall below 0.2% strain are the predominant contribution to the strain history (strain cycling over time) of bone, and are encountered during the body’s regular posture corrections (Fritton et al., 2000). Assuming that focal adhesions of bone cells are distributed along force-bearing members of the surrounding ECM, osteocyte stretch will reflect microstrain-level deformation. However, most in vitro work on osteoctye response to dynamic substrates requires substrate deformations at least 1–2 orders of magnitude larger to induce changes in bone modeling (Han et al., 2004). The contradiction between modeled results using in vivo parameters and in vitro requirements can be rationalized by noting that mechanical loads applied to bone in vivo can cause increased pressure gradients

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and interstitial fluid flow (IFF) through bone channels, as the microcollapsed pores undergo volume changes. Additionally, it is hypothesized that cytoskeletal rearrangements (intracellular actin-bundle strains) can compensate for this distortion (Han et al., 2004). It has been suggested that the bending stresses in bone cause local and opposing tension and compression forces, which generate IFF in addition to the expected microstrains (Turner and Pavalko, 1998). The fluid shear from the IFF is responsible for the larger order disturbances to osteocytes (Cowin and Weinbaum, 1998). Cell Response and Transduction Mechanisms The mechanosensory mechanisms of bone tissues adopt the mechanotransduction models previously discussed, but are complicated by the signaling between bone sensory cells (osteocytes) and their effector cells (osteoblasts and osteoclasts), which are ultimately responsible for bone homeostasis and adaptation to strains. These cells, not including osteoclasts, are linked through what has been called a “connected cellular network” (CCN) (Cowin et al., 1991). Normal bone remodeling involves the creation of canals (bone resorption) by osteoclasts, followed by a filling of surface sites with mineralized osteoid (fibrous organic matrix) by stem-cell-derived osteoblasts. After resorption is triggered, it is believed that receptor-based mitogen (TGF-β, etc.) signaling and production of prostaglandin and nitric oxide (NO) initializes osteoblast activity. From there, the osteoblasts will either maintain their phenotype, resting at the new bone periphery, or differentiate into osteocytes while encapsulated in the surrounding matrix. Canaliculi, or nutrient and biochemical channels between bone cells, connect the embedded osteocytes and surrounding osteoblasts to form the CCN (Huiskes and van Rietbergen, 2005). Intracellular epigenetic mechanisms (cytoskeletal prestress, biochemistry, etc.) in the CCN allow the cells to respond to physical activity, while gap junctions are thought to act as electrical synapses to permit or block bidirectional information exchange (Cowin et al., 1991). Although damage to microstructure appears to be the source of osteoclast recruitment signaling, it is unclear whether the signals stem from breaks in the CCN pathways or from damage to the osteocyte matrix (Huiskes and van Rietbergen, 2005). Varying mechanical factors, including stress, strain-rate, and fatigue microdamage, have been extensively investigated at the macroscopic level, both of tissue explants and in vivo, as to their effect on remodeling and developing bone tissue. In some cases, the effects seem to be time dependent, while other experiments seem to indicate that the amplitude of oscillatory components or peak stress/strain values has the most impact. To parallel the studies on gross loading effects, research is now focused on modeling the CCN, as a number of densely interconnected electrical processing elements, or through computational finite element modeling that are capable of organizing the multitude of mechanical inputs (You, et al., 2001; Huiskes and van Rietbergen, 2005). As a transduction vehicle previously discussed, the distortions of bone matrix under strain may lead to cytoskeletal remodeling around the cell nucleus, via surface adhesions and distortions of the internal cytoskeletal lattice (Shyy and Chien, 1997). The mechanical stimulation of bone cells will also increase intracellular calcium levels and production of prostaglandins and NO within minutes, and has been linked to mechanosensitive ion channels (Turner and Pavalko, 1998). Furthermore, slow pulsating and/or oscillatory flow of interstitial fluids has been proven to be more effective in activating osteoctye cells, over situations of hydrostatic pressure or rapid oscillations (Huiskes and van Rietbergen, 2005). Bioreactor Design Several publications highlight attempts to recreate these mechanical loading regimes for engineered bone within different bioreactor environments. In one study, enhancements to static culture environments were incorporated through spinner-flask and rotating-wall bioreactors, which stimulated mesenchymal stem cells (MSCs) through bulk convective flow stimulation and centrifugal force balance, respectively. The results of

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dynamic culture in the spinner flask were most promising in comparison with rotating-wall vessel culture due to enhanced mixing; however, both showed marked improvement over static culture (Sikavitsas et al., 2001). Another more recent study compared static culture to spinner-flask cultures and perfused cartridge culture, one in which laminar IFF (35 μm/s) is mimicked using a gross fluid pressure differential across the sides of a cell-seeded construct. The results of this study indicated that although spinner-flask culture was the most successful at bulk generation of bone markers, the perfusion bioreactor did so in a randomly distributed manner throughout the construct’s volume (Meinel et al., 2004). To recreate load-induced IFF through indirect measures, one recent study cultured MSC on partially demineralized bone scaffolds subjected to cyclic bending loads in a custom-designed static flow bioreactor. The results of this study showed that mechanical stimulation of this nature promoted osteogenic differentiation of MSC by significantly elevating alkaline phosphatase activity and calcium deposition (known markers for bone) over static controls (Mauney et al., 2004). Cartilage – Chondrocytes Loading Conditions Because of the hydrophilic aggrecan proteins, and thus the large water content (65–85%) within cartilage tissues, the gross mechanical response to compression is somewhat like compressing a pneumatic tire, leading to stress levels varying between 0 and 20 MPa during movements. Collagen fibrils and other matrix proteins compensate for stress-bearing responsibilities under tensile and compressive loads, respectively. The response of cartilage tissue is nonlinear and time dependent (viscoelastic), meaning that deformation will increase with a constant applied stress. Like bone tissue, cartilage also behaves anisotropically and is subject to complex loads, including bending and shear. A thorough review of experimentally derived modeling considerations are outside the scope of this review, and can be found elsewhere (Mow et al., 2005). Cell Response and Transduction Mechanisms As in the two previous examples, chondrocytes within cartilage tissue can sense and respond to mechanical stimuli; however, chondrocytes do not rely on the stimulus/effector relationship with other cell phenotypes. Instead, the interactions of chondrocytes with their matrix seem to be critical, owing to the scarcity of chondrocytes within most cartilage tissues (5%). Cartilage tissue is almost completely avascular and aneural, which somewhat simplifies the study of chondrocytes and their matrix in tissue engineering research. In vivo, static compression of the tissue to physiological strain magnitudes leads to breakdown of cartilage proteoglycan, not renewal; however cyclic loads of a higher magnitude or frequency can also be deleterious (Mow et al., 2005). Therefore, consistent and mid-level stresses appear to create favorable mechanical environments for cartilage regeneration in vitro. The rapid or acute response of chondrocytes to mechanical stimulation was studied in vitro using a twodimensional monolayer model, and revealed that substrate stretch, via deforming pressure gradients of 50 kPa at 0.33 Hz, induced membrane hyperpolarization within 20 min. In vitro, this study confirmed that tyrosine phosphorylation of both paxillin and FAK was induced within 1 min of initiation of stretch, and led to the eventual signaling cascade inducing small conductance potassium channels (Millward-Sadler and Salter, 2004). Bioreactor Design For tissue engineering of cartilage, in vitro cultivation of chondrocytes seeded on biodegradable scaffolds has been pursued using spinner-flask bioreactors. However, some research has shown that the turbulent, highshear mixing environment in spinner flasks can lead to altered, undesirable morphology of the engineered cartilage tissue (Sucosky et al., 2004). The use of a wavy walled bioreactor can produce enhanced mixing at

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low shear levels, leading to increased rates of formation and size of chondrocyte aggregates in suspension cultures (Bilgen et al., 2005). Other research has shown that flow-induced shear stress can be beneficial in increasing ECM component production by chondrocytes. A bioreactor that includes direct perfusion should also exhibit enhanced convective transport of nutrients to the cells and catabolites (waste) away. The level of shear stress applied depends on the culture medium flow rate through the constructs and also the threedimensional scaffold geometry (Cioffi et al., 2005).

CONCLUSION Since mechanical forces play a crucial role in tissue development, function and repair in vivo, the design of novel bioreactors to impart complex mechanical forces to cells and tissues in vitro offers important options to improve functional tissue engineering. These inputs have to be considered within the context of the biomaterial scaffolds used in the bioreactors to transmit the applied forces or to handle fluid flow, to the cells used in these systems, and to the overall system needs to generate functional tissues in vitro for utility in vivo. Full restoration of a mechanical match for tissue grown in vitro to repair needs in vivo may not be required, as long as the engineered tissue satisfies temporary mechanical and related requirements until tissue regeneration and integration is achieved. It is clear that the road ahead is challenging, yet promising results and approaches as summarized here offer a glimpse into the future opportunities and therapeutic benefits.

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Ottensmeyer, M.P. and Salisbury, J. (2001). In vivo data acquisition instrument for solid organ mechanical property measurement. Lect. Not. Comp. Sci. 2208: 975–982. Prendergast, P.J., Van Der Helm, F.C.T. and Duda, G.N. (2005). Analysis of muscle and joint loads. In: Mow and Huiskes (eds.), Basic Orthopaedic Biomechanics and Mechano-Biology, 3rd edn. New York: Lippincott, Williams & Wilkins, pp. 29–89. Satcher, R.L. and Dewey, C.F. (1996). Theoretical estimates of mechanical properties of the endothelial cell cytoskeleton. Biophys. J. 71: 109–118. Sawada, Y. and Sheetz, M.P. (2002). Force transduction by Triton cytoskeletons. J. Cell Biol. 156: 609–615. Shukla, A., Dunn, A.R., Moses, M.A. and Van Vliet, K.J. (2004). Endothelial cells as mechanical transducers: enzymatic activity and network formation under cyclic strain. Mol. Cell Biomech. 1: 279–290. Shyy, J.Y. and Chien, S. (1997). Role of integrins in cellular responses to mechanical stress and adhesion. Curr. Opin. Cell Biol. 9: 707–713. Shyy, Y.-J. and Chien, S. (2002). Role of integrins in endothelial mechanosensing of shear stress. Circ. Res. 91: 769–775. Sikavitsas, V.I., Bancroft, G.N. and Mikos, A.G. (2001). Formation of three-dimensional cell/polymer constructs for bone tissue engineering in a spinner flask and rotating wall vessel bioreactor. J. Biomed. Mater. Res. 62: 136–148. Stephens, L.E., Sutherland, A.E., Klimanskaya, I.V., Andrieux, A., Meneses, J., Pedersen, R.A. and Damsky, C.H. (1995). Deletion of beta 1 integrins in mice results in inner cell mass failure and peri-implantation lethality. Genes Dev. 9: 1883–1895. Sucosky, P., Osorio, D., Brown, J. and Neitzel, G. (2004). Fluid mechanics of a spinner-flask bioreactor. Biotechnol. Bioeng. 85(1): 34–46. Turner, C.H. and Pavalko, F.M. (1998). Mechanotransduction and function response of the skeleton to physical stress: the mechanisms and mechanics of bone adaptation. J. Orthop. Sci. 3: 346–355. Vunjak-Novakovic, G., Altman, G., Horan, R. and Kaplan, D.L. (2004). Tissue engineering of ligaments. Annu. Rev. Biomed. Eng. 6: 131–156. Vunjak-Novakovic, G., Meinel, L., Altman, G. and Kaplan, D.L. (2005). Bioreactor cultivation of osteochondral grafts. Orthod. Craniofac. Res. 8: 209–218. Weaver, V.M., Lelievre, S., Lakins, J.N. Chrenek, M.A., Jones, J., Giancotti, F., Werb, Z. and Bissell, M.J. (2002). B4 integrindependent formation of polarized three-dimensional architecture confers resistance to apoptosis in normal and malignant mammary epithelium. Cancer Cell 2: 205–216. Wolff, J. (1870). Uber der innere Architektur der Knochen und ihre Bedeutung fur die Frage vom Knochenwachstum. Arch. Pathol. Anat. Physiol. Klin. Med. 50: 389–453.

28 Morphogenesis and Morphogenetic Proteins A.H. Reddi

INTRODUCTION Morphogenesis is the developmental cascade of pattern formation, establishment of body plan and the architecture of mirror-image bilateral symmetry of musculoskeletal structures culminating in the adult form. Regenerative medicine is the emerging discipline of the science of design and manufacture of spare parts for the human body including the skeleton to restore function of lost parts due to cancer diseases and trauma. Regenerative medicine and surgery are based on rational principles of molecular developmental biology and morphogenesis and is further governed by principles of bioengineering and biomechanics. The three key elements for regenerative medicine and surgery are inductive morphogenetic signals, responding stem cells, and the extracellular matrix (ECM) scaffolding (Reddi, 1998). Recent advances in molecular cell biology of morphogens will aid in the design principles and architecture for regenerative medicine and surgery. Regeneration recapitulates in part embryonic development and morphogenesis. Among many tissues in the human body, bone has considerable powers for regeneration and therefore is a prototype model for tissue engineering. On the other hand, cartilage is feeble in its prowess for regeneration (Figure 28.1). Implantation of demineralized bone matrix into subcutaneous sites results in local bone induction. The sequential cascade of bone morphogenesis mimics sequential skeletal morphogenesis in limbs and permitted the isolation of bone morphogens. Although it is traditional to study morphogenetic signals in embryos, bone morphogenetic proteins (BMPs), the primordial inductive signals for bone were isolated from demineralized bone matrix from adults. BMPs initiate, promote, and maintain chondrogenesis and osteogenesis and have actions beyond bone. The cartilage-derived morphogenetic proteins (CDMPs) are critical for cartilage and joint morphogenesis. The symbiosis of bone inductive and conductive strategies is critical for regenerative medicine, and is in turn governed by the context and biomechanics. The context in bone is the microenvironment, consisting of ECM scaffolding and can be duplicated by biomimetic biomaterials such as collagens, hydroxyapatite, proteoglycans, and cell adhesion proteins including fibronectins and laminins. The rules of architecture for regenerative medicine and surgery are an imitation and adaptation of the laws of developmental biology and morphogenesis, and thus may be universal for all tissues, including musculoskeletal tissues and a variety of other tissues in the human body. The traditional approach for identification and isolation of morphogens is to first identify genes in fly and frog embryos by genetic approaches, differential displays, substractive hybridization, and expression cloning (Figure 28.2). This information is subsequently extended to mice and men. An alternative approach is to isolate morphogens from bone with known regenerative potential. The principles gleaned from bone morphogenesis and BMPs can be extended to regeneration of bone and cartilage and other tissues.

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Figure 28.1 The spectrum of regeneration potential of musculoskeletal tissues. Bone has the highest and cartilage the lowest. Tissues with intermediate regenerative potential are muscle, tendons, and ligaments.

Figure 28.2 The various approaches to isolation of morphogens. BMPS Bone grafts have been used by orthopedic surgeons for nearly a century to aid in the recalcitrant bone repair. Decalcified bone implants have been used to treat patients with osteomyelitis (Senn, 1989). It was hypothesized that bone might contain a substance osteogenin that initiates bone growth (Lacroix, 1945). Urist made the key discovery that demineralized, lyophilized, segments of rabbit bone when implanted intramuscularly induced new bone formation (Urist, 1965). Bone induction is a sequential multi-step cascade (Reddi and Huggins, 1972; Reddi and Anderson, 1976; Reddi, 1981). The key steps in this cascade are chemotaxis, mitosis, and differentiation. Chemotaxis is the directed migration of cells in response to a chemical gradient of signals released from the insoluble demineralized bone matrix. The demineralized bone matrix is predominantly composed of type I insoluble collagen and it binds plasma fibronectin (Weiss and Reddi, 1980). Fibronectin has domains for binding to collagen, fibrin, and heparin. The responding mesenchymal cells attached to the collagenous matrix and proliferated as indicated by [3H]thymidine autoradiography and incorporation into acid-precipitable DNA on day 3 (Rath and Reddi, 1979). Chondroblast differentiation was evident on day 5, chondrocytes on days 7 and 8, and cartilage hypertrophy on day 9 (Figure 28.1). There was concomitant vascular invasion on day 9 with osteoblast differentiation. On days 10–12 alkaline phosphatase was maximal. Osteocalcin, bone γ-carboxyglutamic acid containing gla protein (BGP), increased on day 28. Hematopoietic marrow differentiated in the ossicle and was maximal by day 21. This entire sequential bone development cascade is reminiscent of bone and cartilage morphogenesis in the limb bud (Reddi, 1981; Reddi, 1984). Hence, it has immense implications for isolation of inductive signals initiating cartilage and bone morphogenesis. In fact, a systematic investigation of the chemical components responsible for bone induction from the demineralized bone matrix was undertaken. The foregoing account of the demineralized bone matrix-induced bone morphogenesis in extraskeletal sites demonstrated the potential role of morphogens in the ECM. A systematic study of the isolation of putative morphogens from the bone matrix was initiated. A prerequisite for any quest for novel morphogens is the establishment of a battery of bioassays for new bone formation. The three key steps in bone morphogenesis are chemotaxis of progenitor stem cells, mitosis, and differentiation (Figure 28.3). A panel of in vitro assays

499

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were established for chemotaxis, mitogenesis, and chondrogenesis, and an in vivo bioassay for bone formation. Although the in vitro assays are expedient, we monitored routinely a labor-intensive in vivo bioassay as it is the only valid bona fide bone induction assay. A major stumbling block in the approach was that the demineralized bone matrix is insoluble and in the solid state. In view of this dissociative extractants such as 4 M guanidine HCl or 8 M urea as 1% sodium dodecyl sulfate (SDS) at pH 7.4 were used (Sampath and Reddi, 1981) to solubilize proteins. Approximately 3% of the proteins were solubilized from demineralized bone matrix, and the remaining residue was mainly insoluble type I bone collagen. The extract alone or the residue alone was incapable of new bone induction. However, addition of the extract to the residue (insoluble collagen) and then implantation in a subcutaneous site resulted in bone induction (Figure 28.4). Therefore, for optimal osteogenic activity it is essential to have a collaboration between soluble signal in the extract and the insoluble substratum of collagenous ECM (Sampath and Reddi, 1981). This bioassay was a critical advance in the ultimate purification of BMPs and led to determination of limited tryptic peptide sequences leading to the eventual cloning of BMPs (Wozney et al., 1988; Luyten et al., 1989; Ozkaynak et al., 1990). The dissociative extraction of soluble signals from the demineralized ECM of bone and its subsequent reconstitution with collagen established the cardinal principle of regenerative medicine. The key principle is that morphogenetic signals stimulate the stem cells to differentiate in the optimal scaffold microenvironment (Figure 28.5). Thus, the triumvirate of signals, stem cells, and scaffolds for regenerative medicine was conceived as a concept. Although the basic description of bone induction was performed in rats, purification requires a larger and more abundant source of bone. A switch was made to bovine bone. Demineralized bovine bone matrix was not osteoinductive in rats and the results were variable. However, when the guanidine extracts of demineralized

Three key steps in bone morphogenesis • Chemotaxis • Mitosis • Differentiation

Figure 28.3 The three key steps in bone morphogenesis.

Dissociative extraction and reconstitution DBM

Activity 

4 M Guanidine



Collagen



Extract

 

Figure 28.4 Dissociative extraction of bone matrix by chaotropic reagents such as 4 M guanidine hydrochloride, and reconstitution of extract with collagenous matrix scaffold. The results indicate that there is a collaboration between soluble signal in the extract and the insoluble ECM of bone.

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bovine bone were fractionated on a S-200 molecular sieve column, fractions less than 50 kD were consistently osteogenic in rats when bioassayed after reconstitution with allogeneic insoluble collagen (Sampath and Reddi, 1983; Reddi, 1994). Thus, protein fractions inducing bone were not species specific and appear to be homologous in several mammals. It is likely that larger molecular mass fractions and/or the insoluble xenogeneic (bovine and human) collagens were inhibitory or immunogenic. Initial estimates revealed 1 μg of active osteogenic fraction in a kilogram of bone. Hence, over a ton of bovine bone was processed to yield optimal amounts for animo acid sequence determination. The amino acid sequences revealed homology to transforming growth factor (TGF)-β1 (Reddi, 1994). The decisive work of Wozney et al. (1988) cloned BMP-2, BMP-2B (now called BMP-4), and BMP-3 (also called osteogenin). Ozkaynak et al. (1990) cloned osteogenic proteins 1 and 2 (OP 1 and OP 2). There are several members of this BMP family (Figure 28.6). The other members of the extended TGF-β/BMP superfamily include inhibins and activins (implicated in follicle stimulating hormone release from pituitary). Müllerian duct inhibitory substance (MIS), growth/differentiation factors (GDFs), nodal, and lefty genes implicated in establishing right/left asymmetry (Cunningham et al., 1995,

Bone morphogenesis and regenerative medicine Signal  Scaffolding

Bone

 Stem cells

Figure 28.5 The key principle of regenerative medicine is that signals stimulate differentiation of stem cells in the appropriate scaffold. BMP family BMP-5 BMP-6 BMP-7/OP-1 BMP-8a/OP-2 BMP-8b/OP-3 BMP-2 BMP-4 BMP-14/CDMP-1/GDF-5 BMP-13/CDMP-2/GDF-6 BMP-12/CDMP-3/GDF-7 BMP-10 BMP-3/osteogenin BMP-3b/GDF-10 GDF-1 GDF-3 GDF-9 BMP-15/GDF-9b GDF-8 BMP-11

Figure 28.6 Members of the BMP family include three main subfamilies: BMP 5, 6, and 7; BMP 2 and 4; BMP 3 and 3b; and GDF 5, 6, and 7.

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Reddi, 1997; Reddi, 1998. BMPs are also involved in embryonic induction (Melton, 1991; Lemaire and Gurdon, 1994; Lyons et al., 1995; Reddi, 1997). BMPs are dimeric molecules and the conformation is critical for biological actions. Reduction of the single interchain disulfide bond resulted in the loss of biological activity. The mature monomer molecule consists of about 120 amino acids, with 7 canonical cysteine residues. There are three intrachain disulfides per monomer and one interchain disulfide bond in the dimer. In the critical core of the BMP monomer is the cysteine knot. The crystal structure of BMP-7 has been determined (Griffith et al., 1996). Morphogenesis is a sequential multi-step cascade. BMPs regulate each of the key steps: chemotaxis, mitosis, and differentiation of cartilage and bone. BMPs initiate chondrogenesis in the limb (Chen et al., 1991; Duboule 1994). The apical ectodermal ridge is the source of BMPs in the developing limb bud. The intricate dynamic, reciprocal interactions between the ectodermally derived epithelium and mesoderm-derived mesenchyme sets into motion the train of events culminating in the pattern of phalanges, radius, ulna and the humerus. The chemotaxis of human monocytes is optimal at femtomolar concentration (Cunningham et al., 1992). The apparent affinity was 100–200 pM. The mitogenic response was optimal at 100 pM range. The initiation of differentiation was in nanomolar range in solution. However, caution should be exercised as BMPs may be sequestered by ECM components and the local concentration may be higher when BMPs are bounded on the ECM. Thus BMPs are pleiotropic regulators that act in concentration-dependent thresholds. It is well known that ECM components play a critical role in morphogenesis. The structural macromolecules and their supramolecular assembly in the matrix do not explain their role in epithelial–mesenchymal interaction and morphogenesis. This riddle can now be explained by the binding of BMPs to heparan sulfate heparin, and type IV collagen (Paralkar et al., 1990, 1991, 1992) of the basement membranes. In fact, this might explain in part the necessity for angiogenesis prior to osteogenesis during development. In addition, the actions of activin in development of the frog, in terms of dorsal mesoderm induction, are modified to neuralization by follistatin (Hemmati-Brivanlou et al., 1994). Similarly, Chordin and Noggin from the Spemann organizer induces neuralization by binding and inactivation of BMP-4. Thus neural induction is likely to be a default pathway when BMP-4 is non-functional (Piccolo et al., 1996; Zimmerman et al., 1996). Thus, this is an emerging principle in development and morphogenesis that binding proteins can terminate a dominant morphogen’s action and initiate a default pathway. Finally, the binding of a soluble morphogen to ECM converts it into an insoluble matrix bound morphogen to act locally in the solid state (Paralkar et al., 1990). Although BMPs were isolated and cloned from bone, recent work with gene knockouts has revealed a plethora of actions beyond bone. Mice with targeted disruption of BMP-2 caused embryonic lethality. The heart development is abnormal indicating a need for BMP-2 in heart development (Zhang and Bradley, 1996). BMP-4 “knockouts” exhibit no mesoderm induction, and gastrulation is impaired (Winnier et al., 1996). Transgenic overexpression of BMP-s under the control of keratin 10 promoter leads to psoriasis. The targeted deletion of BMP-7 revealed the critical role of this molecule in kidney and eye development (Dudley et al., 1995; Luo et al., 1995; Vukicevic et al., 1996). Thus the BMPs are really true morphogens for such disparate tissues as skin, heart, kidney, and eye. In view of the emerging wider role, BMPs may be called body morphogenetic proteins (BMPs). Recombinant human BMP-4 and BMP-7 bind to BMP receptor IA (BMPR-IA) and BMP receptor IB (BMPR-IB) (ten Dijke et al., 1994). CDMP-1 also binds to both the type I BMP receptors. There is a collaboration between type I and II BMP receptors (Nishitoh et al., 1996). The type I receptor serine/threonine kinase phosphorylates a signal-transducing protein substrate called Smad 1 or 5 (Chen et al., 1996). Smad is a term derived from fusion of Drosophila Mad gene and Caenorhabtitis elegans (nematode) Sma gene. Smads 1 and 5 signal in partnership with a common co-Smad, Smad 4 (Figure 28.7). The transcription of BMP-response genes are initiated by Smad 1/Smad 4 heterodimers. Smads are trimeric molecules as gleaned by X-ray crystallography. The phosphorylation of Smads 1 and 5 by type I BMP receptor kinase is inhibited by inhibitory

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BMPs Noggin chordin dan

Extracellular matrix collagens I & IV heparan sulfate BMPR-1A

Cytoplasm

BMPR-1B P

P

SMAD-6

P P

SMAD-7 BMPR-II SMAD-5

SMAD-1 P

P

SMAD-1

SMAD-5

ⴙ

ⴙ

SMAD-4

SMAD-4

Nucleus P

P

SMAD-1

SMAD-5

SMAD-4

SMAD-4

SMAD-6 SMAD-7

BMP response genes

Figure 28.7 BMP receptors and signaling cascades. BMPs are dimeric ligands with cysteine knot in each monomer fold. Each monomer has two β sheets represented as two pointed fingers. In the functional dimer the fingers are oriented in opposite directions. BMPs interact with both type I and II BMP receptors. The exact stoichiometry of the receptor complex is currently being elucidated. BMPR-II phosphorylates the GS domain of BMPR-I. The collaboration between type I and II receptors forms the signal-transducing complex. BMP type I receptor kinase complex phosphorylates the trimeric signaling substrates Smad 1 or Smad 5. This phosphorylation is inhibited and modulated by inhibitory Smads 6 and 7. Phosphorylated Smad 1 or 5 interacts with Smad 4 (functional partner) and enters the nucleus to activate the transcriptional machinery for early BMP-response genes. A novel SIP may interact and modulate the binding of heteromeric Smad 1/Smad 4 complexes to the DNA.

Smads 6 and 7 (Hayashi et al., 1997). Smad interacting protein (SIP) may interact with Smad 1 and modulate BMP-response gene expression (Heldin et al., 1997; Reddi, 1997). The downstream targets of BMP signaling are likely to be homeobox genes, the cardinal genes for morphogenesis and transcription. BMPs in turn may be regulated by members of the hedgehog family of genes such as Sonic and Indian hedgehog (Johnson and Tabin, 1997).

STEM CELLS It is well known that the embryonic mesoderm-derived mesenchymal cells are progenitors for bone, cartilage, tendons, ligaments, and muscle. However, certain stem cells in adult bone marrow, muscle, and fascia can form bone and cartilage (Figure 28.8). The identification of stem cells readily sourced from bone marrow may

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Figure 28.8 The lineages of the putative musculoskeletal stem cell. The BMPs determine the lineage into chondro/osteo progenitor cells and further specialization into articular chondrocytes growth plate chondrocytes and osteoblast lineage. BMPs are critical morphogens to direct the differentiation of cartilage and bone cells.

lead to banks of stem cells for cell therapy and perhaps gene therapy with appropriate “homing” characteristics to bone marrow and hence to the skeleton. The pioneering work of Friedenstein et al. (1968, 1987), and Owen and Friedenstein (1988) identified bone marrow stromal stem cells. These stromal cells are distinct from the hematopoietic stem cell lineage. The bone marrow stromal stem cells consist of inducible and determined osteoprogenitors committed to osteogenesis. Determined osteogenic precursor cells have the propensity to form bone cells without any external cues or signals. On the other hand inducible osteogenic precursors require an inductive signal such as BMP or demineralized bone matrix. It is noteworthy that operational distinction between stromal stem cells and hematopoietic stem cells are getting more and more blurry!

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The stromal stem cells of Friedenstein and Owen are also called mesenchymal stem cells (Caplan, 1991; Pittenger et al., 1999), with potential to form bone, cartilage, adipocytes, and myoblasts in response to cues from environment and/or intrinsic factors. Mesenechymal stem cells are present in synovium (De Bari et al., 2001), periosteum (Nakahara et al., 1991), adipose tissue (Zuk et al., 2001), and blood (Zvaifler et al., 2000). There is very recently considerable hope and anticipation that these bone marrow stromal cells may be excellent vehicles for cell and gene therapy (Prockop, 1997; Kuznetsov et al., 1997). From a practical standpoint these stromal stem cells can be obtained by bone marrow biopsies and expanded rapidly for use in cell therapy after pre-treatment with BMPs. The potential uses in both cell and gene therapy is very promising. There are continuous improvements in the viral vectors and efficiency of gene therapy (Mulligan, 1993; Kozarsky and Wilson, 1993; Bank, 1996; Morsy et al. (1993)). For example, it is possible to use BMP genes transfected in stromal stem cells to target to the bone marrow.

SCAFFOLDS OF BIOMIMETIC BIOMATERIALS The earlier discussion of inductive signals (BMPs) responding stem cells (stromal cells) leads us to the scaffolding (the microenvironment/ECM) for optimal tissue engineering. The natural biomaterials in the composite tissue of bones and joints are collagens, proteoglycans, and glycoproteins of cell adhesion such as fibronectin and the mineral phase. The mineral phase in bone is predominantly hydroxyapatite. In native state the associated citrate, fluoride, carbonate, and trace elements constitutes the physiological hydroxyapatite. The high protein binding capacity makes hydroxyapatite a natural delivery system. Comparison of insoluble collagen, hydroxyapatite, tricalcium phosphate, glass beads, and polymethylmethacrylate as carriers revealed collagen to be an optimal delivery system for BMPs (Ma et al., 1990). It is well known that collagen is an ideal delivery system for growth factors in soft and hard tissue wound repair (McPherson, 1992). During the course of systematic work on hydroxyapatite of two pore sizes (200 or 500 μm) in two geometrical forms (beads or disks) an unexpected observation was made. The geometry of the delivery system is critical for optimal bone induction. The disks were consistently osteoinductive with BMPs in rats; but the beads were inactive (Ripamonti et al., 1992). The chemical composition of the two hydroxyapatite configurations was identical. In certain species the hydroxyapatite alone appears to be “osteoinductive” (Ripamonti, 1996). In subhuman primates the hydroxyapatite induces bone albeit at a much slower rate. One interpretation is that osteoinductive endogenous BMPs in circulation progressively bind to implanted disk of hydroxyapatite. When an optimal threshold concentration of native BMPs is achieved the hydroxyapatite becomes osteoinductive. Strictly speaking most hydroxyapatite substrata are ideal osteoconductive materials. This example in certain species also serves to illustrate how an osteoconductive biomimetic biomaterial may progressively function as an osteoinductive substance by binding to endogenous BMPs. Thus, there is a physiological–physicochemical continuum between the hydroxyapatite alone and progressive composites with endogenous BMPs. Recognition of this experimental nuance will save unnecessary arguments amongst biomaterials scientists about the osteoinductive action of a conductive substratum such as hydroxyapatite. Complete regeneration of baboon craniotomy defect was accomplished by recombinant human osteogenic protein (rhOP-1; human BMP-7) (Ripamonti et al., 1996). Recombinant BMP-2 was delivered by poly(-hydroxy acid) carrier for calvarial regeneration (Hollinger et al., 1996). Copolymer of polylactic and polyglycolic acid in a non-union model in rabbit ulna and the results were satisfactory (Figure 28.9) (Bostrom et al., 1996). An important problem in the clinical application of biomimetic biomaterials with BMPs and/or other morphogens in regenerative medicine is the sterilization. Although gas (ethylene oxide) is used, one always should be concerned about reactive free radicals. Using allogeneic demineralized bone matrix with endogenous native

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BMPS and tissue regeneration • • • • • • • •

Orthopaedics Fractures Spine/fusions Articular cartilage repair Dentistry/oral surgery Periodontal surgery Craniofacial surgery Plastic surgery

Figure 28.9 BMPs have wide ranging roles in regenerative medicine and surgery. The applications include but are not limited to orthopedics, plastic and reconstructive surgery, in dentistry and oral surgery. Recombinant BMP 2 has been approved by the FDA for spine fusions and non-unions of fractures. BMPs, as long as low temperature (4°C or less) is maintained, the samples tolerated up to 5–7 M rads of irradiation (Weintroub and Reddi, 1988; Weintroub et al., 1990). The standard dose acceptable to the Food and Drug Administration (FDA) is 2.5 M rads. This information would be useful to the biotechnology companies preparing to market recombinant BMP-based osteogenic devices. Perhaps, tissue banking industry with interest in bone grafts (Damien and Parson, 1991) could also use this critical information. The various freeze-dried and demineralized allogeneic bone may be used in the interim as satisfactory carriers for BMPs. The moral of this experiment is it is not the irradiation dose but the ambient sample temperature during irradiation is absolutely critical.

CARTILAGE-DERIVED MORPHOGENETIC PROTEINS Morphogenesis of the cartilage is the key rate-limiting step in the dynamics of bone development. Cartilage is the initial model for the architecture of bones. Bone can form either directly from mesenchyme as in intramembranous bone formation or with an intervening cartilage stage as in endochondral bone development (Reddi, 1981). All BMPs induce, first, the cascade of chondrogenesis, and therefore they all sense are cartilage morphogenetic proteins. The hypertrophic chondrocytes in the epiphyseal growth plate mineralize and serves as a template for appositional bone morphogenesis. Cartilage morphogenesis is critical for both bone and joint morphogenesis. The two lineages of cartilage are clear-cut. The first at the ends of bone, forms articulating articular cartilage. The second is the growth plate chondrocytes which hypertrophy synthesize cartilage matrix destined to calcify prior to replacement by bone and are the “organizer” centers of longitudinal and circumferental growth of cartilage setting into motion the orderly program of endochondral bone formation. The phenotypic stability of the articular (permanent) cartilage is at the crux of the osteoarthritis problem. The “maintenance” factors for articular chondrocytes include TGF-β isoforms and the BMP isoforms (Luyten et al., 1992). An in vivo chondrogenic bioassay with soluble purified proteins and insoluble collagen scored for chondrogenesis. A concurrent reverse transcription-polymerase chain reaction (RT-PCR) approach was taken with degenerate oligonucleotide primers. Two novel genes for CDMPs 1 and 2 were identified and cloned (Chang et al., 1994). CDMPs 1 and 2 are also called GDF-5 and GDF-6, respectively (Storm et al., 1994). CDMPs are related to BMPs (Figure 28.6). CDMPs are critical for cartilage and joint morphogenesis (Tsumaki et al., 1999). CDMPs stimulate proteoglycan synthesis in cartilage. GDF-7 initiates tendon and ligament morphogenesis. REGENERATIVE MEDICINE AND SURGERY Unlike bone with its considerable prowess for repair and even regeneration, cartilage is recalcitrant. This part may be due to relative a vascularity of hyaline cartilage, the high concentration of protease inhibitors and

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perhaps even cytokine inhibitors. The wound debridement phase is not optimal to prepare the cartilage wound bed for the optimal regeneration. Although cartilage has been successfully engineered to predetermined shapes (Kim et al., 1994), true repair of the tissue continues to be a real challenge in part due to hierarchical organization and geometry (Mow et al., 1992). The utility of autologous culture-expanded human chondrocytes is gaining (Brittberg et al., 1994). Also gaining increasing attention is mosaicplasty for defects in articular cartilage (Hangody et al., 2001). A continuous challenge in chondrocyte cell therapy is progressive de-differentiation and loss of characteristic cartilage phenotype. The re-differentiation and maintenance of the chondrocytes for cell therapy can be aided by BMPs, CDMPs, TGF-β isoforms, and insulin growth factors (IGFs). It is also possible to repair cartilage using muscle-derived mesenchymal stem cells (Grande et al., 1995). The potential possibility of the problems posed by cartilage proteoglycans in preventing cell immigration for repair was investigated by chondroitinase ABC and trypsin pre-treatment in partial-thickness defects (Hunzinker and Rosenberg, 1996), with and without TGF-β. Pre-treatment with chondroitinase ABC followed by TGF-β revealed a contiguous layer of cells from the synovial membrane hinting at the potential source of “repair” cells from synovium. Multiple avenues of cartilage morphogens, cell therapy with chondrocytes and stem cells from marrow and muscle and a biomaterial scaffolding may lead to an optimal tissue engineered articular cartilage. It is inevitable during aging most humans will confront the challenges of impaired locomotion due to wear and tear in bones and joints. Therefore, the repair and possibly complete regeneration of the musculoskeletal system and other vital organs such as skin, liver, and kidney may potentially need optimal repair or a spare part for replacement. Can we create spare parts for the human body? There is much reason for optimism that tissue engineering can help patients. We are living at an extraordinary time in the biology, medicine, surgery, and computational and related technology. The confluence of advances in molecular developmental biology and attendant advances in inductive signals for morphogenesis, stem cells, and biomimetic biomaterials. The symbiosis of biotechnology and biomaterials has set the stage for systematic advances in tissue engineering (Langer and Vacanti, 1993; Reddi, 1994; Hubbell, 1995). The recent advances in enabling platform technology include molecular imprinting (Mosbach and Ramstrom, 1996). In principle, specific recognition and catalytic sites are imprinted using templates. The applications range from biosensors, catalytic applications to antibody, and receptor recognition sites. For example, the cell binding RGD site in fibronectin (Ruoslahti and Pierschbacher, 1987) or YIGSR domain in laminin can be imprinted in complementary sites (Vukicevic et al., 1990). The rapidly advancing frontiers in morphogenesis with BMPs, hedgehogs, homeobox genes, and a veritable cornucopia of general and specific transcription factors co-activators and repressors will lead to co-crystallization of ligand–receptor complexes, protein-DNA complexes, and other macromolecular interactions. This will lead to peptidomimetic agonists for large proteins as exemplified by erythroprotein (Livnah et al., 1996). To such advances one can add new developments in self-assembly of millimeter-scale structures floating at the interface of perfluorodecalin and water and interacting by capillary forces controlled by the pattern of wettablity (Bowden et al., 1997). The final self-assembly is due to minimization of free energy in the interface. These are truly incredible advances that will lead to man-made materials that mimic ECM in tissues. Superimpose on such chemical progress a biological platform in a bone and joint mold. Let us imagine a head of the femur and a mold is fabricated with computer-assisted design and manufacture. It faithfully reproduces the structural features and may be imprinted with morphogens, inductive signals, and cell adhesion sites. This assembly can be loaded with stem cells and BMPs and other inductive signals with a nutrient medium optimized for optimal number of cell cycles, and then predictably exit into differentiation phase to reproduce a totally new bone femoral head. In fact such a biological approach with vascularized muscle flap and BMPs yielded new bone with a defined shape and has

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demonstrated the proof of principle for further development and validation (Khouri et al., 1991). We indeed are entering a brave new world of prefabricated biological spare parts for the human body based on sound architectural rules of inductive signals for morphogenesis, responding stem cells with lineage control, and with growth factors immobilized on a template of biomimetic biomaterial based on ECM.

ACKNOWLEDGEMENTS This work is supported by the Lawrence Ellison Chair in Musculoskeletal Molecular Biology and the NIH grant AR4 7345-01 A2. I thank Ms. Danielle Neff for outstanding bibliographic assistance and enthusiastic help.

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29 Physical Stress as a Factor in Tissue Growth and Remodeling Robert E. Guldberg, Christopher S. Gemmiti, Yash Kolambkar, and Blaise Porter INTRODUCTION The role of physical stresses and strains in regulating tissue growth and remodeling has been of tremendous interest to investigators for well over 100 years. Although somewhat unfairly to his contemporary colleagues, Julius Wolff is often credited with the concept that tissue structure or form follows from its function (i.e. Wolff ’s Law). At the time, Wolff ’s Law was simply based on the general correspondence noted between anatomical observations of trabecular bone organization and estimations of principal stress directions due to functional loading conditions. The recognition that adaptation of tissue structure and composition is cell mediated was not made until later by other investigators. These early observations spawned the interdisciplinary field of mechanobiology, focused on identifying mechanisms by which mechanical signals are transduced into cellular activity, and emphasized the need to consider the effects of physical factors on tissue growth and remodeling as an important part of strategies for tissue regeneration. Many different cell types from various tissues have been shown to be sensitive to mechanical stimuli in one form or another. The effects of physiological mechanical signals on cells and tissues can be beneficial, playing a central role in the maintenance of tissue structural integrity via remodeling processes. Alterations in mechanical signals can also contribute to the development of pathological conditions. For example, local shear stresses play a key role in the development and localization of atherosclerotic lesions. Likewise, the progression of osteoarthritis is due to a vicious cycle of cartilage matrix degradation and increased local stresses. In bone, the mechanical environment also has important clinical implications in the development of osteoporosis, stress fractures, total joint implant loosening, and bone loss during space flight. Given the ability of cells to sense and respond to mechanical signals, in vitro and in vivo strategies for engineering tissues that serve a mechanical function must consider adaptational responses to physical stresses. For many tissue types, static culture conditions in vitro produce tissue-engineered constructs with deficient mechanical properties, typically due to reduced content and organization of structural protein constituents. Bioreactor systems that deliver tissue appropriate mechanical signals have been designed to overcome this limitation and exploit cellular adaptation responses to produce constructs that more closely resemble native tissue properties. Upon implantation, the interaction between constructs and the in vivo mechanical environment is a critical determinant of whether functional integration is ultimately achieved. This chapter begins by introducing the continuum concept and the idea that structural hierarchy must be considered when studying the effects of physical stresses on cells and tissues. After defining stress and strain, an overview is provided of the role of mechanical factors in tissue growth, repair, and remodeling in vivo. The

512

Structural hierarchy

Tissue level e.g. trabecular bone Microstructural level e.g. osteon

Force transmission

Adaptation

Organ level e.g. whole femur

Ultrastructural level e.g. collagen/mineral

Cellular response e.g. bone formation

Cellular level e.g. osteoblasts

Figure 29.1 Force transmission through the structural hierarchy of bone to the cellular level resulting in cell-mediated adaptation of tissue structure and composition.

fundamental mechanisms by which cells may sense and respond to mechanical signals are then reviewed. Finally, the chapter concludes by considering the application of mechanical stimuli in bioreactor systems to produce larger and stronger tissue constructs for implantation.

STRUCTURAL HIERARCHY AND THE CONTINUUM CONCEPT It is useful to view tissues as a structural hierarchy through which functional loads are transmitted down to the cellular level (Figure 29.1). In bone, for example, applied joint and muscle forces result in stresses and strains within the mineralized tissue that can be defined at different scale levels from the whole bone level down to sub-micron mineral crystals embedded within collagen molecules. At each hierarchical level, it is convenient to assume that everything below that level is a continuum (i.e. there is a finite mass density at every point within the material). This simplification allows material properties to be expressed at a given hierarchical level in terms of constitutive equations. As described in the next section, constitutive equations define the relationship between stresses and strains at each level. Cells sense and respond to local stresses or strains produced by forces transmitted from the macro level down through the complex structural hierarchy to the cellular level. Cell-mediated adaptational changes in tissue structure and composition subsequently alter the local stresses and strains resulting from functionally applied loads, thus providing a regulatory feedback mechanism. It is important to note that the sensitivity of the cellular response to mechanical stimuli can be altered by a variety of non-mechanical factors such as age, disease, as well as numerous biochemical factors.

STRAIN AND STRESS DEFINITIONS Strain Strain is a normalized measure of deformation. Consider the simple case of a thin rectangular piece of tissue being axially loaded by a force, as shown in Figure 29.2a. The axial force increases the length of the tissue, but

513

514 CELLS AND TISSUE DEVELOPMENT

2

(a)

L0

L

1 (b) c

dF

dF dF

F W0 dF W

Figure 29.2 (a) Axial and transverse strains associated with uniaxial tensile loading. (b) Shear strain associated with torsional or shear loading.

at the same time decreases its width and thickness. Engineering strain is defined as the change in a dimension of the tissue normalized by its original dimension, and is given in the axial direction by: ε11 

L  L0 L0

Another important deformation parameter is the Poisson’s ratio ν, which is defined as the ratio of lateral strain to axial strain, and is given in this case by:

ν

ε22 ε11

W  W0 W0  L  L0 L0

The Poisson ratio is a measure of the tendency for a material body to try to retain its total volume as it is deformed. When ν  0.5, the material is said to be incompressible (e.g. water), and does not undergo a volume change after deformation. The typical value of ν for tissues is between 0.2 and 0.45. Thus, a tissue subjected to tensile deformation and strain would increase in volume slightly. In contrast to normal strains, shear deformations and strains due to shear forces dF or from pure torsional loading, for example, produce a change in shape but not volume, as shown in Figure 29.2b. Measurement of the angle of shear deformation, ψ, allows calculation of shear strain, as given by: ε12 

ψ 2

The complex deformations created by forces acting in multiple directions necessitate the generalization of deformation to 3-D space. Deformation in 3-D can be expressed by the deformation gradient F. Consider the body shown in Figure 29.3a undergoing a deformation from the reference state to a deformed configuration. If one follows the particles P1 and P2, they move from position XP1 and XP2 to xP1 and xP2, respectively. There will also be a similar one-to-one mapping of other particles in the reference and deformed configurations. Thus the deformation of the body can be written as a function: x  f(X)

Physical Stress as a Factor in Tissue Growth and Remodeling 515

(a)

2 Reference configuration P1 dS P2 XP1

XP2

p1 ds p2

Xp1

Deformed configuration

Xp2

1 3 (b)

σ22 2

σ23 σ 21

B ΔF

σ32

σ31

σ12

σ11

σ13

σ33

ΔA S

1 3

Figure 29.3 (a) Deformation of a 3-D body from a reference configuration to a deformed configuration. (b) Stress on a surface element, and the nine stress components defining the stress state at a point.

In scalar form, this would involve three equations: x1  f1(X1, X2, X3) x2  f2(X1, X2, X3) x3  f3(X1, X2, X3) where 1, 2, and 3 correspond to the three directions of the coordinate system. The displacement vector is given by: uxX The deformation gradient F is then defined as: F

∂x ∂X

516 CELLS AND TISSUE DEVELOPMENT

In matrix form, the deformation gradient can be written as: ⎡ ∂x ⎢ 1 ⎢ ∂X ⎢ 1 ⎢ ∂x F⎢ 2 ⎢ ∂X1 ⎢ ∂x ⎢ 3 ⎢ ∂X ⎢⎣ 1

∂x1 ∂X 2 ∂x 2 ∂X 2 ∂x 3 ∂X 2

∂x1 ⎤⎥ ∂X 3 ⎥⎥ ∂x 2 ⎥ ⎥ ∂X 3 ⎥ ∂x 3 ⎥⎥ ∂X 3 ⎥⎦⎥

and is related to the gradient of displacement by the following expression in which I is the unit vector: F

∂u I ∂X

The engineering strains as defined above are appropriate to use when the strains in the material are small (typically less than 5%). However the analysis of large deformations, as frequently observed for soft tissues under functional loading conditions, requires use of other strain measures. Consider the segment P1P2 of length dS that has deformed to p1p2 with length ds. When the deformation is large, a useful measure of deformation is the Green (i.e. Lagrangian) strain (E), which is defined as: 1 ⎛ ds 2  dS 2 ⎞⎟ ⎟ E  ⎜⎜ 2 ⎜⎝ dS 2 ⎟⎟⎠ The Green strain in the body can be expressed in terms of the gradient of displacement as: E

1⎡ D  D T  D T D ⎤⎥⎦ , 2 ⎢⎣

⎡ ∂u ⎤ ⎥ and the superscript T stands for the transpose of the matrix form of the second-order tensor. where D  ⎢ ⎢⎣ ∂X ⎥⎦ If the deformation under consideration is small, as is typically the case for bone and most structural-engineering materials, the quadratic term in the Green strain can be neglected to give the infinitesimal (engineering) strain tensor (ε): ε

1⎡ D  D T ⎤⎦⎥ 2 ⎣⎢

This is what gives us the familiar expression of engineering strain in a uniaxial test: ε

L  L0 . L0

To get a feel for the relative values of these strain measures, consider the following example of uniaxial elongation of our rectangular tissue having original length of 5 cm. In one case, the tissue is stretched to a final length of 5.05 cm (small strain), whereas in the second case, it is elongated to 10 cm (large strain).

Physical Stress as a Factor in Tissue Growth and Remodeling 517

Case I (L  5.05 cm)

Case II (L  10 cm)

⎛ L2  L20 ⎞⎟⎟ Green strain ⎜⎜⎜E  1 ⎟ ⎜⎝ 2 L20 ⎟⎟⎠

0.01005

1.5

⎛ L  L0 ⎟⎞⎟ Engineering strain ⎜⎜⎜ ε  ⎟ ⎜⎝ L0 ⎟⎟⎠

0.01

1.0

Thus, we see that for the small deformations, the different strain definitions give approximately the same value and engineering strain is reasonably accurate. Whereas for large deformations, the strain definitions yield very different values due to neglect of the higher-order terms in the engineering strain definition. Stress Stress is a measure of the intensity of internal force developed in a material upon application of an external force. Consider the force ΔF acting on a small surface element of area ΔA in Figure 29.3b. This element lies ΔF on the surface S, which is part of the larger body B. As ΔA tends to zero, the ratio tends to a finite limit ΔA dF , which is defined as the stress on the surface element. dA Consider an infinitesimal cube in the body as shown in Figure 29.3b. Due to the external force applied on the body, internal forces are applied on the surface of the cube. Each internal force can be resolved into its three components and normalized by the area to give three stress components on each face. The volume of the cube can be continuously decreased such that the cube collapses to a point. The nine stress components define the second-order stress tensor, and completely describe the stress state at this point. Using equilibrium conditions, we can show that σij  σji; thus the stress tensor has only six independent components. If a stress component acts in a direction perpendicular to the surface it acts on, it is referred to as a normal stress. On the other hand, if it is parallel to the surface, it is called a shear stress. Thus σ11, σ22, and σ33 are normal stresses, while σ12, σ23, and σ31 are shear stresses. Normal stresses tend to change the volume of the body, while shear stresses tend to modify the shape. If the body is in the original reference configuration, ΔA represents the undeformed area and the stress is called the first Piola–Kirchoff stress tensor (T). In a typical experiment, the force is constantly measured, but the cross-sectional area is not. Thus the first Piola–Kirchoff stress is an easy quantity to compute as the undeformed cross-sectional area can be measured prior to loading. However, while considering force balance in the deformed body at equilibrium after external force is applied, the deformed area Δa of the surface element is required for ΔF the stress definition. The Cauchy stress is thus defined as the limit of σ  as Δa tends to zero. The difference Δa between Δa and ΔA is negligible for small deformations. For large deformations, however, the stress definition again makes a significant difference. Constitutive Equations A constitutive equation is a mathematical model that specifies the relationship between stress and strain. Typically the model is phenomenological in nature, and is not derived from the microstructure. For example, the simplest constitutive equation is that for the linearly elastic materials, where there is a linear relationship

518 CELLS AND TISSUE DEVELOPMENT

between stress and strain. Most engineering materials and stiff biomaterials like bone can be treated this way. These materials follow Hooke’s law, which can be written for the general 3-D case in indicial notation as: σij  Cijkl εkl, where Cijkl is a fourth-order tensor describing the material properties, and contains 81 constants. However, due to symmetry arguments (including symmetry of stress and strain), the number of independent constants is reduced to 21. If the stress and strain are written in the form of a column matrix, the material tensor can be represented by a matrix called the stiffness matrix: ⎡ C11 C12 ⎢ C22 ⎢ ⎢ C⎢ ⎢ ⎢ ⎢ ⎢⎣

C13 C14 C23 C24 C33 C34 C 44

C15 C16 ⎤ ⎥ C25 C26 ⎥ C355 C36 ⎥⎥ C 45 C 46 ⎥ C55 C56 ⎥⎥ C66 ⎥⎦

where the other side of the diagonal is symmetric (i.e. Cij  Cji). The above stiffness matrix represents a fully anisotropic linear elastic material, for which 21 constants must be determined experimentally to fully characterize the material behavior. Fortunately most materials, including tissues, have some degree of material symmetry. For example, trabecular bone has been frequently described using an orthotropic material model, which consists of three mutually perpendicular planes of symmetry that coincide with the chosen reference coordinate system. This reduces the numbers of independent constants to 9, which are related to the Young’s (Y) and shear moduli (G) and the Poisson’s ratio (ν) in the three planes giving: ⎡ 1  23 32 ⎢ ⎢ ΔY Y 2 3 ⎢ ⎢ ⎢ ⎢ ⎢ ⎢ ⎢ C⎢ ⎢ ⎢ ⎢ ⎢ ⎢ ⎢ ⎢ ⎢ ⎢ ⎢ ⎣

21  31 23 ΔY2Y3 1  13 31 ΔY1Y3

31  21 32 ΔY2Y3

32  12 31 ΔY1Y3 1  12 21 ΔY1Y2

0

0

0

0

0

0

2G23

0 2G31

⎤ 0 ⎥ ⎥ ⎥ ⎥ 0 ⎥ ⎥ ⎥ 0 ⎥⎥ ⎥ ⎥ 0 ⎥ ⎥ ⎥ 0 ⎥⎥ ⎥ ⎥ 2G12 ⎥ ⎥ ⎦

The simplest case of material symmetry is the isotropic material, in which all planes are planes of symmetry, that is, the material properties are independent of direction. This material has only two independent constants, a Young’s modulus and a Poisson’s ratio (or shear modulus), that are valid for all directions. Note that in the isotropic case, the shear modulus, Young’s modulus, and the Poisson’s ratio are related and therefore only two of them are independent. Finally, for a uniaxial loading test on an isotropic and linearly elastic material we have the familiar 1-D version of Hooke’s law: σ  Yε. Thus it can be seen that for a linearly elastic material, stress is linearly related to strain. However many soft tissues, especially at large deformations, display non-linearity in the stress–strain relationship. Furthermore, most biological materials display time-dependent behavior, a property known as viscoelasticity. If a constant

Physical Stress as a Factor in Tissue Growth and Remodeling 519

stress is applied to a viscoelastic material, it continues to deform with time (i.e. creep). Alternatively, if the material is subjected to a constant strain, the stresses in the material decrease with time (i.e. stress relaxation). Thus, for a viscoelastic material, the constitutive equation includes the rate of change of stress and strain over time. Textbooks by Fung (1965, 1993) are an excellent resource for additional information on tissue material behavior that deviates from linear elasticity.

TISSUE GROWTH, REPAIR, AND REMODELING The composition and structure of tissues continually change in response to biochemical and biomechanical demands in vivo. While dramatic changes occur during early tissue morphogenesis and growth, alterations in tissue structure and composition may also occur in adulthood via repair or remodeling processes. In concert with genetic and biochemical influences, local stresses and strains help regulate each of these processes. The response to modulation of a specific physiochemical stimulus depends not only on the type and magnitude of the stimulus but also the recent history at that particular site. For example, consider some of the numerous complex changes that occur in humans upon exposure to microgravity conditions. Reduced functional loading in microgravity leads to a rapid loss in bone mass from load-bearing sites within the skeleton at a rate of approximately 1% per month (Cowin, 2004). However, a corresponding fluid shift in the body toward the head may actually thicken bone in the skull due to increased cranial fluid pressure. Effects of Stress on Morphogenesis and Growth Morphogenesis refers to the process by which tissue patterns and structure arise from an initial amorphous collection of cells. Many tissue-engineering strategies seek to recapitulate the events involved during morphogenesis, and therefore an understanding of the effects of physical stresses is important. Although genetic factors clearly play a dominant role in morphogenesis, physical stresses contribute by fine tuning and perhaps optimizing the tissue’s structure and function for its intended function. Muscle contractions and joint movement begin around the sixth week of gestation in humans, producing intermittent stresses and strains that play an important role in the normal development and growth of musculoskeletal tissues. For example, paralysis of chick embryos results in a significant reduction in the recruitment and proliferation of immature growth plate chondrocytes compared to controls with normally functioning muscles (Germiller and Goldstein, 1997). Growth is the process by which tissue volume expands over time due to a net increase in either interstitial (within the tissue) or appositional (on the tissue surface) matrix synthesis. In bones, postnatal growth can be manipulated clinically by altering the local mechanical environment across a given growth plate. Increased pressure or compression slows growth likely due to compromised epiphyseal vasculature, whereas tensile forces applied by distraction devices can be used to accelerate growth (De Bastiani et al., 1986). Both approaches are used clinically to correct angular deformities or limb length discrepancies. Effects of Stress on Repair and Remodeling There is also strong evidence to suggest that alterations in the in vivo mechanical environment affect composition, structure, and mechanical properties of a wide variety of tissues in adults. In blood vessels, hemodynamic forces play multiple important roles in the regulation of vascular cells (Riha et al., 2005). The pulsatile nature of blood flow produces cyclic strain within vessel walls as well as shear stresses on the walls of the lumen. These two types of physical stimuli influence the phenotype and activity of smooth muscle cells and endothelial cells within the vasculature. Tremendous recent research attention has been directed toward studying hemodynamic effects given the potential implications for prevention or treatment of atherosclerosis

520 CELLS AND TISSUE DEVELOPMENT

as well as vascular tissue engineering. Arteries are capable of remodeling their structure in response to changes in their mechanical environment. A chronic increase in systemic blood pressure induces an increase in vessel wall thickness and area, while reduced pressure leads to a decrease in vessel dimensions (Arner et al., 1984). Abnormal joint loads have been shown to induce changes in composition, structure, and mechanical properties of articular cartilage. Disuse studies, for example, that use casting or other means of immobilization have demonstrated a loss of matrix constituents such as proteoglycans and a reduction in tissue thickness and mechanical properties (Akeson et al., 1987). Conversely, moderate exercise may have beneficial effects on maintaining healthy articular cartilage (Lane, 1996). However, high-impact loading or altered joint loading due to instability or injury is recognized as a significant risk factor for the development and progression of osteoarthritis (Buckwalter, 1995; Lane, 1996b). These studies suggest that there is a range of local stresses and strains that promote healthy tissue homeostasis, but loading conditions that are abnormally high or low can trigger catabolic responses and a loss of tissue function. Several theories have been put forth to explain the relationship between mechanical stress and strain distributions and patterns of cellular differentiation and tissue formation. For musculoskeletal connective tissues, Carter et al. (1998) introduced tissue differentiation phase diagrams that shift depending on the local vascular environment. The theory asserts that bone will form directly under conditions of moderate loading and adequate local blood supply. However, high shear or tensile hydrostatic stresses will tend to stimulate fibrous tissue formation, as often seen in unstable fracture non-unions. In addition, high-compressive hydrostatic stresses as well as a poor vascular supply are predicted to shunt tissue differentiation toward a cartilage pathway. The concept of taking advantage of mechanical stimuli to promote tissue repair has been applied clinically. Whereas prolonged rest was once typically prescribed to repair injured tissues, it is now recognized that early resumption of limited physical activity can promote tissue repair and restoration of function (Buckwalter, 1995a). Physical Stresses and Regenerative Medicine Replacing tissues that serve a significant biomechanical function has proven exceptionally challenging (Butler et al., 2000). Musculoskeletal connective tissues such as bone, cartilage, meniscus, tendon, and ligament and cardiovascular tissues such as blood vessels and heart valves are excellent examples of tissues that are subjected to repetitive high stress conditions in vivo. Tissue-engineering strategies designed to replace or regenerate such tissues must provide adequate biomechanical properties and integrate with surrounding native tissues in order to restore local function. Baseline biomechanical data for the tissue targeted for repair or replacement is essential (Butler et al., 2000). For example, the types and magnitudes of stresses and strains applied to the native tissue in vivo during a variety of activities must be determined. Along with measurements of native tissue mechanical properties, stress and strain history data provide design objectives for tissue-engineered constructs or regeneration strategies. Prioritization of desired mechanical properties will likely be necessary since the optimized structure– function relationships in native tissues may be difficult or impossible to duplicate. As such, a critical issue in the field is setting standards for adequate mechanical integrity (Butler et al., 2000). Such standards will certainly be tissue dependent and may even require patient-specific information such as weight or level of physical activity. Few studies to date have attempted to directly assess the effects of in vivo stresses on tissue-engineered constructs following implantation. Case et al. (2003) investigated the effects of controlled intermittent compressive deformation on cellular constructs using a hydraulic bone chamber device implanted into the distal femoral metaphyses of rabbits (Figure 29.4). Constructs receiving 4 weeks of daily mechanical loading at 0.5 Hz were found to have nine-fold more new bone formation compared to contralateral control constructs

Physical Stress as a Factor in Tissue Growth and Remodeling 521

Figure 29.4 Hydraulic bone chamber implant (top) used to apply cyclic compressive loading to tissueengineered constructs in vivo. Implanted constructs receiving the mechanical stimulus (bottom right) had nine-fold more new bone formation than no load controls (bottom left).

that did not receive loading. This study demonstrates the important role that the in vivo mechanical environment can play in the repair and integration of an implanted tissue-engineered construct.

MECHANOTRANSDUCTION MECHANISMS So how are local mechanical signals transduced into cellular responses that affect tissue growth, repair, and remodeling? The process of mechanotransduction can be divided into four stages (Gooch et al., 1998), as shown in Figure 29.5. They are: (1) force transmission, (2) mechanotransduction, (3) signal propagation, and (4) cellular response. The first stage refers to the transmission of the force from the point it is applied to the cell surface. The second corresponds to the sensory action of the cells in sensing mechanical stimuli, and transducing it into a biochemical signal, which is propagated inside the cell in the third stage. Finally the cell responds to the intracellular signal by modulating gene expression, completing the mechanotransduction process. In the first stage of mechanotransduction, applied forces are converted into local stimuli that may be detected by cells. Transmitted forces can cause direct cellular deformation by deforming the surrounding extracellular matrix (ECM). Applied forces may also result in local fluid flow and/or hydrostatic pressures. For example, compression of articular cartilage generates hydrostatic pressure that can regulate chondrocyte metabolism. Dynamic compression of cartilage induces fluid flow through the matrix and exposes cells to local shear stresses. The relative importance of these different types of local stimuli in vivo is not clear due to

522 CELLS AND TISSUE DEVELOPMENT

(1) Matrix α β

Integrin

Cell plasma membrane

(2) Receptor

Ion flux

Structural complex

Mechanosensitive ion channel

Signaling complex

Cytoskeleton

(3)

Gene expression modulation

(4)

Figure 29.5 Schematic showing the four stages of mechanotransduction: (1) force transmission, (2) mechanotransduction, (3) signal propagation, and (4) cellular response. See text for details.

the difficulty of isolating each kind of mechanical stimulus. However extensive research has been done to study the effects of various forms of mechanical stimuli on cells in vitro. These include tensile stretch, compression, hydrostatic pressure, and fluid-flow-induced shear stress, applied either statically or dynamically. These studies have allowed investigators to identify potential mechanotransduction mechanisms. The next stage of mechanotransduction occurs at the plasma membrane of the cell, and it is here that the cell detects the external signal and converts it into an intracellular signal. The plasma membrane contains numerous receptors and ion channels that can serve as sensors of the mechanical stimuli. The key structures in this interaction are the mechanosensitive (also known as stretch-activated) ion channels, integrin receptors, and other plasma membrane receptors. Mechanosensitive ion channels (Sachs, 1991; Hamill and Martinac, 2001; Martinac, 2004) are thought to be important to many cell types including chondrocytes (Wright et al., 1996; Guilak and Hung, 2005), osteoblasts (Charras and Horton, 2002), endothelial cells (Davies, 1995), and cardiac myocytes (Hu and Sachs, 1997). Experiments involving direct perturbation of the chondrocyte membrane have implicated such ion channels in the increase in concentration of cytosolic calcium ion (Guilak et al., 1999), which is a second messenger and has well-known intracellular effects (Rasmussen, 1986; Carafoli, 1987; Faber and Sah, 2003). Recently annexin V, a calcium-dependent phospholipid-binding protein, was proposed as a Ca2channel in osteoblastic cells (Haut Donahue et al., 2004). The flux of ions through these channels also affects the membrane potential (Wright et al., 1992; Gannier et al., 1996; Zabel et al., 1996) that triggers voltage-gated ion channels (Mobasheri et al., 2002), which further change the ion concentrations inside the cell. Two models have been proposed to explain

Physical Stress as a Factor in Tissue Growth and Remodeling 523

the mechanism of gating of these channels: the bilayer (Martinac et al., 1990; Hamill and Martinac, 2001) and the tethered models (Hamill and McBride, 1997; Gillespie and Walker, 2001). In the bilayer model, lipid bilayer tension alone is sufficient to activate the channels directly. The tethered model assumes that molecules in the cortical cytoskeleton and/or the extracellular domains directly interact with the channel protein to open/close the channel. Integrins are heterodimeric transmembrane proteins that bind to ECM proteins and cluster together leading to the assembly of focal adhesions, at which the cell contacts the ECM. Focal adhesions intracellularly associate with α-actinin (Otey et al., 1993), talin (Critchley, 2004), tensin (Bockholt and Burridge, 1993), and other cytoskeletal-binding proteins as well as signaling molecules like focal adhesion kinase (FAK) (Schaller et al., 1995). Due to their associations with both structural and signaling proteins, integrins are well placed to act as transducers of physical stimuli, and have been implicated as a link between the extracellular and intracellular environments for a variety of cell types that allows transmission of inside-out and outside-in signals capable of modulating cell behavior (Wright et al., 1997; Pelham and Wang, 1999; Jalali et al., 2001; Aikawa et al., 2002; Martinez-Lemus et al., 2003). In one study, over-expression of the tumor suppressor PTEN, which inhibits outside-in integrin signaling, strongly suppressed stretch-induced activation of p38 mitogen-activated protein kinase (MAPK) in cardiac myocytes (Aikawa et al., 2002). Jalali et al. (2001) demonstrated that fluid flow over endothelial cells activates integrin-mediated adhesion in an ECM-specific manner. The shear stress-induced mechanotransduction was abolished when new integrin–ECM ligand interactions were prevented by either blocking the integrin-binding sites of ECM ligands or conjugating the integrins to immobilized antibodies. Wright et al. (1997) reported that the transduction pathways involved in the hyperpolarization response of human articular chondrocytes in vitro after cyclical pressure-induced strain involve α5β1 integrin, which they suggest to be an important chondrocyte mechanoreceptor. Externally applied forces would cause changes in the conformations of the ECM molecules that would affect their binding to integrins, and modify the force balance within focal adhesions. It is thought that increased tension within focal adhesions can trigger increased integrin clustering and FAK phosphorylation (Sieg et al., 1999; Katsumi et al., 2004), which initiates a signal cascade resulting in altered gene expression. In addition to integrins, the plasma membrane is host to other receptors for specific ECM proteins like collagen, aggrecan, and hyaluronic acid, which may also be able to sense extracellular forces due to their interactions with their ligands. It is also possible that G-protein-coupled receptors may act as mechanotransducers or be activated secondary to other pathways, as the consequences of G protein stimulation of phospholipase C (PLC)–inositol trisphosphate (IP3) pathway has been observed in mechanically stimulated cells (Davies, 1995). It is very likely that the above-mentioned transducer molecules collaborate in the mechanotransduction response. In fact both integrin function and mechanosensitive ion channel activity were found to be required for chondrocyte response to cyclic pressurization (Lee et al., 2000). It has also been suggested that mechanical stimuli regulate cell behavior by a physical connection from intracellular organelles to the ECM via the cytoskeleton and the adhesion plaque (Guilak and Hung, 2005). The third stage of mechanotransduction is signal propagation, in which the signal generated at the plasma membrane in the second stage is propagated within the cell. This is usually carried out using the same machinery that the cell uses for responding to biochemical stimuli. Signal propagation is initiated by second messengers such as Ca2 , cAMP, and MAPK. Activated kinases subsequently phosphorylate transcription factors leading to changes in gene expression. Cytoplasmic calcium serves as a ubiquitous signal for regulation of important cellular processes such as cell growth, differentiation, protein synthesis, and even cell death. Numerous studies have found an increase in cytosolic Ca2 concentration due to mechanical loading in a variety of cell types (Hung et al., 1997; Edlich et al., 2001; Sharma et al., 2002; Donahue et al., 2003). This may be due to the opening of mechanosensitive

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Ca2 channels as discussed above or secondary to a mechanotransducer. The intracellular Ca2 concentration can also be elevated by release of calcium from intracellular stores through the IP3/diacylglycerol (DAG) pathway (Berridge, 1987). This pathway can be triggered by G-protein-coupled receptors leading to the activation of the enzyme PLC. PLC cleaves the phosphoinositide PIP2 to generate two second messengers: DAG and IP3. After diffusing though the cytosol, IP3 interacts with and opens Ca2 channels in the membrane of the endoplasmic reticulum, causing release of Ca2 into the cytosol. One of the various cellular responses induced by a rise in cytosolic Ca2 is recruitment of protein kinase C (PKC) to the plasma membrane, where it is activated by DAG. The activated kinase can phosphorylate various proteins, including transcription factors, leading to gene activation. Ca2 is also known to bind to the small cytosolic protein calmodulin to form a complex that interacts with and modulates activity of other enzymes and transcription factors. Ca2 influx is known to activate certain K channels thus affecting membrane potential (Wright et al., 1992; Faber and Sah, 2003), and has been shown to be necessary for integrin-dependent tyrosine phosphorylation of focal adhesion associated molecules (Alessandro et al., 1998). The cyclic nucleotide cAMP is produced by adenylyl cyclases which are in turn activated by G-proteincoupled receptors. Protein kinase A (PKA), which consists of two catalytic subunits and two regulatory subunits, is the most well-known cAMP effector. Binding of cAMP to the regulatory subunits releases the catalytic subunits, which are then free to phosphorylate substrates (Dumaz and Marais, 2005). cAMP, along with intracellular Ca2 , has been implicated in the regulation of gene expression in response to static compression of cartilage explants (Valhmu et al., 1998; Fitzgerald et al., 2004). Boo et al. (2002) demonstrated that shear stress stimulates phosphorylation of eNOS and thus nitric oxide (NO) production in bovine aortic endothelial cells in a PKA-dependent manner. As discussed earlier, mechanical stimuli may be able to activate FAK and other signaling proteins via integrin receptors. In chondrocytes, these signaling proteins are known to stimulate docking proteins such as Src-homology collagen (Shc) leading to the activation of the MAPK pathway (Shakibaei et al., 1999). The MAPK family consists of an array of serine/threonine kinases (ERK1/2, p38 MAPK, etc.) that are activated by a variety of physical and biochemical stimuli. However, integrins specifically appear to be involved upstream in this mechanotransduction response, irrespective of the tissue involved. The MAPKs are known to be activated by Ras, a small G protein. Ras is a membrane anchored switch protein that is turned on by certain receptors via docking proteins (Mitin et al., 2005). After being switched on, Ras phosphorylates and consequently activates a cascade of proteins, which ultimately lead to the activation of the MAPKs. The activated MAPKs regulate several regulatory molecules in the cytoplasm and in the nucleus to initiate cellular processes such as proliferation, differentiation, and development (Seger and Krebs, 1995). Many studies have implicated MAPKs in the cellular response to fluid flow and stretch (Hung et al., 2000; You et al., 2001; Plotkin et al., 2005; Torsoni et al., 2005). For example, Hung et al. (2000) showed that fluid-induced shear stress suppression of aggrecan gene expression in culture bovine chondrocytes is mediated in part by calcium-independent MAPK regulation. There is also evidence to show that some of the signal transduction pathways are linked. For example, Ca2 -activated calmodulin activates the enzyme cAMP phosphodiesterase that degrades cAMP and thus terminates its effect (Kakkar et al., 1999). Also as mentioned above, IP3 is an important mediator of cytosolic Ca2 release from intracellular stores. It has also been shown that cAMP inhibits MAPKs in several cell types (Dumaz and Marais, 2005). The final stage of mechanotransduction is the altered response of the cell, which may include changes in matrix synthesis/degradation, proliferation, differentiation, apoptosis, cell alignment, and migration. The effectors of the mechanotransduction pathways are the various transcription factors, which are activated by the events discussed previously. Numerous studies on vascular cells have shown activation of transcription factors

Physical Stress as a Factor in Tissue Growth and Remodeling 525

like AP-1, CRE, and NF-κβ in response to cyclic strain (Kakisis et al., 2004). The activated transcription factors interact with the promoter and enhancer regions of various genes to mediate transcription. This results in an increase in expression of genes like Cox-2, VEGF, TGF-β3, and eNOS (Kakisis et al., 2004), which orchestrate the cellular responses. Lee et al. (2001) demonstrated that vascular smooth muscle cells respond to mechanical strain by increasing specific proteoglycan synthesis and aggregation. It is known that mechanical loading of osteocytes results in anabolic responses such as the expression of c-fos, insulin-like growth factor-I (IGF-I), and osteocalcin (Mikuni-Takagaki, 1999). Elevations in Ca2 activate a Ca2/calmodulin-dependent protein kinase that causes increased c-fos expression, which is a pro-growth transcription factor. Calcineurin, a Ca2/calmodulin-activated phosphatase, dephosphorylates and activates the NF-AT family of transcription factors. Different NF-ATs, expressed in different cells including those of the heart, cartilage, and bone, serve as tissue-specific activators of cell growth and differentiation (Crabtree, 1999; Iqbal and Zaidi, 2005).

IN VITRO MECHANICAL CONDITIONING The replacement of tissues which reside in a complex, dynamic mechanical environment is a daunting challenge. Articular cartilage and blood vessels, for example, must bear tremendous stress and strain over repeated loading cycles in vivo while maintaining normal function. To date, no engineered construct has been developed in vitro possessing the same biomechanical properties as its in situ counterpart. One approach to address this challenge is the use of physiologically inspired mechanical forces to transmit stimuli to developing constructs in vitro. Since these tissues normally experience a dynamic environment in vivo, the rationale is that the application of mechanical forces such as compression or shear stress will stimulate the cells of the engineered construct to secrete and organize the proper matrix proteins required to reproduce the native tissue mechanical function. Delivery of controlled stresses and strains in vitro is achieved through mechanical devices known as bioreactors. Bioreactors have been used extensively as production vessels for engineered tissues. Many of these systems take advantage of the controlled in vitro environment to investigate the effects of specific biochemical or biomechanical factors on construct development. Bioreactor systems are highly diverse, but many are designed to delivery-specific mechanical signals to tissue constructs. Another common feature is they typically function to increase the mass transport of nutrients and waste through constructs via convective fluid flow. Bioreactors are also amenable to large-scale tissue production, as they are inherently scaleable and allow for increased process control, such as on-line measurement of pH or dissolved oxygen. Perhaps the tissues of the body most subjected to mechanical forces are those of the musculoskeletal and cardiovascular origin. Consequently, orthopaedic and cardiovascular tissue-engineered constructs represent the bulk of the research in which mechanical forces have been applied to developing tissues in vitro. Cartilage, bone, tendon, ligament, blood vessels, heart valves, and muscle have been cultured in vitro under the influence of mechanical forces. The remainder of this section will discuss select examples from the orthopaedic and cardiovascular fields which use the in vivo environment as inspiration to mechanically condition tissue-engineered constructs in vitro. Cartilage Bioreactors Articular cartilage is the whitish, low-friction tissue which lines the ends of long bones in a diarthrodial joint. It is a highly hydrated tissue (80% water), with type II collagen and proteoglycans constituting the majority of the solid matrix. These constituents combine to yield resilient mechanical properties which provide the shock absorption and nearly friction-free surface in joints such as the knee, shoulder, and hip. Jointbearing surfaces regularly experience complex high-magnitude mechanical loads through activities such as

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running and walking. In situ, normal joint loading produces compressive, tensile, and shear forces which deform the cells (chondrocytes) and induce interstitial fluid flows and streaming potentials throughout the matrix (Mow and Ratcliffe, 1997). These mechanical, chemical, and electric signals prominently influence the metabolism of the chondrocytes. As articular cartilage in adults is devoid of a blood supply, mechanical deformations are of critical importance to facilitate flow of nutrients waste products into and out of the tissue. Mechanical deformations also serve to maintain the tissue’s proper matrix composition, organization, and mechanical properties. It is generally accepted that static or constant compression/pressure results in loss and/or reduction of synthesis of proteoglycans and DNA in nearly a dose-dependent manner (Li et al., 2001). Dynamic compression has been shown to positively modulate proteoglycan synthesis and this stimulation is heavily influenced by both the frequency and amplitude of the compressive waveform (Li et al., 2001). Importantly, dynamic compression also modulates biomarkers implicated in important disease states (e.g. osteoarthritis) such as cartilage oligomeric matrix protein (COMP) (Piscoya et al., 2005). Similarly, dynamic tissue shear also has a pronounced effect on matrix components in a frequency- and amplitude-dependent manner (Jin et al., 2001). These insights into the role that mechanical deformations play on the cell metabolism, tissue growth, and remodeling in native tissue can be used to more effectively create tissue-engineered constructs. Thus, bioreactors constructed to apply compression and/or shear forces have been developed to modulate construct matrix composition and mechanical properties. While many different tissue-engineering models exist for cartilage (e.g. alginate, agarose, pellet/micro-mass, scaffold, and scaffold-free culture), these in-vitro-grown constructs generally possess similar amounts of proteoglycans compared to native articular cartilage, but lack the organization and amount of type II collagen (Freed et al., 1998; Carver and Heath, 1999; Waldman et al., 2003; Hung et al., 2004). Consequently, the mechanical properties necessary to withstand the complex and demanding in vivo mechanical environment have yet to be recapitulated. For clinical success, it has been suggested that tissue-engineered constructs may need to approximate the matrix composition, organization, and biomechanical properties of native tissue in order to promote construct integration and load-bearing capability in vivo (Hung et al., 2004). Bioreactor systems have produced encouraging results indicating that in vitro mechanical conditioning of tissue-engineered constructs is a promising approach to reproducing native tissue properties. As one example, a novel dual-chambered, parallel-plate flow bioreactor system has been used to apply controlled shear stresses to surface of cartilaginous constructs grown de novo from primary bovine articular chondrocytes without the aid of a scaffold (Figure 29.6). The “parallel-plate” design refers to the top bioreactor surface and tissue-engineered construct face which forms two parallel walls separated by a defined distance that creates a flow channel. Fluid is flowed through the channel, resulting in a parabolic velocity profile. Consequently, a shear stress is applied that is maximal at the upper wall and tissue surface; this is commonly referred to as Poiseuille flow (Fox and McDonald, 1992). One can estimate the wall shear stress (τw) by the following equation: τw 

6μQ bh2

where μ is the media viscosity, Q is the volumetric flow rate, b is the flow chamber width, and h is the fluid gap height. The system is designed to deliver this consistent level of shear stress to more than 95% of the tissue’s length in a laminar flow regime. This is critical as it has been shown that bioreactor-grown constructs cultured under turbulent conditions result in inferior tissues (Martin et al., 2000). Such findings suggest that a welldefined, controlled fluid environment is necessary to encourage proper tissue growth (Williams et al., 2002; Saini and Wick, 2003).

Physical Stress as a Factor in Tissue Growth and Remodeling 527

Entry port

Exit port Cap

Shim

Shim Upper-media chamber Wall shear stress τw

Fluid flow Q

Top

L

h

Top

Cells/tissue Membrane

Frame

Lower media chamber

Frame

Bottom

Figure 29.6 Dual-chambered parallel-plate bioreactor system that applies controlled shear stresses to the surface of cartilaginous construct slabs.

Chondrocytes are seeded on to a semi-permeable membrane that provides nutrients from either the top or bottom media chamber. After 2 weeks of static culture, a thin slab of cartilage has formed and attained a thickness of 250–1000 μm, depending on the number of cells used. Following the static pre-culture period, fluid-induced shear stress is applied to the construct. The application of flow significantly increases type II collagen compared to static (no flow) controls, as well as both Young’s modulus and ultimate strength (Gemmiti and Guldberg, 2006). This study suggests that flow-induced shear stresses may be an effective functional tissue-engineering strategy for modulating matrix composition and mechanical properties in vitro. Other bioreactor systems have used compression as a stimulus for cartilage construct development. Davisson et al. (2002) showed a decrease in sulfated glycosaminoglycans and protein synthesis under static compression, but an enhancement under a dynamic environment. Mauck et al. (2000) have shown that dynamic loading induces an increase in proteoglycan and total collagen content compared to static (free swelling, uncompressed) controls in an agarose gel model. Furthermore, this dynamic loading resulted in an increase in equilibrium aggregate modulus. The concurrent increase in matrix components and mechanical properties under the influence of in vitro mechanical conditioning indicates that bioreactor systems may be an effective approach to producing functional tissue-engineered cartilage constructs in vitro. Bone Bioreactors Without a vascular blood supply in vitro, nutrient delivery to cells throughout 3-D tissue-engineered constructs grown in static culture must occur by simple diffusion alone. As a result, attempts to engineered bone greater than 1 mm in thickness usually result in a thin shell of viable tissue and cells localized at the periphery (Gersbach et al., 2004). It has been theorized that this effect is due to sub-optimal mass transport conditions and a lack of mechanical stimulation in static culture. Therefore, tissue culture systems that provide dynamic

528 CELLS AND TISSUE DEVELOPMENT

media flow around or within tissue-engineered constructs have been designed to enhance nutrient and waste exchange in vitro (Bujia et al., 1995). In addition to enhancing mass transport, fluid flow applies shear stresses to the cells within the scaffolds. The effects of flow-mediated shear on cells have been studied in 2-D monolayer cultures. Continuous fluid flow applied to osteoblasts in vitro has been shown to alter bone-related gene expression and cellular phenotype (Ogata, 2000). Parallel-plate flow experiments have shown that bone cells cultured in monolayer are highly responsive to flow-mediated shear stresses. Shear stresses in the range of 5–15 dynes/cm2 affect osteoblast proliferation as well as production of NO and prostaglandin E2 (PGE2), suggesting that shear stress is an important regulator of osteoblast function (McAllister et al., 2000). Pulsatile and oscillatory flow conditions applied to osteoblasts using in vitro parallel-plate flow chambers have also been shown to increase gene expression, intracellular calcium concentration, and the production of NO and PGE2 in comparison to static controls (Klein-Nulend et al., 1997; Bakker et al., 2001). Furthermore, cell responsiveness has been reported to vary with fluid flow rate and frequency (Jacobs et al., 1998; Edlich et al., 2001). Proposed mechanisms for the stimulation of cells by fluid flow include increased mass transport, generation of streaming potentials, and application of shear stresses to the cell membranes (McAllister and Frangos, 1999; Bakker et al., 2001). Although these studies were performed using 2-D cell culture systems for short-term experiments, they suggest that variable flow conditions may also have differential effects in 3-D tissue culture systems. Such tissue culture systems may be useful to engineer thicker, more uniform bone graft substitutes for implantation or as test bed models that simulate aspects of the in vivo environment. While many different bioreactor systems have been developed, perfusion bioreactors in particular have shown significant increases in both cell viability and mineralized matrix formation on large 3-D constructs in vitro. In a recent study, micro-CT has been used to quantify mineralized matrix production within perfused and statically cultured marrow progenitor cells seeded on large polymer scaffolds (6.35 mm diameter, 9 mm thick) (Porter et al., 2005). Statically cultured constructs were found to have mineralized matrix localized only to the periphery of the constructs. In contrast, perfused constructs were found to have a several fold increase in mineralized matrix production distributed throughout the constructs (Figure 29.7). Blood Vessel Bioreactors Following the same rationale for mechanical conditioning of orthopaedic-engineered tissues, cardiovascular tissues can also be enhanced by in vitro mechanical stimulation. Cardiovascular tissues reside in a dynamic environment which can be mimicked in vitro using bioreactors and mechanical loading systems to deliver the physiologically inspired environmental cues. Small-diameter blood vessels (6 mm) are of particular importance because of their potential use to alleviate complications associated with atherosclerosis. Generally, a blood vessel has three layers (intima, media, and adventitia) in a tubular shape, forming a lumen through which blood passes. The intima is comprised mostly of a confluent, tightly adherent monolayer of endothelial cells (collectively called the endothelium) which is necessary to provide a non-thrombogenic surface for the blood to flow (van Hinsbergh, 2001). The media possesses smooth muscle cells and elastin and is set between the intima and the adventitia. The adventitia contains connective tissue (i.e. collagen) with fibroblasts embedded within. The ECM produced by the smooth muscle cells – the organized, cross-linked network of collagen and elastin – gives rise to the mechanical properties (Bank et al., 1996). These layers come together to form a vital tissue which must respond to the body’s complex and dynamic needs. In vivo, the pulsatile flow of blood imparts cyclic strains and shear stresses to the vessel’s constituents, which respond in a variety of ways to these mechanical signals. Endothelial cells are uniquely situated in the

Physical Stress as a Factor in Tissue Growth and Remodeling 529

Perfusion 1.3 1.1 0.9 0.7 0.5 0.3 0.0

Scaffold Flow rate (mm/s)

0.06 0.05 0.04 0.03 0.02 0.01 0.00 Shear stress (dynes/cm2)

Figure 29.7 Perfusion bioreactor system (left) for production of mineralized constructs for bone defects. Computational fluid dynamics simulation of flow rate and shear stresses within the 3-D scaffold porosity (right). lumen and are directly in contact with the flowing blood, which causes a shear stress to be applied to the cells. Consequently, these rapidly responding, mechanosensitive cells attain an elongated shape, aligning their long axis with the direction of flow. Sensing of the shear via cell surface receptors, ion channels, or integrins leads to secretion and/or activation of a number of signaling molecules, such as NO, endothelial nitric oxide synthase (eNOS), kinases, and transcription factors (Takahashi et al., 1997; Fisslthaler et al., 2000; Fisher et al., 2001). Perhaps most importantly, the fluid-induced shear stress confers a protective effect on the vessel by decreasing the probability of atherosclerosis (Traub and Berk, 1998). Indeed, areas of irregular blood flow (i.e. velocity, direction, and shear stress) have been implicated as sites of increased atherosclerosis (Papadaki et al., 1999). Shear stress also modulates smooth muscle cells’ production of signaling molecules (such as NO) (Papadaki et al., 1998) and gene transcription levels of cell surface receptors (Papadaki et al., 1998). Tissue-engineered vessels aim to reproduce cellular and mechanical properties of the native vessel in order to be an effective replacement. However, similar to other engineered tissues, those cultured in static conditions fall short of native tissue properties. Use of mechanical conditioning inspired by the in vivo environment has been shown in a variety of in vitro systems to modulate and improve engineered constructs. Exposing tissue-engineered vascular grafts to fluid-induced shear stress has been shown to increase endothelial cell adherence (Ott and Ballermann, 1995) and proliferation (Imberti et al., 2002) and alter tissue morphology and mechanical properties (Niklason et al., 2001). Cyclic mechanical strains cause an increase in collagen (types I and III) transcription by smooth muscle cells (Leung et al., 1976), an increase in mechanical properties (strength and stiffness), attributed to an increase in remodeling enzymes such as matrix

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metalloproteinase-2 (Seliktar et al., 2001), and an increase in matrix and cellular organization (Seliktar et al., 2000; Imberti et al., 2002). Subjecting smooth muscle cell impregnated constructs to dynamic mechanical stress not only causes ultrastructural and orientation changes in the cell phenotype and matrix, but can also induce cells to shift from a synthetic to a contractile state (Kanda and Matsuda, 1994). Similar constructs (smooth muscle cells seeded into polyglycolic acid meshes) exposed to pulsatile radial stresses of 165 beats per minute (analogous to fetal heart rates) and 5% radial strain produces constructs with burst pressures in excess of 2000 mm Hg, increased collagen deposition and desirable histological characteristics (Niklason et al., 1999). While great strides have been made in the field of tissue-engineered vascular grafts, a completely successful graft still has yet to be identified. However, as the field continues to progress and learn more about the in vivo environment, those cues can be translated to more realistic conditioning techniques for in-vitro-grown constructs. This mechanical stimulation is critical to remodeling the graft to possess proper mechanical properties as well as matrix composition and organization. The same can be said for cartilage and bone as well. Thus, mechanical conditioning in an in vitro setting has proven to be a powerful technique to increase the similarity of tissue-engineered constructs to the native tissues they aim to replace.

CONCLUSIONS Regenerating or replacing tissues that serve a significant biomechanical function has proven exceptionally challenging (Butler et al., 2000). It is now clear that tissue regeneration strategies must take into consideration the complex and demanding in vivo mechanical environment into which tissue-engineered constructs are implanted. Furthermore, static culture conditions have repeatedly been shown to produce tissues in vitro with vastly inferior mechanical properties compared to native tissue counterparts. Fortunately, a wealth of knowledge is now available to tissue engineers about how local stresses and strains affect cell function within tissues. Integration of this knowledge into strategies for tissue replacement or regeneration will be the key to achieving the goal of long-term functional restoration in patients.

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30 Engineering Cellular Microenvironments Wendy F. Liu, Elliot E. Hui, Sangeeta N. Bhatia, and Christopher S. Chen

INTRODUCTION Engineering cellular environments at the micrometer scale is critical for tissue engineering. The primary strategy for engineering tissue constructs uses a combination of cells and artificial scaffolds. Obtaining an adequate source of cells is a major challenge, since many of the cell types taken from adult tissue have a limited capacity for expansion. Recent developments in stem cell biology suggest that these cells might provide a key source of cells because they have the capacity for self-renewal and differentiation into multiple lineages. While promising, these cells alone cannot form a tissue. Cells must be combined with a scaffold, which provides the initial structural support onto which the cells adhere and organize into a functioning tissue. While simple in concept, forming complex tissues such as liver, which contain many different cell types and a defined tissue architecture, is a formidable task. When cells are removed from their natural in vivo environment, and placed in an artificial environment they often lose their tissue-specific functions. Hepatocytes, for example, are normally rounded and do not proliferate, but when removed from the body and cultured on a plastic culture dish, they spread, dedifferentiate, and reduce their liver-specific functions (Mooney et al., 1992). Mesenchymal stem cells (MSCs), which are derived from the bone marrow, differentiate into osteoblasts or adipocytes depending on their adhesive environment (Pittenger et al., 1999; McBeath et al., 2004). Engineering a functional cellular phenotype in an artificial environment has become a major effort in tissue engineering. A greater understanding of the extracellular cues that control the behavior of cells, stem cells or others, may lead to smarter design of scaffold materials. Biological structure and function are intricately linked at the tissue, cellular, and subcellular scales. Cells interact with soluble factors such as growth factors and cytokines, as well as insoluble factors such as extracellular matrix (ECM) proteins and other cells. The integration of soluble cues with those from both the matrix and neighboring cells plays an important role in regulating cell function. Cells are physically connected to the ECM through adhesion molecules known as integrins, which link the intracellular cytoskeleton to the ECM (Tamkun et al., 1986; Hynes, 1992). Many studies have demonstrated that binding of integrins to ECM leads to their clustering and the formation of focal adhesions, which then trigger intracellular signaling cascades and changes in numerous cellular processes (Schwartz and Ginsberg, 2002). Similarly, cells are physically connected to neighboring cells through cadherin molecules, which also serve as both mechanical linkages to the extracellular environment as well as signaling hubs to relay information to intracellular signaling pathways (Fagotto and Gumbiner, 1996; Wheelock and Johnson, 2003). Both integrin- and cadherin-mediated adhesions have been shown to modulate the ability of specific growth factor receptors to initiate intracellular signaling, induce changes in gene expression, and trigger specific cellular phenotypes. On the multicellular scale, cells within tissues are organized into functional units composed of multiple different cell types and arranged

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in a spatially defined manner. For example, the acinus, which consists of epithelial cells and fibroblasts organized in a spherical geometry, is critical for milk production in mammary glands (Nelson and Bissell, 2005). In fact, most tissues have highly defined structural components, which are indispensable for the functional integrity of the tissue. Thus, designing tissue-engineered constructs is not a simple amalgamation of cells with a scaffold, but instead requires an understanding of how cells behave in response to extracellular cues and the ability to design scaffolds with cellular scale resolution to mimic the architecture of the in vivo cellular environment. Numerous recent advances in microscale fabrication technologies have enabled investigators to control the architecture of biomaterials at the cellular and multicellular scale, and the organization of cells on such materials. These tools, which have been adapted from the microfabrication industry, utilize photolithographic methods to generate microscale features on silicon wafers. Poly(dimethylsiloxane) (PDMS), a biocompatible silicone rubber, is then cast directly on the silicon wafers yielding a rubber stamp with a negative replicate of the original features (a technique termed soft lithography). PDMS stamps are then used in a variety of different applications such as microfluidic delivery of biological agents or microcontact printing of proteins. These methods allow for spatial and temporal control over the presentation of extracellular cues to cells. Furthermore, the ability to miniaturize assays using microscale technologies allows for higher throughput screening of hundreds of thousands of materials and molecules for studying cell–environmental interactions. These tools have utility not only in basic research, where they can help identify the relevant structural cues that stabilize specific cellular phenotypes, but also in applications for producing tissue constructs, where devices to manipulate cellular phenotype by extracellular cues can help to improve overall tissue function. In the following chapter, we will examine recent efforts using microscale technologies to further advance the field of regenerative medicine. We will describe how these tools have been utilized to improve both the cellular and materials components of regenerative medicine. For engineering cells, these tools help investigators understand how the presentation of soluble cues, adhesive cues, and mechanical cues affects cellular behavior. For the biomaterials component, microfabrication can help to create spatially and structurally defined scaffolds that can be used to direct cellular function. We will also describe how these tools are being developed specifically for introducing tissue complexity in engineered cultures, such as in the examination of multiple cells or cell types, or in creating a structurally defined, three-dimensional scaffolds. While far from a complete review, this chapter will provide a glimpse into the ways in which microfabrication tools can be used to study cellular interactions and to create artificially engineered tissues for regenerative medicine.

DEFINING THE CELLULAR MICROENVIRONMENT The ability to control the cellular microenvironment has traditionally been limited by the inability to generate spatially defined structures on the cellular scale. Here, we will describe some of the pioneering studies and recent advances in microfabrication technologies used to engineer the cellular microenvironment, including techniques to spatially control the soluble, adhesive, and mechanical environment. Microfluidics to Spatially Control Soluble Cues Many of the earliest studies in biology focused on understanding the role of soluble factors. Changes in media components dramatically affect simple cellular behaviors such as cell growth and proliferation. These studies were generally performed with bulk changes in the concentration of soluble factors within well-mixed media. However, it has long been known that geometric patterns and gradients of soluble factors have profound effects on cell migration and differentiation. For example, asymmetric growth factor signaling in a developing embryo determines the anterior–posterior layout of the organism. During wound healing, the release of chemokines

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promotes the directed migration of numerous cells to the wound site. Early studies demonstrating the effects of soluble factor gradients on cells in culture used Dunn, Zigmond, or Boyden chambers, which utilized reservoirs or micropipette delivery of soluble agents (Boyden, 1962; Zigmond, 1977; Zicha et al., 1991; Wilkinson, 1998; Weiner et al., 1999). These methods relied on diffusion of molecules from a “source” to a “sink,” and could not provide control over the spatial geometry or dynamic properties of the gradients. The convergence of microfluidic technologies with biocompatible surface chemistries has recently achieved some of these goals (for reviews see Beebe et al., 2002; Sia and Whitesides, 2003). Microfluidic devices fabricated from PDMS have numerous advantages in biological studies, including biocompatibility, reduction in reagent consumption, and versatility in design. Investigators have used microfluidics to demonstrate that embryos cultured within microfluidic channels actually have developmental rates more similar to in vivo development compared to embryos cultured in a large culture dish (Raty et al., 2004). Interestingly, the volume of liquid within these microfluidic channels is comparable to the amount of liquid present near embryos within the crypts of the female reproductive tract in vivo. Importantly, it was found that the increased rate of development was caused by enhanced autocrine signals localized to these cells within the small channels. These devices not only have the advantage of improved cellular function, but also have improved handling and automation, enabling the efficient use of these precious cells. Another important advantage of microfluidics for biological applications is the ability to precisely control solute transport. Within microchannels, laminar flow dominates, thus limiting the lateral transport of molecules primarily to diffusion. Laminar streams flowing side by side will remain unmixed, but will eventually equilibrate if given enough time for diffusion to occur (by increasing the length of the channel and/or by decreasing the flow rate). Using a microfluidic network composed of repeated mixing and recombination of two or more laminar streams, Jeon et al. (2000) demonstrated the formation of arbitrarily defined spatial concentration gradients. The network of serpentine channels (Figure 30.1a) can generate gradients of specific patterns using variations in flow velocities and channel geometry. Recent work has further advanced these methods to include features that expedite mixing of fluids using microfabricated grooves within the channels (Stroock et al., 2002) or form complex flow patterns with the addition of PDMS valves (Unger et al., 2000).

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Figure 30.1 (a) Schematic diagram of microfluidic network used to generate spatially defined gradients of soluble factors. Solutions in the channels are iteratively split, combined, and mixed by diffusion to generate a larger channel with a gradient perpendicular to the direction of flow (Li Jeon et al., 2002). (b) Phase image of neural stem cells cultured in a single microfluidic channel with a gradient of growth factors (top) and immunofluorescence staining of astrocytes in green and nuclei in blue (bottom), demonstrating increased cell density (resulting from higher proliferation) in high concentrations of growth factor and increased differentiation into astrocytes in low concentrations of growth factor (Chung et al., 2005).

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Using microfluidic culture platforms, novel effects of soluble factor gradients have been revealed. Li Jeon et al. (2002) discovered that neutrophils migrate toward increasing concentrations of IL-8, independent of the steepness of the gradient. Interestingly, when migrating cells reached a local peak in the concentration gradient, the direction of migration was reversed after cells moved past a steep concentration drop in IL-8, but gradual decreases in concentration delayed the reversal response. This group has also demonstrated spatial control of differentiation versus growth of neural stem cells using a gradient of growth factor concentration (Chung et al., 2005). Across a channel 2.4 mm wide, neural stem cells on one side of the channel experiencing high concentrations of growth factors proliferated and remained undifferentiated, while cells on the opposite side of the same channel experiencing no growth factors differentiated and did not proliferate (Figure 30.1b). The demonstration of differentiating and proliferating stem cells in proximity allows one to begin to examine the role of crosstalk between these cells in a developing “tissue.” Such devices may be important in generating spatially defined tissue-engineered constructs. Microfluidic platforms have recently been extended to the treatment or analysis of a part or region of a single cell. Previous studies of cellular fractions were limited to fractionation by solubility (e.g. surfactants such as triton are used to separate soluble components from insoluble components) or by density (using an ultracentrifuge). However, it was neither possible to separate cellular fractions by their spatial location, nor to subject parts of a single cell to different treatments. Cells cultured within microfluidic channels may sit across more than one laminar stream, and therefore experience more than one soluble treatment. Therefore, a fraction of a single cell can be treated with a labeling agent, pharmacological drug, or enzyme (Takayama et al., 2003). In a different method, two laminar streams were separated by a compartment containing microgrooves, which were large enough to allow the passage of neural axons but not the cell bodies (Taylor et al., 2005). In addition, surface tension prevented the exchange of fluids between the streams. Using this device, axons and cell bodies could be subjected to different soluble treatments and each compartment could be harvested independently, permitting biochemical analysis of pure axonal fractions. In sum, microfluidic technology has developed substantially in the past decade to allow many researchers to use these tools for cellular studies. However, these technologies must become more widely adopted before investigators can understand how gradients may be applied to assist in engineering tissues. Microengineered Tools to Define the Adhesive Environment The insoluble environment, consisting of both ECM proteins and biomaterials, plays a critical role in determining the behavior of adherent cell types such as endothelial cells, epithelial cells, fibroblasts, bone cells, cartilage cells, and numerous others. Considerable research is currently focused on how both the composition and the spatial arrangement of these insoluble cues affect cell fate and function. In the following section, two different ways microengineered tools have been used to help define the cellular adhesive environment will be discussed. First, we will describe how microarrays of synthetic and natural molecules have been used to screen thousands of different materials for their effects on cell function. We will then examine several different ways microengineered tools can precisely control the geometry of adhesive ligand placement, and how these tools have revealed unique mechanisms of cellular behavior. Micropatterned Screening Arrays One of the major advantages of microengineered tools for biological applications is the ability to miniaturize assays and therefore reduce the total amount of reagents needed. Using traditional cell culture techniques to screen hundreds of thousands of potential ECM protein combinations and synthetic biomaterials for optimal culture conditions is impractical simply because of the cost of the reagents and supplies. Micropatterned

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screening arrays allow practical and efficient screening of these different materials. Robotic spotting technologies enable the deposition of nanoliter volumes of proteins (MacBeath and Schreiber, 2000; Falsey et al., 2001; Flaim et al., 2005), nucleic acids (Ziauddin and Sabatini, 2001), and biomaterials (Anderson et al., 2004), thus allowing for high-throughput screening of the effects of these molecules on cell function. In early studies, cells were seeded across the entire surface of the array, and analysis needed to be accomplished within 24–48 h of the initiation of the experiment, prior to when cells began to migrate away from their original location. Recent advances in these assays utilize a non-adhesive background surface such as poly(hydroxyethyl methacrylate) (pHEMA) (Anderson et al., 2004) or polyacrylamide (Flaim et al., 2005) to prevent cellular migration away from their original spot and therefore maintaining pattern fidelity over long periods of time (days to weeks). Anderson et al. (2004) generated an array of synthetic polymers by depositing different commercially available acrylate monomers that were polymerized with a photoinitiator (activated by light) onto a pHEMAcoated glass slide. After seeding of cells, the substrates were analyzed by typical fluorescence immunoassays. Human embryonic stem cells were cultured on these biomaterial arrays for 6 days. This group found that cells that adhered and spread typically differentiated into epithelial cells as detected by cytokeratin immunostaining. Using the microarray as an initial screen, potential “hits” that generated certain cellular phenotypes could then be further examined for specific mechanisms of adhesion and differentiation. Such arrays can unveil novel materials that yield desirable cellular phenotypes, which can then be tested as potential tissue engineering scaffolds. In addition to arrays of synthetic biomaterials, arrays of ECM proteins have also been fabricated. In this study, Flaim et al. (2005) examined the behavior of hepatic cells and embryonic stem cells on 32 different combinations of five different ECM molecules using a commercial protein array spotter (Figure 30.2a). Importantly, this work introduced a generalized platform technology, allowing arbitrary mixtures of proteins to be bound non-covalently on an otherwise non-adhesive background. Interestingly, the effects of the ECM

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Figure 30.2 (a) Hoffman contrast image of array of primary rat hepatocytes (top, left) and live/dead (red/green) stained hepatocytes (top, right). High magnification phase contrast (bottom, left) and immunofluorescence (bottom, right) images of a single island (Flaim et al., 2005). (b) Nomarski image of bovine adrenal capillary endothelial cells confined on different sized patterned islands of fibronectin (larged square is 40 μm in width) (Chen et al., 1997).

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combinations were not additive – the presence collagen IV in combination with other ECM molecules was sufficient to support hepatocyte function, but the differences observed among the varying combinations did not appear to be caused by differences in collagen IV concentration, since changing the concentration in collagen IV matrices alone had no significant effect on hepatocyte function. Such information can provide an initial screen for optimizing cell culture environments. With the numerous possibilities for adhesive ligands and biomaterials that can be used in engineering tissue constructs, obtaining this knowledge is made feasible and expedited with the use of micropatterned screening arrays. Spatial Patterning of the Adhesive Environment Traditional methods to modify the cellular adhesive environment typically varied the coating density of ECM ligand across the entire culture surface. Using these techniques, it was demonstrated that increasing ligand density increased the degree to which cells are spread, and concurrently increased growth rate. However, these techniques could not isolate the effects of ligand density from cell spreading, and were also limited in their ability to control the geometric placement of the adhesive ligands. Patterning techniques that combined microfabrication techniques with biologically compatible materials have since allowed investigators to create a patterned surface with discrete regions of chemistries that are adhesive or non-adhesive to cells. Thus, the independent manipulation ligand density, total ligand quantity and cell spreading, was possible. Early attempts to direct the location of cells in culture used patterning methods that consisted of depositing metals such as palladium through a nickel mask onto an otherwise non-adhesive surface (Carter, 1967). When cells were seeded onto these substrates, they landed exclusively onto the palladium-coated regions. However, the mechanism of adhesion onto palladium and other metal surfaces was not well defined. Furthermore, these methods required the use of specialized equipment for chemical deposition, preventing their widespread use. To overcome some of these limitations, a number of techniques based on soft lithography have been developed from the microfabrication industry and adapted to a variety of biological systems. Soft lithography requires a photolithographically generated silicon master, which once generated can be used repeatedly to cast PDMS rubber stamps. In a method called microcontact printing, stamps are used to directly transfer ECM ligands. The stamp is first coated with a solution of ECM proteins, and then dried and stamped onto the cell culture surface. The unstamped regions are blocked with a non-adhesive such as bovine serum albumin (BSA) or pluronic. Upon seeding, cells adhere and spread onto the micrometer-sized adhesive islands but are restricted from spreading onto the non-adhesive regions. These patterns are viable for several days to weeks, depending on the type of non-adhesive material used (Nelson et al., 2003). This method has been adapted to a number of different commonly used cell culture substrates such as glass, PDMS, and polystyrene (Tan et al., 2004). An alternative method to pattern using PDMS stamps is via microfluidic delivery of solutions of adhesive ligands (Chiu et al., 2000). Delivery of ECM ligands through microchannels that form upon sealing a stamp against a substrate (typically glass) can be achieved either by capillary action or by fluidic pumping. After the ECM proteins adsorb to the surface, the stamp is removed and the remainder of the surface is blocked with a non-adhesive, yielding a pattern of adhesive and non-adhesive regions. Conversely, a solution of non-adhesive such as agarose or polyacrylamide can be delivered through the channels and upon stamp removal, the remaining regions can be coated with an ECM protein (Nelson and Chen, 2002). While it has long been thought that cell spreading or shape influences a variety of cellular behaviors, micropatterning techniques have definitively demonstrated that cell spreading is a critical mechanism by which cells regulate their behavior. Singhvi et al. (1994) first demonstrated that hepatocytes cultured on islands of increasing sizes exhibited increased proliferation and decreased differentiation. Based on this study, it was still not clear whether the increase in total amount of ECM presented to spread cells was causing the increases in proliferation, or if cell spreading itself could induce proliferation. Chen et al. (1997) explored

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this question using micropatterning tools. They found that cells spread across multiple small islands (3 μm diameter) had similar proliferation to cells that were spread across a solid ECM substrate. Therefore, even though the cells that were spread across multiple small islands were presented with less total amount of ECM, they could still proliferate, suggesting that spreading alone provided a physical cue to regulate cell proliferation. Recently, the role of spreading on differentiation of human MSCs has also been explored (McBeath et al., 2004). MSCs are stem cells derived from adult bone marrow which can differentiate into a number of different lineages such as bone, fat, cartilage, and muscle. McBeath et al. demonstrated that MCSs that were cultured on large islands and well spread were more likely to differentiate into bone, while MSCs cultured on small islands and were rounded were more likely to differentiate into fat. This study demonstrates a systematic way to direct cell fate using the geometric presentation of adhesive ligands, and may provide a way to direct stem cells fate for use in artificially engineered tissues.

Engineering Substrate Mechanics While much effort in developing scaffolds for tissue engineering has been focused on their chemical and adhesive properties, it is also well established that cells are sensitive to their mechanical environment. As a cell adheres to the underlying substrate, forces are generated and transmitted through the intracellular cytoskeleton to adhesive structures formed at the membrane, resulting in cell spreading and changes in intracellular signaling (Geiger and Bershadsky, 2001). Both the mechanical environment surrounding the cell and the intracellular cytoskeletal mechanics play an important role in determining the magnitude of these forces and the resulting changes in cell behavior. Early studies to perturb the cellular mechanics exposed cells to spatially uniform stimuli, for example, by adding a cytoskeletal inhibiting pharmacological agent or by applying a uniform mechanical stimulus to cells seeded on a flexible membrane. Microengineered tools provide a spatially defined mechanical environment and the capacity to detect forces at the cellular, and even subcellular, level. In the following section, we will describe (1) how microscale technologies have provided simple methods to measure cell traction forces and (2) how micropatterning tools are used to create substrates with spatially defined mechanical properties. MEMS Devices to Measure Cellular Forces One of the earliest methods used to measure subcellular forces involved seeding cells on soft materials such as hydrogels or silicone elastomers (Harris et al., 1980). As cells attach and generate forces against the underlying compliant substrate, the substrate deforms and wrinkles. The magnitude and number of wrinkles provided a qualitative estimate of the traction forces. Investigators further advanced this system to enable the quantification of forces by embedding tracking particles within poly(acrylamide) sheets and measuring their displacement (Oliver et al., 1995; Dembo et al., 1996). Using these tracking particles, it was demonstrated that forces exerted at adhesions correlated with the size of the adhesion and that pharmacological agents to disrupt cytoskeletal tension abolished these forces. However, these methods were limited because they are computationally intensive and the movements of discrete particles do not fully describe the deformations of a continuous substrate. To circumvent some of these problems, several MEMS or microfabricated electro-mechanical systems have been developed. These devices have micrometer-scaled, mechanically deformable parts that allow the precise detection and quantification of cell-generated forces. Galbraith and Sheetz (1997) were the first to use a microfabricated device to measure the traction forces of a migrating fibroblast. In this study, they fabricated microscale mechanical cantilevers that could deflect as a cell migrates over it. Each cantilever provided a discrete measure of forces, as opposed to

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Figure 30.3 (a) Schematic diagram of cells adhered to the tips of microneedle arrays, and the deformation of the needles with traction forces (left), and SEM image of cell on microneedles (bottom) (Tan et al., 2003). (b) Phase contrast image of bovine pulmonary artery endothelial cells seeded on acrylamide substrates with patterned stiffnesses. Cells migrate to the stiff regions over the course of 48 h after seeding (Gray et al., 2003).

the previous methods where forces could propagate across the continuous substrate. Using this tool, they demonstrated rearward forces at the leading edge of a migrating fibroblast and frontward forces at the trailing edge. In an approach that combined microfabrication technologies with deformable substrates, Tan et al. (2003) developed a microfabricated post-array detector (mPAD) to measure the traction forces of stationary cells (Figure 30.3a). This device consists of an array of PDMS posts or microneedles, approximately 3 μm in diameter, 11 μm in height, and separated by 9 μm. The tips of the needles are coated using microcontact printing techniques described in section “Spatial patterning of the adhesive environment,” and the remainder of the substrate is blocked with a non-adhesive. Cells adhere solely to the tips of the needles and deform them as they exert forces at their adhesions. Using this system to control the different degrees of cell spreading while measuring cell traction forces, it was demonstrated that the greater the extent of cell spreading, the greater the degree of forces. This microneedle system also enable the control of mechanical properties by changing the substrate geometry (e.g. increasing the length of the post can generate softer posts) without changing the polymer crosslinker density or the substrate chemistry, therefore eliminating the effects of surface chemistry on cell mechanics. Furthermore, the post-geometry allows the measurement of forces in multiple directions, unlike cantilevers that measure only along the vertical axis. While MEMS devices that can measure cellular mechanics are only beginning to emerge, their utility is indisputable. As more investigators begin to delve deeper into this area and improve such devices by increasing their resolution or incorporating active components to apply mechanical forces, a greater understanding of how cells interact mechanically with their environment can be revealed.

Patterning Substrate Stiffness Most conclusions drawn from studies of cell biology are based on cells cultured on very hard surface such as plastic culture dishes or glass substrates. However, several studies have demonstrated that cells respond dramatically to their surrounding substrate stiffness. For example, endothelial cells form capillaries or tube-like structures on soft substrates, but spread out and proliferate on rigid substrates (Ingber and Folkman, 1989; Deroanne et al., 2001). Myocytes differentiate and form striations only on substrates of intermediate stiffnesses, but not very stiff or soft substrates (Engler et al., 2004). Mammary epithelial cells form normal acini on soft substrates, but have a malignant behavior or stiff substrates (Paszek et al., 2005). Interestingly, the

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stiffnesses of the substrates on which differentiated phenotypes were observed in vitro mimicked the physiologic stiffnesses of these tissues. The stiffness of a tissue is likely inhomogeneous in nature – stiffness might vary from region to region, across micrometer length scales. Cells respond to gradients in substrate stiffness, also termed durotaxis, or the migration of cells between regions of different mechanical properties. Wang et al. (2000) demonstrated that fibroblasts migrate from regions of soft rigidity to regions of stiff rigidity, and furthermore that the processes mediating cellular responses to substrate stiffness are regulated by intracellular contractility (Pelham and Wang, 1997; Guo et al., 2006). Grey et al. (2003) extended this to the micrometer scale by adapting microengineering tools. Here, the stiffnesses of PDMS or acrylamide substrates were tuned from 1.8 to 34 kPa by varying the crosslinker density. Stiff islands were patterned among a soft substrate, and cells were observed to migrate predominantly onto the stiff regions, forming islands of cells (Figure 30.3b). These effects were observed in both endothelial cells and fibroblasts. Currently there are only a handful of studies suggesting that cell substrate mechanics or stiffness play an important role in modulating cell behavior, but this concept is quickly gaining widespread support. A deeper understanding of how different cells respond to stiffness of their surroundings may be useful for applications in tissue engineering. Moreover, the design of materials with spatially and temporally controlled mechanical properties may be important for generating functional units of tissue.

DEFINING THE ORGANIZATION OF MULTICELLULAR CONSTRUCTS Tissues and organs are exquisitely ordered three-dimensional structures composed of multiple cells and cell types. To a large extent, the microenvironment is defined by the local organization of cells, which secrete paracrine factors, deposit ECM, present surface ligands, and exert physical force. Therefore, fully understanding and recapitulating the microenvironment involves not only the techniques described in section “Defining the cellular microenvironment,” but also additional methods to organize and study heterogeneous multicellular constructs, in both two and three dimensions. Patterning Multicellular Constructs in Two Dimensions Patterning of adhesive and non-adhesive regions on two-dimensional substrates has been described in section “Microengineered tools to define the adhesive environment.” Once the surface is defined, uniformly seeded cells will selectively adhere to adhesive regions and form the desired pattern. Co-cultures of multiple cell types can be patterned using biochemistries specific to individual cell types. Selective chemistries are not always available, however, thus recent studies have explored more general means to pattern multiple cell types. Using microfluidics, cells can be directly delivered to desired locations on a uniform substrate (Chiu et al., 2000). Additional methods include hydrogel molding (Tang et al., 2003), layer-by-layer deposition of ionic polymers (Khademhosseini et al., 2006), and dynamically regulated surfaces (see section “Dynamically changing the adhesive environment”). Microscale control of multicellular organization has brought an unprecedented ability to study interactions between individual cells or groups of cells within a colony. Using simple but carefully planned geometries, the following examples illustrate the biological insights that can be gained using cell patterning tools. Homotypic Interactions Previous studies examining the role of cell–cell adhesions typically uniformly seeded cells at different densities. Cells seeded at low densities had few cell–cell contacts, while cells seeded at high densities had many

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cell–cell contacts. However, when seeding density is changed, other microenvironmental cues are also varied. At higher density, not only does the number of cell–cell contacts increase, but also the amount of cell spreading decreases as the cells become crowded to fill the culture dish. Furthermore, the amount of soluble paracrine signals secreted from the cells may differ across cultures with varying densities. Based on these studies, it was therefore unclear what respective roles are played by each of these factors in regulating cell function. Microfabricated tools can be used to independently vary microenvironmental factors such as the degree of cell spreading from cell–cell interactions, therefore enabling investigators to tease out the critical parameters leading to changes in cell function. A microfabricated bowtie system was devised to decouple control of cell–cell contact and cell spreading. Bowtie-shaped regions for cell attachment were defined by patterning a non-adhesive agarose gel on a glass substrate. Each half of the bowtie allowed room for a single cell, fixing the amount of cell spreading. Pairs of cells could contact each other through the constriction at the center of the bowtie (Figure 30.4a). Cells were cultured either in pairs or as single cells occupying only half of the bowtie. Paired cells in contact demonstrated significantly higher rates of proliferation in comparison to single cells, implicating contact as an inducer of proliferation. In addition, paired cells in bowties where contact was physically blocked (Figure 30.4a) did not show greater proliferation than single cells, suggesting that paracrine signaling at close proximity was not sufficient. In fact, the authors demonstrated that specific receptors – cadherins – engaged upon cell–cell contact, and this receptor ligation induced the changes in cell function (Nelson and Chen, 2003). Besides the biochemical signaling that occurs within a community of cells, physical forces are another important “signal” that is transmitted through cell–cell interactions. Recently, Nelson et al. (2005) utilized micropatterned cultures to bring new insight into the factors that drive tissue morphogenesis. It was observed that cell proliferation was greatest at the edges of patterned sheets of endothelial and epithelial cells. In addition, the effect was more pronounced along longer edges of a rectangular sheet of cells and was not observed on concave edges. Mechanical modeling of variously shaped cell patterns revealed distributions of tensile stress within the cell sheets that directly correlated to the observed patterns of proliferation, with higher

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Figure 30.4 (a) Phase contrast and fluorescence images of cell pairs plated onto bowtie structures, in contact (left) and without contact (right) (Nelson and Chen, 2002). (b) Plot of strain over FEM models of patterned cell sheets (top) corresponds to regions of rapid proliferation in cultures (bottom) (Nelson et al., 2005). (c) Phase contrast image of patterned hepatocyte islands surrounded by fibroblasts (left), and fluorescence image of albumin expression (green) localized to the periphery of a hepatocyte island (right).

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proliferation in regions of high stress (Figure 30.4b). The modeled stress distributions were verified using the mPADs as described in section “MEMS devices to measure cellular forces.” Furthermore, the patterns of proliferation disappeared upon pharmaceutical disruption of cell tension, further implicating tensile stress in the regulation of cell proliferation. Most importantly, the proliferation at edges required that these contractile forces in individual cells were transmitted through cell–cell junctions to allow the multicellular sheet to act as a single mechanical unit; disrupting cell–cell adhesions caused the cells to no longer proliferate only at the edges, and instead stochastically throughout the sheet. It was concluded, therefore, that the shape of a tissue dictates the internal distribution of stress, which in turn drives asymmetries in cell proliferation. Tissue form, therefore, is not simply a consequence of growth, but is itself an active regulator of growth. Heterotypic Interactions While the generation of patterns of cells of the same type can be achieved by using patterns of the appropriate geometry, the patterning of multiple cell types with controlled placement of each of the different cell types is experimentally more challenging. Bhatia et al. (1999) employed micropatterned cultures to examine heterotypic cell interactions in a liver culture model. Typically, primary hepatocytes rapidly lose their phenotype in culture, however, co-cultivation of hepatocytes with non-parenchymal cells has been found to stabilize liver-specific function for a period of weeks. In order to explore the optimization of these co-cultures, cell patterning was employed to control precisely the interactions between different cell types. Microfabrication was used to define collagen regions on a glass substrate. Hepatocytes preferentially attached in collagen regions, while subsequently seeded non-parenchymal cells adhered in the remaining glass regions via adsorbed serum proteins (Figure 30.4c). By varying the size of the hepatocyte islands, it was possible to vary the interfacial area between the two cell types, and thus the amount of heterotypic contact, while holding constant the overall ratio of hepatocytes to non-parenchymal cells in the culture dish to eliminate the effects of paracrine signaling. In another experiment, the cell ratio was varied while interfacial area was held constant. Significantly, it was observed that liver-specific function increased as heterotypic contact increased. In addition, using an in situ assay, it was demonstrated that hepatocytes near the periphery of islands exhibited higher function, indicating that it was important for hepatocytes to be in close proximity to non-parenchymal cells. Finally, function also increased as the ratio of non-parenchymal cells to hepatocytes increased. These studies demonstrated that heterotypic interactions between hepatocytes and neighboring non-parenchymal cells within the liver are critical to liver function. Patterning in Three Dimensions While most studies engineering the cellular microenvironment have been performed in two-dimensional cultures, cells in vivo exist within a three-dimensional environment. Importantly, studies have demonstrated that cells cultured in a three-dimensional environment may have distinct phenotypes from the same cells cultured in two dimensions (Mueller-Klieser, 1997; Cukierman et al., 2002). Of the many strategies that have been devised for fabricating three-dimensional tissue constructs (Tsang and Bhatia, 2004), hydrogel-based constructs offer some of the greatest potential for precise control of the microenvironment (Lee and Mooney, 2001). In particular, recent advances in synthetic hydrogels offer the ability to tailor the presentation of bioactive ligands and proteolytic remodeling in response to cell-secreted factors (Lutolf and Hubbell, 2005). However, most studies have examined a bulk mixture of cells in a gel, without spatial control over where the cells are located within the gel. A number of recently reported methods therefore focus on patterning three-dimensional hydrogel cell cultures. Liu and Bhatia (2002) used photopatterning to construct three-dimensional structures of hydrogels containing encapsulated living cells. Poly(ethylene glycol) diacrylate (PEGDA) was dissolved and combined with

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Figure 30.5 (a) Three-layer patterned cell–hydrogel hybrid construct, shown with individual layers and stacked (right). (b) Articular chondrocyte clusters of varying size, with deposited sulfated glycosaminoglycans (sGAG) stained in blue (left). Plot of sGAG deposition as a function of cluster size (filled squares) compared to experimental controls in which cells were clustered and redispersed (open circles) (Albrecht, 2006).

cells and a photoinitiator, and the mixture was then polymerized by exposure to ultraviolet (UV) light. Patterned structures were formed by exposing through a photomask and polymerizing locally in the regions exposed to UV light. The process could be repeated multiple times using different cells and mask patterns to generate multilayered constructs of multiple cell types. Structural patterning becomes particularly important for larger tissue constructs, where diffusive transport of nutrients through the bulk hydrogel is limited. Branched structures within a hydrogel can ensure that cells will receive the appropriate nutrients (Figure 30.5a). Furthermore, complex structures can be formed with multilayer patterns that contain varied cell types and hydrogel formulations. The photopatterning method was able to form cell-containing structures with minimum features on the order of 100 μm, however the arrangement of individual cells within the hydrogel was not controllable. In a complementary method reported by Albrecht et al. (2004, 2006), cells were positioned within a similar PEGDA hydrogel with near single-cell resolution using dielectrophoresis, by which polarizable objects (such as cells) experience electrokinetic forces in the presence of an electric field. Cell viability and differentiated markers were maintained for over 2 weeks following electropatterning. To study the effect of cell proximity on function, articular chondrocytes were patterned in clusters of varying size, and the biosynthesis of sulfated glycosaminoglycans (sGAG) was measured. It was found that the rate of sGAG deposition per cell was highest for unclustered cells and decreased in a dose-dependent manner with increasing cluster size, reaching a plateau for clusters of more than five cells (Figure 30.5b). In a related study, combining the photopatterning and electropatterning methods, live cells were first positioned by dielectrophoresis and then immobilized by local photopolymerization through a mask (Albrecht et al., 2005). Thus, hierarchal patterning control was achieved over a length scale ranging from

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microns to centimeters. Future advances in three-dimensional micropatterning methodology promise further elucidation of tissue biology as well as continued functional improvements in engineered constructs.

DYNAMICALLY CHANGING THE MICROENVIRONMENT While most studies have examined the cellular microenvironment in a static setting, cells are constantly experiencing dynamic changes in their natural environment. For example, during angiogenesis, or the development of new blood vessels, endothelial cells detach and migrate away from their neighboring cells, form new adhesions with the surrounding ECM, proliferate, and form tube-like structures. While numerous researchers have sought to understand the various environmental cues that affect this process, very little is known about its temporal regulation. A greater understanding of blood vessel formation could not only have broad scientific impact, but also have practical applications such as help to find new methods to vascularize engineered tissues. Studies that can modulate a temporal component are only in their infancy since methods to control the dynamics of extracellular cues are technically challenging. In the following sections, we will describe some of the recent developments in microscale technologies that have not only spatial, but also temporal control over the cellular microenvironment.

Dynamic Regulation of the Soluble Environment Experimentally, it is difficult to dynamically regulate the soluble environment at a physiologically relevant frequency and to provide controlled, reproducible dynamic changes for systematic studies. With bulk changes in the media, the frequency is limited by the researcher’s ability to change the medium. However, cells in the body experience dynamic change with frequencies that cannot be attained by manual changes in medium. For examples, chondrocytes experience changes in osmotic loading due to mechanical forces on the charged ECM; the frequency of these changes are on the order of 0.01–0.1 Hz. An advantage of microfluidic technology described earlier in section “Microfluidics to spatially control soluble cues” is that computers, pumps, and valves control the changes in media, therefore allowing much higher frequencies of loading. Chao et al. (2005) applied changes in osmotic pressures to chondrocytes using a microfluidic device with two input liquid streams of different osmotic pressures. Here, the dynamic changes in cell morphology response to osmotic loading were dependent on frequency of loading. In addition to the ability to generate geometrically defined soluble gradients described earlier, microfluidic technology also has the ability to temporally control these gradients. Irimia et al. examined neutrophil migration response to changes in gradients of IL-8. They tracked neutrophil migration response to “step up” (increased steepness), “step down” (decreased steepness), or “flip” (reversed) changes in gradient that were achieved in less than 5 s. Neutrophils changed their velocity but not direction in response to “step up” and “step down,” and changed direction in response to the “flip.” These findings may provide further insight to the mechanism of neutrophil chemotaxis.

Dynamically Changing the Adhesive Environment A variety of methods to define the adhesive microenvironment are discussed in section “Microengineered tools to define the adhesive environment,” however these procedures are only applicable prior to the introduction of cells into the system. Once cells are plated, little adhesive regulation is experimentally possible short of global application of an enzymatic cleaving agent to release all cells. Only recently have groups begun to report methods to dynamically modify substrate surfaces during cell culture.

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Okano et al. (1995) developed a thermally responsive material, poly(N-isopropyl acrylamide) (pNIPA), which switches from hydrophobic at 37°C to hydrophilic at 20°C. Cells can attach and spread on the surface at the higher temperature, but detach when the surface is cooled. Cheng et al. (2004) extended this strategy by fabricating a microheater array underneath a pNIPA layer to locally regulate cell adhesion. Significantly, Yeo et al. (2003) have been able to achieve dynamic, molecular-level control of a substrate surface. Self-assembled monolayers (SAMs) were modified to present an electrically active ester, which can release and bind ligands via electrochemical redox reactions. Using this method to control the expression of the RGD peptide on a substrate, fibroblast adhesion to the surface was dynamically regulated. The RGD peptide mediates cell adhesion via integrin binding sites, thus cells were able to attach and spread on surfaces expressing this ligand. Upon application of an electrical potential, the RGD ligand was electrochemically cleaved from the SAM and released, along with the attached cells. Electrically active and non-active chemistries could be patterned together, enabling selective patterned release. In addition, following release, another RGD conjugation could be introduced to bind onto the vacated sites, rendering those regions cell adhesive once again. Although dynamic surfaces are rapidly increasing in capability, this field is still relatively new, and to this point there has yet to be much success in applying these tools to study the biology of the cellular microenvironment. One recent example, reported by Jiang et al. (2005), employed a patterned SAM to constrain adhered cells to a teardrop shape. After applying an electric potential to electrochemically desorb the SAM, the cells were observed to migrate in the direction of the blunt end of the teardrop. Studies such as these will help to provide insight to how cells respond to dynamic changes in their local environment.

CONCLUSIONS AND FUTURE DIRECTIONS While the field of regenerative medicine has blossomed in the past decade, there are still major obstacles that must be overcome before the dream of functional artificially engineered tissues can be achieved. Understanding how to use artificial environments to control cell function and finding suitable scaffolds to provide this control are keystones for future endeavors in tissue engineering. Microscale technologies undoubtedly will provide some of the tools necessary to achieve these goals. Appropriately directing cell fate and function remains a critical challenge in engineering tissue constructs. Microfabricated systems as those presented here will provide an important tool in elucidating the mechanisms underlying how extracellular cues can be used to drive cell function. While much is known about how the chemical properties of these cues affect various intracellular signaling pathways, microfabricated systems have only recently revealed that physical and mechanical cues are also equally important. It is now being realized that a cell can sense the physical and mechanical parameters of its surroundings through the cytoskeleton and through numerous intertwined intracellular signaling pathways. Understanding how the spatial presentation of soluble, adhesive, and mechanical cues is integrated within cells will be a critical challenge of the near future of regenerative medicine. In addition, one must not overlook the fact that the body is composed of many different types of cells, each of which has a distinctive response that defines the phenotype of that particular cell type. A deeper understanding of how each type of cell behaves in the context of other cells and in response to multiple cues remains an enormous task that may be partially simplified by the miniaturization and screening approaches offered by microscale technologies. Moreover, the recent shift of the focus of the biomedical community toward stem cells for regenerative applications further highlights the need for microculture systems to study these rare and valuable cells. Thus, microfabricated systems may provide a critical set of tools to engineer stem cells for regenerative medicine applications.

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While understanding the cellular component of tissue engineering is a major endeavor, this only constitutes one-half of the challenge, as cells are combined with different types of scaffolds to yield the desired tissue product. Microengineered tools may impact tissue engineering by increasing the physical complexity and spatial resolution of scaffold materials. As described in this chapter, microfabrication technologies can easily generate features with spatial resolution on the micrometer scale. With improvements of these tools and the advent of nanotechnologies, generation of devices with subcellular-scale resolution is on the immediate horizon. Here, we have reviewed how these tools can be used to control the geometric presentation of soluble, adhesive, and mechanical cues. In addition, microfabrication technologies can also be used to include other features such as substrate topology, and mechanically or electrically activated components that can interact with cells. It remains to be seen how and when these functionalities can be applied to scaffold engineering for regenerative medicine. The integration of all of the above elements into a microfabricated scaffold that can support the growth and maintenance of specified cellular phenotypes, and most importantly a desirable multicellular functionality will be critical to the design of novel, serviceable tissue-engineered constructs. While currently only a budding area of study, this field offers exciting new potential to engineer devices on a level of complexity that would otherwise not be possible.

ACKNOWLEDGMENTS The authors declare no competing financial interests. This work was supported in part by the NIH, NSF, David and Lucile Packard Foundation, and Desphande Foundation. W.F.L. acknowledges the NSF for financial support and E.E.H. acknowledges a NIH NRSA postdoctoral fellowship.

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31 Applications of Nanotechnology Benjamin S. Harrison

INTRODUCTION Regenerative medicine represents one of the greatest challenges in modern day science and medicine. With the goal of repairing diseased or damaged tissue to restore normal function, it has become increasingly apparent that our current understanding of biology is insufficient to reach such a lofty goal. Numerous implants, cell therapies, and engineered tissues that have been developed so far indicate that our current understanding of the superstructure and the microstructure of tissues is no longer adequate to create truly regenerative therapies. Understanding and controlling the underlying nanostructures in cells and the extracellular matrix represent key pieces in reaching the goals of regenerative medicine. Nanotechnology is a bottom-up approach that focuses on assembling simple elements to form complex structures. At the nanometer scale, where many biological processes operate, nanotechnology can provide the tools to probe and even direct these biological processes. This means that nanotechnology can be used for repairing damaged parts, curing diseases, and even monitoring and responding to the needs of the body. Cells and the extracellular matrix possess a multitude of nanodimensionality that interplays with one another. Cells, typically microns in diameter, are composed of numerous nanosized components all working together, creating a highly organized, self-regulating machine. For example, the cell surface is composed of ion-channels that regulate the coming and going of ions such as calcium and potassium in and out of the cell. Enzyme reactions, protein dynamics, and DNA all possess some aspect of nanodimensionality. These nanodimensional components control how cells produce the extracellular matrix (ECM) including the ECM composition and architecture. The ECM that the cell interacts with also abounds with nanosize features that influence the behaviors of other cells and tissues. These nanosized features, such as fiber diameter and pores, along with the intrinsic properties of the matrix itself, control the mechanical strength, the adhesiveness of the cells to the matrix, cell proliferation, and the shape of the ECM. Nanotechnology will provide regenerative medicine with the new multifunctional tools for imaging and monitoring the regenerative process and controlling the structure of the ECM. This is an exciting feature of nanotechnology in that it should not be thought of as a single object that has only one function. The size of nanomaterials allows multiple components to be combined and contained in a single nanocarrier unit. In addition, the small size allows nanomaterials to probe biological processes with minimal intrusion. Included in or on this nanocarrier can be therapeutic, targeting, contrast, and/or biocompatibilizing components which can be designed to meet a particular need. A description of the various components of a nanocarrier can be found in Table 31.1. Individual components, such as a therapeutic agent, can be exchanged or removed to create the desired effect without necessarily compromising the remaining components. This is significantly different compared to drug synthesis, for example, in which a single change can dramatically influence the pharmacological kinetics and potency of the drug. Besides realizing the potential of the small size, exploration into the nano-world is revealing unique quantum phenomena that only occur on the nanoscale. These quantum effects could be exploited to provide

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Table 31.1 A typical nanocarrier of image contrast and/or therapeutic agents is composed of six components Binder

All the different components are held together using a binder. The binder may be an inert piece of the nanocarrier, however many times it also serves another purpose. The binder may also be the image contrasting agent. For example, iron nanoparticles and QD serve as the core for the attachment of the other components. Polymers such as polyglycolic acid may serve as the binder of the therapeutic and also the biocompatiblizing agent.

Biocompatibilization

This component makes the nanocarrier compatible with the biological environment. It does this by minimizing aggregation of the nanocarrier and increasing the lifetime of the carrier by avoiding the defense mechanisms of the biological systems such as the reticuloendothelial system.

Imaging contrast

This component provides the means for imaging modalities to observe the nanocarrier. These contrasting agents may be observed using optical, magnetic, ultrasound, and scintillating methods.

Sensor

The sensor or trigger is used to alter the behavior of the nanocarrier once it has been deployed. For example, near-infrared light or electromagnetic radiation may be used to accelerate the release of a therapeutic or cause rapid localized heating as part of a therapy. Chemical sensors such as polymers that are pH or ion sensitive may also provide feedback to the nanocarrier in the delivery of its payload.

Targeting

This component provides the means of driving the nanocarrier to its desired location. There are two types of targeting: passive and active. Passive targeting incorporates only nonspecific targeting agents which may be useful for determining microenvironment permeability or areas of increased angiogenesis. Active targeting uses ligands or antibodies that bind to specific receptors at the target site. Active targeting aids in obtaining higher concentrations of therapeutics and contrasting agents at the desired site. Also, multiple targeting agents can be bound to the nanocarrier, allowing lower binding affinity molecules to be used to increase binding probabilities.

Therapeutics

Bioactive agents such as drugs or DNA are typical payloads of the nanocarrier. Drugs that are incapable of penetrating cellular membranes or hydrophobic drugs which cannot be administrated systemically by themselves can be contained within the nanocarrier awaiting release in a controlled manner. Other novel properties of nanoparticles have also shown promise as hyperthermic agents.

new approaches to regenerative therapies. Such quantum effects result in high optical absorptivities coupled with large photostabilities, or unusually magnetic properties within nanomaterials. Already such nanomaterials are being explored to enhance cellular imaging (Zhang et al., 2002; Medintz et al., 2005). Besides imaging, these quantum effects will allow novel methods of drug delivery, using light, electric or magnetic fields as drug delivering triggers. While these may involve exotic materials or elements which never would be found to naturally occur in the body, the expectation of nanotechnology should only be to serve as a temporary aid to direct the regeneration process and so should be developed with a relatively short-term use in mind. Nanotechnology’s impact on regenerative medicine will be through the development of multifunctional tools to enhance the performance or capabilities of implants, cell therapies, and tissue engineering. These advances will be the result of understanding and exploiting the underlying nanodimensionality of life. Nanotechnology will play a role in the ongoing development of tools for controlling the cell and its support matrix. Since nanotechnology is at the interface of modern physical science and medicine, new and unconventional ideas will develop, capable of bringing about major revolutions in science and medicine. Therapies developed using nanotechnology will someday minimize or eliminate the side-effects of drugs through targeted delivery and will

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provide real-time, and even non-invasive, monitoring of the disease and tissue repair. In this chapter, we will examine the impact nanotechnology will have on regenerative medicine related to cellular therapies and biomaterial control which play an important role for implant design and tissue engineering.

NANOTECHNOLOGY AS A MULTI-FUNCTIONAL TOOL FOR CELL-BASED THERAPIES Cell-based therapies, in particular those based on stem cells, have generated much excitement, both in the media and scientific communities, and are one of the most promising and active areas of research in regenerative medicine. One method of accelerating the pace of research is through the creation of multi-functional tools which allow the improved monitoring and modifying of cell behavior. While cancer-related research is a large part of the nanomedicine effort, there is great potential for applying nanotechnology in cell-based therapies for regenerative medicine. For example, with the enormous self-repair potential of stem cells, it is important to be able to locate, recruit, and signal these cells to begin the regeneration process. Improving non-invasive monitoring methods is particularly desirable since current methods of evaluating cell treatments typically involve destructive or invasive techniques such as tissue biopsies. Traditional non-invasive methods such as magnetic resonance imaging (MRI) and positron emission tomography (PET), which rely heavily on contrast agents, lack the specificity or resident time to be a viable option for cell tracking. However, in vitro and in vivo visualization of nanoscale systems can be carried out using a variety of clinically relevant modalities such as fluorescence microscopy, single photon emission computed tomography (SPECT), PET, MRI, ultrasound, and radiotracing such as gamma scintigraphy. Nanoparticulate imaging probes include semiconductor quantum dots (QD), magnetic and magnetofluorescent nanoparticles, gold nanoparticles, and nanoshells among others. While there are currently few examples of nanotechnologies being applied to the understanding of important processes in tissue regeneration, relevant uses of nanoparticles for regenerative medicine such as monitoring angiogensis (Winter et al., 2003) and apoptosis are appearing (Jung et al., 2004). QD is one type of nanomaterial that is receiving special attention. QD are inorganic nanocrystals that possess physical dimensions between 2 and 10 nm. The emission wavelength is controlled by the size of the nanocrystal and can be tuned throughout the visible spectrum to the near-infrared region (670 nm). Early live cell experiments using fluorescent QD sparked interest in using nanoparticles for immunocytochemical and immunohistochemical assays as well as for cell tracking (Akerman et al., 2002; Tokumasu et al., 2003; Sukhanova et al., 2004). A significant advantage for QD is their increased photostability (typically 10–1,000 times more stable) compared to organic dyes. This allows QD and the cells or proteins attached to them to be tracked over longer periods of time. Tumor cells labeled with QD have been intravenously injected into mice and successfully followed using fluo-rescence microscopy (Gao et al., 2004; Voura et al., 2004). As passive imaging agents, QD can be used for image microvasular in animals since polyethylene glycols (PEG)-coated QD injected into mice have shown good tissue perfusion and appear to be biocompatible (Ballou et al., 2004). QD represent just one novel class of nanomaterials whose ability to aid in long-term imaging of cells would help develop better regenerative therapies. Other nanoparticles are showing promise for optical cell tracking and imaging. For instance, nanosized tubes of carbon known as carbon nanotubes possess optical transitions in the near-infrared that can be utilized for tracking cells. The infrared spectrum between 900 and 1,300 nm is an important optical window for biomedical applications because of the lower optical absorption (greater penetration or depth of light) and small auto-fluorescent background. Like QD, carbon nanotubes possess good photostability and can be imaged over long periods of time using Raman scattering and fluorescence microscopy. However, unlike QD, which are typically composed of heavy metals such as cadmium, carbon nanotubes are made of carbon, an abundant element in nature. Carbon nanotubes possess large aspect ratios with nanometer diameters and lengths ranging

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from submicron to millimeters. These tubes can contain a single wall of carbon or multiple walls (typically 3–10) of carbon, commonly called single-wall carbon nanotubes (SWNT) or multi-wall carbon nanotubes (MWNT), respectively. SWNT dispersed in a pluronic surfactant can be readily imaged through fluorescence microscopy after being ingested by mouse peritoneal macrophage-like cells. The small size of the SWNT makes it possible for 70,000 nanotubes to be ingested where they can remain stable for weeks inside 3T3 fibroblasts and murine myoblast stems cells (Cherukuri et al., 2004; Heller et al., 2005). Having such a high concentration of carbon nanotubes within a cell without distributing the cell behavior means such probes could be used for studying cell proliferation and stem-cell differentiation, even through repeated cells. While such nanomaterials have yet to reach clinical applications, it does show the potential for non-invasive optical imaging. Along with optical contrast agents, magnetic nanoparticles also have been used to track cells and report on cell behavior. Many nanoparticle contrasting agents are based on superparamagnetic iron oxide nanoparticles and some have already been approved as clinical MRI contrast agents. When placed into a magnetic field, magnetic nanoparticles create perturbations of the external field that significantly reduce the spin–spin relaxation time (T2) of the nearby environment generating MRI contrast. Typically, these probes consist of a magnetic iron oxide core that is surrounded by a biocompatibilizing material such as dextran. Sizes of these particles can range from one nanometer to hundreds of nanometers in diameter. When used in conjunction with HIV-Tat and polyArginine peptides, these particles are readily taken up by many cell types (Dodd et al., 2001; Zhao et al., 2002). For example, stabilizing pressure input orthosis (SPIO) labeled rat mesenchymal stem cells injected into rats could be imaged and tracked to the liver and kidneys (Bos et al., 2004). Apoptosis is commonly detected by using the binding of annexin V to externalized phosphatidylserine. This binding event is the basis of optical and radiolabels methods for detecting apoptotic cells and can be bound to iron nanoparticles for sensing using MRI. It has been demonstrated that tumor-bearing mice injected with SPIO particles bearing apoptotic sensing proteins showed a sharp decrease in the T2* weight image corresponding to the location of the tumor (Zhao et al., 2001). This demonstrated that nanomaterials can be used to create high specificity MRI contrast agents for apoptotic cells. Such results are encouraging because they show that nanomaterials can be used for not only imaging the physical location of cells, but also providing information on the biological state of cells. While MRI has revolutionized our way of visualization in vivo, allowing cells to be tracked non-invasively, it is difficult to quantify the MRI signals and provide real quantification of cell numbers. The difficulty arises because MRI contrasting agents that are based on paramagnetic gadolinium and iron metals are not directly detected by the scanner but are indirectly detected by their influence on surrounding water molecules. However, the use of perfluoronated nanoparticles has recently been shown to be a new way to provide quantitative numbers to MRI since the fluorine nuclei (19F) can be directly detected (Morawski et al., 2004; Ahrens et al., 2005). Since endogenous fluorine is negligible in the body, 19FMRI is capable of directly detecting fluorine against a dark background similar to radiotracers and fluorescent dyes. While this has been demonstrated with dendritic cells, similar results should be obtainable using other cell types. Nanotechnology can provide powerful new tools for non-invasive tracking of cells in engineered tissues. As was also mentioned in the outset, the real benefits of nanotechnology are the multifunctional tools that it can bring. Besides imaging enhancements, nanotechnology can produce carriers for delivery of therapeutics for aiding the regeneration process. For example, biodegradable nanoparticles can deliver drugs, growth factors, and other bioactive agents to cells and tissue (Panyam et al., 2003). Nanodelivery vehicles possess three distinct advantages over conventional drug delivery methods. First, nanoparticles, due to their small size, are able to bypass biological barriers such as cell membranes and the blood brain barrier (BBB) allowing greater concentrations of therapeutics to be delivered. Second, nanocarriers can be functionalized with active targeting agents to allow selective delivery of bioactive active agents. Third, drug delivery systems can incorporate nano-triggers for non-invasive delivery of therapeutic agents. These sensitive triggers can be activated using in vivo signals such as

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pH, ion concentration, and temperature or external sources such as near-infrared light, ultrasound, and magnetic fields. As nanotechnology progresses, new nanomaterials and techniques are being developed regarding cellular imaging and drug delivery which will better equip those practicing regenerative medicine to reach their goals. Cellular therapies for regenerative medicine would benefit from nanotechnology since tracking of implanted cells would provide the means to better evaluate the viability of engineered tissues and help in understanding the biodistribution and migration pathways of transplanted cells. Nanotechnology would also allow better and more intelligent control of the bioactive factors which can influence cellular therapies. The potential of nanotechnology for impacting regenerative medicine is great, creating the hope of individualized and targeted therapies.

NANOTECHNOLOGY AS A MULTI-FUNCTIONAL TOOL FOR BIOMATERIAL CONTROL Biomaterials play an important role in regenerative medicine because they make up a large component of implants and tissue scaffolds. Increasing evidence shows that the nature of the biomaterial greatly affects the long-term success of biomedical implants and the short-term wound healing response. Substrate features such as the chemical composition and surface morphology affect the viability, adhesion, morphology, and motility of cells. Therefore, controlling the three-dimensional structure and surface composition of a biomaterial is important to promoting normal tissue growth or minimizing foreign body response. To illustrate the importance of controlling the biomaterial surface, one can examine the use of implants to repair bone defects. Currently, there are several strategies for repairing large bone defects including using implants made of metal, plastic, ceramics, or graphing of tissue. However, there are limitations to these biomaterials. Autographs can be expensive, difficult to handle, and may have physical limitations in their use. Allographs are also expensive and carry additional risks of an autoimmune response and disease transmission. While metal and plastics mitigate many of the aforementioned risks, implants made from these materials instead of integrating with bone often form soft undesirable fibrous tissue. This is especially true with surfaces that are uniform and non-porous. This mechanical mismatch between tissue leads to the wear and tear of the implant that either aggravates or in some cases leads to cell death in nearby tissue causing implant failure. However, the inclusion of nanosized particles into implant materials, for example, has shown to increase osteoblast adhesion (Kay et al., 2002). While this may be partially due to increased surface area, other factors may be involved, such as controlling protein adsorption. For instance, on carbon nanofiber surfaces, osteoblast adhesion was greater than other competitive cell types; possible due to the nanofibers’ high surface energy and small diameter fibers and aligned structure (Price et al., 2003). Taking advantage of the electroactive properties of carbon nanotubes blended into a biomaterial, new cell behaviors can be obtained. For example, this has been accomplished by applying an alternating current to a nanocomposite of polylactic acid and multi-walled carbon nanotubes, resulting in an increase in osteoblast proliferation by 46% and a greater than 300% increase in calcium production (Supronowicz et al., 2002). Also, upregulation of collagen I (a major component in organic bone formation), osteonectin, and osteocalcin was observed. Such results suggest that nanocomposites would accelerate the bone regeneration process. Nanomaterials, like carbon nanotubes, are part of a growing new class of multifunctional biomaterial– smart biomaterials. Unlike passive structural biomaterials, smart biomaterials are designed to actively interact with their environment either by responding to changes in their surroundings or by stimulating or suppressing specific cellular behavior. They can change their shape, porosity, or hydrophilicity based on changes in temperature (Gan et al., 2001), pH (Bulmus et al., 2003), or external stimuli such as electric (Lahann et al., 2003) or magnetic fields (Jordan et al., 1999). Such control of the biomaterial behavior through nanotechnology could create a major shift in the way one uses biomaterials. Examples of some techniques used for creating

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Table 31.2 Examples of tissue scaffolds created using nanofabrication techniques Technique

Tissue scaffold prepared

Lithography Electrospinning

Nerve (Gabay et al., 2005) Heart (Zong et al., 2005) Nerve (Yang et al., 2005) Bone (Fujihara et al., 2005) Nerve (Ellis-Behnke et al., 2006) Bone (Du et al., 1999; Kikuchi et al., 2001; Liao et al., 2004; Kim et al., 2006) Bladder (Thapa et al., 2003; Thapa et al., 2003; Pattison et al., 2005) Bladder (Thapa et al., 2003; Thapa et al., 2003; Pattison et al., 2005)

Self-assembly Polymer demixing Solvent casting Salt leaching

nanostructured surfaces for tissue engineering are shown in Table 31.2. The current paradigm to tissue regeneration is to isolate a patient’s cells and then incubated outside the body and finally place or seed the cells onto scaffold-like biomaterials before implantation. This method of engineered tissue using two different cell types has met with great success (Atala et al., 2006). Ideally, one would want to directly implant a biomaterial into the patient that would then selectively recruit the correct cell types to the correct location in the tissue. This method would be especially important for organs with very elaborate structures. A smart biomaterial would allow the correct cells and supporting vascular to grow onto the scaffold in the correct orientation without permitting inflammatory cells and fibroblasts, which typically wall off any implants, to become established on the biomaterial. Such smart biomaterials would be a boon to regenerative medicine. Another area where controlling biomaterial surfaces through nanotechnology can make an impact on regenerative medicine is stem-cell differentiation for engineered tissue. Currently, concoctions of expensive growth factors are used to guide the differentiation of stem cells down certain lineages. With the ability to control the surface morphology and chemistry at the nanoscale, nanobiomaterials may eliminate the need to culture different cell types for reassembly into an engineered tissue as they can recruit the body’s own stem cells and differentiate them into the correct phenotype (Silva et al., 2004). Biomaterials play an important role in regenerative medicine through their use in implants and tissue scaffolds. Nanotechnology is posed to provide the tools for rapidly increasing the pace of biomaterials development. Through the ability to control the nanostructure of a biomaterial, better understanding and control of cell behaviors will result, creating better regenerative therapies. The timeline of the impact of nanotechnology on biomaterial development as it relates for regenerative medicine will first be felt through betterperforming, longer-lasting implants and will eventually give way to smart biomaterials, which can be implanted and can direct the regenerative process at the cellular level.

CONCLUSION As nanotechnology continues to grow, it will provide new and powerful tools which will revolutionize regenerative medicine. The most significant impact nanotechnology will have on regenerative medicine is that it will help in providing a detailed understanding and control of biology. Already, nanotechnology, albeit a young technology, has demonstrated significant advances over traditional imaging, sensing, and structural technologies. Many of these advantages stem from the capability of nanomaterials to be multi-functional. These advances help in tackling one of most significant challenges we face in designing new biomedical technologies – targeting biological functions while at the same time avoiding nonspecific effects. While there have been challenges for some time, nanotechnology provides us with the means to successfully negotiate these challenges and create new innovations in regenerative medicine.

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32 GeneChips in Regenerative Medicine Jason Hipp and Anthony Atala

INTRODUCTION Stem cells will offer revolutionary therapeutics for regenerative medicine that is curative, rather than delaying the disease progression. In addition, they will serve as models to studying the genetic mechanisms of regeneration and provide novel insights into cancer and degenerative diseases such as diabetes, Alzheimer’s, and Parkinson’s. While much is known about the potential of stem cells, very little is known about how they grow and differentiate. With an inability to control their growth and differentiation, their unique ability to form any tissue type also becomes their limitation. While there are only a handful number of genes that are known to be specific to stem cells (Brivanlou et al., 2003), an understanding of their signaling networks responsible for differentiation is essential for their therapeutic application. Although the sequencing of the human genome was recently accomplished, the function of a majority of these genes is unknown. Of the estimated 20,000–25,000 genes (Stein, 2004), an understanding of which one provide stem cells with their unique properties of “stemness” (self-renewal and pluripotentiality – the ability to become any tissue of the body) would be of tremendous value for clinical applications. While there are millions of patients that are suffering from degenerative diseases that could potentially be cured by stem cells, new technologies must be applied to quickly and efficiently answer these questions and provide for this unmet medical need. With their ability to monitor the gene expression levels of almost every known and unknown, GeneChip technology could provide the answers (Lockhart et al., 1996; Lipshutz et al., 1999). GeneChips are miniature platforms with approximately 1 million 25 base nucleotide sequences that measure the transcriptional expression levels of 47,000 transcripts and variants including 38,500 well characterized human genes (HG-U133 Plus 2.0, www.affymetrix.com). This comprises of almost every known gene in the human transcriptome – the mRNA equivalent of the human genome (we will use the Mendellian definition of gene as a unit of inheritance often inferring it in its transcriptional form (non-coding/coding RNA), and will further define it when necessary). Performing a GeneChip experiment is like snapping a picture of a cell’s mRNA (transcripts), thus giving one a static view and measurement of gene expression inside the cell. By taking multiple “pictures” and comparing them to one another, one can gain insight into the kinetics of a cellular process such as differentiation, or can create a “transcriptional signature” to be used to compare and contrast different stem and somatic cell types. What we may lose in specificity when compared to reverse transcriptase polymerase chain reaction (RT-PCR), we do gain in sensitivity considering we are able to monitor the expression levels of over 660,000 genes (over 30 GeneChips  22,000 genes/GeneChip were used in the experiments described below) – what would take years to analyze by RT-PCR now takes a few days. Thus, by making the appropriate comparisons

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GeneChip databases Experimental design GeneChip platform Sample preparation Expression analysis Interpretation Application

Figure 32.1 An overview of the chapter, illustrating the multiple components of a GeneChip experiment.

between stem and somatic cells, we can begin to efficiently, in a high throughput manner, ascribe function by association to thousands of known and unknown genes. The ultimate goal of regenerative medicine is to use these cells as a resource to unlock the potential of regeneration, whether it is by directly regenerating tissue through the differentiation of stem cells, or indirectly, by regenerating the organ itself in vivo through genetic manipulation, new chemical entities (NCE), or with biomaterials. In the first half of this chapter, we will discuss the many components involved in a GeneChip experiment (Figure 32.1), illustrating the many variables at each step, and describing our protocol for analysis. We will then describe how we and others are applying GeneChips to regenerative medicine.

PROTOCOL GeneChip Databases We believe the first step of a GeneChip experiment should begin with a thorough search of the publically available GeneChip databases. We usually begin by searching the following databases: NCBI GEO (Barrett et al., 2005), ArrayExpress (Parkinson et al., 2005; Sarkans et al., 2005), Stanford (Ball et al., 2005), and Stembase (Ball et al., 2005; Perez-Iratxeta et al., 2005), Public Expression Profiling Resource (Chen et al., 2004). These databases contain hundreds of GeneChip and other microarray data files and often provide raw and analyzed data files for download. This will not only prevent you from repeating an experiment, but will provide tremendous insight into how others are designing and implementing GeneChip technology. Another advantage of this technology is that a GeneChip file is like a permanent archive which can be constantly re-analyzed and re-interpreted in the context of new computer or biological advances. Furthermore, we recommend even designing your experiments to build off the existing publically available GeneChips for direct and indirect comparisons. This will also serve as an efficient and inexpensive way to cross validate your own data, a method we refer to as in silico post hoc. Because in silico comparisons are not limited by biological, time, and financial constraints, we have found that the more novel the comparisons, the more exciting and unexpecting are the results. One is therefore not limited by biological, time, or financial constraints, but rather by one’s own creativity. We believe this ability to perform diverse comparisons “in silico” and integrate data sets across multiple disciplines is the true benefit of this technology.

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Experimental Design Defining the Biological Difference If the cornerstone of a good GeneChip experiment is in the experimental design, and the experimental design is only as good as the biological question, then the first and most important step in a GeneChip experiment is in defining a biological question. The power of GeneChips lies in their ability to provide a global genetic explanation for a biological difference – the larger the genetic difference the better. It is important to remember that the answer provided by a GeneChip experiment may consist of hundreds if not thousands of genes, therefore valuable answers can easily be hidden within the mass of data. To ensure that your genetic explanation correlates with your biological question, we recommend to clearly define your biological difference with functional studies, if applicable or with other methods such as RT-PCR, Western’s, or immunocytochemistry (ICC) before running your GeneChips. For example, if we are comparing the gene expression profiles of stem cell and stem cell derived osteoblasts, we made sure that they were osteoblasts by measuring their calcium production with Von Kossa while others such as Palmqvist et al. (2005) are using functional measures of pluripotency before performing their GeneChip experiments. Hence, we like to think of this as performing the traditional GeneChip “post hocs” before, not after our experiments. Not only does this help us better refine our experimental design before we run our GeneChips, but it provides inherent controls in analyzing the quality of our data sets. The question of how many replicates to do is often dependent on the difficulty of generating enough sample RNA and one’s budget. The more GeneChips used in the analysis, the more statistical significance there will be in identifying differentially expressed genes. Usually, a minimum of three is recommended. Our philosophy has been to plan on running more GeneChip experiments, harvesting more RNA than necessary, and be willing to run the additional samples if necessary, depending on the quality and types of genes identified. GeneChip Platform There are many factors that influence which platform to use for GeneChip analysis. Depending on the expertise of your laboratory, one can print your own microarray platforms or use commercially available GeneChip platforms. The benefit of printing your own GeneChips is that it can be cheaper in the long run if you perform many microarrays and require less sample RNA. Or, one can use serial analysis of gene expression (SAGE) techniques and make cell type specific platforms where RNA is isolated from the cell type of interest and a cDNA library is created and printed on a platform (Velculescu et al., 1995; Saha et al., 2002). However, the disadvantages of these techniques are in reproducibility. Minor errors can be introduced generating the probes, their printing, hybridization, labeling of RNA, and validation, thus, many samples must be run to identify potential errors. With the commercially available platforms, these problems have already been addressed. The advantage of using commercially available platforms is that they can be performed by most every laboratory disregarding experience and expertise – all one needs to do is generate the sample (tissue/cells). While most groups isolate their own RNA, there are even some commercial enterprises that will do this for you such as Genome Explorations (www.GenomeExplorations.com). The choice of GeneChip platform is influenced by the level of depth you seek in your genetic explanation and how much you are willing to pay. The next question one has to decide is if you want to screen the entire transcriptome or specific pathways. If one is already interested in a particular pathway, Superarrays would be ideal because they are pathway specific (www.superarray.com). If one is interested in covering as many genes as possible, then Affymetrix, Agilent, or GE Healthcare platforms would be ideal (www.agilent.com, www.GEHealthcare.com). The important difference between the platforms that scan the whole transcriptome is single (Affymetrix) versus dual channel

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chips (Agilent, GE Healthcare). Dual channel means that the two samples are labeled with different color dyes; however, the disadvantage is that one can make only one comparison and this has to be decided upfront. Affymetrix uses single channel meaning that the sample is labeled with only one dye, and the difference in intensity of this dye is measured. This allows one to perform infinite sample comparisons amongst other Affymetrix chips and particular for the field of regenerative medicine, is one of the reasons why we chose the Affymetrix platform, as we will demonstrate later in chapter. Another advantage of the Affymetrix platform is the number of programs that are built based on the Affymetrix chip design. Lastly, another advantage of the Affymetrix GeneChip is its platform design that allows for the absolute detection of whether a gene is identified as present or absent (described below). Sample Preparation It is important to have a pure cellular or tissue population for GeneChip analysis to limit contaminating RNA from other cell types. Cell lines are much easier to use than biopsies because of the possibility of contaminating cell types. Laser microdissection can be used to remove contaminating cell types, but with tumors, where there are mixtures of cells (with different ratios) and cells cycling at different rates for the genetic difference identified by the GeneChip maybe due to the different cell types rather than true biological difference. The quantity and quality of RNA isolated will directly influence the quality of GeneChip generated. Because RNA has a free 2-OH which can serve as a nucleophile, one must be very careful to prevent RNA degradation during isolation. There are commercially available kits that help purify the RNA, and solutions to spray on the equipment (RNA ZAP) to remove proteases. It is recommended to check the quality of RNA by spectrophotometric analysis (A260/280 should be 1.8–2.1 in TE buffer or 1.6–1.9 in water) and by running an RNA degradation gel for the identification of two bands (18s and 28s peaks with a ratio of 1:2). This extra effort will save you money in the long run because samples submitted for GeneChip analysis are usually checked by gel capillary electrophoresis for RNA quality, which costs a couple hundred dollars, if your RNA does not pass, you will have to pay the fee and re-submit your sample and repeat. One must be extra careful when dealing with small amounts of RNA and using RNA amplification techniques. In these circumstances, even a small amount of RNA degradation will be amplified. In many experiments, the amount of RNA is the limiting factor; the ability to generate large number of stem cells in vitro is therefore another advantage and makes applying GeneChip technology more advantageous. The advantage of using Affymetrix is that their labeling and hybridization protocol has been standardized and thoroughly tested. It is rare for a laboratory to perform its own labeling and hybridization. This is usually done by a core facility. In addition, many prefer to use core facilities rather than doing it themselves because of the importance in having experience which introduces less variability. We have even noticed that within some core facilities, others are more experienced and sometimes have to re-submit samples to re-do the analysis. Briefly, the core facility will then reverse transcribe the RNA (1–15 μg) into double stranded cDNA with a T7 olig(dT) promoter primer (Affymetrix). It is then transcribed with biotin labeled nucleotides making cRNA. This is then fragmented and hybridized to the GeneChip. Because the RNA is reverse transcribed, 11 probes (25-mer nucleotide sequences) are chosen to match (perfect match, PM) the gene transcript. Each of these PM probes is paired with a miss match (MM) probe that differs in that the 13th nucleotide sequence is mutated. The core facility will do a quality control check by looking at the number of present genes, 5 and 3 ratio, and signal to background ratio and provide you with a set of files: the .DAT file is the scanned array image file, .CEL file is the cell intensity data file derived from the .DAT and contains the PM–MM values, .CHP is the analyzed and saved .CEL file, and a summary report file, .RPT file (Affymetrix).

566 CELLS AND TISSUE DEVELOPMENT

Expression Analysis Overview of Programs Data analysis by the biologist with minimal computer training can become overwhelming with the many different ways GeneChip data can be analyzed, combined with complex computer programs (sometimes requiring code writing, in addition to converting and transferring files between different programs); which often becomes a self-limiting task. While there are commercially available packages, these often are associated with high costs and the different methods of analysis can be difficult to understand. Here, we will provide an overview of some of the more commonly used data analysis programs and discuss their application in a relatively inexpensive manner, catering to the biologists with limited computer expertise. The most commonly used programs to analyze GeneChip data are MAS5, dCHIP, and Robust Multichip Analysis (RMA). These programs differ in the way they normalize the data, which adjust for non-biological variability that could have occurred during the labeling and hybridization procedure. MAS5 is the program created by Affymetrix, and because of its size, usually runs on its own workstation (Affymetrix;Affymetrix). It determines gene expression intensity by applying a Tukey’s Biweight algorithm to determine probe set intensities based on the PM/MM ratio. It then calculates an expression value, and based on the p-value, whether a gene is present, absent, or marginally expressed. Before GeneChips can be compared to one another, they first need to be normalized, and MAS5 normalizes by picking specific regions within the GeneChip and adjusting the signal intensities for each probe to a user defined value. GeneChips that have been normalized to the same value can then be compared to determine differential gene expression (Affymetrix). Li and Wong (2001) noticed that some probes had an outlying expression value that was consistent across GeneChips and developed dCHIP. dCHIP uses a non-linear normalization method to remove outlying probe effects (noise) by first normalizing all GeneChips to the median intensity and then computes a model based expression index (MBEI) to estimate gene expression levels. The program is easy to use with a graphical user interface (GUI), available to the public for free, can be run off a laptop. Irizarry et al. (2003) compare RMA to dCHIP and MAS5 and demonstrated that RMA has better precision, particularly for genes with low expression levels – RMA provides a greater than five-fold reduction of the within-replicate variance as compared to dCHIP and MAS5, provides more consistent estimates of fold change, and provides higher specificity and sensitivity when using fold change analysis to detect differential expression. However, they noticed that MAS5 is more accurate than RMA, but believe this modest loss is worth the gains in precision (Irizarry et al., 2003a). RMA is unique in that it adjusts for background noise, performs a quantile normalization, transforms the data into log based 2, and then summarizes the multiple probes into one intensity (Bolstad et al., 2003, 2004; Irizarry et al., 2003a, b; Cope et al., 2004). It only uses the PM probes and ignores the MM probes which cause exaggerated variance (Cope et al., 2004). Quantile normalization was chosen because it has been shown to have the best performance and works by making the distribution of intensities at the probe level (rather than choosing a baseline GeneChip or standard intensity; Bolstad et al., 2003). RMA is available to the public for free and can be run off a laptop computer (www.bioconductor.org). GCRMA is an adaptation of RMA but differs in that it uses a model based background correction based on the G–C content and PM–MM (Wu and Irizarry, 2005). GCRMA was shown to be even more accurate than MAS5 without losing much precision. However, in terms of both accuracy and precision, RMA is best at high concentrations, and GCRMA is better at low concentrations (Wu and Irizarry, 2005). AffyPLM is another program that is similar to RMA but the summarization method differs in that it uses a probe linear model instead of a median polish (Affymetrix;Bolstad;Bolstad). A detailed comparison of different GeneChip analysis methods can be found at http://affycomp.biostat.jhsph.edu (Cope et al., 2004; Irizarry et al., 2005).

GeneChips in Regenerative Medicine 567

Data Analysis We believe the strength of MAS5 is in its present/absent call detection. Using a Turkey-Biweight formula, it determines if the PM intensities are greater than MM and then assigns a p-value to determine if a gene is either “present” (p  0.04), marginal (0.04  p  0.06), or absent (p  0.06) (Affymetrix;Affymetrix). We like to use this program to identify a “transcriptional signature” of genes that are flagged as present. Because of the large number of genes that are screened, we often refer to this as a “transcriptome,” and often use this as an absolute comparison amongst similar and different cell types – kind of like a large-scale RT-PCR experiment (we used this to assess the differentiation of human embryonic stem cell (HESC) into rat pancreatic extract (RPE), described below). Core facilities usually provide or may charge a fee for data analysis using the Affymetrix workstation. Some laboratories have their own workstations; however, since we are only using the present/absent/marginal detection calls, we ask the core facility for the MAS5 output and make our comparisons using Excel and MS Access. To quantify the relative differences in gene expression, we use AffylmGUI, the sister program of LimmaGUI (Wettenhall, 2004; Wettenhall and Smyth, 2004). The advantage of this program is that it reads the raw Affymetrix CEL files directly, summarizes gene expression values with either RMA, AffyPLM, or GCRMA, and has a built-in statistical program, Linear Modeling of Microarray data (Smyth) to identify differentially expressed genes, all using a GUI. LIMMA fits a linear model for every gene (like an analysis of variance (ANOVA) or multiple regression), then hypothesis tests and adjusts p-values for multiple testings (Smyth, 2004, 2005). It computes a moderated t-statistic for each gene where the standard errors are shrunk to a common value using a Bayesian model (Smyth, 2005). Although multiple methods of moderation can be chosen, the most common is the Benjamini and Hochberg’s method to control false discovery rate (Benjamini and Hochberg, 1995). In addition, it computes a B-statistic which is similar to Lonch and Speed (Lockhart et al., 1996), however is reformulated using a moderated t-statistic in which posterior residual standard deviations are used in place of ordinary standard deviations (Smyth, 2005). Running the program is relatively simple and quick. One begins by creating a text file that contains a contrast matrix (which identifies the CEL files) and a design matrix (which associates the CEL files with contrast groups). Once the CEL files are incorporated, you then have the choice of summarizing the gene expression values with RMA, GCRMA, or AffyPLM. A linear model is then computed, and the desired contrasts are made and the results can then be opened in an Excel. The output file consists of columns representing the Affymetrix probe set (gene identification), M-value (fold change in log based 2); A-value (average signal intensity), t-statistic, p-value, and B-value (Figure 32.2). An FDR adjusted p-value of less than 0.001 means that those genes that are selected are expected to have a proportion of false discovery that is controlled to be less than 0.1% (REF). A B-value of 2 is an odds of differential expression of 4, which means that there is an 80% chance of differential expression. Interpretation of GeneChip Data Once lists of differentially expressed genes are identified, one faces the most difficult task of a GeneChip experiment – deriving biological meaning and its application. The immense challenge of this is analogous to putting together the pieces of a puzzle in which there is no picture as a guide, or taking apart a car and a train and trying to describe how they are different. An initial analysis should begin with those genes that had the greatest B-value and/or FC. We like to annotate these genes with EASE (Dennis Jr. et al., 2003; Hosack et al., 2003) which is a GUI program downloaded onto your desktop. Affymetrix probe set (gene ids) can be pasted directly into the window, and then the different databases for annotation are chosen. We like to annotate our lists with the gene name, gene symbol, alias symbol, chromosomal location, geneRIF, and OMIM. In less than 30 s, an HTML file is generated which

M

A

t

p -value

B

fc

1/x

Gene name

Gene symbol

206268_at

5.906204

9.279683

17.31139

3.56E-05

7.724784

0.016675

59.97146

left–right determination, factor B

LEFTB

221245_s_at

3.913827

9.355386

15.45639

5.46E-05

7.458079

0.066347

15.0723

hypothetical protein DKFZp434E2135

DKFZP43 4E2135

203798_s_at

3.741892

7.154159

12.5054

0.000121

5.867624

0.074744

13.37894

visininlike 1

VSNL1

205626_s_at

3.581726

7.026003

18.44256

2.82E-05

8.667054

0.08352

11.97311

calbindin 1, 28 kDa

CALB1

210265_x_at

3.527591

10.19973

10.25064

0.000294

4.269039

0.086714

11.53216

POU domain, class 5, transcription factor 1

POU5F1

206012_at

3.433276

9.196597

9.390249

0.000436

3.760723

0.092572

10.80237

endometrial bleeding associated factor (left-right determination, factor A; transforming growth

EBAF

210852_s_at

3.384071

7.705703

12.69315

0.000114

5.979122

0.095784

10.44015

aminoadipate-semialdehyde synthase

AASS

204469_at

3.36158

8.065143

13.03747

0.000104

6.180218

0.097289

10.27866

protein tyrosine phosphatase, receptor–type, Z polypeptide 1

PTPRZ1

205625_s_at

3.330555

7.170094

19.18256

2.43E-05

8.975543

0.099404

10.05998

calbindin 1, 28 kDa

CALB1

206023_at

3.324269

8.132626

9.564309

0.000403

3.683753

0.099838

10.01624

neuromedin U

NMU

206653_at

3.3216

7.703198

12.49027

0.000121

5.856694

0.100023

9.997723

220184_at

3.262387

9.563476

19.20401

2.43E-05

9.044987

0.104213

9.595696

Nanog homeobox

NANOG

Figure 32.2 The output from affylmGUI, modified by adding fold and inverse fold change, gene name and gene symbol annotated with EASE. From left to right, ID (Affymetrix probe set which corresponds to the gene identification), M-value (fold change in log based 2), A-value (average signal intensity), t-statistic, p -value (adjusted for FDR), and B-statistic (log based 2- odds of differential expression), fc (fold change), 1/x (inverse of fold change), Gene Name and Gene Symbol using EASE annotation. Shown are the top 12 genes with the greatest fold change (B  0) upon HESC differentiation (Sato et al., 2003).

568 CELLS AND TISSUE DEVELOPMENT

ID

GeneChips in Regenerative Medicine 569

consists of a table of your genes with all the desired annotations. The advantage of this program is that it can be used to annotate thousands of genes in less than 30 s (a process that would take years to do by hand). We then take those genes that were identified as differentially expressed, usually a B-value greater than 0, and we try to identify biological themes using the “find over-represent gene categories” function. We do this by loading our list of genes into EASE, like above, and choosing a categorical system such as gene ontology (GO), chromosomal location, protein domains (there are over twenty different categories to chose from). GO is a collaborative effort to develop a controlled vocabulary (ontologies) that describe gene products in terms of their biological processes, cellular components, and molecular functions in a species independent manner (Ashburner et al., 2000). A reference file is then chosen, which contains all the possible categories for every gene on the GeneChip. EASE then compares your list to this reference file and depending on which statistical method you chose (Bonferroni or EASE score – a conservative variant of the Fisher exact probability) and will generate a p-value for the most over-represented themes, in less than 30 s. When comparing multiple data sets, we like to use the program GenMapp. GenMapp is a program which consists of hundreds of pre-made pathway maps (Dahlquist et al., 2002; Doniger et al., 2003). Lists of genes (in the Affymetrix probe set format) are loaded directly into GenMapp and assigned a color based on its expression. Hundreds of pathways can then be viewed where each gene is color-coded based on its cell type and direction of change (up- or down-regulated). This allows one to efficiently identify pathways with significant genetic changes (Figure 32.3). Information on each gene is integrated from multiple databases and is easily accessed by right or left clicking the name.

STEM CELL DIFFERENTIATION The most widely used application of GeneChips for regenerative medicine is in stem cell biology to identify “stemness” genes. Stemness genes will uncover the secrets of human development and regeneration and could potentially provide insights into degenerative and chronic diseases. A knowledge of these genes would have a significant impact on the field of regenerative medicine, for their potential ability to reprogram somatic cells (Cowan et al., 2005), or to design of novel cell culture conditions for the ex vivo expansion of progenitor cells for tissue engineering. With the proper experimental design, GeneChips can begin to explain how stem cells are capable of regenerating themselves, how they differentiate into any tissue in the body, or even explain the pathogenesis of cancer. By comparing stem cell data sets of diverse origins, we can begin to identify novel stem cell functions. Since little is truly known of how stem cells self-renew and differentiate, GeneChips, because of their ability to screen almost every known and unknown gene, can potentially make rapid advancements in not our understanding and their clinical application. HESC Using the Affymetrix U133a GeneChip, Brivanlou et al. compared undifferentiated HESC H1 line to their progeny that were differentiated for 26 days on matrigel with non-conditioned medium (Sato et al., 2003). They made their CEL files publically available and we analyzed them using the methodology described above. Briefly, the raw CEL files were incorporated into AffylmGUI, normalized with PLM, p-values were adjusted using FDR, and those genes with a B-value of greater than 0, we identified as differentially expressed. This identified 897 genes as being down-regulated, 1,269 genes as being up-regulated. As expected, those genes that were most significantly down-regulated were known embryonic stem cell (ESC) genes such as Oct-4, TDGF1, SOX2, Nanog, and telomerase (Figure 32.2). However, there were many other genes that had similar expression profiles such as LeftyB, Thy1, FGF13, Galanin, and DNMT3B. Since there are too many genes to analyze individually, we then clustered these genes based on their GO. The most over-represent theme using GO biological process is mitotic cell cycle, using SwissProt keyword is mitochondrion, GO cellular component is mitochondrion, and

Cell cycle Gene DNA damage checkpoint ARF

Growth factor

Growth factor withdrawal

SMC1L1 BUB1

hesup

MDM2

MAPK signaling pathway

e

SKP2

pgesod

RB1

TP53

hafsol

p27,57

CDKN1A

hesup

p

e

e

u p21

u

CDKN2A hafsol CDKN1B

hesdow

BUB3 MPEG1

p

e

GADD45A

hesup

BUB1B

CHEK1

hesdow

CHEK2

hesdow

YWHAG

hesdow

MAD1L1

hafsol

MAD2L1

hesdow

CCND3 CCND2

CCNE2 hafsol

CDK4

hesdow

CDK2

hafsol

u

CDK6

pgesod

CDK2

hesdow hafsol

CDC6

p

p

ABL1

CCNA2

CCNH CDK2

hesdow

CCNA1

hafsol

CCNA2

p

CDC2

SCF SKP2

CCNA1

hafsol

CCNE1

RB1

RBL1

hafsol

pgesol

C45L MCM

ORC

hafsol

p

p

RB1

hesdow

CCNB1

hesdow

CCNB2 CCNB3

hesdow

CDC2

hafsol

hafsol

p p

WEE1

E2F

CDC7

TFDP1

ASK

DNA

p

PLK1

hesdow

APC/C p

u

MCM (Mini-Chromosome Maintenance) complex

ORC1L

hesdowORC2L hesdow

MCM2

hesdowMCM3 hesdow

ORC3L ORC5L

hesdowORC4L hesdow

MCM4 MCM6

hafsol

ORC6L

hafsol

MCM5

CDH1

hafsol

pgesod

p

CDC14A CDC14B

hafsol

hafsol

TBC1D8 hesdow

ORC (Origin Recognition Complex)

hesdow

Ubiquitin mediated proteolysis

hesdow

p

HDAC

Securin

p

p

PKMYT1

hesdow

14-3-3

CDC25C CDC25B

hesdow

hafsol

PTTG2 PTTG1

CDC20

p

CDC25A

PTTG3

Separin

APC/C

hafsol

p

hafsol

u

MAD2L2 PCNA

ESPL1

hafsol

Apoptosis

e

p16,15,18,19

ATM

p

SCF

hesup

SMAD3 SMAD4

ATR

PRKDC

GSK3B

R-point (Start)

Condensin

EP300 TGFB1

p

hafsol

MEN

DNA biosynthesis DNA S-phase proteins

hafsol

hesdowMCM7 hesdow

S G1 Histone deacetylases HDAC1 HDAC2 HDAC3 HDAC4

hafsol pges

Transcription factor E2F

HDAC5 HDAC6

HDAC7A HDAC8

Legend

hesu

E2F1 E2F2

hafsol

E2F3

pges

E2F4 E2F5

RBL1 E2F6 UBE2F

hafsol

G2

M

HESC Up HESC Down PGESC Up PGESC Down HAFSC Up HAFSC Down No criteria met Not found

Figure 32.3 The GenMapp output of the genes that were up-regulated (lighter shade), down-regulated (darker shade) in HESC (red), PGESC (yellow), HAFSC (green). It demonstrates how active pathways can be easily identified, compared and contrasted amongst multiple disparate data sets.

570 CELLS AND TISSUE DEVELOPMENT

Author: Adapted from KEGG Maintained by: GenMAPP.org E-mail: [email protected] Last modified: 10/02/2002 Right-click here to see notes

GeneChips in Regenerative Medicine 571

GO molecular function is catalytic activity. Clustering your data set in this manner serves as a way to validate your data set in its biological context; for example, we would expect to identify numerous genes involved in mitosis as being down-regulated upon differentiation. However, we also identified other processes that might not have been as obvious such as genes involved in mitochondria and ATP metabolism which provide insight into energy demands of cell division, chromatin assembly, and remodeling. Furthermore, using EASE to cluster genes based on their ontology allows one to efficiently identify tissue engineering targets. For example, we loaded this list of down-regulated genes into EASE and selected those genes that clustered under GO biological process “Growth” and SwissProt “Growth Factor” and identified numerous growth factors that HESC secrete in an auto/paracrine manner such as FGF2, FGF13, EGAF, TDGF1, GDF3, and nucleostemin. PGESC We next applied a similar approach to understand non-human primate parthenogenetic embryonic like stem cells (PGESC) differentiation (Cibelli et al., 2002; Vrana et al., 2003). Parthenogenesis entails the in vitro activation of oocytes which stimulates their growth and development as if fertilized by sperm. These parthenotes cannot develop past the blastocyst stage, even if transferred in vivo. ESCs can be isolated from the inner cell mass (ICM). These cells express HESC markers, express telomerase, and are capable of differentiating into all three germ layers (Cibelli et al., 2002; Vrana et al., 2003). We also showed that they are capable of differentiating into neuronal progenitors and then onto neurons that express functional voltage dependent sodium channels. We then used the human U133A GeneChip to profile undifferentiated PGESC, their neuronal progenitor progeny. The CEL files were analyzed as described above. We identified 658 genes as down-regulated and 647 as up-regulated. As an inherent control, we identified the ES markers Oct-4, TDGF1, telomerase (we do not identify SOX2 in this data set because it also serves as a neural progenitor marker). Clustering this data set reveals many of the same ontological processes as identified in the HESC data set. To further our understanding between HESC and PGESC, we identified 160 genes when these down-regulated data sets were intersected. This means that a number of genes that were down-regulated upon differentiation in PGESC and not HESC and vice-versa. Does this mean that there are many distinct genes responsible for pluripotency and selfrenewal? Or can pluripotentiality result from unique genetic combinations? Since this data set was differentiated along a particular lineage, we can analyze those genes that were upregulated to assess differentiation or identify new genetic markers. For example, the most significantly upregulated gene is neurofilament 68 KD, neurofilament 3, neural cell adhesion molecule. When this data sets is clustered using GO biological function, the most over-represent processes are antigen processing major histocompatibility type I (MHC-I) and neurogenesis. However, when analyzing the large list of genes that are up-regulated upon differentiation, it is difficult to distinguish those genes responsible for “differentiation” and those differentiation genes that are neuronal lineage specific. AFSC Pluripotent stem cells can be isolated from amniotic fluid between 14–18 weeks of gestation and comprise approximately 0.8–1.4% of the cells present in amniotic fluid (in submission). These cells are grown in basic medium supplemented with serum, have a high self-renewal capacity (300 population doublings), with a doubling time of less than 36 h, do not require a feeder layer for undifferentiated expansion, and are autologous with the fetus. In addition, human amniotic fluid stem cell (HAFSC) maintain their telomeres and normal karyotype throughout late passaging. Early passage HAFSC express ESC markers, but do not form teratomas when injected into severe combined immunodeficient (SCID) mice. HAFSC can be differentiated in vitro into bone, muscle, fat, endothelia, beta-islet cells, liver, and neurons. When mouse chimeras were created by injecting AFSC into blastocysts, AFSC derived cells were found throughout the embryo.

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We used GeneChips to understand HAFSC differentiation, to identify AFSC specific genetic markers. HAFSC were profiled with GeneChips after 30 days of differentiation along four lineages – bone, muscle, endothelia, and liver. When comparing the differentiated to the undifferentiated we identified hundreds of genes that were up- and down-regulated. We noticed many genes were shared amongst the four lineages, genes involved in cell cycle and proliferation intermixed with tissue specific genes. We then treated all four lineages as one data point, and created a signature of genes that were “universally” conserved amongst their differentiation. We then took the genes that were up-regulated upon each lineage and subtracted out those genes that were “universally” conserved and were then able to identify tissue specific ontological processes. We then took the list of genes that were “universally” down-regulated in HAFSC which comprised of 1,498 genes, and compared it to the signature of genes we identified as being down-regulated upon HESC differentiation (from above) and identified 243 genes in common which probably represent pluripotential or self-renewal genes. Furthermore, we identified 1,160 genes as being down-regulated in HAFSC and not in HESC or PGESC and can think of those genes as potentially representing HAFSC specific genetic markers. We are currently mining this data set as a way to identify the origin of these cells. Sartorelli profiled C2C12, skeletal muscle cell line, that was pre-cultured with trichostatin and then after differentiation with GeneChips (NCBI GEO # GSE1984). Trichostatin is a histone deacetylase inhibitor, and based on our GeneChip data from above which identified histone deacetyltransferase 1 (HDAC1) as being down-regulated upon differentiation, we hypothesized the HDAC inhibition might induce differentiation, and if so, along a particular lineage. After 36 h of incubation with trichostatin, we identified 586 genes being up-regulated and 682 genes as down-regulated with the method described above. We then intersected this data set with those genes that were “universally” up- and down-regulated and identified 423 genes in common between the former and 135 genes between the later. Furthermore, after subtracting out this “universally” upregulated from those that were up-regulated by trichostatin, we clustered them based on their GO biological process and the most over-represent process were genes involved in “response to chemical substance.” When we clustered based on its SwissProt ontology, we identified “neurone” as the most over-represented process, which may indicate that trichostatin is inducing the up-regulation of neuronal genes. Furthermore, we also identified many genes that are known to be imprinted and after re-analyzing the C2C12 data set exposed to trichostatin, we are finally many of the same and believe this approach will be useful in identifying genomic imprinting’s role in stem cell differentiation. Meta-Analysis We believe many of the answers promised by stem cell biology will be uncovered when stem cells are analyzed as a whole. Our most interesting data comes from comparing and contrasting the genes that are identified as being up- and down-regulated upon differentiation amongst stem cells of different origins. GenMapp is a program that consists of hundreds of pre-made pathways. Affymetrix probe sets can be loaded directly into the program and color coded based on their biology and expression. We then color coded those genes that we identified as being up- (lighter shade) or down-regulated (darker shade) in HESC (red), PGESC (yellow), HAFSC (green). Analyzing pathways in this context provides a unique way to identify pathways that are shared or uniquely involved in stem cell differentiation, such as genes involved in extracellular matrix (ECM) production (Figure 32.3). From a tissue engineering perspective, one particular data set that we are particularly interested in are those genes that are up-regulated upon differentiation. These genes could potentially be easily applied to improve the efficiency and quality of differentiation. After analyzing many pathways, we identified a pathway involved in ECM production as predominantly consisting of genes that were up-regulated upon differentiation. If one thinks of the in vitro differentiation of stem cell as a model or organogenesis, it is of no surprise to find a number of genes that are universally up-regulated upon differentiation are involved in ECM/scaffold

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development. We then too the genes that were up-regulated in HESC and HAFSC and clustered them based on their GO molecular function and the most over-represented pathway was “ECM structural constituent.” Furthermore, our laboratory and others have shown the importance of scaffold selection on tissue formation. Thus, the genes that are up-regulated and conserved amongst lineages could represent the basic core of an in vivo scaffold, and those genes that are unique to a particular lineage as providing specificity in guiding differentiation. Thus, these studies have direct application to the fields of chemical engineering and biomaterials in the creation of novel tissue specific synthetic scaffolds.

CELL SCREENING GeneChip can be used as a screening tool to monitor the expression of cellular population as a way to assess a stem or progenitor cell population before early on in expansion or after before transplantation. They can also be used to determine when a differentiated derivative is truly representative of its in vivo counterpart. This would be of significant value in predicting clinical efficacy because we are currently limited by the tissue specific expression of genetic markers. GeneChips can potentially be used to predict whether tissue engineering would be feasible. Screening Differentiation Derivatives HESC Derived Retinal Pigmented Epithelial In a more clinically applicable manner, we are using GeneChips as a screening tool to assess “differentiation.” The question we wanted to address is when is a differentiation derivative fully differentiated? HESC were differentiated into retinal pigmented epithelial cells (HESC-RPE) (Klimanskaya et al., 2004). After determining the expression of many known markers, we ran a GeneChip and identified many more RPE using present calls. We then compared our signature to publically available RPE cell lines, ARP19 and D407. We identified many genes in common between them, but we noticed that there were a number of genes that were expressed in our HES-RPE and not in the RPE cell lines. As a control, we used a bronchial epithelial data set that was publically available, and chose this cell type as the ideal control because of its similar bronchial epithelial origin but lack of RPE specific genes. We then analyzed this data set and identified many known RPE genes and did not identify the retention of “stemness” genes such as Oct-4, Sox2, and TDGF1. We concluded that the cell lines lost some of their RPE specific genes, as is common when working with cell lines. We then decided to compare our HES-RPE to freshly isolated human fetal-RPE (fe-RPE). We decided on comparing our cells to these because these are the cells that are currently used as transplantation therapies. Our data demonstrates that our HES-RPE looked more genetically similar to fe-RPE than the existing cell lines, one of which has been used clinically. While the D407 line has been used in transplantation studies and resulted in clinical improvement. Currently, the ideal transplantation resource is fe-RPE. While effective, the disadvantage of this treatment is the scarcity of donor tissue. We therefore compared our HESC-RPE to RPE that have been successfully used clinically. Figure X not only demonstrates their transcriptional similarity to fe-RPE but also a greater similarity to fe-RPE than D407. We believe there is great potential in the use of GeneChips to assess the clinical potential of stem cells. Here we used GeneChips to assess the quality of our differentiation by comparison to its in vivo counterpart and other lines that have been clinically effective. We also believe that this can be used to efficiently assess the clinical potential of new stem cell lines or the ability to use diseased progenitor cells for tissue reconstruction (see below). Although we screened almost the entire transcriptome, follow-on studies will identify a transcriptional signature that could be used to print custom arrays which would be less expensive and requires less RNA and more desirable to use on a therapeutic scale. Furthermore, this signature could be compared to a “stemness” signature to assess the potential to form teratomas.

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Screening HESC for Genetic Variability In an analogous manner, Abeyta et al. (2004) assessed the variability in gene expression of HESC. They profiled HESC lines HSF6 (p46-female), HSF1 (p36-female), H9 (p51-male) using the HG-U133 GeneChip. They used MAS5 for call detection and identified 7,385 genes present in all three lines, and after analyzing with RMA, 52% of these genes had greater than a two-fold difference. They also identified 2,279 present only in H9, 337 only in HSF1, and 641 only present in HSF6 and since all three lines were grown and handled in a similar manner, they believe these results could be due to DNA sequence variability.

RE-ENGINEERING HEALTHY FROM DISEASED TISSUE Bladder Smooth Muscle Progenitors Our laboratory has shown that bladders can be reconstructed using a small 1 cm  1 cm biopsy (Oberpenning et al., 1999). The urothelial cells and smooth muscle cells can separated, expanded, and seeded onto a scaffold and then transplanted. To enhance smooth muscle cell expansion, we ran GeneChips on human smooth muscle cells. Triplicates were performed, and we created a genetic signature of genes that were present in all three replicates. We then clustered these genes based on their ontology to identify receptors, growth factors, and ECM components to improve their culture conditions. Our laboratory has also shown that bladders can be engineered from diseased bladders (Lai et al., 2002). To understand the genetic abnormalities of smooth muscle from bladder exstrophy and neurogenic bladders, we used GeneChips to compare their genetic signatures to healthy bladder smooth muscle. Triplicate microarrays were performed and we created a genetic signature by identifying genes that were present in all three replicates. After comparing these genetic signatures, it appears that exstrophic and neurogenic bladder smooth muscle cells express many of the core genetic smooth muscle cell components. We then compared the relative expression amongst the three and it appears that exstrophic and neurogenic smooth muscle cells look like a relatively immature phenotype. Both neurogenic and exstrophic bladder smooth muscle cells also over-express a number of ECM related genes, and this is also seen in pathologic sections. Lastly, in the exstrophic data set, we identified an up-regulation of inflammatory genes (MHC, chemokines) and believe this to be because of their incomplete closure during development and are thus fused with the abdomen and exposed to the peritoneal fluid of the abdomen. Interestingly, this inflammatory response is still maintained throughout multiple passages in vitro. We are currently interested in correlating these genetic signatures with expansion and differentiation potential as a way to predict tissue engineering potential. When interpreting the difference between exstrophic and normal bladder smooth muscle cells in the context of our stem cell data, we hypothesized that trichostatin could be used to further differentiate these smooth muscle cells, and would thus decrease the amount of ECM production (based on the GeneChip data from above that identified the over-expression of developmental genes). Our initial studies have shown that trichostatin induces a significant decrease in collagen production of exstrophic bladder smooth muscle cells and demonstrates how we are applying GeneChip data to re-engineering healthy from diseased tissue.

APPLICATIONS TO CANCER Because of the similarities between cancer and stem cells, many believe that cancer is a de-differentiation of a somatic cell back into a stem cell (Pardal et al., 2003; Al-Hajj et al., 2004). Not only do cancers and stem cells have similar growth properties, but germ cell tumors such as embryonal carcinomas (Andrews, 1998) and benign ovarian teratomas (Linder et al., 1975) are even capable of differentiating into tissues of different germ

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layers. Therefore, insights into the genetic mechanisms of stem cell growth and differentiation might provide information in designing therapeutics for cancer. A benign ovarian teratoma, a germ cell tumor, is the result in the “activation” of a diploid (meioses II arrested) oocyte through an unknown mechanism. What is unusual about this tumor is its pluripotentiality, not only have many different tissue types been identified, but can sometimes result in a homunculus (Latin for “little man”) (Abbot et al., 1984). Thus, it appears that an oocyte alone is capable of differentiating into all three germ layers and can form the proper three-dimensional tissue structures and architecture. In addition, we have shown that we can mimic this effect in vitro with the use of monkey oocytes. The activation of oocytes in this context is referred to parthenogenesis. Many lower species are capable of this form of reproduction such as bees, fish, serpents, monotremes, but not eutharians. However, when mammalian oocytes are activated, they usually do not progress past the blastocyst stage, the stage at which stem cells can be isolated from their ICM. We have shown that these cells express ESC markers and are capable of differentiating into all three germ layers. Furthermore, injection into the peritoneum of SCID mice results in a benign teratoma. Lastly, when patients with benign ovarian teratomas are treated with chemotherapy, the prognosis grading is based on the percentage of neuronal tissue. Thus, it appears that the induction of cell cycle arrest results in a default differentiation into neurons. This is also seen with our in vitro with PGESC whose differentiation into neuronal progenitor cells is predominant. Thus, the difference between a benign ovarian teratoma and a parthenogenetic stem cell is that the former is “activated” in vivo while the latter is activated in vitro. Thus, if there is thought to be a “cancer stem cell,” we believe that PGESC are the stem cells of benign ovarian teratomas. The advantage of studying PGESC instead of benign ovarian teratomas is that we have a single cell in which we can study its self-renewal and differentiation. In addition, an understanding of these genetic mechanisms will not only give us a unique insight into cancer, but also provide us with potential therapeutic targets. We believe that identifying the genes that are unique to PGESC, “stemness” genes, can serve as therapeutic targets. We therefore performed GeneChips on undifferentiated PGESC and those that have undergone neuronal differentiation to create a data set of 233 genes. What is the best way of applying this data set? Successful approaches have involved the creation of small molecule inhibitors (however, they are often associated with side effects), or monoclonal antibodies to genes such as Her2/Neu (herceptin) (Cobleigh et al., 1999; Vogel et al., 2001, 2002; Tripathy et al., 2004). siRNA offers tremendous potential due to its specificity in its capability of knocking down single genes. We are interested in creating cancer vaccines by engineering a cytotoxic T cells (CTLs) specific response using a patient’s own immune system. CTLs are the T cells responsible for the cell mediate immune response and recognize specific peptide sequences (8–12 amino acids in length) that are expressed in MHC-I. This approach is promising in its specificity and ability to target micrometastasis, but is limited by the number of tumor specific antigens (peptide sequences). In a previous study, we reported the identification of peptides derived from the enzyme telomerase (responsible for maintaining telomeric ends), and were able to generate CTLs from patients with prostate cancer to recognize and lyse 7/8 different types of tumor lines with one peptide sequence and 7/8 lines of another peptide sequence, non-overlapping (Minev et al., 2000). This antigen was chosen based on the idea that cancer cells over-express this enzyme which allows them to replicate indefinitely. Although, other cells like bone marrow and skin cells express telomerase, we showed that the expression level is not significant enough to generate an immune response. Tumor antigens have also been identified from other proteins such as Her2/Neu. Non-coincidentally, both of these genes were identified in our 233 gene “stemness” data set. We believe that our parthenogenetic “stemness” genes and other “stemness” genes identified from other “stem cell data sets” could serve as targets in this manner. Furthermore, screening of these antigens different cancer types (breast, lung, melanoma) can be used to create a cocktail of tumor specific antigens.

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CONCLUSION Future Direction The component that has the most potential for improvement in a GeneChip experiment is in the interpretation of data. Trying to ascribe stem cell functions or classifying different stem cells from gene lists is analogous to taking a train apart and trying explaining how it travels or trying to explain the difference between a car and a train by comparing its parts. These GeneChip data files contain the cell’s genetic networks. However, understanding the genetic networks of stem cells is analogous to a computer chip; however, it is complicated by differentiation which allows it to alter its “motherboard.” Because there are over 300 cell types in the body, there are over 300 types of “motherboards” or genetic networks for lineage directed differentiation. To begin to address this goal, we are trying to draw upon multiple comparisons amongst disparate stem cells to try to identify what are conserved or uniquely expressed. We will need to be performed with other types of stem cells – adult, multipotent, unipotent, alternative, pluripotent stem cells to identify conserved genetic networks that are stem cell specific and to compare their presence in somatic cells. This will then allow one to associate gene lists with particular stem cell characteristics, that is their differentiation potential or identify core genetic networks that are housekeeping functions. Future experiments will need to profile stem cell differentiation in a lineage and temporal specific manner that will enable one to dissect out the genetic wiring of differentiation and its commitment to a lineage. To accomplish such goals, many more GeneChip experiments will need to be performed and made publically available.

SUMMARY The goal of this chapter was to discuss the many components involved in a GeneChip experiment and discuss its applications to regenerative medicine. Since others have demonstrated the importance of methodology in GeneChip experiments (REF), we wanted to demonstrate the many variables at each step, and describe a protocol for GeneChip analysis that requires minimal computer skills, is efficient, reliable, quick, inexpensive, and results in meaningful data. In discussing the application of GeneChip data to regenerative medicine, we first described how we and others are using GeneChips to understand the genetics of how different stem cells grow and differentiate. Next, we showed how GeneChips can be used as a screening tool to assess differentiation and variability between different stem cell lines. Then we demonstrated how GeneChips can be used to re-engineer healthy from diseased tissue and concluded with how stem cell GeneChip data can be applied to understanding and treating cancer.

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Part IV Biomaterials for Regenerative Medicine

33 Design Principles in Biomaterials and Scaffolds Hyukjin Lee and Tae Gwan Park INTRODUCTION Tissue or organ transplantation is severely limited by the problems of donor shortage and immune rejection from the patients. The development of tissue engineering allows the transplantation of cells from a patient’s own tissue to regenerate damaged tissue or organ without causing immune responses. For the cell transplantation, extracted cells are often required to cultivate in a large scale to attain a sufficient cell seeding density. In culturing the cells, the in vitro culture conditions play pivotal roles in proliferation and differentiation. Three-dimensional biomaterial scaffolds are firstly developed for the temporary substrate to grow cells in an organized fashion. Although direct injection or implantation of in vitro cultured cells is often performed, using cell suspension is doubtful for the successful regeneration of impaired tissues. It is also well established that the three-dimensional organization of cells often related with cellular attachments affects the fate of cellular development. As a result, biodegradable and biocompatible polymers have been widely used to fabricate threedimensional scaffolds for tissue engineering. In the past, biomaterial scaffolds were mainly used for temporary prosthetic devices to fill the void spaces after tissue necrosis or surgery. However, current biomaterials pursue to mimic the role of natural extracellular matrix (ECM) which can support cell adhesion, differentiation, and proliferation. ECM mimicking biomaterial scaffolds should be designed considering the following requirements. First, suitable biomaterials are selected for particular applications (Mikos and Langer, 1993; Athanasiou and Agrawal, 1996). This is analogous to the effort to build up the target-specific biological scaffolds. Second, biomaterial scaffolds require a highly open porous structure with good interconnectivity, yet possessing sufficient mechanical strength for cellular in- or outgrowth (Cima and Langer, 1991). Third, the surface of fabricated scaffolds must be able to support cellular attachment, proliferation, and differentiation (Varkey and Uludag, 2004; Peattie and Prestwich, 2006; Vasita and Katti, 2006). Fourth, drug or cytokine releasing scaffolds are ideal for modulating tissue regeneration since cytokines such as growth factors and other small molecules have fundamental roles on growing functional living tissues (Niemann, 2005; Raghunath and Seifalian, 2005; Keilhoff and Wolf, 2006). The harmony of the above considerations is essential to fulfill the requirements of excellent biological scaffolds, thereby inducing synergic effects on successful tissue repair. This chapter focuses on recent developments on fabricating biomimetic, ECM-like porous scaffolds useful for tissue engineering. Our experiences on designing novel biomaterials and innovating scaffold fabrication techniques are highlighted here as well as other leading researchers’ works. Novel fabrication methods and designing strategies are elucidated such as generating the macroporous biodegradable scaffolds, the surface modification of biodegradable scaffolds to enhance cellular attachment and biological activity, and the incorporation of bioactive molecules within the scaffold systems. A number of excellent reviews are available for synthetic biomaterials for medical applications and tissue engineering (Peppas and Langer, 1994; Ratner, 1996; Uhrich, 1999).

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SELECTION OF BIOMATERIALS Natural biomaterials have been extensively used for tissue engineering since they have advantages over synthetic materials such as similarity with natural ECM. For example, alginate, chitosan, collagen and its derivatives, fibrin, and hyaluronic acid (HA) were investigated for the fabrication of three-dimensional scaffolds (Rosso and Barbarisi, 2005). However, the properties of natural biomaterials are difficult to adjust and the source-related immunogenicity is still unsolved. In contrast, synthetic biomaterials are man-made materials mainly composed of synthetic polymers. Although synthetic polymers often reveal poor biocompatibility, the building up process of synthetic polymer provides precise control of the properties of synthetic materials and even can give better performance over naturally occurring biomaterials. For instance, aliphatic polyesters and polyanhydrides are common synthetic polymers for tissue engineering and drug delivery system. These polymers have distinct biodegradability and biocompatibility. The combinations of hydrophilic and hydrophobic segments in the structure generate a variety of synthetic biomaterials with different mechanical properties and degradation behaviors.

BIODEGRADABLE SYNTHETIC POLYMERS Aliphatic Polyesters Aliphatic polyesters are Food and Drug Administration (FDA) approved synthetic biomaterials which have been widely used for biodegradable applications such as surgical sutures and bone fixing screws. Poly(α-hydroxyl esters) such as poly(L-lactic acid) (PLLA), poly(lactic-co-glycolic acid) (PLGA), and polycarprolactone can be synthesized by the ring-opening polymerization of monomers and have hydrolytically cleavable bonds along the polymer backbone. When these synthetic polymers are implanted in the body, hydrolysis of polymer backbone reduces the molecular weight of polymer and their degraded products such as lactic and glycolic acid can be metabolized in the body (Figure 33.1). In addition, based on their biocompatibility and safety record in humans, these polyesters have been used extensively in drug delivery systems and tissue engineering applications (Saltzman, 1999; Putman, 2001).

O

CH3

CH3

O

O

H

OH



C

HO CH3

O

OH HO

n

O

CH3

Lactic acid

Poly(L-lactic acid)

O CH3

O

O H

O

O O

HO O

CH3

n

OH HO

OH O

CH3 O

m

Poly(lactic-co-glycolic acid)

OH HO

Lactic acid and glycolic acid

Figure 33.1 Structure of PLLA and PLGA and their degradation products; acid hydrolysis of PLLA and PLGA to give lactic and glycolic acid.

581

582 BIOMATERIALS FOR REGENERATIVE MEDICINE

Aliphatic polyesters typically lack a chemical functionality for modification with biological molecules. As an example for the introduction of functional groups in the polymer backbone, Barrera and Langer (1993) reported the use of a novel monomer to incorporate functional amine groups into polylactic acid (PLA) polymers. Poly(lactic acid-co-lysine) was synthesized by the copolymerization of cyclic lactide and its analog containing the lysine. This novel amine containing PLA showed similar biocompatibilty while providing additional sites for further chemical modifications. Polyanhydrides Another class of degradable biopolymers is polyanhydrides. Unlike polyesters which predominately show a bulk-erosion process, polyanhydrides exhibit a surface-erosion process which is particularly useful for sustained drug delivery systems. Langer et al. demonstrated the use of polyanhydrides based on sebacic acid (SA) and p-carboxyphenoxyproane (CPP) (Leong et al., 1985). By combining hydrophilic SA and hydrophobic CPP, the rate of surface erosion can be controlled from days to years. In addition, these polyanhydrides exhibit great biocompatibility and excellent in vivo performance for potential biomedical applications.

DESIGN PRINCIPLES OF BIOLOGICAL SCAFFOLDS Fabrication of Macroporous Biodegradable Scaffolds Along with the material selection, fabrication methods are also critical for designing biological scaffolds. For tissue regeneration, highly open porous polymer scaffolds are often required for high density cell seeding, efficient nutrient and oxygen transport. There have been multiple methods to fabricate highly porous biodegradable polymer scaffolds which are listed in Table 33.1. Briefly illustrating a few techniques, the compressed polyglycolic acid (PGA) meshes are made out of non-woven PGA fibers and these meshes have been widely used for soft tissue regeneration (Freed and Langer, 1993). Random coiling and heat treatment of PGA fibers can generate highly open porous and interconnected structures with a high surface to volume ratio. However, the mechanical strength of these meshes is insufficient for hard tissue regeneration (Mikos and Langer, 1993). To enhance the mechanical properties of compressed PGA meshes, Mooney and Langer (1993) demonstrated that a mixed solution of PLLA and PLGA can be applied to the compressed PGA meshes. Mixture of PLLA and PLGA dissolved in organic solvent was sprayed throughout the compressed PGA meshes. As the organic solvent evaporated, dried PLLA/PLGA strengthened the cross regions of fibers and enhanced mechanical properties

Table 33.1 List of fabrication methods for preparation of highly porous biodegradable scaffolds. Fabrication methods

Materials

References

Compressed mesh of non-woven fibers

PGA, PLGA

Solvent casting/salt leaching CO2 expansion Emulsion freeze drying Phase separation

PLLA, PLGA PLGA PLGA PLLA, PLGA

Three-dimensional imprinting

PLLA, PLGA

Freed and Langer (1993), Mikos and Langer (1993), Mooney and Langer (1996) Mikos and Langer (1993), Mikos and Vacanti (1994) Mooney and Langer (1996), Harris and Mooney (1998) Whang and Nuber (1995) LO and Leong (1996), Schugens and Teyssie (1996), Nam and Park (1999) Park and Griffith (1998)

Design Principles in Biomaterials and Scaffolds 583

of compressed meshes. Despite the mechanical result, this method exhibited the reduction of high surface to volume ratio of compressed meshes and the difficulty of matching the degradation rate of surface coated materials and bulk materials. In addition, the solvent casting/salt-leaching technique has been extensively exploited for fabricating scaffolds for tissue engineering (Mikos and Langer, 1993; Mikos and Vacanti, 1994). PLGA dissolved in an organic solvent with salt particles is placed in a mold to produce a polymer/salt mixture, which is immersed in water to remove salt particles to generate open pore structures. The scaffolds prepared by this method often demonstrate a dense surface layer and poor interconnectivity between macropores, which reduces cell seeding into the scaffolds in vitro and causes non-uniform distribution of seeded cells. Thus, poor cell viability and tissue ingrowth after in vivo implantation are observed. In order to resolve the problems from salt-leaching techniques, Nam and Park (2000) utilized PLLA paste containing ammonium bicarbonate salt particles which acts as a gas-foaming agent as well as a salt-leaching porogen to fabricate highly interconnected porous biodegradable scaffolds (Figure 33.2). Sodium bicarbonate salt with acidic excipients has been widely used for effervescent gas evolving oral tablets. Since ammonium bicarbonate salt upon contact to an acidic aqueous solution such as citric acid and/or incubated at elevated temperature produces gaseous ammonia and carbon dioxide by itself, ammonium bicarbonate salt particles could be incorporated into a biodegradable gel paste prepared by dissolving high molecular weight PLLA in an organic solvent. The resultant putty paste was easy of shaping into different geometry and could be immersed in hot water solution and directly dried under vacuum oven to remove or leach out the salt particles while concurrently generated gaseous ammonia and carbon dioxide provide highly interconnected pores within a solidifying polymer scaffold. Thus, the formation of dense surface skin layer was not found on either sides of the surface of the scaffolds (Figure 33.3). Macroporous PLGA scaffolds using gas-foaming/salt-leaching method with controlled degradation rate was also investigated (Yoon and Park, 2001). Unlike semi-crystalline PLLA, amorphous PLGA could form a gel-like paste in an organic solvent even at high concentration. PLGA was dissolved in an organic solvent such as chloroform then precipitated in a non-solvent, ethanol. Resulting precipitates exhibited a gel-like property such that the paste can be molded or hand-shaped in any desirable dimensions. In this study, instead of incubating scaffolds in hot water bath or vacuum oven, citric acid solution was used to control the porosity of scaffolds as well as mechanical property. Using citric acid, carbon dioxide and ammonia gases could be generated

Solvent Polymer gel prepared by non-solvent precipitation

Semi-solidified polymer/salt complex

 Sieved salt particles

Polymer gel paste

Teflon mold

CO2

NH3

Freeze dry Polymer scaffold

Gas foaming in acidic aqueous solution

Figure 33.2 Schematic of gas-foaming and salt-leaching process to fabricate macroporous scaffolds.

584 BIOMATERIALS FOR REGENERATIVE MEDICINE

(a)

(b)

(c)

(d)

Figure 33.3 SEM images of macroporous scaffolds fabricated by gas-foaming and salt-leaching process. Uniform interconnectivity and high porosity are observed on both surface (a, c) and cross-section of scaffolds (b, d).

at room temperature and changing the concentration of citric acid in the solution enabled to control the porosity of scaffolds. The result supported that the increase in porosity of scaffolds was observed with increased citric acid concentration as high concentration citric acid are more gas generating. In addition, degradation and swelling behaviors of PLGA scaffolds with different compositions were investigated. The macroporous scaffolds with three different compositions of lactic and glycolic acid were incubated in phosphate buffered solution (pH 7.4) at 37°C. During the incubation period, significant swelling of the scaffolds was observed depending on the composition, and the change in dimension and morphology was caused by the accelerated degradation of PLGA scaffold which could generate more water adsorbing small molecular weight PLGA oligomers within the degrading scaffolds (Figure 33.4). As an alternative to salt-leaching and gas-forming fabrication, electrospinning has received much attention for fabricating polymeric ultrafine nanofibers to build three-dimensional tissue engineering scaffolds. Nanofibrous biodegradable scaffolds would have definitive advantages for cell attachment, proliferation, and differentiation because they resemble an ECM structure. Recently, Kim and Park (2006) demonstrated ECM mimicking nanofiber mesh for tissue engineering applications. The amine terminated PLGA dissolved in a mixture of DMF/THF solvent was ejected through a nozzle by an electrostatic force, resulting in the formation of non-woven fabrics. During the electrospinning, the solvent evaporates and the charged polymer nanofibers were deposited on a grounded collector. The resultant structure was a three-dimensional, randomly oriented nanofiber network mesh with a highly nanoporous architecture (Figure 33.5). The in vitro cell culture revealed that the resulting nanofiber ranged from 300 to 1,000 nm provided an excellent environment for cellular attachment, proliferation, and differentiation.

Design Principles in Biomaterials and Scaffolds 585

D0

D3

D 10

D 21

D 35

D 49

D 63

D 84

PLGA 75/25

PLGA 65/35

PLGA 50/50

Figure 33.4 Photographs of different PLGA scaffolds after hydrolytic degradation in phosphate buffered saline (PBS) at 37°C. With increasing composition of glycolic acid, rapid degradation and swelling of scaffolds are observed.

(a)

(b)

High voltage power supply Polymer solution

Syringe pump Grounded collection drum

Figure 33.5 Schematic of electrospinning (a) and an SEM image of electrospun PLGA nanofiber (b).

Surface Immobilization of Bioactive Molecules on Macroporous Biodegradable Scaffolds The surface modification of scaffolds is essential since the microenvironment of the body cannot see the bulk property of biomaterials, but the surface of biomaterials. In the past, major issues concerned with biomaterials are their biocompatibility upon the injection or implantation of materials in vivo. In the case of material selections, a few biomaterials are known for free of causing acute inflammation. As a result, the surfaces of fouling devices were modified with non-protein adsorbing materials such as polyethylene glycol (PEG) to stealth the implants from the body. Since many cell adhesive peptides present abundantly in the ECM dictate cellular behaviors, the immobilization of various bioactive ligands on the surface of biomaterials was attempted for actively mimicking physiological conditions, thereby increasing cytocompatibility and biological functionality when the biomaterials are implanted in the body. A number of surface modification methods were developed such as chemical oxidation and etching, plasma and corona discharge, radiation and UV grafting, partial hydrolysis, protein adsorption, and conjugation/immobilization of bioactive ligands (Rasmussen and Whitesides, 1977; Ramsey and Binkowski, 1984; Weisz and Schnaar, 1991; Gao and Langer, 1998; Nam and Park, 1999; Otsuka and Kataoka, 2000).

586 BIOMATERIALS FOR REGENERATIVE MEDICINE

As an example of PLGA scaffolds modified with bioactive ligands, we demonstrated galactose modified PLGA macroporous scaffolds for culturing hepatocytes in vitro (Park, 2002; Yoon and Park, 2002). When selecting bioactive molecules for immobilization, ligands for cell membrane receptors have a pivotal role since these ligands are associated with cellular signaling pathways and activities such as cell migration, proliferation, and differentiation. Moreover, cell-specific ligands help to initiate binding and attachment of cells on modified scaffolds. For instance, galactose is a specific ligand for asialoglycoprotein receptor in hepatocytes. Galactose modified PLGA was prepared by conjugation of end aminated PLGA with lactobionic acid using dicyclohexylcarbodiimide/N-hydroxysuccinimide (DCC/NHS) coupling agents (Figure 33.6). The galactosylated PLGA was then processed to form films and macroporous scaffolds to examine hepatocyte-specific cellular binding to the modified surface. Albumin secretion was quantified as well for validating cellular functionality. For the cell-specific binding to galactose, glucose modified films were also fabricated and the hepatocyte attachment on films was observed. In the result, hepatocytes were selectively attached to the galactose modified films compared to the non-specific glucose modified films. Additionally, it was demonstrated that conjugation of galactose on PLGA surface supported higher cell viability as compared to control PLGA films. The idea of mimicking an in vivo system using peptide amphiphiles such as arginine–glycine–aspartic acid (RGD) was realized long ago and the surface modification with RGD sequences has been widely used for enhancing cellular attachment and growth (Yoon and Park, 2004). Cell adhesive ligands such as RGD are abundantly present in collagen and their roles are vital for cellular attachment via integrin mediated binding to ECM. There are a number of excellent reviews demonstrating the effects of RGD in tissue engineering. For instance, Langer and coworkers published a comprehensive review for creating biomimetic microenvironment

O

CH3

O

CH3

OH

O

O

DCC, DMSO

OH

CH3

O

O O

NH C N

O

O O

n

CH3

O O

OH O

n

O

O NHS, DMSO

O

CH3

OH

CH3

O O

O

O

N

O O

O

n

O HO HO

HO H2N-AGA, DMSO

O

CH3

OH OH OH H O N[H2C]2NH2 C OH OH

O

O

OH OH OH

H2N-AGA

Figure 33.6 Synthesis of galactosylated PLGA.

H

n

OH

OH N

O

H

OH

O

O

N O

O

HO

CH3

O O

O OH

OH

Design Principles in Biomaterials and Scaffolds 587

using adhesive peptides (Shakesheff and Langer, 1998). Continuing the mimicking of biological surface, selecting bioactive ligands is crucial for each application. For cartilage tissue engineering, microenvironment similar to native cartilage such as highly water swollen environment is required. HA is a naturally occurring non-sulfated glycosaminoglycan (GAG) composed of N-acetyl-D-glucosamine and D-glucuronic acid which is a major constituent of ECM and abundantly expressed in cartilage. In addition, HA is known to have vital roles in various biological functions of chondrocytes such as regulating adhesion and motility, and mediating cell proliferation and differentiation (Larsen and Balazs, 1992). There are a number of publications on effects of HA on chondrocyte proliferation and maintaining their original phenotype (Chow and Knudson, 1995; Lindenhayn and Sit, 1999). From the reasons above, Yoo et al. fabricated the HA modified PLGA macroporous scaffold (Yoo and Park, 2005). As previously described, the macroporous structure of PLGA can be obtained from gas-foaming/ salt-leaching process and the surface of these materials was chemically conjugated with HA. Amine end-capped PLGA was synthesized and mixed with PLGA to foam biodegradable scaffolds. To expose the amine groups on the surface, fabricated scaffolds were purged into the HA solution with EDC/NHS coupling agents (Figure 33.7). The resulting HA coated PLGA macroporous scaffolds exhibited higher chondrocyte proliferation probably via CD44 interaction with HA and initiated increased production of GAG, as compared to PLGA alone, while enhancing Type II collagen and aggrecan gene expression.

Figure 33.7 Schematic of surface modification of PLGA biodegradable scaffold with HA.

588 BIOMATERIALS FOR REGENERATIVE MEDICINE

Sustained Release of Bioactive Molecules from Macroporous Scaffolds In many tissue engineering applications using stem cells, specific cellular differentiation is often required to achieve the expression of desirable phenotypes and the secretion of functional proteins and carbohydrates. To satisfy the above requirements, the in situ local delivery of cytokines such as growth factors and molecular drugs within cell seeded scaffolds has been pursued since the sustained release of bioactive molecules is known to stimulate cell proliferation, differentiation, and the secretion of desirable proteins. There have been multiple reports on local delivery of growth factors within the scaffold such as epidermal growth factors (Mooney and Langer, 1996), transforming growth factors (TGF) (Behof and Jansen, 2002), vascular endothelial growth factors (VEGF) (Wissink and Feijen, 2000; Richardson and Mooney, 2001), basic fibroblast growth factors (b-FGF) (Royce and Marra, 2004), and bone morphogenic growth factors (Lee and Battle, 1994; Whang and Healy, 2000). These scaffolds are able to stimulate embedded cells to express tissue-specific phenotypes in mRNA level and induce to produce functional ECM corresponding to the desirable applications. In addition, the sustained release of plasmid DNA for transfecting neighboring cells was also investigated (Chun and Park, 2004, 2005). One of the emerging fields of drug delivery is a local delivery of small drug molecules such as steroid analogs from biodegradable scaffolds in a sustained manner. Dexamethasone is a family of glucocortiocoids that exhibits various inhibitory effects on inflammation process and proliferation of smooth muscle cells (Reil and Gelabert, 1999; Hickey and Moussy, 2002). As well, dexamethasone is commonly used along with specific growth factors to induce stem cell differentiation toward osteoblasts or chondrocyte-like cells (Peter and Mikos, 1998). To investigate the effects of the sustained release of dexamethasone, Yoon and Park (2003) fabricated the dexamethasone encapsulating macroporous scaffolds composed of PLGA. Hydrophobic dexamethasone was incorporated into the PLGA polymer solution and the macroporous scaffolds were fabricated by gas-foaming/salt-leaching method. Due to bulk degradation of PLGA, dexamethasone was slowly released out in a zero order fashion without an initial burst effect. The bioactivity of released dexamethasone was established by culturing smooth muscle cells with/without dexamethasone releasing scaffolds. The results strongly supported a large decrease in smooth muscle cell proliferation with increase in the concentration of dexamethasone. The suppression of lymphocyte activation or anti-inflammation activity by dexamethasone released from the scaffolds was also validated with different concentrations of dexamethasone. With continued development of synthetic biomaterials for drug delivery system, biodegradable scaffolds can also be utilized as a gene carrier for sustained release of plasmid DNA, oligodeoxyribonucleotides (ODN), and siRNA. By delivering growth factor and other cytokine-related genes, transfected cells can be genetically controlled and used in tissue repair. In addition, transfected cells can trigger neighboring cells to proliferate and differentiate to cells with specific phenotypes for specific tissue engineering applications. Common gene delivery carriers usually express highly positive charge that the charge–charge interaction between negatively charged DNA molecules and the carriers can form a tight ionic complex. However an excess use of highly positive polymer species such as polyethyleneimine (PEI), poly(L-lactide) (PLL), and positively charged fatty acids can cause severe cytotoxicity and reduces the biocompatibility of gene carriers. Although a single injection of naked plasmid DNA can induce appreciable protein expression, increasing the transfection efficiency and sustained release of plasmid DNA are still a challenge. To achieve a high level of specific protein synthesis, sustained release of naked DNA is a promising approach to overcome the low transfection efficiency. Therefore the PLGA macroporous scaffolds for sustained release of plasmid DNA was fabricated by thermally induced phase separation method (TIPS) (Chun and Park, 2004). In this study, homogeneous polymer solution at elevated temperature was phase separated

Design Principles in Biomaterials and Scaffolds 589

Figure 33.8 Cross-sectional SEM images of PLGA scaffolds fabricated by TIPS methods, quenching in liquid nitrogen (a) and annealing at –20°C (b). Note that increasing annealing temperature generates larger pores for rapid release of encapsulated plasmid DNA.

into polymer rich and polymer poor domain by lowering the solution temperature while subsequent lyophilization of solvent generated microcellular structure (Figure 33.8). In order to encapsulate plasmid DNA within scaffolds, PLGA was dissolved in 1,4-dioxane and mixed with plasmid DNA dissolved deionized water followed by quenching in liquid nitrogen and solvent lyophilization. To control the release encapsulated plasmid DNA, effects of higher quenching temperature (annealing) and the addition of PLGA grafted PLL were subsequently examined. The resulting scaffolds revealed that encapsulated DNA within the PLGA scaffolds was slowly released out over 20 days and the structure of release DNA was intact. Furthermore, higher quenching temperature produced larger pore formation within the scaffolds giving a rapid release of plasmid DNA while addition of PLGA grafted PLL lowered the release profiles. Lastly, the bioactivity of release plasmid DNA was established by high level of luciferase expression in cells. As described above, biomimetic scaffolds have received much interest (Park, 2002; Yoo and Park, 2005). Since natural ECM plays pivotal roles in various biological events, functions of ECM component such as HA and heparin have been investigated. For tissue engineering, angiogenesis, sprouting of microvessel from existing ones, is crucial for cell–scaffold implantation since a lack of blood supply results poor delivery of oxygen and nutrient causing necrosis of implanted cells. To enhance angiogenesis at implanted sites, angiogeneic growth factor has been applied in various fashions (Wissink and Feijen, 2000; Richardson and Mooney, 2001). A common way of incorporating growth factor is mixing them with polymer solution and cast them to form scaffolds or films. However, the use of organic solvent is a critical problem in maintaining the bioactivity of growth factors. Heparin is a negatively charged polysaccharide and widely used for anticoagulation agents to enhance biocompatibility of implanted devices. In natural ECM, heparin plays a role as a reservoir for controlled secretion of growth factors since it has a high binding affinity with various growth factors such as VEGF, TGF-β, bFGF. Heparin stabilizes the released growth factors and concentrates them in the local areas of demand. Exploiting the unique biological functions of heparin, heparin modified injectable PLGA microscaffolds were fabricated for the sustained release of b-FGF (Figure 33.9). By synthesizing PLGA microspheres with free surface amine groups, carboxylic groups of heparin can covalently conjugated on the surface of PLGA scaffolds. Soluble b-FGFs were readily bound to the heparin resulting in high loading efficiency. At last, in vitro study revealed that the sustained release profile of b-FGF was obtained and the bioactivity of released b-FGF was confirmed (Yoon and Park, 2006).

NH2

NH2

COOH O OH

CONH

EDC/NHS COOH

HOOC

NH2

NH2

H2N

CONH

C

CONH

O

H

HO

CO O

COOH

590 BIOMATERIALS FOR REGENERATIVE MEDICINE

H2COSO3H O O

OH O

CO O

Microsphere surface

H

COOH

CO O H

HNSO3H OH Heparin (Mw 12,000)

ED

C/

NH

PLGA microsphere

S

n

Immobilized heparin

Released b-FGF Heparin bound b-FGF g

in

Growth factor release

ad

F

lo

FG

b-

Figure 33.9 Schematic of heparin immobilized porous PLGA microsphere for local delivery of angiogenic growth factors.

SUMMARY AND CONCLUSION Design of biomaterials and scaffolds is a complex interdisciplinary subject. Biodegradable and erodible biomaterials serve as scaffolds and drug delivery devices for applications in regenerative medicine. Natural biomaterials are already been used for many years by trial-and-error material selection and we are just beginning to understand how synthetic biomaterials can be applied to our body. The use of biomaterials requires the understanding of the differences in structure and properties between these implanted materials and that of the host. In vivo tolerance of early biomaterials helped to initiate a rapid development of more complex biomimetic systems. Especially, the development of synthetic polymers allows us to engineer and build new properties exceeding naturally occurring biomaterials. Applying these biomaterials for in vivo use, application-specific fabrication and scaffold design are essentially required. Since implanted biomaterials interacted with physiological environment, each scaffold needs specific requirements for specific applications. Since tissue engineering and regenerative medicine is composite of cells, cytokines, and scaffold, we already emphasized the importance of each selection in different applications. Continuing with development of synthetic biomaterials, aliphatic polyesters have been utilized for many years and offer excellent design versatility and biocompatibility. The complicated requirements of scaffold allowed developing more sophisticated designs of scaffolds such as highly macroporous scaffolds for facilitating nutrient and oxygen transfer, addition of specific biological ligands on the surface for promoting cell attachment, proliferation, and differentiation, and finally the cytokine releasing scaffolds for the manipulating cellular functions of encapsulated cells. The combination of complex requirements will envision the creation of ultimate biomimetic scaffolds for tissue regeneration.

Design Principles in Biomaterials and Scaffolds 591

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Nam, Y.S. and Park, T.G. (1999b). Adhesion behaviors of hepatocytes cultured onto biodegradable polymer surface modified by alkali hydrolysis process. J. Biomater. Sci. Polymer Ed. 10: 1145–1158. Nam, Y.S. and Park, T.G. (2000). A novel fabrication method for macroporous scaffolds using gas foaming salt as porogen additive. J. Biomed. Mater. Res. 53: 1–7. Niemann, C. (2005). Controlling the stem cell niche: right time, right place, right strength. Bio Essays 1: 1–5. Otsuka, H. and Kataoka, K. (2000). Surface characterization of functional polylactide through the coating with heterobifunctional poly(ethylene glycol)/polylacide block copolymers. Biomacromolecules 1: 29–48. Park, T.G. (2002). Perfusion culture of hepatocytes within galactose-derivatized biodegradable poly(lactide-co-glycolide) scaffolds prepared by gas foaming of effervescent salts. J. Biomed. Mater. Res. 59: 127–135. Park, A. and Griffith, L.G. (1998). Integration of surface modification and 3D fabrication techniques to prepare patterned poly(L-lactide) substrates allowing regionally selective cell adhesion. J. Biomater. Sci. Polymer Ed. 9: 89–110. Peattie, R.A. and Prestwich, G.D. (2006). Dual growth factor induced angiogenesis in vivo using hyaluronan hydrogel implants. Biomaterials 9: 1868–1875. Peppas, N.A. and Langer, R. (1994). New challenges in biomaterials. Science 263: 1715–1720. Peter, S.J. and Mikos, A.G. (1998). Osteoblastic phenotype of rat marrow stromal cells cultured in the presence of dexamethasone, b-glycerolphosphate, and L-ascorbic acid. J. Cell Biochem. 71: 55–62. Putnam, D. (2001). Polymer-based gene delivery with low cytotoxicity by a unique balance of side-chain termini. Proc. Natl. Acad. Sci. USA 98: 1200–1205. Raghunath, J. and Seifalian, A.M. (2005). Advancing cartilage tissue engineering: the application of stem cell technology. Current Opinion in Biotechnology 15: 503–509. Ramsey, W.S. and Binkowski, N.J. (1984). Surface treatments and cell attachment. In Vitro 20: 802–808. Rasmussen, J.R. and Whitesides, G.M. (1977). Introduction, modification, and characterization of functional groups on the surface of low density polyethylene films. J. Am. Chem. Soc. 99: 4736–4745. Ratner, B.D. (1996). Biomaterials Science. San Diego: Academic Press, 11–35. Reil, T.D. and Gelabert, H.A. (1999). Dexamethasone suppress vascular smooth muscle cell proliferation. J. Surg. Res. 85: 109–114. Richardson, T.P. and Mooney, D.J. (2001). Polymeric system for dual growth factor delivery. Nat. Biotechnol. 19: 1029–1034. Rosso, F. and Barbarisi, A. (2005). Smart materials as scaffolds for tissue engineering. J. Cell Physiol. 203: 465–470. Royce, S.M. and Marra, K.G. (2004). Incorporation of polymer microspheres within fibrin scaffolds for controlled delivery of FGF-1. J. Biomater. Sci. Polymer Ed. 15: 1327–1336. Saltzman, W.M. (1999). Delivering tissue regeneration. Nat. Biotechnol. 17: 534–535. Schugens, C. and Teyssie, P. (1996). Poly-lactide macroporous biodegradable implants for cell transplantation. II. Preparation of polylactide foams by liquid–liquid phase separation. J. Biomed. Mater. Res. 30: 449–461. Shakesheff, K. and Langer, R. (1998). Creating biomimetic micro-environment with synthetic polymer–peptide hybrid molecules. J. Biomater. Sci. Polym. Ed. 9:507–518. Uhrich, K.E. (1999). Polymeric systems for controlled drug release. Chem. Rev. 99: 3181–3198. Varkey, M. and Uludag, H. (2004). Growth factor delivery for bone tissue repair: an update. Expert. Opin. Drug Deliv. 1: 19–36. Vasita, R. and Katti, D.S. (2006). Growth factor delivery systems for tissue engineering: a materials perspective. Expert. Rev. Med. Dev. 1: 29–47. Weisz, O.A. and Schnaar, R.L. (1991). Hepatocyte adhesion to carbohydrate-derived surfaces II. Regulation of cytoskeletal organization and cell morphology. J. Cell Biol. 115: 495–504. Wissink, M.J.B. and Feijen, J. (2000). Improved endothelialization of vascular grafts by local release of growth factor from heparinized collagen matrices. J. Contr. Release 64: 103–114. Whang, K. and Nuber, G.A. (1995). Novel methods to fabricate bioabsorbable scaffolds. Polymer 36: 837–842. Whang, K. and Healy, K.E. (2000). A biodegradable polymer scaffold for delivery of osteotropic factors. Biomaterials 21: 2535–2551. Yoon, J.J. and Park, T.G. (2001). Degradation behaviors of biodegradable macroporous scaffolds prepared by gas foaming of effervescent salts. J. Biomed. Mater. Res. 55: 401–408.

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Yoon, J.J. and Park, T.G. (2002). Surface immobilization of galactose onto aliphatic biodegradable polymers for hepatocyte culture. Biotech. Bioeng. 78: 1–10. Yoon, J.J. and Park, T.G. (2003). Dexamethasone releasing biodegradable polymer scaffolds fabricated by a gas foaming/salt leaching method. Biomaterials 24: 2323–2329. Yoon, J.J. and Park, T.G. (2004). Immobilization of cell adhesive RGD peptide onto the surface of highly porous biodegradable polymer scaffolds fabricated by gas foaming/salt leaching method. Biomaterials 25: 5613–5620. Yoon, J.J. and Park, T.G. (2006). Heparin immobilized biodegradable porous scaffolds for sustained release of angiogenic growth factors. Biomaterials 79A(4): 934–942. Yoo, H.S. and Park, T.G. (2005). Hyaluronic acid modified biodegradable scaffolds for cartilage tissue engineering. Biomaterials 26: 1925–1933.

34 Naturally Occurring Scaffold Materials Stephen F. Badylak

INTRODUCTION Most regenerative medicine approaches to the restoration and replacement of damaged or missing tissues require a scaffold material upon which cells can attach, migrate, proliferate, and/or differentiate, hopefully into a functionally and structurally appropriate tissue. A variety of scaffold materials are available including synthetic polymers, and naturally occurring polymers that are produced during the course of tissue development in both vertebrate and invertebrate species. These various materials are characterized by unique physical and mechanical properties and each material is associated with a distinctive tissue response when implanted in a mammalian host. Synthetic scaffold materials such a poly(L)-(lactic acid) and poly(glycolic acid) have received considerable attention for tissue engineering applications and have shown promise in preclinical animal studies. Synthetic materials have predictable mechanical and physical properties and can be manufactured with great precision. However, synthetic materials tend to elicit a chronic active inflammatory response within the host tissue, which limits constructive remodeling and tissue regeneration, and promotes the deposition of fibrous connective tissue. The present chapter will not deal further with synthetic materials, but will instead focus upon naturally occurring scaffold materials. Naturally occurring scaffold materials are defined as those that occur in nature and are produced by the cells of living organisms. These materials typically occupy an extracellular location; that is, they become part of the extracellular matrix (ECM). Individual components of the ECM such as collagen or the intact matrix itself can be harvested and prepared for use as a scaffold for a variety of regenerative medicine applications. The present chapter will describe the use of three such materials as scaffolds; specifically purified collagens, chitosan, and intact extracellular matrix. Other naturally occurring materials such as hyaluronic acid and alginates have also shown potential as useful scaffold materials, but will not be discussed herein. COLLAGEN The most common and abundant naturally occurring scaffold material is the structural protein collagen. Collagen is a highly conserved protein that is ubiquitous among mammalian species and accounts for approximately 30% of all body proteins (Nimni et al., 1987). Inherent common amino acid sequences and epitope structures exist within collagen molecules across species lines (Boyd et al., 1991; Garrone et al., 1993; Beier et al., 1996). These common antigens appear to account for the lack of an adverse immune response when

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xenogeneic collagen is used as an implantable scaffold material. Bovine and porcine type I collagen provide readily available sources of scaffold material for numerous clinical applications and have proven quite compatible with most human systems. Examples of collagen scaffolds include Autologen® (Collagenesis Corporation, Beverly, MA), Contigen® (C.R. Bard, Inc., Covington, GA), Zyplast® and Zyderm I® and II® (INAMED Aesthetics, formerly McGhan Medical Corporation, Fremont, CA), and the Collagen Meniscal Implant (CMI) (ReGen Biologics, Inc., Franklin Lakes, NJ). In its native state, collagen is a natural substrate for cellular attachment, growth, and differentiation. In addition to its desirable structural properties, collagen has inherent functional properties such as the stimulation or inhibition of angiogenesis (Cornelius et al., 1998; Maeshima et al., 2000; Brennan et al., 2006), and the promotion of cellular proliferation and differentiation. For the above-mentioned reasons and others, collagen has become a favorite substrate for many tissue engineering and regenerative medicine applications. Collagen can be extracted from tissues such as tendons and ligaments, solubilized, and then reconstituted into fine strands that can, in turn, be fashioned into a variety of shapes and sizes that mimic body structures such as heart valves, blood vessels, and skin (Berthiaume et al., 1995). The reconstituted collagen is usually stabilized by chemical cross-linking methods and must be sterilized prior to surgical use. Allogeneic and xenogeneic collagen is generally recognized as “self ” tissue when used as a biologic scaffold material regardless of its species of origin, and it is subjected to the fundamental biological processes of degradation and integration into adjacent host tissues when left in its native ultrastructure. Certain processing methods, however, can alter the mechanical and physical properties of collagen-based materials and may negatively affect the processes of host-cell attachment, proliferation, differentiation, and tissue remodeling. These methods include glutaraldehyde treatment, carbodiimide treatment, dye-mediated photooxidation, exposure to polyepoxy compounds, and glycerol treatment. Commonly used methods of terminal sterilization include gamma or electron beam irradiation, or ethylene oxide treatment. Exposure to chemical crosslinking agents can change a biocompatible collagen-based material into a form that incites a host foreign body response (Sato, 1983). Most methods of chemical cross-linking alter (i.e. usually decrease) the rate of in vivo degradation and change the mechanical properties (i.e. usually strengthen) of collagen. Collagen provides considerable mechanical strength in its natural polymeric state. The necessary and required mechanical and physical properties of tissue engineered products for use in cardiovascular, orthopedic, and other body systems often depend upon the chemical manipulation of collagen-based scaffolds to achieve the desired mechanical properties. The tissue and species source of collagen and its treatment prior to use are important variables in the design of tissue engineered devices.

CHITOSAN Chitosans are the second most abundant biopolymer in nature and represent a family of biodegradable cationic polysaccharides consisting of glucosamine and randomly distributed N-acetylglucosamine linked in a β(1–4) manner (Dornish et al., 2001), and have a chemical structure similar to hyaluronic acid. Chitosans are derived by the alkaline N-deacetylation of chitin, a component of the protective layer of shellfish. The molecular weight of chitosan ranges from 300 to over 1,000 kD, depending on the preparation procedure and the degree of deacetylation, where the degree of deacetylation is defined as the ratio of glucosamine and N-acetylglucosamine (Madihally and Matthew, 1999; Dornish et al., 2001). The degree of deacetylation of commercially available chitosan can vary from 50% to 90%, while degrees of deacetylation higher than 95% can be achieved using acetylation chemistry methods (Mima et al., 1983; Madihally and Matthew, 1999; Cao et al., 2005). Chitosan, in its crystalline form, is generally insoluble in solutions with a pH of 7 and above; however, in dilute acids of pH less than 6, the free amino groups are protonated, allowing chitosan to form a viscous

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solution which can then be molded into various structures (e.g. blocks, tubes, beads, membranes) (Aiedeh et al., 1997; Madihally and Matthew, 1999; Dornish et al., 2001; Cao et al., 2005; Freier et al., 2005). The formation of a porous structure is generally achieved by freezing and then lyophilizing a chitosan construct, leaving pores in the space originally occupied by frozen solvent crystals. The size, distribution, and orientation of the pores can be controlled by varying the freezing method (ice crystal size, temperature gradient, and freezing rate) (Madihally and Matthew, 1999). Pore size in chitosan scaffolds is easily controlled in the range of 40–250 μm and porosities greater than 80% can be achieved (Madihally and Matthew, 1999). Porous chitosan scaffolds can also be formed by various processes which do not involve lyophilization (Chow and Khor, 2000; Ho et al., 2004; Geng et al., 2005). One such method is the freeze-gelation method, in which frozen chitosan solution is placed in an NaOH/ethanol solution at –20 C in order to adjust the pH so that gelation of the chitosan can occur at a temperature less than the freezing point of the solution, thus allowing for the formation of a chitosan scaffold while retaining a porous structure, without the necessity of lyophilization (Ho et al., 2004). In general, porous chitosan membranes possess a low elastic modulus (0.1–0.5 MPa) and tensile strength (30–60 kPa), while the extensibility can range from 30% to 110% based on pore size and orientation (Madihally and Matthew, 1999; Suh and Matthew, 2000; Di Martino et al., 2005). The mechanical properties of other configurations of chitosan (tubes, blocks, and beads) vary depending on the size, shape, and pore characteristics of the scaffold. The design and production of porous chitosan scaffolds have been extensively reviewed in Madihally and Matthew (1999). Chitosan can be enzymatically degraded in vitro using chitinase, chitosanase, lysozyme, and pectinase. Some other proteolytic enzymes have also been shown to have low-level degradation effects (Tomihata and Ikada, 1997; Jolles and Muzzarelli, 1999; Khor, 2001). In vivo, chitosan is degraded primarily by lysozyme into oligosaccharides through the hydrolysis of acetylated residues (Tomihata and Ikada, 1997; Zhang and Neau, 2001; Huang et al., 2004). The in vivo degradation products of chitosan are non-toxic and non-immunogenic (Muzzarelli, 1993). The degree of deacetylation has been shown to play an important role in the rate of degradation of chitosan materials; an important consideration for tissue engineering applications. Studies have shown that, when implanted subcutaneously in a rat model, chitosan materials with a degree of deacetylation of less that 70% were readily degraded in vivo while those with a degree of deacetylation of greater than 70% degraded more slowly (Tomihata and Ikada, 1997; Zhang and Neau, 2001). The degree of deacetylation of chitosan materials has also been shown to be directly related to the ability of the material to support cell attachment, with a higher degree of deacetylation being more favorable for cell attachment (Mao et al., 2004; Cao et al., 2005). Chitosan, due to its cationic nature and high charge density in solution, is able to interact with glycosaminoglycans and other negatively charged particles, including various water soluble anionic polymers (Denuziere et al., 1998; Gaserod et al., 1998; Di Martino et al., 2005; Raman et al., 2005; Chen et al., 2006; Mi et al., 2006). This property has been shown to allow the immobilization of glycosaminoglycans on the surface of chitosan (Denuziere et al., 1998; Madihally and Matthew, 1999; Raman et al., 2005; Mi et al., 2006). These glycosaminoglycans can then, via various pathways, influence cell adhesion, migration, proliferation, and differentiation (Takahashi et al., 1990). Furthermore, the N-acetylglucosamine moiety on chitosan is analogous to that on glycosaminoglycans and suggests that additional biological activity may be attributed to this naturally occurring scaffold. The in vivo tissue response to various chitosan-based implant materials is consistent with an acute to subacute inflammatory reaction (Nishimura et al., 1984; Muzzarelli et al., 1988; Muzzarelli et al., 1989; Damour et al., 1994; Peluso et al., 1994; Muzzarelli, 1997). Chitosan oligosaccharides have been shown to modulate macrophage response through interactions via their acetylated residues (VandeVord et al., 2002). Both chitosan and chitin have been shown to be chemoattractants for neutrophils in vitro and in vivo (Leuba and

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Stossel, 1986; Iida et al., 1987; Muzzarelli et al., 1990), resulting in a high concentration of neutrophils at the site of implantation during the first 7 days post implantation. However, the neutrophil population dissipates thereafter, and a chronic inflammatory response does not develop (Chen et al., 2006). In most cases, when chitosan is used in vivo, little or no fibrous encapsulation is observed nor does chitosan elicit the multinucleate giant cell or chronic mononuclear cell presence that is typically associated with an adverse foreign body response (Suh et al., 2000). Granulation tissue accompanied by accelerated angiogenesis in response to chitosan implantation has been reported (Chen et al., 2006). The in vivo response to chitosan in tissue engineering applications has been reviewed (Suh et al., 2000; Khor and Lim, 2003; Di Martino et al., 2005). Chitosan has been used as a conduit for guided peripheral nerve regeneration (Jenq and Coggeshall, 1987; Aebischer et al., 1990; Knoops et al., 1990; Guenard et al., 1991; Kim et al., 1993; den Dunnen et al., 1995; Rodriguez et al., 1999; Wang et al., 2005) and as a scaffold for the treatment of experimentally induced skin wounds with good results (Ueno et al., 1999; Ueno, 2001a, b; Chen et al., 2002; Tanabe et al., 2002; Mizuno et al., 2003). Cartilage repair (Di Martino et al., 2005) and bone tissue engineering applications (Lee et al., 2002; Bumgardener, 2003a, b; Wang et al., 2004) of chitosan have also been investigated. In summary, a significant body of work has been conducted with chitosan as a naturally occurring scaffold for tissue engineering applications and perhaps more is known about its chemistry, degradation, and host tissue response than any of the synthetic or naturally occurring scaffold materials.

INTACT EXTRACELLULAR MATRIX AS A SCAFFOLD MATERIAL The use of intact ECM, derived via the decellularization of various tissues and organs, has received considerable attention in the past 15 years. The ECM consists of the naturally occurring milieu of structural and functional molecules that are secreted by the resident cells of each tissue and organ; thus, there is a unique ECM composition and ultrastructure for each tissue and organ. The ECM even varies from location to location within various tissues such as the endocrine versus exocrine loci within the pancreas, or the valvular versus mural loci within the heart. The molecular motifs for cell attachment, migration, and differentiation are tissue specific, and attempts to mimic this compositional and structural complexity by synthetic methods have achieved very limited success. Naturally occurring ECM is one of the scaffold materials that has achieved commercial success for tissue engineering applications. ECM scaffolds derived from human dermis (Wainwright, 1995; Isch et al., 2000; Clemons et al., 2003), porcine and human urinary bladder (Duel et al., 1996; Atala, 1998; Dahms et al., 1998), porcine small intestinal submucosa (SIS) (Oelschlager et al., 2003; Wang et al., 2003; Badylak, 2004; Derwin et al., 2004; Musahl et al., 2004), porcine heart valves (Cohn et al., 1989; Hammermeister et al., 1993; Simon et al., 2003), and bovine dermis (Barber et al., 2006; Coons and Barber, 2006), among others, have all been used in human clinical applications. Methods for the decellularization of these tissues have recently been reviewed (Gilbert et al., 2006), and although complete elimination of all cellular remnants from any tissue is unlikely, the biologic response to scaffold materials composed of ECM is not characterized by immune-mediated rejection, even when the ECM is of xenogeneic origin (Allman et al., 2001; Allman et al., 2002; Palmer et al., 2002). Few studies have examined the host immune response to ECM scaffolds. Such studies have probably been the most extensive with allogeneic and xenogeneic heart valves and SIS. These studies have shown that host immune recognition of the ECM material does indeed occur, but is of Th-2 (accommodation) type of response rather than a Th-1 (cell-mediated rejection) type of response (Allman et al., 2002). In addition, although small amounts of the galactosyl 1,3 galactose (i.e. “GAL-epitope”) can be found in ECM scaffolds of porcine origin, they are not of sufficient amount to activate a complement in human serum (McPherson et al., 2000; Raeder et al., 2002). T-lymphocyte suppression has been found in some in vitro studies and this

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phenomenon has been suggested as a contributing factor to the absence of an adverse immune response when ECM is used as a biologic scaffold material. Preclinical studies have shown that immune challenges with inactivated influenza virus, bovine serum, albumin, and other antigens cause identical responses in mice that have been implanted with SIS versus mice that have not been exposed to SIS; that is, no systemic immune suppression was found. The in vivo degradation of ECM scaffolds has been most thoroughly evaluated with porcine SIS. 14 C-labeling studies have shown that rapid degradation occurs following in vivo implantation of SIS that has not been chemically cross-linked. Approximately 60% of the SIS-ECM is degraded and removed from the implantation site by 28 days and virtually 100% of the SIS-ECM has been eliminated (mostly via urinary excretion) within 60 days (Badylak et al., 1998). The rapid degradation of the ECM scaffold material is likely to be responsible for the absence of a chronic inflammatory or foreign body type of tissue response when it is used as a scaffold for tissue reconstruction. Protein–protein cross-linking agents such as glutaraldehyde, carbodiimide, and diisocyanate convert degradable ECM scaffolds into non-degradable or slowly degradable scaffolds and, thus, elicit a chronic inflammatory or foreign body type of tissue response when implanted in mammalian hosts. Although mechanical properties (strength) can be enhanced by the use of such agents, this benefit occurs at the cost of diminished constructive remodeling in many applications (Valentin et al., in press). ECM-based scaffolds have been extensively evaluated in preclinical animal studies for numerous applications, including lower urinary tract reconstruction (Kropp et al., 1995; Badylak et al., 1998; Kropp et al., 1998), the treatment of dermal wounds (Lindberg and Badylak, 2001), and musculoskeletal tissue reconstruction (Hodde et al., 1997; Valentin et al., in press). Human clinical studies with ECM scaffolds have also included a broad range of clinical uses (De Ugarte et al., 2004; Alpert et al., 2005; Dedecker et al., 2005; Helton et al., 2005; Jones et al., 2005a, b; Sievert et al., 2005; Smith et al., 2005; Zalavras et al., 2006). The host response to ECM scaffolds includes angiogenesis, mononuclear cell infiltration, and the deposition of new ECM by host cells that assume residence at the site of scaffold degradation (Voytik-Harbin et al., 1997; Badylak et al., 1999; Hodde et al., 2000; Badylak et al., 2002; Badylak, 2002; Valentin et al., in press). These biologic phenomena are thought to be the result of released growth factors and cytokines during scaffold degradation and the response to biologically active degradation products of the parent molecules within ECM (Sarikaya et al., 2002; Li et al., 2004). In summary, biologic scaffolds composed of extracellular matrix show promise for numerous surgical applications. Of the scaffolds reviewed in this chapter, both collagen and ECM biomaterials have been successfully translated into devices currently used in human patients. Optimization of the applications will depend upon a more thorough understanding of the mechanisms of action and the effect of various processing methods upon the in vivo remodeling outcomes.

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35 Synthetic Polymers M.C. Hacker and A.G. Mikos

INTRODUCTION Regenerative medicine is an emerging, interdisciplinary approach to repairing or replacing damaged or diseased tissues and organs. In order to re-establish tissue and organ function impaired by disease, trauma, or congenital abnormalities, regenerative medicine employs cellular therapies, tissue engineering strategies, and artificial or biohybrid organ devices. Typically, these techniques rely on combinations of cells, genes, morphogens, or other biological building blocks with bioengineered materials and technologies to address tissue or organ insufficiency. Materials used in these approaches range from metals and ceramics, to natural and synthetic polymers, as well as micro- and nanocomposites thereof. When used in a three-dimensional context, these materials are processed into micro- and/or nanoporous cell carriers, typically addressed as scaffolds, of various structures and properties, a topic that is discussed elsewhere in this book. This chapter focuses exclusively on synthetic polymers used in regenerative medicine. Some synthetic derivatives of natural materials are briefly discussed where appropriate. Accompanying the various facets of regenerative medicine, a plethora of synthetic polymers with different compositions and physicochemical properties have already been developed and investigated; however, research is still ongoing. Synthetic materials play a key role in many applications of regenerative medicine, including implants, tissue engineering scaffolds, and orthopedic fixation devices. In a broader sense, sutures, drug delivery systems, non-viral gene delivery vectors, and sensors made from synthetic polymers are further examples. This chapter provides a structural overview of these synthetic polymers and discusses their physicochemical characteristics, structure property relationships, applications, and limitations. Synthetic polymers that are hydrolytically labile and erode (biodegradable polymers) as well as those that are bioinert and remain unchanged after implantation (non-degradable polymers) are considered. It is the authors’ intention to provide a thorough overview over the synthetic material classes available. Some polymer classes are briefly mentioned and their chemical structures are provided, other more relevant materials are discussed in more detail. For most polymer classes and properties, reviews are referenced to present guidance to further reading. Biomaterial history in general can be best organized into four eras: prehistory, the era of the surgeon hero (first generation biomaterials), designed biomaterials and engineered devices (second generation biomaterials), and the contemporary era leading into the new millennium (third generation biomaterials) (Hench and Polak, 2002; Ratner, 2004). As far back as 600 AD, the use of dental implants made from materials like seashells or iron was reported. Also, there is evidence that sutures have been used for as long as 32,000 years to close large wounds. The word “biomaterials,” however, was first introduced within the last 50 years. Almost at the same time, aided by rapid advancements in industrial polymer development and synthesis, the exploration of synthetic polymers for biomedical applications began. The development of plastic contact lenses, utilizing primarily poly(methyl methacrylate) (PMMA), started around 1936, and the first data on implantation of nylon as a suture was reported in 1941. This development was accompanied by studies on the biocompatibility of the new materials. From the beginning, differences in foreign body reaction to materials like nylon and Teflon®,

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which exhibited a very mild response, became apparent. Additives such as plasticizers, unpolymerized reactants, and degradation products were discussed as possible causes leading to awareness of polymer quality for biomedical applications and biocompatibility testing. At the end of World War II, a wide variety of durable high-performance metal, ceramic, and especially polymeric materials was available inspiring surgeons to break new grounds in replacing diseased or damaged body parts. Materials including silicones, polyurethanes (PUs), Teflon®, nylon, methacrylates, titanium, and stainless steel were available “off-the-shelf” for surgeons to apply to medical problems (Ratner, 2004). Primarily medical and dental practitioners, driven by the vision to replace lost organ or tissue functionality, made use of minimal government regulatory activity and negligible human subject protections to develop and improvise replacements, bridges, conduits, and even organ systems based on such materials. Those pioneering approaches laid the foundation for novel procedures and engineered biomaterials. Such early implants made from materials available “off-the-shelf” in part proved to be either pathogenic or toxic. With a developing understanding of the immune system and foreign body reaction, a first generation of materials was developed during the 1960s and 1970s by engineers and scientists for use inside the human body. The primary goal of early biomaterial development was to achieve a suitable combination of physical properties to match those of the replaced tissue with a minimal toxic response in the host (Hench, 1980). Following this paradigm, more than fifty implanted devices made from forty different materials were in clinical use in 1980. In the early 1980s, research began to shift from materials that exclusively exhibited a bioinert tissue response to materials that actively interacted with their environment. Another advance in this second generation was the development of biodegradable materials that exhibited controllable chemical breakdown into non-toxic degradation products, which were either metabolized or directly eliminated. Biodegradable synthetic polymers were designed to resolve the interface problem, since the foreign material is ultimately replaced by regenerating tissues and eventually the regeneration site is histologically indistinguishable from the host tissue. Resorbable polymers were routinely used clinically as sutures by 1984. Other applications in fracture fixation aids or drug delivery devices emerged quickly. Despite considerable clinical success of bioinert, bioactive, and resorbable implants, there is still a high long-term prostheses failure rate and need for revision surgery (Ratner, 2004). Improvements of first and second generation biomaterials have been limited for one main reason: unlike living tissue, artificial biomaterials cannot respond to changing physiological loads or biochemical stimuli. This limits the lifetime of artificial body parts. To overcome these limitations, a third generation of biomaterials is being developed that involves molecular tailoring of resorbable polymers for specific cellular responses. By immobilizing specific biomolecules, such as signaling molecules or cell-specific adhesion peptides or proteins, onto a material it is possible to mimic the extracellular matrix (ECM) environment and provide a cell-adhesive surface (Hench and Polak, 2002; Drotleff et al., 2004; Lutolf and Hubbell, 2005). Biomimetic surfaces are promising tools to control cell adhesion, implant integration, cell differentiation, and tissue development. Synthetic polymer matrices can also be tailored to deliver drug, signaling molecules, and genetic code and thus provide versatile technologies for regenerative medicine (Saltzman and Olbricht, 2002; Segura and Shea, 2002; Tabata, 2003). Constantly expanding knowledge of the basic biology of stem cell differentiation and the corresponding signaling pathways as well as tissue development provide the basis for molecular design of scaffolds. In tissue engineering attempts, which aim at regenerating lost or defective tissue by transplanting in vitro engineered tissue constructs based on a patient’s own cells, one no longer attempts to closely match scaffold mechanical properties to those of the replaced tissue. It is rather considered important that the transplanted construct is engineered to be steadily remodeled in vivo to resemble the histological and mechanical properties of the surrounding tissue (Nerem, 2006). Due to this paradigm shift, mechanically labile hydrogels, especially injectable systems that can be used to directly encapsulate cells, have gained great importance as basis for biomimetic cell carriers. Hydrogels are characterized by a high water content that allows encapsulated cells to survive and enables

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sufficient passive transport of nutrients, oxygen, and wastes. Hydrogel-forming materials typically offer functional groups for chemical modifications, and their degradation can be controlled by chemical composition and crosslinking content. In the following sections inert and biodegradable synthetic polymers representative of all three generations will be presented. Their structure, synthesis, physicochemical properties, and applications will be described.

NON-DEGRADABLE SYNTHETIC POLYMERS A common characteristic of most non-degradable synthetic polymers is their biological inertness (Hench and Polak, 2002). These materials were developed to reduce to a minimum the host response to the biomaterial. Non-degradable synthetic polymers provide the basis for a plethora of medical devices as diverse as suture materials, orthopedic implants, fracture fixation devices, and catheters and dialysis tubing. These materials are also applied as implantable carriers for the long-term delivery of drugs (e.g. contraceptive hormones). Despite their excellent biological inertness and well adjustable mechanical properties, orthopedic implants made from non-degradable synthetic polymers and non-degradable bone cements ultimately fail at a high rate from problems at the interface arising from a lack of integration with the surrounding tissue, infections, or bone resorption caused by stress shielding (Bobyn et al., 1992; Jacobs et al., 1993). Major groups of non-degradable synthetic polymers are highlighted in the following paragraphs. Polymers with a 9C9C9 Backbone Polyethylene and Derivatives Poly(ethylene), Poly(propylene), and Poly(styrene) Poly(ethylene) (PE) (Figure 35.1a), poly(propylene) (PP) (Figure 35.1b), and poly(styrene) (PS) (Figure 35.1c) are ubiquitous industrial polymers and have been applied as biomaterials. All three thermoplastic polymers, which only consist of carbon, are synthesized by direct polymerization of their corresponding monomers. While PE can be synthesized by radical or ionic polymerization of ethylene, special organometallic catalysts are required to polymerize propylene to useful PP. PE and PP are classified into several different categories based on their density, branching, and molecular weight. These parameters significantly influence the crystallinity and mechanical properties of the polymers. PE has been used for the production of catheters. High-density PE, which is characterized by a low degree of branching and thus strong intermolecular forces and tensile strength, has been processed into highly durable hip prostheses. A three-dimensional fabric comprising PE fibers and coated with hydroxyapatite was used to regenerate hyaline cartilage in osteochondral defects in rabbit knees and showed successful biocompatibility (Hasegawa et al., 1999). The best-known application for PP is its use for syringe bodies. Copolymers of PE and vinyl acetate (poly(ethylene-co-vinyl acetate), PEVAc) (Figure 35.1d) are widely used in non-degradable drug delivery devices (Langer, 1990). PEVAc is one of the most biocompatible implant materials (Langer et al., 1981a) and has been approved by the FDA for use in implanted and topically applied devices. Ocusert® and Progestasert® are prominent examples for PEVAc-based drug delivery systems (Chandrasekaran et al., 1978). PS is a hard and brittle polymer used for the fabrication of tissue culture flasks and dishes. By copolymerization with butadiene, copolymers with improved elasticity are synthesized that are used for the fabrication of catheters and medical devices for perfusion and dialysis. Poly(tetrafluoroethylene)

Poly(tetrafluoroethylene) (PTFE) (Figure 35.1e), well known as Teflon® (DuPont), can be synthesized from liquid tetrafluoroethylene by radical polymerization and through fluorination of PE. Among known

Synthetic Polymers 607

H

H CH3

H H n

*

*

n

*

H H

H n

*

*

H

H H

(a) Poly(ethylene) H H

(b) Poly(propylene) H

n H

* H

(c) Poly(styrene)

H m H O

*

F

*

F n

* F

O

*

F

CH3 (d) Poly(ethylene-co-vinyl acetate)

(e) Poly(tetrafluoroethylene) H H

H

H n

*

*

H H

* O

H3C O

H3C

n

*

n

* H HN

O

O

O CH3

CH3 H3C

OH (f) Poly(methyl methacrylate)

*

(g) Poly(2-hydroxyethyl methacrylate)

(h) Poly(N-isopropylacrylamide)

Figure 35.1 Chemical structures of non-degradable synthetic polymers (I). polymers, PTFE has the lowest coefficient of friction, has excellent resistance to chemicals, and is well hemocompatible. Porous PTFE fiber meshes (Goretex®) have become a popular synthetic vascular graft material (Xue and Greisler, 2003). Poly(meth)acrylates and Polyacrylamides Poly(meth)acrylate hydrogels have found applications in medical devices, especially for ocular applications (e.g. contact lenses and intraocular lenses), as drug delivery systems and as cell delivery systems (Langer and Peppas, 1981; Peppas et al., 2000; Lloyd et al., 2001). Three major types, PMMA, poly(2-hydroxyethyl methacrylate) (PHEMA), poly(N-isopropylacrylamide), are discussed in more detail. For (meth)acrylic ester and acrylamide monomers, the typical monomers used for poly(meth)acrylate and polyacrylamide synthesis, respectively, a number of functional derivatives are available that, together with the free carboxylic acid group of (meth)acrylic acid, allow for the presentation of different functional groups along the polymer chains or within crosslinked hydrogels. Using an imprinting technique, these moieties can be oriented in a way that pouches are created which interact non-covalently with molecules (e.g. drugs or therapeutic peptides and proteins), by ionic interactions, hydrogen bonds, ππ interactions, and hydrophobic interactions (Mosbach and Ramstrom, 1996; Tunc et al., 2006). Besides intelligent hydrogels for controlled drug release this technology has impact on micro-fluidic devices, biomimetic sensors, intelligent polymeric membranes (Ulbricht, 2006), and analyte-sensitive materials (Byrne et al., 2002). Poly(methyl methacrylate)

PMMA (Figure 35.1f) is a non-degradable polyacrylate and is the most commonly applied non-metallic implant material in orthopedics. After being used as an essential ingredient in making dentures, PMMA was

608 BIOMATERIALS FOR REGENERATIVE MEDICINE

introduced to orthopedic surgery in the mid-1950s (Saha and Pal, 1984). PMMA tissue biocompatibility became further apparent when Plexiglas fragments were accidentally implanted in the eyes and other body tissues of World War II fighter pilots during aircraft crashes. PMMA can be in situ polymerized and crosslinked from a slurry containing PMMA and MMA monomers and is so used as a common bone grafting material mainly in the fixation of orthopedic prosthetic materials for hips, knees, and shoulders (Kenny and Buggy, 2003). PMMA-based bone cements can be mixed with inorganic ceramics or bioactive glass to modulate curing kinetics and enforce mechanical properties. Antibiotics can be loaded within the cement to reduce the risk of prosthesis-related infection. Significant drawbacks of self-curing PMMA cements include that they are not degraded, that their high curing temperatures and toxic monomers can cause necrosis of the surrounding tissue, and that the cements show limited interactions with the surrounding bone (Hendriks et al., 2004). Therefore, development of alternative injectable bone cements is directed toward biodegradable materials with improved curing properties and osteoconductive interfaces (Yaszemski et al., 1996; Hendriks et al., 2004). Due to its excellent bio- and hemocompatibility and ease of manipulation, PMMA is used in many medical devices, including blood pumps and dialyzers. Its optical properties make it a candidate material for implantable ocular lenses and hard contact lenses (Lloyd et al., 2001). PMMA also offers physical and coloring properties that are beneficial for denture fabrication (Hendriks et al., 2004). Poly(2-hydroxyethyl methacrylate)

PHEMA (Figure 35.1g) was the first hydrogel successfully employed for biological use (Wichterle and Lim, 1960). PHEMA has become the major component of most soft contact lenses and is also part of intraocular lenses (Lloyd et al., 2001). Due to their free hydroxyl groups, PHEMA gels contain relatively high amounts of water, facilitating the diffusion of solutes and oxygen. PHEMA has excellent biocompatibility which initiated the development of a plethora of HEMA-containing copolymers. Hydrogels fabricated from PHEMA and copolymers have been intensively characterized for controlled drug delivery applications (Mack et al., 1987; Lu and Anseth, 1999) and employed for biomedical uses. PHEMA gels, which have limited mechanical properties, have been used in attempts to reconstruct female breasts, nasal cartilages, and as artificial corneas as well as wound dressings (Young et al., 1998). In a subcutaneous rabbit model, porous PHEMA sponges promoted significant cellular ingrowth and neovascularization in combination with good cytocompatibility (Chirila et al., 1993). Recently, a mineralization technique has been demonstrated that exposes carboxylate groups on crosslinked PHEMA hydrogel scaffolds, promoting calcification (Song et al., 2003). Poly(N-isopropylacrylamide)

Poly(N-isopropylacrylamide) (PNiPAAm) (Figure 35.1h) has gained great significance for injectable applications in drug and cell delivery using minimally invasive techniques due to its unique physicochemical properties (Hoffman, 2002). PNiPAAm undergoes (lower critical) phase separation resulting in the formation of an opaque hydrogel in response to a temperature above 32°C, the material’s lower critical solution temperature (LCST). This thermoresponsive behavior is the result of strong hydrogen bonds between the polymer and water molecules and the specific molecular orientations of these bonds due to the molecular structure of the polymer. The formation of hydrogen bonds between the polymer and the solvent lowers the free energy of the solution. Due to the hydrophobic N-isopropyl residues in PNiPAAm, the hydrogen bonds between water and the amide functionality require specific molecular orientations, which lead to negative entropy changes and positive contributions to the free energy. Since the enthalpic contribution to the free energy is temperature dependent, the formation of strong but specifically oriented hydrogen bonds is no longer thermodynamically favored above a

Synthetic Polymers 609

O H3C CH3 H

O

n

OH

(a) Poly(ethylene glycol)

Si *

O n

n*

*

(b) Poly(dimethylsiloxane)

*

O

O

O (c) Poly(ethylene terephthalate)

Figure 35.2 Chemical structures of non-degradable synthetic polymers (II). certain temperature. Consequently, PNiPAAm dissolves in water below the LCST. At and above the LCST, the polymer chains partially desolvate and undergo a coil-to-globule transition resulting in colloidal aggregation that may lead to gel formation or polymer precipitation (Schild and Tirrell, 1990; Schild, 1992). Hydrogels formed by linear PNiPAAm at 32°C are instable and collapse substantially as the temperature is increased above the LCST. The synthesis of crosslinked networks and copolymers, typically with hydrophilic building blocks, has resulted in materials that demonstrate reversible thermogelation and form hydrogels without significant syneresis at body temperature. Different PNiPAAm-containing copolymers for cell delivery have been synthesized with acrylic acid, poly(ethylene glycol) (PEG), hyaluronic acid, and gelatin (Stile et al., 1999; Ohya et al., 2001; Hoffman, 2002; Morikawa and Matsuda, 2002). Detailed information is available for the in vitro and in vivo use of gelatin–PNiPAAm conjugates for the regeneration of articular cartilage (Ibusuki et al., 2003a, b). Polyethers PEG (Figure 35.2a), often also called poly(ethylene oxide) (PEO), is a non-degradable polyether of the monomer ethylene glycol. Technically, PEG and PEO should not be used as synonyms, since PEO is synthesized from the monomer ethylene oxide and typically terminated by only one hydroxyl group and an initiator fragment. Commonly, PEG is often used to refer to the polymer with molecular weight less than 50,000 Da while PEO is used for higher molecular weights. PEG is water soluble and solutions of its high molecular weight form can be categorized as a hydrogel. PEG hydrogels for biomedical applications are typically comprised of polymer chains that are crosslinked. These crosslinked networks frequently contain chemical bonds between the PEG chains and the crosslinkable moieties, which are prone to aqueous hydrolysis and are therefore characterized as biodegradable system. The molecular weight of the PEG chains crosslinked in such hydrogels is below a threshold molecular weight to allow for complete resorption by renal elimination of the individual chains. Consequently, these systems are discussed with biodegradable polymers in section “Biodegradable crosslinked polymer networks.” Favorable characteristics of PEG and PEO are their high hydrophilicity, bioinertness, and outstanding biocompatibility, which make them candidate biomaterials. PEG and PEO are frequently used as hydrophilic polymeric building blocks in copolymers with more hydrophobic degradable or non-degradable polymers for drug delivery (Jeong et al., 1997), gene delivery, tissue engineering scaffolds, medical devices, and implants. PEG has also been immobilized on polymeric biomaterial surfaces to make them resistant to protein absorption and cell adhesion. These effects are attributed to highly hydrated PEG chains on the polymer surfaces that exhibit steric repulsion based on an osmotic or entropic mechanism. Attempts to benefit from this phenomenon include the design of long-circulating nanoparticles or liposomes (Gref et al., 1997, 2000; Photos et al., 2003; Vonarbourg et al., 2006) and PEGylated enzymes or proteins with prolonged functional residence time in vivo compared to unmodified biomolecules (Roberts et al., 2002; Harris and Chess, 2003). A variety of PEG-containing block copolymers for injectable drug delivery have been developed over the last decades. The most prominent class are triblock copolymers composed of two hydrophilic PEO blocks and one hydrophobic poly(propylene oxide) (PPO), also known as Pluronics® or poloxamers. These materials are

610 BIOMATERIALS FOR REGENERATIVE MEDICINE

designed to show similar phase transition behavior as the thermogelling PNiPAAm-containing materials (section “Poly(N-isopropylacrylamide)”). Poloxamers have been intensively investigated for the delivery of drugs and proteins (Jeong et al., 2002). Since poloxamers are non-degradable, biodegradable structural analogs have been synthesized and are described within the next chapter on biodegradable synthetic polymers (section “Biodegradable synthetic polymers for regenerative medicine.”) Polysiloxanes Polysiloxanes, or silicones, are a general category of polymers consisting of a silicon–oxygen backbone with organic groups, typically methyl groups, attached to the silicon atoms (Colas and Curtis, 2004). Certain organic side groups can be used to link two or more chains together. By varying the 9Si9O9 chain length, side groups, and crosslinking extent, silicone with properties ranging from liquids to hard plastics can be synthesized. Silicone synthesis typically involves the hydrolysis of chlorosilanes into linear or cyclic siloxane oligomers, which are then polymerized into polysiloxanes by polycondensation or polymerization, respectively. The most common polysiloxane is linear poly(dimethylsiloxane) (PDMS) (Figure 35.2b). Polysiloxanes, which are characterized by unique material properties combining biocompatibility and biodurability, have found widespread application in health care (Curtis and Colas, 2004). The material’s high biodurability is a result of other material properties such as hydrophobicity, low surface tension, and chemical and thermal stability. Silicone surfaces have been found to inhibit blood from clogging for many hours and have been therefore used for the fabrication of silicone coated needles, syringes, and other blood-collecting instruments. Silicone materials have also been employed as heart valves and as components in kidney dialysis, bloodoxygenator, and heart-bypass machines due to their hemocompatibility. Silicone elastomers have found application in numerous catheters, shunts, drains, and tubular implants, such as artificial urethra. Significant orthopedic applications of silicone are hand and foot joint implants. The most prominent application of silicones is their extensive use as cosmetic implants in esthetic and reconstructive plastic surgery. Prosthetic silicone implants are available for the breast, scrotum, chin, nose, cheek, calf, and buttocks. Different silicone materials, including slightly crosslinked silicone gels, are combined to achieve a natural feel. Controversy aroused regarding the safety of popular silicone gel-filled breast implants in early 1990s. These discussions initially involved increased risk for breast cancer, then progressed to autoimmune connective tissue disease, and continued to evolve to the frequency of local or surgical complications such as rupture, infection, or capsular contracture. To date, no epidemiology study has indicated that the rate of breast cancer has significantly increased in women with silicone breast implants (Silverman et al., 1996). Similarly, studies on autoimmune or connective tissue disease agreed on a lack of causal association between breast implants and these diseases (Sanchez-Guerrero et al., 1995; Lewin and Miller, 1997). A safety concern that has been controversially discussed recently involves the amount of platinum (part of catalysts used during silicone synthesis) that is released from silicone implants and accumulated in the host organism (Arepalli et al., 2002; Brook, 2006). Other mentioned complications, especially implant rupture, are persisting problems; in 1992, the FDA restricted the use of silicone gel-filled implants. Since that time, the implants may be used only under certain controlled conditions. The pre-market approval, an application for marketing a device, has only been approved for two saline-filled breast implants and no silicone gel-filled implants by the FDA as of 2004 (US FDA, 2004). Polysiloxane gels, combining the high oxygen permeability of silicone and the comfort and clinical performance of conventional, polyacrylate hydrogels, enabled the fabrication of soft, gas permeable contact lenses for extended wear. In contrast to conventional hydrogels, silicone gels make the lens surface highly hydrophobic and less “wettable,” which frequently results in discomfort and dryness during lens wear. Surface modifications of the silicones or the addition of conventional hydrogels are suitable strategies to compensate for the hydrophobicity.

Synthetic Polymers 611

Overall, polysiloxanes have displayed expanded medical application since the 1960s and today are one of the most thoroughly tested and important biomaterials. Other Non-degradable Polymers Poly(ethylene terephthalate) Poly(ethylene terephthalate) (PET) (Figure 35.2c), a linear polyester synthesized by polycondensation of terephthalic acid and ethylene glycol, is typically processed into fiber meshes. These meshes are applied as vascular grafts (Xue and Greisler, 2003) or used to reinforce prostheses. Hydrolytically Stable Polyurethanes

PUs are a heterogeneous class of polymers that consist of organic units joined by urethane links (Figure 35.3). Generally, PUs can be synthesized from a bischloroformate and a diamine or by reacting a diisocyanate with a dihydroxy component. PUs used in biomedical applications typically have a segmented structure that results in useful physicochemical properties (Boretos and Pierce, 1967). Such segmented PUs or PU copolymers are elastomers composed of alternating polydispersed “soft” and “hard” segments. These two segments are thermodynamically incompatible and phase-segregate, resulting in discrete, crystalline domains of the associated “hard” segments surrounded by a continuous, amorphous phase of “soft” segments. The segregated domains

Components: P = (HO-RP-OH): D = (OCN-RD-NCO): C = (X-RC-X; X = OH, or NH2): dihydroxy terminated oligomer diisocyanate chain extender (diol or diamine)



Step1:

2 prepolymer 

Step2:



soft segment hard segment –(O–RP–O–(CO–NH–RD–NH–CO–X–RC–X)m–CO–NH–RD–NH–CO)x– –(P–(DC)mD)x– (a) Polyurethane synthesis P

D methylenebisphenyldiisocyanate

poly(tetramethyleneoxide)

C ethylenediamine

O *

O

soft segment

n

O N H

N H

N H

H N

H N

H N

m O

hard segment (b) Biomer® a polyurethaneurea

Figure 35.3 General synthesis scheme (a) and an example structure (b) for polyurethanes.

O O

x

*

612 BIOMATERIALS FOR REGENERATIVE MEDICINE

are stabilized by interchain hydrogen bonds and are responsible for the materials’ mechanical properties (Gunatillake et al., 2003). Segmented PUs are synthesized in a two-step process that provides control over polymer architecture (Figure 35.3a). The first step involves the synthesis of an isocyanate-terminated prepolymer from a diisocyanate (D in Figure 35.3) and a hydroxyl group terminated polyether or polyester (P in Figure 35.3). The prepolymer and excess diisocyanate is then reacted with a hydroxy or amine group terminated chain extender (C in Figure 35.3) to generate the final PU (Figure 35.3a). A chain extender terminated with hydroxy groups yields segmented PUs, while a diamine extender yields polyurethaneurea (Figure 35.3b). The “hard” segment of the PU copolymer is comprised of the diisocyanate and the chain extender, while the “soft” segment contains the polymeric segment introduced during the first step. The extent of phase separation is dependent on molecular weights, chemistry, and relative percentages of the building blocks (Fromstein and Woodhouse, 2006). After almost 50 years of use in biomedical applications, PUs remain one of the most popular group of biomaterials for the fabrication of medical devices. Their popularity results from a wide range of versatility with regard to tailoring their physicochemical and mechanical properties, blood and tissue compatibility, and degradative properties by altering block copolymer composition. PUs are traditionally applied as synthetic polymers in numerous medical devices, such as breast implants, catheters, vascular, and aortic grafts, pacemaker leads, artificial heart valves, and artificial hearts. For such applications, traditional PUs, such as Biomer® (P: polytetramethylene oxide; D: methylene bisphenylenediisocyanate; C: ethylenediamine) (Figure 35.3b), were materials of first choice. However, the assumption of polyetherurethane non-degradability had to be revised following well-documented failures of pacemaker leads and breast implant coatings containing PUs in the late 1980s. Although PUs can be designed to be stable against hydrolysis, these materials have been shown to degrade in the biological environment by mechanisms including oxidation and enzyme and cell-mediated degradation (Howard, 2002; Santerre et al., 2005; Fromstein and Woodhouse, 2006). Oxidation of PUs can be initiated by peroxides, free radicals, and enzymes. Metal-catalyzed oxidation was found to be most frequently associated with pacemaker lead failure. Another important oxidation driven problem with long-term PU implants is environmental stress cracking. It has also been found that PU surfaces become coated with a protein layer that enhances the adhesion of macrophages. The macrophages, activated by proteins of the complement family, release oxidative factors that accelerate degradation of the polymer (Stokes et al., 1995). Chemical design criteria for biostable PUs have been identified. To increase the degree of interchain hydrogen bonding, on which biostability depends in part, low molecular weight oligomeric diols (P) are preferred as building blocks. To avoid oligomer hydrolysis, oligoethers are favored over oligoesters. Aromatic diisocyanates (D) have been found to yield more biostable PUs than aliphatic diisocyanates. The use of a diamine chain extender (C) instead of a dihydroxy-terminated one typically results in stronger polyurethaneurea, but polymer fabrication is often hampered due to solubility problems. Using soft segment building blocks with high crystallinity, such as polycaprolactone, or employing silicone-based oligomers are also assumed to improve polymer biostability (Fromstein and Woodhouse, 2006). Biomedical PUs were found to perform well in a variety of in vivo applications and to generally have better blood and tissue compatibilities in comparison to numerous other synthetic polymers. The efficient removal of impurities from the polymer synthesis, such as catalyst residues and low molecular weight oligomers, has been found to critically determine PU biocompatibility (Gogolewski, 1989). PUs can be surface modified to reduce the risk of thrombosis or improve the interactions with cells and tissues. Different strategies, including adsorption, covalent grafting, or the use of self-assembled monolayers, have been applied to distribute proteins, such as fibronectin, or adhesion peptides, which contain the integrinbinding peptide motif RGD, across the PU surface (Lin et al., 1994; Fromstein and Woodhouse, 2006).

Synthetic Polymers 613

BIODEGRADABLE SYNTHETIC POLYMERS FOR REGENERATIVE MEDICINE Biodegradable synthetic polymers offer a number of advantages over non-degradable materials for applications in regenerative medicine. Like all synthetic polymers, they can be synthesized at reproducible quality and purity and fabricated into various shapes with desired bulk and surface properties. Specific advantages include the ability to tailor mechanical properties and degradation kinetics to suit various applications. Clinical applications for biodegradable synthetic polymers are manifold and traditionally include resorbable sutures, drug delivery systems, and orthopedic fixation devices such as pins, rods, and screws (Behravesh et al., 1999). More recently, synthetic biodegradables were widely explored as artificial matrices for tissue engineering applications (Seal et al., 2001; Nguyen and West, 2002; Salgado et al., 2004). For such applications, the mechanical properties of the scaffolds, which are determined by the constitutive polymer, should functionally mimic the properties of the tissue to be regenerated. Ultimately, the polymeric support is designed to degrade while transplanted or invading cells proliferate, lay down ECM, and form coherent tissue that, in the ideal case, is functionally, histologically, and mechanically indistinguishable from the surrounding tissue. To engineer scaffolds suitable for different applications, a wide variety of biodegradable polymers is required ranging from pliable, elastic materials for soft tissue regeneration to stiff materials that can be used in load-bearing tissues such as bone. In addition to the mechanical properties, the degradation kinetics of polymer and ultimately scaffold also have to be tailored to suit various applications. The major classes of synthetic, biodegradable polymers are briefly reviewed and their potential in regenerative medicine is discussed below. Polyesters Polyesters have been attractive for biomedical applications because of their ease of degradation by primarily non-enzymatic hydrolysis of ester linkages along the backbone. Additionally, degradation products can be resorbed through the metabolic pathways in most cases, and there is the potential to tailor the structure to alter degradation rates (Gunatillake and Adhikari, 2003). A vast majority of biodegradable polymers studied belong to the polyester family (Middleton and Tipton, 2000). Polyester fibers, which also became popular with the textile industry, were used as resorbable sutures (Freed et al., 1994). Promising observations regarding biocompatibility of the materials lead to applications in drug delivery, orthopedic implants, and most recently tissue engineering scaffolds, particularly for orthopedic applications (Heller, 1984; Amecke et al., 1992; Hubbell, 1995; Behravesh et al., 1999; Webb et al., 2004). Poly(α-hydroxy acids) The family of polyesters can be subdivided according to the structure of the monomers. In poly(α-hydroxy acids) each monomer carries two functionalities, a carboxylic acid and a hydroxyl group, located at the carbon atom next to the carboxylic acid (α-position), that form ester bonds. Poly(α-hydroxy acids) are linear thermoplastic elastomers that are typically synthesized by ring-opening polymerization of cyclic dimers of the building blocks. Poly(lactic acid) (PLA) (Figure 35.4a), poly(glycolic acid) (PGA) (Figure 35.4b), and a range of their copolymers (poly(lactic-co-glycolic acid), PLGA) (Figure 35.4c) are prominent representatives of not only biodegradable polyesters but of biodegradables in general. Poly(α-hydroxy acids) have a long history of use as synthetic biodegradable materials in a number of clinical applications. Initially, resorbable sutures were made from these materials (Cutright et al., 1971). Later, poly(α-hydroxy acids) were the basis for controlled release systems for drugs and proteins (Juni and Nakano, 1987; Brannon-Peppas, 1995; Jain, 2000) and orthopedic fixation devices. Langer and coworkers have pioneered the development of these polymers in the form of porous scaffolds for tissue engineering (Langer and Vacanti, 1993).

614 BIOMATERIALS FOR REGENERATIVE MEDICINE

O O

O

O

n*

*

O

O

CH3

yn O

O O

xO

O

y CH3

*

(c) Poly(L-lactic-co-glycolic acid)

(b) Poly(glycolic acid)

O H

O

CH3

(a) Poly(D,L-lactic acid)

O

x

*

n*

*

O *

n*

*

O

O

n*

CH3 (d) Poly(D,L-lactic acid)-block-poly(ethylene glycol) monomethyl ether CH3 *

O

O

CH3

O

O

O

O

R

(f) Poly(p-dioxanone)

(e) Poly(ε-caprolactone)

O

OR

O

O

n*

O

N H

n*

O

* (g) Poly(ortho ester)

O

O O

(h) Poly(amide carbonate)derived from desaminotyrosine and a tyrosine alkyl ester (R: alkyl) O

R

O

*

4 O O (i) Poly(anhydride), here: poly(SA–DPP)

O

n

*

N P

n*

R' (j) Poly(phosphazene)

Figure 35.4 Chemical structures of biodegradable synthetic polymers.

Due to the chiral nature of lactic acid, several forms of poly(lactid acid) exist: poly(L-lactic acid) (PLLA), for example, is synthesized from dilactide in the L form. The polymerization of racemic dilactide leads to poly(D,L-lactic acid) (PD,LLA), which is an amorphous polymer. PLLA, in contrast, is a semicrystalline polymer with a crystallinity of around 37%. PLLA is characterized by a glass transition temperature between 50°C and 80°C and a melting temperature between 173°C and 178°C. Amorphous PD,LLA is typically used in drug delivery applications, while semicrystalline PLLA is preferred in applications where high mechanical strength and toughness are required (e.g. for sutures and orthopedic devices). PGA is also a semicrystalline polymer with a higher crystallinity of 46–52%. Thermal characteristics of PGA are glass transition and melting temperatures of 36°C and 225°C, respectively. Because of its high crystallinity, PGA unlike PLA is not soluble in most organic solvents; the exceptions are highly fluorinated and highly toxic organic solvents such as hexafluoroisopropanol. Consequently, common processing techniques for PGA include melt extrusion, injection, and compression molding. PLA, PGA, and PLGA undergo homogeneous erosion via ester linkage hydrolysis into the degradation products lactic acid and glycolic acid, which are both natural metabolites that are excreted as carbon dioxide and water. Degradation of poly(α-hydroxy acids) was found to show typical characteristics of bulk erosion. Bulk erosion occurs when water penetrates the entire structure, and the device degrades simultaneously (Goepferich, 1996). During the initial stages of degradation almost no mass loss can be detected. Analysis of the average molecular weight of the polymer bulk over the same period, however, reveals a steady decrease in molecular weight. Once the polymer chains throughout the bulk are degraded below a certain threshold, the water-soluble degradation products are washed out and the system collapses accompanied by significant mass loss. Due to its well accessible ester group, PGA degrades rapidly in aqueous media. PGA sutures typically lose

Synthetic Polymers 615

their mechanical strength over a period of 2–4 weeks postoperatively (Reed and Gilding, 1981). In order to adapt these properties to a wider range of applications, copolymers with more hydrophobic PLA were synthesized and investigated. The two main series are those of PLLGA (Figure 35.4c) and PDLLGA. It has been shown that compositions in the 25–75% range for L-LA/GA and 0–70% for the DL-LA/GA are amorphous (Miller et al., 1977; Sawhney and Hubbell, 1990; Li, 1999; Middleton and Tipton, 2000; Gunatillake and Adhikari, 2003). For the PLLGA copolymers, the rate of hydrolysis was found to be slower at either extreme of the copolymers compositions range. It is generally accepted that intermediate PLGA copolymers are more unstable than either homopolymer. Besides polymer composition, the rate of degradation is affected by factors such as configurational structure, copolymer ratio, crystallinity, molecular weight, morphology, stresses, amount of residual monomer, bulk porosity, and site of implantation (Gunatillake and Adhikari, 2003). Multiple in vitro and in vivo studies that were conducted on the biocompatibility of PLA, PLGA, and PGA generally revealed satisfying results (Athanasiou et al., 1996). Consequently, PLA, PLGA copolymers, and PGA are among the few biodegradable polymers with FDA approval for human clinical use. Concerns with poly(α-hydroxy esters) typically focus on the accumulation of acidic degradation products within the polymer bulk that can have detrimental effects on encapsulated drugs in delivery applications (Brunner et al., 1999; Lucke et al., 2002) or can cause late non-infectious inflammatory responses when released in a sudden burst upon structure breakdown (Simon et al., 1997). This adverse reaction can occur weeks and months postoperatively and might need operative drainage. This is a major concern in orthopedic applications, where implants of considerable size would be required, which may result in release of degradation products with high local acid concentrations. Inflammatory response to poly(α-hydroxy acids) were found to be also triggered by the release of small particles during degradation that were phagocytized by macrophages and multinucleated giant cells (Anderson and Shive, 1997; Xia and Triffitt, 2006). In general, implant size as well as surface properties appear to be critical factors with regard to biocompatibility. Fewer concerns seem to exist toward the application of poly(α-hydroxy acids) in soft tissues compared to hard tissue applications (Athanasiou et al., 1996). Poly(α-hydroxy acids) were the materials of choice when one of the key concepts of tissue engineering, the de novo engineering of tissue by combining isolated cells and three-dimensional macro-porous cell carriers in vitro, was first realized and developed (Langer and Vacanti, 1993; Freed et al., 1997; Mooney and Mikos, 1999). Polymers based on lactic and glycolic acid are still popular scaffold materials especially for orthopedic applications, such as bone, cartilage, and meniscus, as outlined in several reviews (Agrawal et al., 2000; Hutmacher, 2000; Seal et al., 2001). Limitations of this class of materials include insufficient mechanical properties with regard to load-bearing applications (Webb et al., 2004) and inflammatory or cytotoxic events due to above-mentioned accumulation of acidic degradation products. In order to cover a broader range of mechanical and physicochemical properties, such as water absorption, polymer degradation, and polymer–drug interactions, block copolymers containing PLA and hydrophilic PEO or PEG were synthesized for drug delivery applications (Bouillot et al., 1998). Solid particulate systems from these block copolymers were found to be almost invisible to the immune system due to the hydrophilic PEG chains that swell on the surface (Gref et al., 1994; Bazile et al., 1995) (section “Polyethers”) (Figure 35.4d). The stealthiness of such surfaces is mainly caused by the suppression of protein adsorption, which also inhibits cell adhesion. Investigations of cell adhesion to PEG–PLA diblock copolymer surfaces revealed that cell adhesion can be controlled and cell differentiation can be modulated by the PEG content (Lieb et al., 2003). With the objective to specifically control cell–polymer interactions, PEG–PLA copolymers were further developed to allow for the covalent attachment of signaling molecules (Cannizzaro et al., 1998; Tessmar et al., 2003). Since these polymers were insoluble in water, they could be processed into macroporous scaffolds for tissue engineering applications (Hacker et al., 2003).

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Polylactones The most prominent and thoroughly investigated polylactone is poly(ε-caprolactone) (PCL) (Figure 35.4e), an aliphatic, semicrystalline polyester with an interestingly low glass transition temperature (–60°C) and melting temperature (59–64°C) (Middleton and Tipton, 2000). PCL is considered biocompatible (Matsuda et al., 2003). PCL is prepared by the ring-opening polymerization of the cyclic monomer ε-caprolactone, and is compatible with a range of other polymers. Catalysts, such as stannous octoate, are used to catalyze the polymerization and low molecular weight alcohols can be used as initiator and to control the molecular weight of the polymer. ε-caprolactone can be copolymerized with numerous other monomers. Copolymers with PLA and PEG are probably the most noteworthy and have been investigated extensively (Pitt C.G. et al., 1979; Pitt G.G. et al., 1981; Cerrai et al., 1994; Petrova et al., 1998). PCL degrades at a much slower rate than PLA and is therefore most suitable for the development of long-term, implantable drug delivery systems. Aforementioned copolymers of caprolactone with dilactide were synthesized to accelerate degradation rates (Middleton and Tipton, 2000). Tubular, highly permeable poly(L-lactide-co-ε-caprolactone) guides were found to be suitable for regeneration and functional reinnervation of large gaps in injured nerves (Rodriguez et al., 1999). While this study focuses on tissue regeneration, the application of PCL in drug-delivery devices is still far more common (Sinha et al., 2004). With increasing popularity of electrospinning, a laboratory-scale technique that allows for the fabrication of non-woven meshes composed of nano- and/or micro-fibers (Pham et al., 2006), PCL might find its way into cell-based therapies since slowly degrading polymers are preferred for this technique to ensure sufficient stability of the fibers (Yoshimoto et al., 2003). Poly(p-dioxanone) (Figure 35.4f), another polylactone, and its copolymers with lactide, glycolide, and/or trimethylene carbonate are synthesized by catalyzed ring-opening polymerization and have been used in a number of clinical applications ranging from suture materials to bone fixation devices (Wang et al., 1998; Yang et al., 2002). Polyorthoesters Polyorthoesters (POEs) (Figure 35.4g) have been developed by the Alza Corporation and SRI International in the 1970 in search of a new biodegradable polymer for drug delivery applications (Heller et al., 2002). Since then, polymer synthesis has been improved over four generations. POEs are synthesized by condensation or addition reactions typically involving dialcohols and monomeric orthoester or diketene acetals, respectively. The use of triethylene glycol as the diol component produced predominantly hydrophilic polymers, whereas hydrophobic materials could be obtained by using 1,10-decanediol. Orthoester is a functional group containing three alkoxy groups attached to one carbon atom. In POEs two of the three alkoxy groups are typically part of a cyclic acetal (Figure 35.4g). POEs were synthesized that degrade by surface erosion, which is characterized by a constant decrease of bulk mass while polymer molecular weight within the polymer bulk is preserved (Burkersroda et al., 2002). It is known that materials built from functional groups with short hydrolysis half lives and low water diffusivity tend to be surface eroding. Polymers that exhibit surface erosion can be used to fabricate drug delivery systems that, at a high aspect to volume ratio (e.g., as for wafers), release loaded drugs at a constant rate. The addition of lactide segments to the POE structure resulted in self-catalyzed erosion and allowed for tunable degradation times ranging from weeks to months (Ng et al., 1997). POEs provide the material platform for a variety of drug delivery applications including the treatment of postsurgical pain, osteoarthritis, and ophthalmic diseases as well as the delivery of proteins, and DNA. Block copolymers of POE and PEG have been prepared, and their use as drug delivery matrices or as colloidal structures for tumor targeting are being explored (Heller et al., 2002).

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Initial biocompatibility studies revealed that POEs provoked little inflammation and were largely absorbed by 4 weeks. In contrast, PD,LLA degraded slower and provoked a chronic inflammation with multinuclear giant cells, macrophages with engulfed material, and proliferating fibroblasts within the same model. Ossicles with bone marrow had formed in the implants of PEO in combination with demineralized bone. In PLA/demineralized bone implants the bone formation was inhibited (Andriano et al., 1999; Solheim et al., 2000). Polycarbonates Polycarbonates have become interesting biomaterials due to their excellent mechanical strength and good processability. Since pure polycarbonates degrade extremely slowly under physiological conditions, polyiminocarbonates (Kohn and Langer, 1986) and tyrosine-based polycarbonates (Pulapura and Kohn, 1992) (Figure 35.4h) have been engineered to yield biodegradable polymers of good mechanical strength (Engelberg and Kohn, 1991) for use in drug delivery and orthopedic applications. Degradation of most polycarbonates is controlled by the hydrolysis of the carbonate group which yields two alcohols and carbon dioxide thus alleviating the problem of acid bursting seen in polyesters (Gunatillake and Adhikari, 2003). Structural variation of the pendant side groups allows for the preparation of polymers with different mechanical properties, degradation rates, as well as cellular response. Polycarbonates that contain a pendant ethyl ester group have been shown to be osteoconductive and to possess mechanical properties sufficient for load-bearing bone fixation. Long-term (48 week) in vivo degradation kinetics and host bone response to tyrosine-derived polycarbonates were investigated using a canine bone chamber model (Choueka et al., 1996). Histological sections revealed intimate contact between bone and the tested polycarbonates. It was concluded that, from a degradation–biocompatibility perspective, the tyrosine-derived polycarbonates appear to be comparable, if not superior, to PLA in this model. Amino Acid-Derived Polymers, Poly(amino acids), and Peptides Amino acids are an interesting building block for polymers due to the biocompatibility of the degradation products and the degradability of the amide or ester bonds by which amino acids are typically polymerized or integrated in copolymers. Early studies on pure poly(amino acids)s revealed significant concerns with the materials immunogenicity and mechanical properties (Bourke and Kohn, 2003). To improve those unfavorable properties, amino acids have been used as monomeric building blocks in polymers that have a backbone structure different from natural peptides. Based on polymer structure and chemistry, four major groups have been used to classify such “non-peptide amino acid-based polymers.” As for the above described tyrosine-derived polycarbonates, L-tyrosine is the predominantly employed amino acid for the formation of tyrosine-derived polyarylates and polyesters. These polymers exhibit excellent engineering properties, and polymer systems can be designed whose members show exceptional strength (polycarbonates), flexibility and elastomeric behavior (polyarylates), or water-solubility and self-assembly properties (copolymers with PEG). Poly(DTE carbonate) (DTE: desaminotyrosyl-tyrosine ethyl ester) (Figure 35.4h, R: 9CH2CH3) exhibits a high degree of tissue compatibility and is currently being evaluated for possible clinical uses by the US Federal Drug Administration (Bourke and Kohn, 2003). Solid-phase peptide synthesis, pioneered by Merrifield, and genetic engineering allow for the automated and highly efficient synthesis of peptides of a predefined sequence. In contrast to synthetic poly(amino acid)s, which are traditionally composed of a single amino acid and were found to be highly immunogenic in most cases, synthetic peptides have become an important polymer class for biomedical applications. Specifically peptides and peptide-amphiphiles that undergo self-assembly-driven in situ gelation in response to temperature, pH, or chemical stimuli are of interest as these materials can be minimally invasively implanted starting from aqueous solutions (Stupp et al., 1997; Meyer and Chilkoti, 1999; Hartgerink et al., 2001).

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Genetically engineered elastin-like polypeptides, which are composed of a pentapeptide repeat and undergo inverse temperature phase transition, have been used to encapsulate chondrocytes. The cell culture studies showed that cartilaginous tissue formation, characterized by the biosynthesis of sulfated glycosaminoglycans (GAGs) and collagen, was supported (Betre et al., 2002). Self-assembled peptide-amphiphiles, which form hydrogels composed of nanofibers resembling the native ECM components, have been demonstrated to be cytocompatible in cell encapsulation studies (Beniash et al., 2005). Recently, peptide-amphiphile nanofibers were shown to promote in vitro proliferation and osteogenic differentiation of marrow stromal cells (MSCs) (Hosseinkhani et al., 2006). Polyurethanes As outlined in section “Hydrolytically stable polyurethanes,” PUs represent a major class of synthetic elastomers that have excellent mechanical properties and good biocompatibility. PUs have been evaluated for a variety of medical devices and implants, particularly for long-term implants. Knowledge gained about the mechanisms of PU biodegradation in response to implant failures throughout the 1990s has been translated to form a new class of bioresorbable materials (Santerre et al., 2005). Recent research has utilized the flexible chemistry and diverse mechanical properties of PUs to design degradable polymers for a variety of regenerative applications. Segmented PUs with varied molecular structure have been synthesized to control rates of hydrolysis (Skarja and Woodhouse, 2001; Santerre et al., 2005). To obtain biodegradable, segmented PUs significant changes were required to the structural components historically used for their synthesis. Traditional aromatic diisocyanates (D in Figure 35.3) can yield toxic or carcinogenic degradation products when part of a degradable PU; therefore, linear diisocyanates, such as lysinediisocyanate that yields the non-toxic degradation product lysine, are preferred. The soft segment, typically comprised of an oligomeric diol (P in Figure 35.3), is typically the block of the PU used to modify the degradation rate. Biodegradable PUs have been synthesized with a variety of soft segments including PEO, degradable polyesters such as PLA, PGA, or PCL, and combinations thereof. Other strategies focus on the copolymers’ hard segments. PUs were synthesized that contain enzyme sensitive linkages introduced with the chain extender (C in Figure 35.3). For example, the use of a phenylalanine diester chain extender yielded a PU that showed susceptibility to enzyme-mediated degradation upon exposure to chymotrypsin and trypsin. Saad et al. investigated cell and tissue interactions with a series of degradable polyesterurethanes. In vivo investigations showed that all test polymers exhibited favorable tissue compatibility and degraded significantly during the course of 1 year (Saad et al., 1997). Polyurethaneurea matrices were shown to allow vascularization and tissue infiltration in vivo (Ganta et al., 2003). The flexible chemistry and diverse mechanical properties of PU materials allowed researches to design degradable polymers for the regeneration of tissues as varied as neurons, vasculature, smooth muscle, cartilage, and bone (Xue and Greisler, 2003; Zhang et al., 2003; Santerre et al., 2005). Block Copolymers of Polyesters or Polyamides with PEG Amphiphilic block copolymers of biodegradable polymers with PEG have become popular materials for injectable drug delivery applications (Jeong et al., 2002). Inspired by the thermoresponsive behavior observed for non-degradable A–B–A type triblock copolymers composed of hydrophilic PEO (block A) and hydrophobic PPO (block B), polymer development focused on synthesizing biodegradable analogs of these poloxamers (or Pluronics®) that were water soluble at ambient temperature and formed stable hydrogels at body temperature. Biodegradable block copolymers were synthesized by substituting the hydrophobic PPO block with a biodegradable polymer block, such as PLA or PCL (Jeong et al., 1997; Lee et al., 2001; Ruel-Gariepy and Leroux, 2004).

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Biodegradable, physically crosslinkable block copolymers of inverse structure, that is, B–A–B triblock copolymers with two biodegradable hydrophobic polymer blocks (block B) and a hydrophilic PEO block, have also been investigated as protein delivery systems (Kissel et al., 2002). Polyanhydrides Drug delivery technologies rely on engineered polymers that degrade in a well controllable and adjustable fashion (Langer, 1990). Increasing understanding of erosion mechanisms led to a demand for synthetic polymers that contain a hydrolytically labile backbone while limiting water diffusion within the polymer bulk significantly to confine erosion to the polymer–water interface. Such surface eroding polymers allow for the fabrication of drug delivery devices that erode at constant velocity at any time during erosion, thereby, releasing incorporated drugs at constant rates (Gopferich and Tessmar, 2002). Polyanhydrides were engineered following this paradigm by selecting the anhydride linkage, one of the least hydrolytically stable chemical bonds available, to connect the building hydrophobic monomers. Polyanhydrides (Figure 35.4i) have been synthesized by various techniques, including melt condensation, ring-opening polymerization, interfacial condensation, dehydrochlorination, and dehydrative coupling agents (Kumar et al., 2002). Solution polymerization traditionally yielded low molecular weight polymers. Different dicarboxylic acid monomers have been polymerized to yield polyanhydrides with various physicochemical properties. Examples are linear, aromatic, fatty acid-based dicarboxylic acid monomers, and fatty acid terminated polyanhydrides. Polyanhydrides made from linear sebacic acid (SA) and aromatic 1,3-bis(p-carboxyphenoxy) propane (CPP) (Figure 35.4i) have been engineered to deliver carmustine (1,3-bis(2-chloroethyl)-1-nitrosourea (BCNU)), an anticancer drug, to sites in the brain following primary resection of a malignant glioma (Westphal et al., 2003). Poly(SA–CPP) hydrolyzes into non-toxic degradation products and the local chemotherapy with BCNU wafers was shown to be well tolerated and to offer a survival benefit to patients with newly diagnosed malignant glioma. The chemical composition of a polyanhydride can be used to custom-design its degradation properties. While polyanhydrides from linear monomers, such as poly(SA), degrade within a few days, polymerized aromatic dicarboxylic acids, such as poly(1,6-bis(p-carboxyphenoxy) hexane), degrade much more slowly (up to a year) (Temenoff and Mikos, 2000). The structural versatility of polyanhydrides in combination with their unique degradation and erosion properties make them precious materials for numerous medical, biomedical, and pharmaceutical applications in which degradable polymers that allow for a perfect erosion control are needed (Gopferich and Tessmar, 2002). With regard to tissue engineering applications, polyanhydrides have also been interesting polymers due to their degradative properties and their good biocompatibility (Katti et al., 2002). The use of polyanhydrides in load-bearing orthopedic applications, however, is restricted due to limited mechanical properties. Poly(anhydrides-co-imides) which were developed in order to combine the good mechanical properties of polyimides with the degradative properties of polyanhydrides were shown to meet compressive strengths comparable to human bone (Uhrich et al., 1995) and displayed good osteocompatibility (Ibim et al., 1998). Photopolymerizable polyanhydrides have been synthesized with the objective to combine high strength, controlled degradation, and minimal invasive techniques for orthopedic applications and were shown to be osteocompatible (Anseth et al., 1999). Depending on the chemical composition, these materials reached compressive and tensile strengths similar to those of cancelleous bone (Muggli et al., 1999). Polyphosphazenes Polyphosphazenes (Figure 35.4j), which are polymers containing a high molecular weight backbone of alternating phosphorus and nitrogen atoms with two organic side groups attached to each phosphorus atom, is a

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relatively new heterogenic class of biomaterials. Because different synthetic pathways allow for a tremendous variety of substituents, phosphazene polymers exhibit a very diverse spectrum of chemical and physical properties. This spectrum makes them suitable for many biomedical applications ranging from templates for nerve regeneration, cardiovascular, and dental uses to implantable and controlled release devices (Langone et al., 1995; Schacht et al., 1996; Andrianov and Payne, 1998). The best studied and most important route to polyphosphazenes, whose synthesis is generally more involved than that for most petrochemical biomaterials but offers unique flexibility, is a macromolecular substitution route. A reactive polymeric intermediate, poly(dichlorophosphazene), is typically synthesized by a thermal ring-opening cationic polymerization of hexachlorocyclotriphosphazene in bulk at 250°C that yields a polydisperse high molecular weight product. The intermediate is reacted with low molecular weight organic nucleophiles resulting in stable, substituted polyphosphazenes, which in this case are also addressed as poly(organo)phosphazenes. Depending on the substituent chemistry, the polyphosphazene is more or less susceptible to hydrolysis. Biodegradable hydrophobic polyphosphazenes have been synthesized using imidazolyl, ethylamino, oligopeptides, amino acid esters, and depsipeptide groups (dimers composed of an amino acid and a glycolic or lactic ester) as hydrolysis sensitive side groups. Hydrolytic degradation products include free side group units, phosphate, and ammonia due to backbone degradation (Andrianov and Payne, 1998). Hydrogelforming, hydrophilic polyphosphazenes can be synthesized through the introduction of small, hydrophilic side groups, such as glucosyl, glyceryl, or methylamino side groups. Ionic side groups yield polymers that form hydrogels upon ionic complexation with multivalent ions (Allcock and Kwon, 1989). Hydrophilic, water-soluble polyphosphazenes with amphiphilic side groups, such as poly(bis(methoxyethoxyethoxy)phosphazene) (Figure 35.4j, R,R: 9OCH2CH2OCH2CH2OCH3), display a LCST (section “Poly(N-isopropylacrylamide)”) and are responsive to changes in temperature and ionic strength (Lee, 1999). Both hydrophilic and hydrophobic polyphosphazenes have demonstrated their potential as biocompatible materials for controlled protein delivery. Ionic polyphosphazenes have been explored as vaccine delivery systems and poly(di(carboxylatophenoxy)phosphazene) has demonstrated a remarkable adjuvant activity on the immunogenicity of inactivated influenza virions and commercial trivalent influenza vaccine in the soluble state (Andrianov and Payne, 1998). Porous scaffolds from biodegradable polyphosphazenes have been shown to be good substrates for osteoblast-like cell attachment and growth with regard to skeletal tissue regeneration (Laurencin et al., 1996). Tubular polyphosphazene nerve guides were investigated in a rat sciatic nerve defect. After 45 days, a regenerated nerve fiber bundle was found bridging the nerve stumps in all cases (Langone et al., 1995). Biodegradable Crosslinked Polymer Networks The chemical crosslinking of individual, linear polymer chains results in networks of increased stability. This concept has been extensively explored for applications in regenerative medicine and most likely represents the concept of choice for modern biomaterial research, especially if polymer crosslinking can be conducted inside a tissue defect (Temenoff and Mikos, 2000). The crosslinking of hydrophobic polymers or monomers results in tough polymer networks that can be used for orthopedic fixation. PMMA (Figure 35.1f), the main component in injectable bone cements, is the most prominent example. Due to their hydrophobicity, the precursors are typically injected as a moldable liquid or paste free of additional solvents. In situ crosslinking can be initiated thermally or photo-chemically by UV-rich light. Both ways of initiation are also applicable to hydrophilic injectable systems that form highly swollen gels (hydrogels) as a result of precursor crosslinking. In contrast to hydrophobic networks that scarcely swell in the presence of water, injectable hydrogels are characterized by a high water content and diffusivity, which allow for the direct encapsulation of cells and sufficient transport of oxygen, nutrients, and waste. Hydrophobic networks, however, often require the addition of a leachable porogen, such as salt particles, to facilitate cell migration and tissue ingrowth. Generally, injectable polymer

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systems have considerable advantages over pre-fabricated implants or tissue engineering scaffolds, which include the ability to fill irregularly shaped defects with minimal surgical intervention (Peter et al., 1998a). A number of demanding requirements have to be fulfilled by synthetic materials for applications in regenerative medicine. Not only do the physicochemical properties have to be adjusted to the application site, but also the polymer and any adjuvant component required to formulate an in situ crosslinkable system have to be biocompatible. Ideally, the resulting network should also have the ability to support cell growth and proliferation early in the tissue regeneration process (Temenoff and Mikos, 2000). The crosslinkable synthetic polymers that will be discussed in the following sections are reactive polyesters. The main chemical functionality involved in the chemical crosslinking mechanisms is the polarized, electron-poor double bond, such as in vinylsulfones and in esters of acrylic acid, methacrylic acid, and fumaric acid. Other chemically or thermally crosslinkable macromonomer functional groups are styryl, coumarin, and phenylazide and will not be discussed here (Hou et al., 2004). Crosslinked Polyesters Fumarate-based polymers: The development of fumarate-based polyesters for biomedical applications started

around 20 years ago. Fumaric acid is a naturally occurring metabolite, which is found in the tri-carboxylate cycle (Krebs cycle), and is comprised of a reactive double bound available for chemically crosslinking reactions. These characteristics make fumaric acid a candidate building block for crosslinkable polymers. The first and most comprehensively investigated fumarate-based copolymer is the biodegradable copolyester poly(propylene fumarate) (PPF) (Figure 35.5a). PPF was first polymerized from fumaric acid and propylene oxide (Domb et al., 1990). Mikos and coworkers optimized the synthesis of PPF and broadly investigated tissue

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Figure 35.5 Chemical structures of synthetic polymers for the fabrication of crosslinked biodegradable networks.

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compatibility and applications of PPF both in vitro and in vivo. Synthesis progressed to copolymerizing fumaryl chloride with 1,2-propanediol (propylene glycol) (Peter et al., 1999b) and now involves the transesterification of diethylfumarate with propylene glycol and subsequent polycondensation of the diester intermediate bis(2-hydroxypropyl) fumarate (PF) (Shung et al., 2003). A variety of methods to synthesize PPF have been explored, and each results in different polymer molecular weights and properties (Peter et al., 1997a). PPF has been developed as an alternative to PMMA bone cements. PPF can be injected as a viscous liquid and thermally crosslinked in vivo eliminating the need for direct exposure of the defect site to light. Typically, PPF is crosslinked with either MMA or N-vinyl pyrrolidone (NVP) monomers and benzoyl peroxide as a radical initiator (Gresser et al., 1995; Frazier et al., 1997). Depending on the ratio of initiator, monomer, and PPF, the curing time can be controlled between 1 and 121 min. Compared to PMMA, which is not resorbable and suffers from the fact that its high curing temperatures (94°C) can cause necrosis of the surrounding tissue, the curing temperature of PPF has been shown to never exceed 48°C (Peter et al., 1997b, 1999a). PPF can also be photocrosslinked along the electron-poor double bonds along the backbone. Typical formulations include NVP, diethylfumarate, or PF-diacrylate (DA) as co-monomers together with a photoinitiator, such as bis(2,4,6trimethylbenzoyl) phenylphosphine oxide (Fisher et al., 2001, 2002a; He et al., 2001). The mechanical properties of PPF, which are dependent on composition, synthesis condition, and crosslinking density, are already promising. However, these materials are probably not sufficient for load-bearing applications, especially when used as macro-porous scaffolds (Peter et al., 1998a; Fisher et al., 2002a; Timmer et al., 2003). One strategy to further strengthen PPF scaffolds includes the incorporation of nanoparticulate fillers. Reinforced PPF composites have been synthesized using aluminum oxide-based ceramic nanoparticles and modified single walled carbon nanotubes. For just 0.05 wt% loading with the latter, a 74% increase was recorded for the compressive modulus and a 69% increase for the flexural modulus as compared to plain PPF/PF–DA (Shi et al., 2005). The chemical integrations of alumoxane nanoparticles in crosslinked PPF/PF–DA networks resulted in a significantly increased flexural modulus (Horch et al., 2004). Micro-particulate ceramic materials, such as β-tricalcium phosphate (β-TCP), have also been employed as inorganic filler to improve mechanical properties of composite scaffolds and to improve the material’s osteoconductivity (Peter et al., 2000). The composite scaffolds exhibit increased compressive strengths in the range of 2–30 MPa, and β-TCP reinforcement delayed scaffold disintegration significantly in vivo (Peter et al., 1998b). This subcutaneous rat implantation study also revealed a mild initial inflammatory response and formation of a fibrous capsule around the implant at 12 weeks. A deleterious long-term inflammatory response was not observed. Rabbit in vivo studies also revealed biocompatibility of photo-crosslinked PPF scaffolds in both soft and hard tissues (Fisher et al., 2002b). PPF hydrolytically degrades along the ester bond in its backbone. Degradation time was found to be dependent on polymer structure as well as other components, such as fillers. In vitro studies identified the time needed to reach 20% original mass ranging from around 84 (PPF/β-TCP composite) to over 200 days (PPF/CaSO4 composite) (Temenoff and Mikos, 2000). In order to broaden the application spectrum for in situ crosslinkable PPF, block copolymers with hydrophilic PEG of different compositions were synthesized. Poly(propylene fumarate-co-ethylene glycol) (P(PF-co-EG)) (Figure 35.5b) was synthesized from PPF and PEG in a transesterification reaction catalyzed by antimony trioxide; propylene glycol was removed by condensation (Suggs et al., 1997). Behravesh et al. have modified the synthesis to yield well-defined ABA-type triblock copolymers from two moles monomethoxyPEG and one mole PPF (Behravesh et al., 2002a). Generally, P(PF-co-EG) copolymers are hydrophilic polymers with specific properties including crystallinity and mechanical characteristics being dependent on the molecular weights of the individual blocks and the copolymer. As a result, platelet attachment to P(PF-co-EG) hydrogels was significantly reduced as compared to the PPF homopolymer making these copolymers candidate materials when direct biomaterial–blood contact is inevitable, such as for vascular grafts (Suggs et al., 1999b).

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Most P(PF-co-EG) copolymers are soluble in water making them candidate materials for injectable applications. ABA-type copolymers were found to show thermoreversible properties, comparable to other PEGcontaining triblock copolymers discussed above. The thermogelling properties of P(PF-co-EG) were dependent on the PEG molecular weight and salt concentration and the physical gelation temperature could be adjusted to values below body temperature (Behravesh et al., 2002a). In addition, the hydrophobic PPF block is highly unsaturated and available for additional chemical crosslinking, which could result in stiff crosslinked networks suitable for the fabrication of prefabricated cell carriers. In vitro degradation studies of macro-porous, crosslinked P(PF-co-EG) scaffolds revealed considerable mass loss and swelling over 12 weeks. In these studies the degradation rate was mainly dependent on content of the PEG–DA crosslinker and almost unaffected by construct porosity. Overall, the results indicated a bulk degradation mechanism of the macro-porous constructs (Behravesh et al., 2002b). In a subcutaneous rat model, P(PF-co-EG) hydrogels demonstrated good initial biocompatibility, showing an acute inflammatory response characterized by infiltration of neutrophils, followed by development and maturation of a fibrous capsule, characteristic of biomaterial implants (Suggs et al., 1999a). Overall, the reported in vitro cytotoxicity and in vivo biocompatibility assays suggest that P(PF-co-EG) hydrogels have potential for use as injectable biomaterials. Fisher et al. have demonstrated the suitability of thermoresponsive P(PF-co-EG) hydrogels for chondrocyte delivery toward the regeneration of articular cartilage defects (Fisher et al., 2004). Similar to previously discussed, stealthy, PEG-containing biodegradables, PEG-content and hydrophilicity of crosslinked P(PF-co-EG) hydrogels are critical factors affecting cell adhesion (Tanahashi and Mikos, 2002). Low-adhesive hydrogels allow for a controlled surface or bulk modification with adhesion molecules to specifically enhance cell adhesion. P(PF-co-EG) hydrogels have been modified by covalent integration of agmatine (Tanahashi and Mikos, 2003) and the adhesion peptide GRGDS (Behravesh et al., 2003). Significantly increased numbers of smooth muscle cells and MSCs were found adhered as compared to the unmodified networks. An exclusively hydrophilic fumarate-based macromer is oligo(poly(ethylene glycol) fumarate) (OPF) (Figure 35.5c). OPF macromers have been synthesized from PEG and fumaryl chloride by a simple condensation reaction in the presence of triethylamine. OPF crosslinking, with or without the addition of crosslinker such as PEG–DA, can be initiated photo-chemically (Jo et al., 2001) or thermally (Temenoff et al., 2002). In contrast to chemically crosslinked PPF and P(PF-co-EG), which both form rigid scarcely swelling polymer networks, crosslinked OPF gels exhibit typical properties of hydrogels, which were dependent on the molecular weight of PEG and reactant ratio (Jo et al., 2001). Crosslinked OPF hydrogels degrade hydrolytically along the ester bonds between fumaric acid and PEG resulting in increased polymer swelling and decreased dry weight. The weight loss of OPF hydrogels was dependent on their crosslinking density (Shin et al., 2003c). Studies investigating the mechanical properties revealed that crosslinked OPF hydrogels made from low molecular weight PEG (1,000 Da), swelled less, were stiffer, and elongated less before fracture when compared to hydrogels comprised of longer PEG chains. OPF hydrogels can also be combined in layers to form biphasic gels, with each phase having different material properties (Temenoff et al., 2002). In vitro investigation of the cytotoxicity of each component of OPF hydrogel formulations and the resulting crosslinked network were conducted employing MSCs. After 24 h, the MSCs maintained more than 75% viability except for OPF concentrations higher than 25% (w/v). A high molecular weight (3,400 Da) PEG–DA crosslinker demonstrated significantly higher viability compared to lower molecular weight (575 Da) PEG–DA. Leachable products from crosslinked OPF hydrogels were found to have minimal adverse effects on MSC viability (Shin et al., 2003a). The in vivo bone and soft tissue compatibility of OPF hydrogels was demonstrated using a rabbit model (Shin et al., 2003c). Based on these promising biocompatibility data, OPF-based hydrogels were investigated as injectable drug, DNA, and cell delivery devices. Crosslinked OPF hydrogels which encapsulated gelatin microparticles were developed as a means of simultaneously delivering two chondrogenic proteins,

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insulin-like growth factor-1 (IGF-1) and transforming growth factor-β1 (TGF-β1) (Holland et al., 2005b). Similar systems were implanted into osteochondral defects in the rabbit model. No evidence of prolonged inflammation was observed, and hyaline cartilage was found filling the chondral region of the defect at 14 weeks. The subchondral region was filled with bony tissue and completely integrated with the surrounding bone. The newly formed surface tissue stained positive for Safranin O and displayed promising chondrocyte organization (Holland et al., 2005a). Kasper et al. developed and characterized composites of OPF and cationized gelatin microspheres that release plasmid DNA in a sustained, controlled manner in vivo (Kasper et al., 2005). In order to control cell adhesion to the hydrophilic hydrogels, RGD adhesion peptide modified OPF hydrogels have been developed (Shin et al., 2002). OPF hydrogels have also been shown useful as injectable cell delivery vehicles for bone regeneration. MSCs were directly combined with the OPF hydrogel precursors and encapsulated during thermal crosslinking. In the presence of osteogenic supplements, MSC differentiation in these hydrogels was apparent by day 21. At day 28, mineralized matrix could be seen throughout the hydrogels (Temenoff et al., 2004a). Hydrogel properties have been identified to affect osteogenic differentiation within these systems (Temenoff et al., 2004b). Recent studies focused on the combination of cell and growth factor delivery using injectable OPF formulations (Park et al., 2005). Polymers-containing acrylate, methacrylate, or vinylsulfone functionalities: Precursors for crosslinked biodegradable polyester networks that bear vinylsulfone, acrylate, or methacrylate functionalities include PEG–DA (Figure 35.5d), PEG–dimethacrylate (Figure 35.5e), PEG vinylsulfones, diacrylated PLA–PEG–PLA block copolymers, acrylic modified polyvinyl alcohol (PVA), methacrylate-modified dextran, and acrylated chitosan (Hoffman, 2002; Nguyen and West, 2002; Hou et al., 2004). Since the last two are synthetic derivatives of natural macromolecules, they are not discussed further. Besides such hydrophilic, natural macromolecules, which are considered candidate building blocks based on their inherent biocompatibility, PEG is the most prominent synthetic component of crosslinked polymer networks due to its biocompatibility and inertness. As described above, PEG is hydrophilic and does not promote cell adhesion. To improve cell adhesion to crosslinked PEG hydrogels, adhesion peptides containing the tripeptide motif RGD have been incorporated (Hern and Hubbell, 1998; Burdick and Anseth, 2002; Gonzalez et al., 2004). Recent research on engineered hydrogels has been focused on mimicking the invasive characteristics of native ECMs by including substrates for matrix metalloproteinases (MMP) in addition to integrin-binding sites. PEG hydrogels crosslinked in part by MMP sensitive linkers were made degradable and invasive for cells via cell-secreted MMPs (Lutolf et al., 2003a). Critical-sized defects in rat crania were completely infiltrated by cells and were remodeled into bony tissue within 5 weeks when above-mentioned gels were loaded with recombinant human bone morphogenetic protein2 and implanted in the defect site. As in natural ECMs, that sequester a variety of cellular growth factors and act as a local depot for them, invading cells were presented with a mitogen that, in this case, specifically promoted bone regeneration (Lutolf et al., 2003b). The PEG-based hydrogels used in these studies were fabricated by a conjugate addition reaction between vinylsufone-functionalized branched PEG and thiol-bearing peptides under almost physiological conditions. In order to enhance the initial mechanical stability and biodegradability of crosslinked PEG-based hydrogels, oligomeric biodegradable lipophilic blocks, such as oligo(lactic acid) (Burdick et al., 2001) (Figure 35.5f) and oligo(ε-caprolactone) (Davis et al., 2003), were included in the crosslinkable polymeric precursors. In a critical size cranial defect model, porous crosslinked poly(ethylene glycol(2)-lactic acid(10)) scaffolds in combination with osteoinductive growth factors have shown potential as an in situ forming synthetic bone graft material (Burdick et al., 2003). Photopolymerized (meth)acrylated biodegradable hydrogels have been used in a wide range of biomedical applications. As described above, limited interactions with proteins are characteristic for hydrophilic

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surfaces. Consequently, applications such as the use of crosslinked hydrogels as barriers following tissue injury in order to improve wound healing and as cell encapsulation materials to immunoisolate transplanted cells capitalize of this property (Cruise et al., 1999; Nguyen and West, 2002). Islets of Langerhans encapsulated in PEG–DA hydrogels and transplanted in order to develop a bioartificial endocrine pancreas are a prominent example for the later application. The hydrogels are permeable for nutrients, oxygen, and metabolic products allowing for the entrapped islets to survive and to secrete insulin that is released by diffusion. Hydrophilic tissue barriers from crosslinked polyesters, such as poly(ethylene glycol-co-lactic acid) DA, have been used to prevent thrombosis and re-stenosis following vascular injury and postoperative adhesion formation following many abdominal and pelvic surgical procedures. Crosslinked hydrophilic polyesters are also promising depots for local drug delivery because of their compatibility with hydrophilic, macromolecular drugs, such as proteins or oligonucleotides. The materials’ good tissue and hemocompatibility even allows for intravascular applications (An and Hubbell, 2000). Drug release from crosslinked hydrogels generally can be well controlled by adjusting swelling, crosslink density, and polymer degradation (Peppas et al., 1999, 2000; Davis and Anseth, 2002). Photopolymerized (meth)acrylated polymer networks have also been widely explored for injectable tissue engineering (Hoffman, 2002; Varghese and Elisseeff, 2006). Elisseeff and coworkers employed PEG–DA scaffolds for cartilage engineering by encapsulating chondrocytes, MSCs, and embryonic stem cells. In these studies, the crosslinked PEG-based hydrogels served as an efficient scaffold for anchorage-independent cells and promoted tissue formation. Photogelation, which offers good spatial and temporal control of hydrogel curing, has been used to control the spatial organization of different cell types within a three-dimensional system for osteochondral defect regeneration by sequentially polymerizing multiple cell/hydrogel layers. In an attempt to promote hydrogel–tissue integration, a tissue-initiated polymerization technique has been developed that utilizes in situ generated tyrosyl radicals to initiate photogelation of an injectable macromer solution (Varghese and Elisseeff, 2006). Traditionally, photopolymerization occurs by directly exposing materials to UV or visible light in accessible cavities or during invasive surgery. For PEG–dimethacrylate hydrogels, it has been shown that light, which penetrates tissue including skin, can cause a photopolymerization indirectly (transdermal photopolymerization). In vivo studies revealed that gels can be polymerized in 3 min with no harm to imbedded chondrocytes and subsequent cartilaginous tissue formation as indicated by increasing GAG and collagen contents (Elisseeff et al., 1999). In deep crevices, as they may be found in larger orthopedic defects, problems are expected to arise from limited light penetration and inconsistent photopolymerization. For those applications, thermally induced crosslinking techniques appear to be advantageous (Temenoff and Mikos, 2000).

APPLICATIONS OF SYNTHETIC POLYMERS Synthetic polymers play a vital role in biomedical applications, including nano-, micro-, and macroscopic drug and gene delivery devices (Brannon-Peppas, 1995; Hubbell, 1998; Uhrich et al., 1999; Panyam and Labhasetwar, 2003), orthopedic fixation devices (Bostman and Pihlajamaki, 2000), cosmetic, and prosthetic implants (Behravesh et al., 1999), and as artificial matrices for tissue engineering applications (Seal et al., 2001). The interested reader may be directed to the referenced reviews that provide in-depth insight in current trends and technologies. Researches have sought to develop and clinically explore third generation biomaterials (Hench and Polak, 2002) that are designed to control protein adsorption, cell adhesion, and differentiation, implant integration, foreign body reaction, and to develop biomimetic synthetic materials (Shin et al., 2003b; Drotleff et al., 2004; Lutolf and Hubbell, 2005).

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CONCLUSION/SUMMARY Synthetic biomaterials have progressed from testing “off-the-shelf ” plastics not developed for biomedical purposes, to a field of synergistic research by engineers, scientists, and physicians dedicated to tailoring material properties for specific applications. Most recent trends shift the focus toward biology in order to first understand and then mimic physiological interactions and signaling. Hydrogels, especially injectable systems, enjoy increasing attention due to the comfort of their application, their structural similarity to native ECM, and their good compatibility for direct cell encapsulation due to high water contents. It is no longer believed in tissue engineering that the biomaterial itself has to provide mechanical properties comparable to the diseased tissue; the polymer rather has to promote defect site remodeling and tissue regeneration in vivo in a way that the regenerated tissue is histologically and functionally indistinguishable from the surrounding tissue. Hydrogels might be superior to hydrophobic polymers in that regard, as they can degrade faster resolving the problem of non-functional fibrous tissue formation on the polymer–tissue interface. Also, hydrogel breakdown can be synchronized with cell proliferation and migration by using enzymatically cleavable crosslinker. Besides providing tailored degradative properties, synthetic materials for regenerative medicine should allow for minimally invasive application techniques, integrate well with the surrounding tissue, and promote cell adhesion, migration, and finally differentiation. The development and thorough characterization of injectable biodegradables provides the foundation for injectable tissue regeneration. In situ gelation or polymerization concepts will still have to be developed and optimized with regard to cytocompatibility and stability of the resulting construct. The implementation of biomimetic design strategies will allow to control and custom-design cell–biomaterial interactions in order to guide tissue formation from transplanted cells. Strategies based on gene delivery or gene-activating biomaterials also have great potential in regenerative medicine but the long-term safety of such therapies remains to be proven. Overall, the advances that have been made in the field of biomaterial synthesis and design of physicochemical properties during the last 50 years in conjunction with the rapidly increasing knowledge in adult and stem cell biology concerning adhesion, migration, differentiation, and signaling will reveal design concepts for improved injectable, biomimetic polymer-based formulations for tissue engineering applications.

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36 Hybrid, Composite, and Complex Biomaterials for Scaffolds Gilson Khang, Soon Hee Kim, Moon Suk Kim, and Hai Bang Lee

INTRODUCTION It has been recognized that tissue engineering offers an alternative technique to whole organ and tissue transplantation for diseased, failed, or malfunctioned organs. Millions of patients have been suffered by end-stage organ failure or tissue loss annually. In the United State alone, at least 8 million surgical operations had been carried out each year, requiring a total national health care cost exceeding $400 billion annually (Khang et al., 2006). In order to avoid the shortage of donor organ and these problems, a new hybridized method combined with cell and biomaterials has been introduced as tissue engineering very recently. To reconstruct a new tissue by tissue engineering, triad components such as (i) cells which are harvested and dissociated from the donor tissue including nerve, liver, pancreas, cartilage, and bone as well as embryonic stem, adult stem, or precursor cell; (ii) biomaterials as scaffold substrates to which cells are attached and cultured resulting in the implantation at the desired site of the functioning tissue; and (iii) growth factors which are promoting and/or preventing cell adhesion, proliferation, migration, and differentiation by up-regulating or downregulating the synthesis of protein, growth factors, and receptors must be needed. This chapter reviews four categories on the focus of hybrid, composite and complex biomaterials for the application of hybrid scaffolds such as (i) poly(α-hydroxyester) family with natural polymer and bioceramics, (ii) bioceramic scaffolds with other biomaterials, (iii) natural polymer with other biomaterials, and (iv) miscellaneous in order to approach to a more natural three-dimensional environment and support biological signals for tissue growth and reorganization.

BIOMATERIALS FOR TISSUE ENGINEERING Importance of Scaffold Matrices in Tissue Engineering Scaffolds might be played a very critical role in tissue engineering. The function of scaffolds is to direct the growth of cells seeded within the porous structure of the scaffold or of cells migrating from surrounding tissue. The majority of mammalian cell types are anchorage-dependent resulting in dying if an adhesion substrate is not provided. Scaffold matrices can be used to achieve cell delivery with high loading and efficiency to specific sites. Therefore, the scaffold must provide a suitable substrate for cell attachment, cell proliferation, differentiated function, and cell migration. The prerequisite physicochemical properties of scaffolds are as

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follows (i) to support and deliver for cells; (ii) to induce, differentiate, and conduit tissue growth; (iii) to target cell-adhesion substrate; (iv) to stimulate cellular response; (v) wound healing barrier; (vi) biocompatible and biodegradable; (vii) relatively easy processability and malleability into desired shapes; (viii) highly porous with large surface/volume; (ix) mechanical strength and dimensional stability; (x) sterilizability, and so on (Khang et al., 2001, 2006; Lee et al., 2003). Generally, three-dimensional porous scaffolds can be fabricated from natural and synthetic polymers, ceramics, metals as very few case, composite biomaterials and cytokine release materials. Natural Polymers Many naturally occurring scaffolds can be observed as biomaterials for tissue engineering purposes. One of the typical examples is the extracellular matrix (ECM) that is very complex biomaterial and controls cell function. For the ECM of tissue engineering, natural and synthetic scaffolds are designed to mimic specific function. The natural polymers are alginate, proteins, collagens (gelatin), fibrins, albumin, gluten, elastin, fibroin, hyarulonic acid, cellulose, starch, chitosan (chitin), sclerolucan, elsinan, pectin (pectinic acid), galactan, curdlan, gellan, levan, emulsan, dextran, pullulan, heparin, silk, chondroitin 6-sulfate, small intestine submucosa (SIS), acellular dermis, polyhydroxyalkanoates, and so on. Much of the interest in these natural polymers comes from their biocompatibility, relatively abundance and commercial availability, and ease of processing (Khang et al., 2004a). Synthetic Polymers and Poly(α-Hydroxy Ester)s One of the most significant shortages of natural polymers is typical expensive, suffering from batch-to-batch variation, and the possibility of cross-contamination from unknown virus or unwanted disease due to the isolation from plant, animal, and human tissue. On the contrary, synthetic polymeric biomaterials might be easily controlled by physicochemical properties and quality and without immunogenecity. Also, it can be processed with various techniques and supplied consistently in large quantity. In order to adjust the physical and mechanical properties of tissue engineered scaffold at desired place in the human body, the molecular structure, molecular weight, and so on are easily adjusted during the synthetic process. These are largely divided into two categories such as (i) biodegradable and (ii) nonbiodegradable. Some nondegradable polymers include polyvinylalcohol, poly(hydroxylethylmethacryalte) (PHEMA), and poly(N-isopropylacrylamide). Some synthetic degradable polymers are the family of poly(α-hydroxy ester)s such as polyglycolide (PGA), polylactide (PLA) and its copolymer poly(lactide-co-glycolide) (PLGA), polyphosphazene, polyanhydride, poly(propylene fumarate), polycyanoacrylate, polycaprolactone, polydioxanone, biodegradable polyurethanes and so on. (Khang et al., 2006) Among these two polymers, the synthetic biodegradable polymers were preferred for the application of tissue engineered scaffolds to minimize the chronic foreign body reaction and lead to the formation of the completely natural tissue. That is to say, they can form a temporary scaffold for mechanical and biochemical support. The family of poly(α-hydroxy acid)s such as PGA, PLA and its copolymer PLGA that are among the few synthetic polymers approved for human clinical use by US Food and Drug Administration (FDA) are extensively used or tested for the scaffold materials as a bioerodible material due to good biocompatibility, controllable biodegradability, and relatively good processability. It has been used for three decades as suture of PGA, bone plate, screw and reinforced materials for PLA, and drug delivery devices of PLGA in surgical operation and whose safety has been proved in many medical applications. The synthetic methods and physicochemical properties such as melting temperature, glass transition temperature, tensile strength, Young’s modulus, and elongation were reviewed elsewhere (Khang et al., 2006).

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(a)

(b)

Figure 36.1 Inflammatory reaction of (a) PGA nonwoven ( 40 magnifications) and (b) PLGA microspheres ( 200 magnifications) after 1 week implantation in nude mouse.

The mechanism of biodegradation of poly(α-hydroxy acid)s is bulk degradation which is characterized by a loss in a polymer molecular weight while mass is maintained. Mass maintenance is useful for tissue engineering applications of specific shapes. However, loss in molecular weight causes a significant decrease in mechanical properties. Degradation is depending on chemical history, porosity, crystallinity, steric hindrance, molecular weight, water uptake, and pH. Degradable products as lactic acid and glycolic acid decrease the pH in the surrounding tissue resulting in inflammation (Figure 36.1) and potentially poor tissue development. PGA, PLA, and PLGA scaffolds were applied for regeneration of all tissues as skin, cartilage, blood vessel, nerve, liver, dura mater, bone, and other tissue. Bioceramic Scaffolds Bioceramic is a term introduced for biomaterials that are produced by sintering or melting inorganic raw materials to create an amorphous or a crystalline solid body that can be used as an implant. Porous final products have been mainly used scaffolds. The components of ceramics are calcium, silica, phosphorous, magnesium, potassium, and sodium. Bioceramic used in the fabrication for the tissue engineering might be classified as nonresorbable (relatively inert), bioactive or surface active (semi-inert), and biodegradable or resorbable (noninert). Alumina, zirconia, silicone nitride, and carbons are inert bioceramics. Certain glass ceramics are dense hydroxyapatites (HA, 9CaO Ca(OH)2 3P2O5), semi-inert (bioactive), and calcium phosphates, aluminum– calcium-phosphates, coralline, tricalcium phosphates (3CaO P2O5), zinc–calcium–phosphorous-oxides, zinc–sulfate–calcium-phosphates, ferric–calcium–phosphorous-oxides, and calcium aluminates are resorbable ceramics. Among these bioceramics, synthetic apatite and calcium phosphate minerals, coral-derived apatite, bioactive glass, and demineralized bone particle (DBP) will be introduced in this section since they are widely used in hard tissue engineering area (Khang et al., 2004). The porosity like size of mean diameter and surface area is a critical factor for the growth and migration of a tissue into the bioceramic scaffolds. Several methods were introduced to optimize the fabrication porous ceramics such as dip casting, starch consolidation, polymeric sponge method, foaming method, organic additives, gel casting, slip casting, direct coagulation consolidation, hydrolysis-assisted solidification, and freezing methods. Therefore, it is very important to choose the appropriate preparation methods for the physical properties of desired organs.

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HYBRID AND COMPOSITE SCAFFOLDS BIOMATERIALS FOR TISSUE ENGINEERING Poly(α-Hydroxyacid) Family Hybrid Scaffolds Although a poly(α-hydroxyacid) family have been extensively tested as scaffolding materials for tissue engineering due to relatively good mechanical properties, low toxicity and predictable biodegradation kinetics, its poor mechanical strength, small pore size and hydrophobic surface properties for cell seeding have limited its usage. In order to solve these problems, several techniques as a surface treatment, an introduction of bioactive molecules, a development of manufacturing method for porous structure, a hybrid with bioactive materials and so on have been developed. Table 36.1 listed various types of natural polymers, and ceramics impregnated with PLGA, PGA, and PLA scaffolds for the improvement of physicochemical properties with new preparation method to the desired target organ. Wu et al. (2006) proposed that PLGA scaffold with 125–500 μm pore size was coated by the combination of three natural biomaterials solution such as collagen, chitosan, or N-succinyl-chitosan. Collagen coated PLGA surface increased cell attachment and proliferation, but chitosan and N-succinyl-chitosan decreased them. Chitosan and N-succinyl-chitosan increased differentiation, but collagen decreased it. This approach could provide a good strategy for modifying microenvironments to increase osteoblast adhesion, proliferation, and differentiation on PLGA scaffolds surface. Hybrid sponge of PLGA/collagen was used as the porous scaffold (Sato et al., 2001), and then chondrocytes were seeded in vitro and tested in vivo. Results showed hybridization of the PLGA/collagen sponge facilitated cell seeding and promoted the in vivo formation of cartilage tissue since the mechanically strong PLGA sponge functioned there as a skeleton and prevented the embedded collagen sponge from collapsing. A biodegradable hybrid scaffold was prepared for fibrin/PGA fiber (Hokugo et al., 2006). Mixed fibrinogen and thrombin solution homogeneously dispersed in PGA fiber was freeze-dried to obtain fibrin sponges with or without PGA fiber incorporation. The shrinkage of sponges after L929 fibroblasts cell seeding was suppressed by fiber incorporation. The PLGA mesh/collagen hybrid scaffolds were prepared by introducing collagen sponge or gel into the PLGA mesh (Nakanishi et al., 2003). Urothelial and smooth muscle cells were obtained from porcine urinary bladder and seeded on these hybrid scaffolds. Ex vivo construction of urinary bladder wall using hybrid scaffolds prepared by combining PLGA mesh with collagen sponge or gel was successful. This study demonstrated the importance of strengthening of collagen sponge or gel by the composite with PLGA mesh. Cellular responses of ligament cells to PLA/collagen hybrid braids were evaluated both in vitro and in vivo for the ligament tissue engineering (Ide et al., 2001). Hybridization with collagen facilitated cell seeding and spatial cell distribution and promoted cell migration and neoangiogenesis. Electrospinning of PLGA/chitin was investigated to fabricate a biodegradable nanostructured composite matrix for skin tissue engineering (Min et al., 2004). Chitin nanoparticles were distributed uniformly in the PLGA nanofibrous structure and appeared to adhere strongly to PLGA nanofibers by simultaneous electrospinning. Results indicate that the PLGA/chitin composite matrix may be a better candidate than the only PLGA matrix in terms of cell adhesion and spreading. Chen et al. (2005) developed that PLA skeleton covered with bone-like apatite or apatite/collagen composite using phase separation techniques and an accelerated biomimetic coating process for tissue engineered bone. The apatite/collagen composite coating was more effective than apatite coating in improving osteoblastic. Yao et al. (2005) reported on the optimal synthesis parameters and the kinetics of formation of calcium phosphate phase at the surface of PLGA/bioglass composites. PLGA/30% bioglass microspheres based porous scaffolds for bone tissue engineering were examined for their ability to promote osteogenesis of mesenchymal stem cells (MSC). This porous scaffold supported both MSC proliferation and promoted MSC differentiation into cells expressing the osteoblastic phenotype due to ability of bioglass to stimulate osteoblastic differentiation of osteoprogenitor cells. Jung et al. (2005) prepared PLA/calcium metaphosphate composite scaffolds for effective bone tissue engineering using a novel sintering method. Superior characteristics of the novel sintering

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Table 36.1 Lists of hybrid scaffolds for poly(α-hydroxyester) (PLGA, PLA, PGA, and PCL) family with bioceramics and natural polymers Materials

Fabrication methods

Target organ

Pore size (μm)

PLGA/collagen/chitosan

Solvent casting/saltleaching method Phase-separation techniques

Bone

125–500

PLGA/bioactive glass microsphere Poly(propylene fumarate)/ PLGA–PEG microparticles PLA/β-dicalcium silicate

Lay-down pattern (honey comb-like pore) Microsphere, heating mold, porous scaffold Salt leaching

Osteochondral (cartilage) Bone

Particle leaching

Bone

PLA/calcium metaphosphate

Sintering method

Bone

poly(L-lactic acid) (PLLA)/HA/ collagen  chitin fibres PLGA/HA

Ultrasonication and lyophilized

Bone

Gas forming/particle leaching

Bone

Lower: PLGA/TCP Upper: 90% porous PLGA/PLA

TheriForm™ 3D-printing process

Osteochondral defect

PLGA/collagen

Knitted mesh (PLGA), forming microsponge collagen Solvent casting/salt leaching, immersing in collagen solution Conventional freeze-drying method Collagen sponge or gel into the PLGA knitted mesh Collagen solution containing braids, freeze drying Fiber-filled polymerization, etched acetone Electrospinning

Cartilage

PLA/apatite/collagen composite PCL/fibrin glue

PLGA/collagen hybrid sponge PGA fiber/fibrin PLGA mesh/collagen gel PLLA braid/collagen coating PCL fiber/crosslinked pHEMA gel PLGA/chitin PGA mesh/bioactive glass

PLGA/Bioglass® tubular foam

Bioglass particle in distilled water (DW), immersed PGA mesh Dispersion, freeze drying

PLGA/HA/collagen

Phase separation

Bone

Bone

Cartilage Skin Urinary bladder wall Ligament Neural tissue engineering Keratinocyte/ fibroblasts Soft tissue

Intestine, trachea, and blood vessel Guided tissue regeneration

References

Wu et al. (2006) 100–320 Chen et al. (2005) 380  430  Shao et al. 540 μm3 (2006) 350–500 Yao et al. (2005) – Hedberg et al. (2005) 100–500 Cheng et al. (2005) 100–400 Jung et al. (2005) 200 Li et al. (2005) 100–250 Kim et al. (2006) 40–150 Sherwood et al. (2002) – Chen et al. (2003) 355–425 Sato et al. (2001) 300 Hokugo et al. (2006) – Nakanishi et al. (2003) 50–100 Ide et al. (2001) 100–400 Flynn et al. (channel) (2001) – Min et al. (2004) – Day et al. (2004) 100

–

Boccaccini et al. (2005) Pan et al. (2005)

Hybrid, Composite, and Complex Biomaterials for Scaffolds 641

method should have resulted from the fact that the calcium metaphosphate particles could contact directly with cell/tissues to stimulate the cell proliferation and osteogenic differentiation, while the calcium metaphosphate particles would be coated by polymers and hindered to interact with cells/tissue in the case of a solvent casting method. Kim et al. (2006) developed a novel method for fabricating PLGA/nano-HA composite scaffolds by the gas forming and particulate leaching method to high exposure of the bioceramics to the scaffold surface for efficient bone tissue engineering. Compared to the conventional solvent casting/particulate leaching scaffolds, the enhanced bone formation on the gas forming and particulate leaching scaffold may have resulted from the higher exposure of HA nanoparticles at the scaffold surface which allowed for direct contact with the transplanted cells and stimulated the cell proliferation and osteogenic differentiation. Li et al. (2005) developed nano-HA/collagen/PLA composite reinforced by chitin fiber for bone tissue engineering. To enhance the strength of the scaffolds further PLA was linked with chitin fibers by dicyclohexylcarboimide. It showed better mechanical properties than that of the composite without linking. For the regeneration of periodontal tissues, bone around natural teeth and dental implants, nano-HA/collagen/PLA composite with various ratio of each component as guided tissue regeneration membrane was investigated the biodegradability and mechanic behavior in vitro. The optimal nano-HA impregnated with collagen/PLA ratio of the novel membrane is 0.4:1. There is an active dissolution and deposition process of crystals which is propitious to the bone formation on the surface of the composite membrane. In order to improve the physicochemical properties of poly(α-hydroxy acid)s for scaffold materials, the chemical modification on the both end groups of PLA and PGA, that is to say, the addition reaction of moieties to control biological and/or physical properties of biomaterials. For examples, poly(lactic acid-co-lysine-coaspartic acid) (PLAL–ASP) was synthesized in order to endow with cell adhesion property. Similarly, a copolymer of lactide and ε-caprolactone was synthesized to improve the elastic property of PLA. The PLA–poly(ethylene oxide) (PEO) copolymers were synthesized to have the degradative and mechanical properties of PLA and the biological control endow with PEO and its functionalization (Seo et al., 2005). One of the unique characteristics of PLA–PEO block copolymers is the temperature sensitive because of PLA hydrophobicity and PEO hydrophilicity (i.e. sol–gel property that can be applied to injectable cell carriers). Also, nanohybrid composite with other materials have been developed for the application of all organs in the body. Besides poly(α-hydroxyester) family, many synthetic hybrid polymers either degradable or nondegradable are newly launched and tested to mimic the natural tissue and wound healing environment. Examples are hybrid materials of PHEMA hydrogel, injectable poly(N-isopropylacryamide) hydrogel, and polyethylene for neocartilage, poly(iminocarbonates) and tyrosine based poly(iminocarbonates) for bone and cornea, crosslinked collagen/PVA films and an injectable biphasic calcium phosphate/methylhydroxypropylcellulose composite for bone regeneration materials, a PEO-co-polybutylene terephthalate for bone bonding, poly(ortho-ester) and its composites with ceramics for tissue engineered bone, synthesized conducting polymer polypyrrole/hyaruronic acid composite films for the stimulation of nerve regeneration and peptide-modified synthetic polymers for the stimulation of cell and tissue. It is very important for the design and synthesis of more biodegradable and biocompatible scaffold biomaterials to mimic the natural ECM in terms of bioactivity, mechanical properties, and structures. The more biocompatible biomaterials tend to elicit less of an immune response, and to reduce inflammatory response at the implantation site combined with scaffolds manufacturing methods. Ceramic Hybrid Scaffolds Ceramic hybrid scaffolds with synthetic and natural polymers are listed in Table 36.2. In order to improve bioactivity and processability, natural polymers have been mainly impregnated. In order to endow with bioactivity

642 BIOMATERIALS FOR REGENERATIVE MEDICINE

Table 36.2 Lists of hybrid scaffolds for bioceramics with synthetic and natural polymers Materials

Fabrication methods

Target organ

Pore size (μm)

Calcium phosphate/fibrin

Simple mixing

Bone

–

Keratin/HA

Carboxyl-sponge methods

Bone

HA/Collagen

Ice crystal growth method

Bone

HA/chitosan/geltin  MSCs

3D: solid–liquid phase separation

Bone

Chitosan/collagen

Gas forming/freeze drying

Bone

βTCP/collagen

Bone

HA/starch

Suspension, GA crosslinking/ freeze drying Co-precipitation of HA within a gelatin sol, freeze drying Phase-inversion and salt-leaching technique Composite

Calcium phosphate/chitosan

Mannitol salt leaching

Bone

Bone-like hydroxycarbonate apatite/PLA PLA–PEG/HA  BMP

Particle leaching combined with a biomimetic processing Polymer solution dropped on the IP–CHA Laminating

Bone

Gelatin/HA PCL/CAp

Collagen/calcium phosphate layer TCP matrix/HA nanofiber

Hard tissue Bone Bone

Bone Fibroblast

Gel casting/polymer sponge method

Bone

Bone

PDLLA/bioglass

Thermally induced phase separation Simple mixing

Bone/lung

HA/polyamide

Co-solution/co-precipitation

Bone

HA/PCL

Coating

Bone/DDS

Poly(VA–VCI)–HA

Spin coat, photo-patterning

PEEK/HA Calcium phosphate/pHEMA

Selective laser sintering rapid prototyping system Mineralization technique

Hard tissue engineering Bone Bone

HA/Chitosan–gelatin

Phase separation

Bone

βTCP/chitosan

Solid–liquid phase separation

Bone

HA/PLLA

References

Nihouannen et al. (2005) 1–5 Tachibana et al. (2005) 40.1  11  Yunoki et al. 110  21.8 (2006) 70–110 Zhao et al. (2006) 150 Gravel et al. (2006) 100 Zou et al. (2005) 400–500 Kim et al. (2005) – Taddei et al. (2005) – Marques et al. (2005) 52.2–75.2% Xu et al. (2005) 75% Maeda et al. (2005) Hydrogel Kaito et al. 10–40 (2005) 6–8 Yamauchi et al. (thickness) (2004) 300–400 Ramay et al. (20 nm (2004) diameter fiber) 50–100 Wei et al. (2004) 10–100 Verrier et al. (2004) 300 Jie et al. (2004) 150–200 Kim et al. (2004) 6–11 Tsutsumi et al. (thickness) (2003) 100–500 Tan et al. (thickness) (2003) – Song et al. (2003) 300–500 Zhao et al. (2002) 100 Zhang et al. (2001)

Hybrid, Composite, and Complex Biomaterials for Scaffolds 643

Table 36.2 (Continued) Materials

Fabrication methods

Target organ

Pore size (μm)

References

βTCP/PPF  marrow stromal osteoblast Tetraethoxysilane/PDMS

Simple mixing

Bone

–

Sol–gel method

Hepatic reactor

130–200

Gelation/GPSM

Sol–gel process

Bone

300–500

Apatite/polypyrrole

Bioactive coating

Bone

–

BCP/collagen/HCA

Sol–gel and freeze-drying technique 3D printing

Bone

400

Bone

100–250

Peter et al. (2000) Kataoka et al. (2005) Ren et al. (2001) Jiang et al. (2005) Yang et al. (2005) Weinand et al. (2005)

βTCP/hydrogel  stem cell

to calcium phosphate ceramics, fibrin glue was mixed due to its hemostatic, chemotatic, and mitogenic properties for the application of bone tissue engineering (Nihouannen et al., 2005). Zhao et al. (2006) proposed that two types of biomimetic composite materials, chitosan–gelatin and HA/chitosan–gelatin were fabricated and compared to examine the effects of HA on hMSC adhesion and three-dimensional construct development. Results demonstrate that favor osteogenic differentiation upon induction as well as maintain the progenicity of the three-dimensional hMSC constructs of bone tissue engineering. Macroporous composites made of coralline/chitosan were studied for their scaffolding potential in in vitro bone regeneration (Gravel et al., 2006). By using different ratios of natural coralline powder, as in situ gas forming agent and reinforcing phase, followed by freeze drying, scaffolds with controlled porosity, and pore structure were prepared and cultured with MSC. Results suggest that those having a high coralline content, may enhance adhesion, proliferation, and osteogenic differentiation of MSCs in comparison with pure chitosan. Kim et al. (2005) demonstrated that collagen-derived HA/gelatin nanocomposites were synthesized for hard tissue engineering scaffold. In vitro experiment was assessed in comparison with those conventionally mixed gelatin–HA composites. The cell attachment, alkaline phosphatase activity, and osteocalcine were significantly higher on the nanocomposite scaffolds than on the conventional composite scaffolds. This work showed the importance of the nanoscale orientation on the scaffold surface as well as the manufacturing process. Marques et al. (2005) proposed the HA reinforcement of different starch-based polymer such as blends of corn starch and ethylene vinyl alcohol, corn starch and cellulose acetate, corn starch and poly ε-caprolactone (PCL) and its composites with increasing percentages of HA for the application of bone tissue engineering. They concluded that starch-based biomaterials might be good substrates for osteoblast adhesion and proliferation for the potential to be used in orthopedic application and as bone tissue engineering scaffolds. Maeda et al. (2005) developed a novel sponge composed of a PLA composite skeleton covered with bone-like apatite by particle leaching techniques combined with a biomimmetic processing. The scaffold has a large porosity of ~75% with large pores and shows mechanical ductility. Kaito et al. (2005) investigated bone morphogenic proteins (BMPs)/interconnected porous calcium HA/PLA–PEG composite for the construction of a carrier/scaffold system for BMPs. At 8 weeks after implantation, all bone defects in groups treated with 5 or 20 μg of BMP were completely repaired with sufficient strength. Furthermore, the reduction of necessary BMP amount was achieved about a tenth of the amount needed using this carrier scaffold system probably due to the superior osteoconduction ability of interconnected porous

644 BIOMATERIALS FOR REGENERATIVE MEDICINE

calcium HA and the optimal drug delivery system provided by PLA–PEG, inducing new bone formation in the connected pores. Another good example for the application of drug delivery system to the tissue engineered bone was HA/PLA composite coating on HA porous bone scaffold for controlled release of antibiotic drug as tetracycline hydrochloride (Kim et al., 2004). The HA scaffold obtained by a polymeric reticulate method, possessed high porosity (87%) and controlled pore size (150–200 μm). To improve the osteoconductivity and bioactivity of the coating layer, HA powder was hybridized with PCL solution to make the HA  PCL composite coating. Although initial burst release of 20–30% revealed in initial period within 2 h, the release rate was sustained for prolonged periods with controlled HA  PCL coating condition. Kataoka et al. (2005) developed a novel organic–inorganic hybrid scaffold for the culture of HepG2 cell in a bioreactor. The scaffold was made from tetraethoxysilane and poly(dimethylsiloxane) (PDMS) by a sol–gel method using sieved sucrose particles as a porogen (130–200 μm). When HepG2 cell cultivated in hybrid porous scaffold, HepG2 cells secreted a  three-fold greater amount of albumin than that secreted in a monolayer culture. To overcome limited shapes and sizes of conventional scaffolds, a direct hydrogel injection system combining bone marrow derived differentiated MSCs/hydrogel/β-TCP has been proposed (Weinand et al., 2005). The scaffolds provided support for the formation of bone tissue in collagen I, fibrin, alginate, and pluronic F127 hydrogels during culturing in oscillating and rotating dynamic condition. Expression of bone-specific genes was significantly higher in the collagen I samples. Pluronic F127 hydrogel did not support formation of the bone tissue. All samples cultured in dynamic oscillating revealed slightly higher mechanical strength than under rotating conditions. Another specific manufacturing processes were proposed for the desired physicochemical properties at specific target organ such as (i) a selective laser sintering of polyetheretherketone/HA biocomposite blends for the application of bone tissue engineering (Tan et al., 2003), (ii) a photo patterned polyvinylalcohol bearing transcinnamate moieties as chromophoric groups/HA composites for the hard tissue engineering (Tsutsumi et al., 2003), (iii) a biomimetic process/bioactive coating using simulated body fluid of apatite/polypyrrole composite for bone tissue engineering (Jiang et al., 2005), (iv) the multilayer sheets (2–10 layers) of collagen/calcium phosphate using enzymatic mineralization for soft tissue (Yamauchi et al., 2004) and so on. In summary, hybridization techniques with various types of biomaterials might improve the physicochemical properties. Therefore, appropriated manufacturing process for scaffolds must be developed. Natural Polymers Hybrid Scaffold The synthetic biodegradable polymers are easily formed into desired shapes with good mechanical strength and the duration of degradation can be estimated. Despite these advantages, the scaffolds derived from synthetic polymers are insufficient for cell recognition signal, and their hydrophobic properties obstruct smooth cell seeding. In contrast, naturally derived polymers have the potential advantages of specific cell interactions and a hydrophilic nature, but possess poor mechanical properties. Thus, these two kinds of biodegradable polymers have been hybridized to combine the advantageous properties of both constituents (Khang et al., 2004a). In this section, hybridization of natural polymers with another natural polymer and synthetic polymer is mainly reviewed as lists in Table 36.3. Lee et al., (2004) developed β-chitin/collagen hybrid scaffold by means of combining salt leaching and freeze-drying method with 260–300 μm pore size. The mechanical strength and the rate of biodegradation increased with the porosity controlled by the salt concentration. After 14 days, the fibroblasts showed a good affinity to and proliferation on all collagen-coated chitin. Collagen/alginate and collagen/hyaluronan composite hydrogels were investigated for their ability to support ECM synthesis by vocal fold fibroblasts with limited hydrogel compaction and/or resorption (Hahn et al., 2006). Among these two composite scaffolds, collagen/alginate hydrogels appear the better biomaterials for vocal fold restoration. Daamen et al. (2003)

Hybrid, Composite, and Complex Biomaterials for Scaffolds 645

Table 36.3 Lists of hybrid scaffolds for natural polymers with other biomaterials Materials

Fabrication methods

Target organ

Pore size (μm)

References

Chitosan/alginate

Freeze drying, crosslinked by CaCl2 Neutralization technique

Bone

100–200

Li et al. (2005)

Bone

120–250

Freeze-drying technique

Liver

150–200

Salt leaching/freeze drying

Fibroblast

260–330

Collagen–alginate, collagen– hyaluronan

BoneSave®, Ostin®

–

Collagen/elastin/gly cosaminoglycan Chitosan/gelatin

EDC crosslinking

Vocal fold lamina propria Soft tissue

Tampieri et al. (2005) Seo et al. (2006) Lee et al. (2004) Hahn et al. (2006)

20–100

Freeze drying/ice microparticle

–

20–102

Gelatin/siloxane

Sol–gel/post-gelation soaking/ freeze drying Collagen scaffold (Lyostypt®)

Bone

5–500

Bone

–

Liver

127–833

Skin

95.0–150.5

Hyaluronic acid/PEG hydrogel

Simple mixing/freeze-drying method Suspension, freeze-drying, crosslinking Photopolymerization

Protein delivery

–

Collagen/hyaluronic acid

EDC crosslinking-freezing drying

Soft tissue

40–230

Chitosan/hyaluronic acid

Wet spinning method

Cartilage

–

Gelatin/chondroitin/hyaluronic acid Hyaluronan/gelatin hydrogel

Powder mixing, crosslinking, freeze drying Centrifugal casting

Cartilage

–

Vascular graft

–

HA/alginate Alginate/galactosylated chitosan β-chitin/collagen

Collagen/PRP Fibroin/collagen Collagen–GAG

Daamen et al. (2003) Mao et al. (2003) Ren et al. (2002) Sarkar et al. (2006) Lv et al. (2005) O’Brien et al. (2005) Leach et al. (2005) Park et al. (2002) Yamane et al. (2005) Chang et al. (2006) Mironov et al. (2005)

designed molecularly defined collagen/elastin/glycosaminoglycan (GAG)/chondroitin sulfate scaffolds with 20–100 μm by carborimide crosslinking for soft- tissue engineering. Type I collagen provides adhesive properties and tensile strength. Elastin provides elasticity to tissue/organs and is crucial for blood vessels in order to cope with the variations in blood pressure. GAGs are negatively charged polysaccharides with biocharacteristics-like hydration of the ECM and binding of effector molecules. The attachment of chondroitin sulfate increased the water-binding capacity to up to 65%. Scaffolds with higher collagen content had a higher tensile strength whereas addition of elastin increased elasticity. This work showed the importance in the design and application of tailor-made biomaterials for tissue engineering. Chitosan/gelatin hybrid polymer network scaffolds with monolayer and bilayer were prepared via the freeze-drying techniques by using the ice microparticle as a porogen (Mao et al., 2003). The porosity and the pore size of the scaffold could be modulated with thermodynamic and kinetic parameters of ice formation.

646 BIOMATERIALS FOR REGENERATIVE MEDICINE

Porous and bioactive gelatin/siloxane hybrids were proposed by using a combined sol–gel processing, postgelation soaking and freeze-drying process for bone tissue engineering (Ren et al., 2002). The pore size of the hybrid scaffolds can be well controlled by varying the freezing temperature such as 17°C for 300–500 μm, 80°C for 30–50 μm, and 196°C for 5–10 μm. Sarkar et al. (2006) investigated a platelet rich plasma loaded collagen scaffold for bone formation in a long bone defect model. Lv et al. (2005) proposed fibroin/collagen hybrid scaffold with 127–833 μm pore size by freeze-drying method and cultured HepG2 cell. O’Brien et al. (2005) studied the effect of pore size on the cell adhesion collagen/GAG scaffolds. This work reported the strong correlation between the scaffold specific surface area and cell attachment indicates that cell attachment and viability are primarily influenced by scaffold specific surfacearea over range of 95.9–150.5 μm of pore size for MC3T3 cell. Seo et al., (2006) investigated alginate/galactosylated chitosan scaffold with 150–200 μm to enhance liverspecific function of hepatocytes cocultured with NIH3T3 and showed the potential for bioartificial liver devices. Chitosan/alginate hybrid scaffold using coacervation method was observed to enhance mechanical and biological properties for bone tissue engineering (Li et al., 2005). HA/alginate hybrid composite was prepared through bioinspired nucleation and confirmed the ability to favor cell growth and to maintain their osteoblastic functionality. Hyaruronic acid presents a unique combination of advantages for biomaterial formulations; nonimmunogenic, non-adhesive, bioactive GAGs that has been associated with several cellular process, including angiogenesis and the regulation of inflammation. But one of significant drawbacks is the weak of water. In order to improve the water-resistance, composites as well as chemical modification were conducted. Leach et al. (2005) prepared photocrosslinkable hyaluronic acid–PEG hydrogel for protein release in tissue engineering scaffold. Also, porous collagen/hyaluronic acid hybrid scaffold modified by 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide crosslinking with 40–230 μm (Park et al., 2002). Chitosan/hyaluronic acid hybrid scaffold was developed for the application of cartilage tissue engineering (Yamane et al., 2005). Gelatin/chondroitin/hyaluronic acid hybrid scaffold prepared by successive powder mix/crosslinking/freeze drying was proposed for tissue engineered cartilage (Chang et al., 2006). The results for the tissue engineering-treated group were significantly satisfactory, the repair tissue being hyaline cartilage and/or fibrocartilage. Mironov et al. (2005) fabricated tubular tissue constructs by centrifugal casting of cells suspended in an in situ crosslinkable hyaluronic acid/gelatin hybrid hydrogel. Centrifugal casting in this hybrid hydrogel would enable rapid fabrication of tissue engineered vascular grafts, as well as other tubular and planar tissue engineered construct. Natural polymers discussed in this section are proteins, albumin, gluten, elastin, fibroin, cellulose, starch, sclerolucan, elsinan, pectin (pectinic acid), galactan, curdlan, gellan, levan, emulsan, dextran, pullulan, heparin, silk, chondroitin 6-sulfate, and so on. Although their hybrid scaffolds were not explained in this section, they are of interest due to the most abundant biopolymers on earth, and their unusual and useful functional properties. Typical properties are (i) biocompatibility and nontoxic, (ii) easily processing as film and gel status, (iii) heat stability and thermal processability over broad temperature range, and (iv) water solubility. Miscellaneous Scaffolds Cytokines are polypeptides that transmit signals to modulate cellular activity and tissue development such as cell patterning, motility, proliferation, aggregation, and gene expression. As in the development of the tissue engineered organs, regeneration of functional tissue requires maintenance of cell viability and differentiated function, encouragement of cell proliferation, modulation of the direction, and speed of cell migration, and regulation of cellular adhesion. For example, transforming growth factor-β1 (TGF-β1) might be required to induce osteogenesis and chondrogenesis from bone marrow-derived mesenchymal stem cell (DeFail et al.,

Hybrid, Composite, and Complex Biomaterials for Scaffolds 647

Table 36.4 Lists of miscellaneous hybrid scaffolds Materials

Fabrication methods

Target organ

Pore size (μm)

References

Peptide hydrogel/Poly high internal phase emulsion (HIPE) PEG–fibrinogen hydrogel

Polymerization of a HIPE

Bone

100

Bokhari et al. (2005)

PEGylated hydrogel

Starch-based polymers

Composite

Smooth muscle – cell Bone –

SIS, UBS, UBM, UBS  UBM, SS Oligo(poly(ethylene glycol) fumarate)/gelatin  growth factor PLGA  TGF-β1/PEG

Natural biomaterials

Soft tissue

–

Hydrogel/microparticle

Cartilage

–

PPF/TCP Chitosan–gelatin/TGF-β1 PLGA/VEGF

Almany et al. (2005) Marques et al. (2002) Freytes et al. (2004) Holland et al. (2005)

86.64 DeFail et al. 76.88 (2006) (microsphere) Image-based designed and solid Bone 300–800 Schek et al. freeform fabricated scaffold (2006) Freeze-drying method Cartilage defect 300–700 Guo et al. (2006) Gas forming Angiogenesis 250–400 Jang et al. (2005) Microsphere in hydrogel

Cartilage

UBS: urinary bladder submucosa; UBM: urinary bladder matrix; SS: stomach submucosa

2006). Also, brain-derived neurotrophic factor (BDNF) can be enhanced to regenerate spinal cord injury (Khang, et al., 2004b). The easiest method for the delivery of the growth factor is the injection near the site of cell differentiation and proliferation. The most significant problem of the direct injection method of growth factors is the relatively short half-life, the relatively high molecular weight and size, very low tissue penetration, and potential toxicity of systemic level. One promising way of the improvement technique of their efficacy is the locally controlled release of bioactive molecules for desired release period by the impregnation into a biomaterial scaffold as listed in Table 36.4. The duration of cytokine release from a scaffold can be controlled by the types of biomaterials used, the loading amount of cytokine, the formulation factors, and the fabrication process. The release mechanisms are largely divided into three categories; (i) diffusion controlled, (ii) degradation controlled, and (iii) solvent controlled release mechanism through the selection of biomaterials. Mechanism of biodegradable scaffolds materials was controlled by degradation controlled whereas that of nondegradable one was regulated by diffusion and/or solvent controlled. Desired release pattern such as constant, pulsatile, and time programmed behaviors along the specific site and the type of injury can be achieved by the appropriate combination of these mechanisms. Also, cytokine release system might be designed in a variation with geometries and configurations such as scaffold, tube, microsphere, injectable forms, fiber, and so on (Holland et al., 2005). Another available emerging technology is the tethering to the surface that is, immobilization of protein on the surface of scaffold matrix. For the enhancement of cytokine activity, PEO chain was applied as a short spacer between the surface of scaffold and the cytokine. Tethered epidermal growth factor (EGF), immobilized

648 BIOMATERIALS FOR REGENERATIVE MEDICINE

to the scaffold through PEO chain, showed more improved DNA synthesis or cell rounding compared to the physically adsorbed EGF surface (Khang et al., 2006). Conjugation of cytokine with inert carrier prolongs the short half-life of protein molecules. Inert carriers are albumin, gelatin, dextran, and PEG. Especially, PEGylation that means PEG conjugated cytokine is most widely used for the release. It appears to decrease the rate of cytokine degradation, attenuate the immunological response, and reduce clearance by the kidneys. Also, this PEGylated cytokine can be impregnated into scaffold materials by physical entrapment for the sustained release. This conjugation method can be applied to the delivery of proteins and peptides. Immobilized RGD and YIGSR which are typical ECM proteins onto the biomaterials can enhance cell viability, function, and recombinant products in cell. Gene-activating scaffolds are being designed to deliver to targeted gene resulting in the stimulation of specific cellular responses at the molecular level. Modification of bioactive molecules with resorbable biomaterials systems obtains specific interactions with cell integrins resulting in cell activation. These bioactive bioglasses and macroporous scaffolds can also be designed to activate genes that stimulate regeneration of living tissue (Guo et al., 2006). Gene delivery would be accomplished by complexation with positively charged polymers, encapsulation, and gel by means of scaffold structure (Schek et al., 2006). Methods of gene delivery for gene-activating scaffolds are almost same manner with those of protein, drug, and peptides.

SYNTHETIC/NATURAL HYBRID SCAFFOLD FOR TISSUE ENGINEERED INTERVERTEVERAL DISK Since there is no optimal treatment for the persistent pain associated with intervertebral disk (IVD) degeneration, focus has been shifted toward replacement using tissue engineered IVD. An alternative approach has been carried out by a functional IVD composed of disk cells seeded to various scaffolds using tissue-engineering principles. Several kinds of scaffolds were manufactured such as PLGA, PLGA/SIS, PLGA/DBP, PLGA/SIS/DBP, crosslinked SIS sponge, and PGA nonwoven mesh as shown in Figure 36.2.

(a)

(b)

(c)

(d)

(e)

(f)

Figure 36.2 SEM microphotographs of various types of scaffolds. (a) PLGA only, (b) SIS/PLGA, (c) DBP/PLGA, (d) SIS/DBP/PLGA, (e) SIS sponge and (f) PGA nonwoven.

Hybrid, Composite, and Complex Biomaterials for Scaffolds 649

Natural/synthetic hybrid scaffolds as PLGA, SIS/PLGA, DBP/PLGA, and SIS/DBP/PLGA were manufactured by solvent casting/salt-leaching method, and crosslinked SIS sponge was fabricated by freeze-drying method. Scaffolds were characterized by scanning electron microscope (SEM) and porosimetry. (Kim et al., 2006) It was evaluated cell proliferation by 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay (Figure 36.3) and DNA quantification in vitro and in vivo. Scaffolds seeded rabbit disk cells were implanted into nude mouse and then confirmed by the histological staining by H&E (Figure 36.5), Safranin-O, Masson’s

Optical density (590 nm)

1.2

A B C D E F

0.8

0.4

0.0 1

2

3

Time (weeks)

Figure 36.3 Activity of proliferation rate of disk cells analyzed by MTT assay after 1, 2, and 3 weeks in in vitro. (a) PLGA only, (b) SIS/PLGA, (c) DBP/PLGA, (d) SIS/DBP/PLGA, (e) SIS sponge and (f) PGA nonwoven.

(a)

(b)

Figure 36.4 Hybrid type of tissue engineered IVD composed of (a) annulus fibrous as PLGA/DBP composite scaffold with annulus fibrous cell and (b) nucleus pulpose as thermosensitive MPEG–PCL hydrogel with nucleus pulpose cell for rabbit model.

650 BIOMATERIALS FOR REGENERATIVE MEDICINE

1 week

4 weeks

6 weeks x p

(a)

s x

(b) p d v

(c)

p

s (d) d s

s

(e) s

(f)) g

Figure 36.5 Photomicrographs from H&E histological sections of disk cell seeded various types of scaffolds implanted on the back of nude mice after 1, 4, and 6 weeks. (a) PLGA only, (b) SIS/PLGA, (c) DBP/PLGA, (d) SIS/DBP/PLGA, (e) SIS sponge and (f) PGA nonwoven. (200 ). p:undegraded PLGA interconnected area, s:SIS particle and interconnected area, d:DBP particle, g:undegraded PGA nonwoven, v:newly formed vascular capillary, and x:disk cells. trichrome and type II collagen immunochemical staining. SIS sponge appeared better DNA production and proliferation of disk cell but the formation of disk tissue was incomplete due to fast rate of degradation. Natural biomaterials impregnated PLGA scaffolds have better potential for the application of tissue engineered disk due to its bioactive molecules. Figure 36.4 shows that hybrid type of tissue engineered IVD composed of annulus fibrous as

Hybrid, Composite, and Complex Biomaterials for Scaffolds 651

PLGA/DBP composite scaffold and nucleus pulpose as thermosensitive MPEG–PCL hydrogel. This experiment indicated that porosity, bioactive materials, and biodegradation duration play an important role for the formation of tissue engineered disk (Khang et al., 2006).

CONCLUSIONS Tissue engineering including regenerative medicine shows tremendous potential as a revolutionary research push. Also, many successful results have reported the potential for regenerating tissues and organs such as skin, bone, cartilage, nerve of peripheral and central, tendon, muscle, corneal, bladder and urethra, and liver as well as composite systems like a human phalanx and joint on the basis of scaffold biomaterials from polymers, ceramic, metal, composites, and its hybrids. As previously emphasized, scaffold materials must contain the site of cellular and molecular induction and adhesion and must allow for the migration and proliferation of cell through porosity. It should also maintain strength, flexibility, biostability, and biocompatibility to mimic a more natural, three-dimensional environment. From this point of view, the control over precise biochemical signal must be needed by the combination of scaffold matrix and bioactive molecules including genes, peptide molecules, and cytokines. Moreover, the combination of the cells and redesigned bioactive scaffolds has attempted to expand to a tissue level of hierarchy. In order to achieve this goal, the novel hybrid scaffold biomaterials, the novel scaffolds fabrication methods, and the novel characterization methods must be developed.

ACKNOWLEDGMENTS This work was supported by grants from KMOWH(0405-BO01-0204-0006) and KMOST (SC3100).

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37 Surface Modification of Biomaterials Andrés J. García

INTRODUCTION Biomaterial Interfaces in Regenerative Medicine Biomaterials, either synthetic (e.g. polymers, metals, ceramics) or natural (e.g. proteins, polysaccharides), play central roles in tissue engineering and regenerative medicine applications by providing (i) three-dimensional scaffolds to support cellular activities; (ii) matrices for delivery of therapeutic agents (e.g. drugs, proteins, DNA, siRNA); and (iii) functional device components (e.g. mechanical supports, sensing/stimulating elements, non-thrombogenic surfaces, diffusional barriers). The bulk properties of the biomaterial are critical determinants of the biological performance of the material (Ratner et al., 2004). For example, the mechanical properties of a vascular substitute, including elastic modulus, ultimate tensile stress, and compliance, dictate the ability of this tissue construct to support the applied mechanical loads associated with blood flow. On the other hand, the biological response to a biomaterial is governed by the material surface properties, primarily surface chemistry and structure. Protein adsorption/activation and cell adhesion, events that regulate host responses to materials, occur at the biomaterial–tissue interface, and the physicochemical properties of the material surface modulate these biological events (Anderson, 2001). For instance, the chemical properties of the surface of a vascular substitute control blood compatibility (i.e. protein adsorption, platelet adhesion, thrombogenicity, patency). Hence, modification of biomaterial surfaces represents a promising route to engineer biofunctionality at the material–tissue interface in order to modulate biological responses without altering material bulk properties. Overview of Surface Modification Strategies Numerous surface modification approaches have been developed for all classes of materials to modulate biological responses and improve device performance. Applications include reduction of protein adsorption and thrombogenicity, control of cell adhesion, growth and differentiation, modulation of fibrous encapsulation and osseointegration, improved wear and/or corrosion resistance, and potentiation of electrical conductivity (Ratner et al., 2004). Surface modifications fall into two general categories: (i) physicochemical modifications involving alterations to the atoms, compounds, or molecules on the surface; and (ii) surface coatings consisting of a different material from the underlying support. Physicochemical modifications include chemical reactions (e.g. oxidation, reduction, silanization, acetylation), etching, and mechanical roughening/polishing and patterning (Figure 37.1). Overcoating alterations comprise grafting (including tethering of biomolecules), non-covalent and covalent coatings, and thin film deposition (Figure 37.2). While the specific requirements of the surface modification approach vary with application, several characteristics are generally desirable. Thin surface modifications are preferred for most applications since thicker

656

CF3

CF3

CF3

CF3

CF3

C O C O C O C O C O OH

OH

OH

OH

OH (CF3C O)2O

O

O

O

O

O

Surface chemical reaction (e.g. fluorination of hydroxylated surfaces via tri-fluoroacetic anhydrides) TiO2

HNO3 Ti

Ti Conversion coating (e.g. passivation of titanium to yield titanium oxide layer)

sandblasting

Mechanical roughening (e.g. sandblasting)

Figure 37.1 Schematic representations of common physicochemical surface modifications of biomaterials. coatings often negatively influence the mechanical and functional properties of the material. Ideally, the surface modification should be confined to the outermost molecular layer (10–15 Å), but in practice, thicker layers (10–100 nm) are used to ensure uniformity, durability, and functionality. Stability of the modified surface is a critical requirement for adequate biological performance. Surface stability not only refers to mechanical durability (i.e. resistance to cracking, delamination, debonding) but also chemical stability, especially in aggressive, chemically active environments such as biological milieu. Several types of surface rearrangements, such as translation of surface atoms or molecules in response to environmental factors and mobility of bulk molecules to the surface and vice versa, readily occur in polymers and ceramics following exposure to biological fluids. Given the uniquely reactive nature and mobility/rearrangement of surfaces, as well as the tendency of surfaces to readily contaminate, rigorous analyses of surface treatments are essential to surface modification strategies. Surface analyses technologies generally focus on characterizing topography, chemistry/composition, and surface energy (Woodruff and Delchar, 1994) (Table 37.1). Important considerations for these surface analysis technologies include operational principles (impact of high-energy particles/X-rays under ultrahigh vacuum, adsorption or emission spectroscopies), depth of analysis, sensitivity, and resolution. For most applications, several analysis techniques must be used to obtain a complete description of the surface.

PHYSICOCHEMICAL SURFACE MODIFICATIONS Physicochemical modifications involve alterations to the atoms, compounds, or molecules on the material surface (Figure 37.1). Chemical Modifications Countless chemical reactions, including UV/laser irradiation and etching reactions to clean, alter or cross-link surface groups, have been developed to modify biomaterial surfaces (Ratner and Hoffman, 2004). Non-specific reactions yield a distribution of chemically distinct groups at the surface, and the resulting surface is complex and difficult to characterize due to the presence of different chemical species in various concentrations.

657

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Non-covalent overcoats (e.g. vapor deposition, solvent casting) dipping in alternating polyelectrolyte solutions

Layer-by-layer deposition of polyelectrolytes

monomer

Grafting of overcoats (e.g. radiation and photografting, plasma deposition) X X X X X X X X X

Self-assembled films (e.g. Langmuir–Blodgett, self-assembled monolayers)

Surface-modifying additives

Biomolecule immobilization (e.g. passive adsorption, tethering)

Figure 37.2 Schematic representations of common overcoating technologies for surface modification. Nevertheless, non-specific chemical reactions are widely used in biomaterials processing. Examples of non-specific reactions include radio-frequency glow discharge (RFGD) in different plasmas (e.g. oxygen, nitrogen, argon), corona discharge in air, oxidation of metals, and acid–base treatments of polymers. In contrast, specific chemical reactions target particular chemical moieties on the surface to convert them into another functional group with few side (unwanted) reactions. Acetylation, fluorination of hydroxylated surfaces via tri-fluoroacetic anhydrides, silanization of hydroxylated surfaces, and incorporation of glycidyl groups into polysiloxanes are examples of specific chemical reactions. In addition, various chemical methods exist to tether biomacromolecules onto available anchoring groups on surfaces, as described in section “Biological Modification of Surfaces.” Reaction of metal surfaces to produce an oxide-rich layer that conveys corrosion resistance, passivation, and improved wear and adhesive properties (also referred to as conversion coatings) are common surface modifications in metallic biomaterials. For example, nitric acid treatment of titanium and titanium alloys to generate titanium oxide layers is regularly performed on titanium-based medical devices, and the excellent

Table 37.1 Common surface analysis techniques Principle

Operation

Spatial resolution

Information depth

Sensitivity

Texture

Chemical composition information Elements

Contact angle AFM

SEM

EDXA AES

SIMS FTIR

Isotopes

Air Liquid

NA

3–20 Å

NA

Indirect

Air Aqueous

Atomic

NA

Single atom

Yes

No

No

No

Vacuum

40 Å

5–10 Å

High

Yes

No

No

No

Vacuum

40 Å

1 μm

107 g/cm2

No

Z5

No

No

Vacuum

100 Å

15–50 Å

101 0 g/cm2 0.1 atom%

No

Z3

Chemical shift

No

Vacuum

10 μm

10–150 Å

101 0 g/cm2 0.1 atom%

No

Z3

No

Vacuum

3–10 μm

10 Å

101 3 g/cm2

No

All

Chemical shift (excellent) Yes

Air Aqueous (ATR)

10 μm

1 μm

1 mol%

No

Indirect

Vibration frequency

No

Additional Surface energy

Crystallinity

Yes Monolayer orientation

Surface Modification of Biomaterials 659

XPS

Liquid wetting of surfaces Records interatomic forces between tip and sample. Secondary electron emission caused by electron bombardment is imaged X-ray emission caused by electron bombardment Auger electron emission caused by electron bombardment X-rays cause emission of photoelectrons with characteristic energies Ion bombardment causes secondary ion emission Molecular vibrations resulting from adsorption of IR radiation

Compounds

660 BIOMATERIALS FOR REGENERATIVE MEDICINE

surface roughness

surface topography

Figure 37.3 Surface roughness and topography. biocompatibility properties of titanium are attributed to this oxide layer (Albrektsson et al., 1983). Implantation of ions into surfaces via a beam of accelerated ions has been applied to modify the surface properties of mostly metals and ceramics. For example, ion beam implantation of nitrogen into titanium and boron and carbon into stainless steel improves wear resistance and fatigue life, respectively (Sioshansi, 1987). In addition, recent evidence suggests that ion beam implantation of silicone and silver can also enhance the blood compatibility and infection resistance of silicone rubber catheters (Bambauer et al., 2004). Topographical Modifications The size and shape of topographical features on a surface influence cellular and host responses to the material. For example, surface macro- and micro-texture alters cell adhesion, spreading, and alignment (Curtis and Wilkinson, 1998; Flemming et al., 1999) and can regulate cell phenotypic activities, including neurite extension and osteoblastic differentiation (Boyan et al., 1996; Jansen and von Recum, 2004). Moreover, surface topography can have significant in vivo effects. For instance, implant porosity modulates bone and soft tissue ingrowth (Yamamoto et al., 2000; Pilliar, 2005), and surface texture alters epithelial downgrowth responses to percutaneous devices and inflammatory reactions and fibrous encapsulation to materials implanted subcutaneously (Chehroudi et al., 1989; Brauker et al., 1995; Chehroudi and Brunette, 2002). While specific surface texture parameters that elicit particular biological responses have been identified in several cases, the mechanisms generating these behaviors remain poorly understood. Methods for generating surface texture can be grouped into approaches for engineering either roughness or topography (Figure 37.3). Surface roughness indicates a random or complex pattern of features of varying amplitude and spacing, typically on a scale smaller than a cell (10–20 μm). On the other hand, surface topography refers to patterns of well-defined, controlled features on the surface. Surface roughness has been traditionally modified via sandblasting, plasma spraying, and mechanical polishing, and it is the non-specific nature of these processes that renders surfaces with random or complex topographies. Ion beam and electric arc (for conductive materials) texturing approaches have also been applied to modulate surface roughness. For generating controlled topographies, micro- and nano-machining techniques have been exploited using silicon, glass, and polymers as substrate materials (Flemming et al., 1999). Photolithography combined with reactive plasma and ion etching has been extensively applied to generate surfaces with well-defined topographies. This technique allows the preparation of machined silicon and polymeric substrates and silicon templates which can then be used as molds to transfer features to polymers via solvent casting or injection molding. Similarly, LIGA (German for “Lithographie, Galvanoformung, Abformung”), electron beam, and laser machining have been used to manufacture defined topographical features on various materials. More recently, hot embossing imprint lithography has been applied for low cost and rapid fabrication of micro- and nano-scale features on biomedically relevant polymers (Charest et al., 2004).

OVERCOATING TECHNOLOGIES Coating strategies rely on the deposition of a surface layer consisting of a different composition from the underlying base material (Figure 37.2). These surface modification approaches include non-covalent and covalent coatings (Ratner and Hoffman, 2004).

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Non-covalent Coatings Major advantages of non-covalent coatings include simple application and the ability to coat a variety of different base materials. Examples of common non-covalent coating methods are solvent casting, and vapor deposition of metals, parylene, and carbons. In the Langmuir–Blodgett deposition method, one or more highly ordered layers of surfactant molecules (e.g. phospholipids, amphiphiles) are placed at the surface of the base material via assembly at the air–water interface and compression of the surfactant molecules. Langmuir–Blodgett films exhibit high order and uniformity and provide flexibility in incorporating a wide range of chemistries. The stability of these films can be improved by cross-linking or internally polymerizing the surfactant molecules following film formation. Another surface modification strategy that takes advantage of intermolecular interactions is the deposition of multilayer polyelectrolytes (e.g. poly(styrenesulfonate)/poly(allylamine), hyaluronic acid/chitosan). In this simple layer-by-layer method, a charged surface is sequentially dipped into alternating aqueous solutions of polyelectrolytes of opposite charge in order to deposit multilayers of a polyelectrolyte complex. Another elegant strategy for surface modification is the use of surface-modifying additives. These molecules are blended in the bulk material during fabrication but will spontaneously rise to and concentrate at the surface due to the driving force to minimize interfacial energy. Covalent Coatings Covalent coating methodologies rely on direct tethering of overcoats onto the base material to improve film stability and adherence. Radiation grafting, both with ionizing radiation and high-energy electron beams, and photografting have been extensively pursued to modify polymer substrates in order to introduce chemically reactable groups into inert hydrophobic polymers and polymerize overcoats onto the base support (Ratner and Hoffman, 2004). In principle, the radiation breaks chemical bonds in the base material into free radicals and other reactive species, which are then exposed to a monomer. The monomer reacts with the reactive species at the surface and propagates as a free radical chain reaction into a surface grafted polymer. These strategies allow for generation of a wide range of surface chemistries, and unique graft co-polymers can be synthesized by combining different monomers. Plasma deposition (also referred to as glow discharge deposition) via radio frequency or microwave has also been extensively applied to biomaterial surface modification (Hoffman, 1988). In particular, RFGD plasma deposition has received considerable attention because it can generate continuous (relatively free of pin holes and voids) conformal coatings that can be applied to many different types of supports (metals, ceramics, polymers) with complex geometries. In addition, these films exhibit good adherence to the substrate and can be engineered to present different functionalities, although the resulting chemistry is complex and ill-defined. In contrast to these relatively low-energy/low-temperature plasmas, high-energy/high-temperature plasmas have also been used to apply inorganic surface modifications onto inorganic substrates. For example, calcium phosphate ceramic particles, such as hydroxyapatite, have been deposited via flame spraying onto titanium and cobalt chrome orthopedic implants to improve osseointegration (Gruner, 2005). Coatings consisting of self-assembled monolayers (SAMs) have gained significant attention as robust surface modification agents (Ulman, 1991; Mrksich and Whitesides, 1995). These films spontaneously assemble, form highly ordered, well-defined surfaces with excellent chemical stability, and provide a wide range of available surface functionalities. The basic structure of molecules that form SAMs is an anchoring “head” group, organic chain backbone, and functional “tail” group. Common SAM systems are alkanethiols on coinage metals (gold, silver), n-alkyl silanes on hydroxylated supports (glass, silica), and phosphoric acid or phosphate groups on titanium or tantalum surfaces. Assembly of these organic chains into highly ordered structures is driven by the strong adsorption of the anchoring “head” group of the monolayer constituent to the surface and van der Waals interactions of the backbone chains. The order and stability of the SAMs are strongly influenced by the length of the backbone chain, and in the case of alkanethiols, molecules with backbones between

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9 and 24 methylene groups assemble well on gold. Importantly, the terminal functional group is presented at the surface–solution interface and controls the physicochemical properties of the SAM.

BIOLOGICAL MODIFICATION OF SURFACES Biomolecules (e.g. cell receptor ligands, enzymes, antibodies, pharmacological agents, lipids, nucleic acids) have been immobilized onto and within biomaterial supports for numerous therapeutic, diagnostic, and bioprocess applications. Table 37.2 lists several examples of biological modifications to surfaces for biomedical and biotechnological applications. The rationale for these hybrid materials integrating synthetic and biological components is to convey biofunctionality and hence engineer materials that elicit desired biological responses or have attributes associated with biosystems. One of the earliest examples of this strategy is the immobilization of heparin onto polymer surfaces to improve blood compatibility. More recently, drug eluting stents (stents coated with a polymeric layer loaded with anti-hyperplasia drugs) have been developed to reduce restenosis and improve patency. Another example of a widely used biological modification strategy is the immobilization of adhesive ligands, either adsorbed proteins (e.g. fibronectin, laminin) or tethered synthetic oligopeptides (e.g. RGD), on synthetic and natural supports to promote cell adhesion and function in various tissue engineering and regenerative medicine applications (Lutolf and Hubbell, 2005). Three major methods are used to immobilize biomolecules onto biomaterial surfaces: physical adsorption, physical entrapment, and covalent immobilization (Figure 37.4) (Hoffman and Hubbell, 2004). Passive physisorption of biomacromolecules (i.e. proteins, polysaccharides, nucleic acids) is a simple yet efficient method to render surfaces biologically active. Everyday applications include coating of synthetic materials with extracellular matrix proteins, such as fibronectin and collagen, to improve cell adhesion. As discussed in Chapter 56, protein adsorption is a complex, dynamic energy-driven process involving hydrophobic interactions, electrostatic interactions, hydrogen bonding, and van der Waals forces. Protein parameters such as primary structure, size, and structural stability as well as surface properties including surface energy, and chemistry influence the biological activity of the adsorbed biomacromolecules. It is important to point out these biologically modified surfaces can undergo further modifications, such as displacement of adsorbed proteins and cell-mediated deposition and remodeling of matrix components, in the biological milieu. As an approach to improve the stability of these modified surfaces, the biological molecules can be cross-linked following adsorption. Finally, the use of highaffinity interactions, for example avidin–biotin and antibody–antigen, represents a special case of these physical immobilization methods that is particularly important in diagnostics and bioprocessing. Table 37.2 Biomedical and biotechnological applications of immobilized biomolecules Biomolecule

Applications

Heparin Fibronectin, collagen RGD peptides Antibodies DNA plasmids anti-sense oligonucleotides siRNA Growth factor proteins and peptides Enzymes Drugs and antibiotics

Blood-compatible surfaces; growth factor immobilization Cell adhesion and function in biosensors, arrays, devices and tissue-engineered constructs Biosensors; bioseparations; anti-cancer treatments Gene therapy for a multitude of diseases; DNA probes

Polysaccharides

Anti-cancer treatments; treatments for auto-immune and inflammatory conditions; enhanced wound repair Biosensors; bioreactors; anti-cancer treatments; anti-thrombotic surfaces Anti-thrombotic agents; anti-cancer treatments; anti-hyperplasia treatments; anti-infection/inflammation treatments Non-fouling supports for biosensors and bioseparations

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Physical entrapment methods rely on diffusive barriers or matrix systems to control the transport or availability of the biomolecule. For example, entrapment of enzymes within sol-gels with nano-scale porosites and drug or protein therapeutics within encapsulation matrices provides technologies for enhanced stability, separation or recovery of the biological agent, and regulated delivery kinetics. The encapsulation systems can be engineered to permanently isolate the biomolecule or degrade in non-specific (e.g. hydrolysis) or specific (e.g. enzymatic degradation) fashions for controlled release kinetics. An extensive and diverse group of strategies has been developed to covalently immobilize or tether biomolecules to soluble or solid supports (Figure 37.4) (Weetall, 1976; Hoffman and Hubbell, 2004). Soluble polymers functionalized with biomolecules can then be polymerized into a network or grafted onto a solid support. These strategies rely on coupling reactions between groups in the biomolecule (9NH2, 9COOH, 9SH) and the biomaterial support, and often involve cross-linkers or coupling agents such as CNBr, carbodiimides, and N-hydroxysulfosuccinimide. In many instances, the biomolecule is covalently immobilized via a spacer arm

Physical adsorption

immobilization via high affinity interaction (e.g. antibody–antigen)

spontaneous adsorption

Physical entrapment

encapsulation

dispersion in matrix

Covalent immobilization coupling agent tether arm direct tethering to support network formation

grafting conjugation to monomer followed by polymerization

network formation

+

+

grafting

tethering to pre-formed polymer

Figure 37.4 Schematic diagram of methods for immobilizing biomolecules onto and within biomaterials.

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(e.g. polyethylene glycol) that provides increased steric freedom and activity. Additionally, the tether arm can be designed to be hydrolytically or enzymatically labile in order to allow for release of the tethered biomolecule. As expected, the properties of the underlying biomaterial support play central roles in the tethering efficiency and resulting biological activity of the immobilized biomolecule. In some cases, the surface needs to be modified via the techniques described above to introduce reactive groups for the subsequent immobilization step. For example, inert surfaces can be modified by overcoating with a polymeric adlayer that then presents anchoring groups suitable for immobilization of biomolecules. For many biomedical and biotechnological applications, it is desirable to tether biomolecules within a protein adsorption-resistant (non-fouling) background in order to eliminate effects associated with non-specific protein adsorption. This is particularly important in biomaterials and regenerative medicine applications in which inflammatory responses to non-specifically adsorbed proteins limit biological performance. Poly(ethelyne glycol) (PEG) (9[CH2CH2O]n) groups have proven to be the most protein-resistant functionality and remain the standard (Hoffman, 1999). A strong correlation exists between PEG chain density and length and resistance to protein adsorption, and consequently cell adhesion. Other hydrophilic polymers, such as poly(2-hydroxyethyl methacrylate), polyacrylamide, and phosphoryl choline polymers, also resist protein adsorption. In addition, mannitol, oligomaltose, and taurine groups have emerged as promising moieties to prevent protein adsorption.

SURFACE CHEMICAL PATTERNING While the surface chemical and biological modification strategies described above were presented in the context of a uniform surface, many of these technologies can be used to generate surfaces that present chemical or biological functionalities in distinct geometrical patterns. Important applications of patterned surfaces include protein and oligonucleotide arrays, biosensors, and cell-based arrays (Hubbell, 2004). In many instances, these patterned substrates contain spatially defined domains presenting biomolecules surrounded by a non-fouling background. Photolithography and other techniques relying on exposure through masked patterns or direct surface exposure (e.g. laser or electron beam) in combination with chemical reaction or grafting are often used to generate chemically patterned surfaces. Recently, “soft” lithography methods such as microcontact printing and microfluidic fluid exposure have been applied to produce micropatterned substrates in high-throughput, low cost, and without the need of a cleanroom environment (Whitesides et al., 2001). CONCLUSION AND FUTURE PROSPECTS Surface modifications of biomaterials represent promising routes to engineer biofunctionality at the material–tissue interface in order to modulate biological responses without altering material bulk properties. Countless technologies have been developed to create (i) physicochemical modifications involving alterations to the chemical groups on the surface and (ii) coatings consisting of a different material from the underlying support, including immobilized biomolecules. These approaches hold tremendous promise to enhance biomaterial performance in regenerative medicine. Future structure–function analyses on the effects of specific surface chemistries, topographies, and biological modifications on in vivo responses in particular healing and regenerative environments will further advance the understanding of host responses to implanted devices. These insights will result in the identification of surface modifications that synergize with biological elements (e.g. cells, growth, and differentiation factors) to enhance tissue repair and regeneration. It is anticipated that technical breakthroughs in synthetic chemistry, biofunctionalization, micro- and nano-fabrication, and surface characterization will lead to the engineering of advanced, bioactive materials. In particular, complex patterns of bioligand presentation, such as clusters, gradients, temporal exposure, and multiple ligands, are expected to provide unparalleled control over cellular activities and healing responses.

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REFERENCES Albrektsson, T., Branemark, P.I., Hansson, H.A., Kasemo, B., Larsson, K., Lundstorm, I., McQueen, D.H. and Skalak, R. (1983). The interface zone of inorganic implants in vivo: titanium implants in bone. Ann. Biomed. Eng. 11: 1–27. Anderson, J.M. (2001). Biological responses to materials. Annu. Rev. Mater. Res. 31: 81–110. Bambauer, R., Latza, R., Bambauer, S. and Tobin, E. (2004). Large bore catheters with surface treatments versus untreated catheters for vascular access in hemodialysis. Artif. Organs 28: 604–610. Boyan, B.D., Hummert, T.W., Dean, D.D. and Schwartz, Z. (1996). Role of material surfaces in regulating bone and cartilage cell response. Biomaterials 17: 137–146. Brauker, J.H., Carr-Brendel, V.E., Martinson, L.A., Crudele, J., Johnston, W.D. and Johnson, R.C. (1995). Neovascularization of synthetic membranes directed by membrane microarchitecture. J. Biomed. Mater. Res. 29: 1517–1524. Charest, J.L., Bryant, L.E., Garcia, A.J. and King, W.P. (2004). Hot embossing for micropatterned cell substrates. Biomaterials 25: 4767–4775. Chehroudi, B. and Brunette, D.M. (2002). Subcutaneous microfabricated surfaces inhibit epithelial recession and promote long-term survival of percutaneous implants. Biomaterials 23: 229–237. Chehroudi, B., Gould, T.R. and Brunette, D.M. (1989). Effects of a grooved titanium-coated implant surface on epithelial cell behavior in vitro and in vivo. J. Biomed. Mater. Res. 23: 1067–1085. Curtis, A.S. and Wilkinson, C.D. (1998). Reactions of cells to topography. J. Biomater. Sci. Polymer. Ed. 9: 1313–1329. Flemming, R.G., Murphy, C.J., Abrams, G.A., Goodman, S.L. and Nealey, P.F. (1999). Effects of synthetic micro- and nano-structured surfaces on cell behavior. Biomaterials 20: 573–588. Gruner, H. (2005). Thermal spray coating on titanium. In: Brunette, D.M., Tengvall, P., Textor, M. and Thomsen, P. (eds.), “Titanium in Medicine.” Berlin: Springer-Verlag, pp. 375–416. Hoffman, A.S. (1988). Biomedical applications of plasma gas discharge processes. J. Appl. Polymer Sci. Appl. Polymer Symp. 42: 251–267. Hoffman, A.S. (1999). Non-fouling surface technologies. J. Biomater. Sci. Polymer Ed. 10: 1011–1014. Hoffman, A.S. and Hubbell, J.A. (2004). Surface-immobilized biomolecules. In: Ratner, B.D., Hoffman, A.S., Schoen, F.J. and Lemons, J.E. (eds.), “Biomaterials Science: An Introduction to Materials in Medicine”. San Diego: Academic Press, pp 225–233. Hubbell, J.A. (2004). Biomaterials science and high-throughput screening. Nat. Biotechnol. 22: 828–829. Jansen, J.A. and von Recum, A.F. (2004). Textured and porous materials. In: Ratner, B.D., Hoffman, A.S., Schoen, F.J. and Lemons, J.E. (eds.),“Biomaterials Science: An Introduction to Materials in Medicine.”San Diego: Academic Press, pp. 218–225. Lutolf, M.P. and Hubbell, J.A. (2005). Synthetic biomaterials as instructive extracellular microenvironments for morphogenesis in tissue engineering. Nat. Biotechnol. 23: 47–55. Mrksich, M. and Whitesides, G.M. (1995). Patterning self-assembled monolayers using microcontact printing: a new technology for biosensors? Trends in Biotechnology 13: 228–235. Pilliar, R.M. (2005). Cementless implant fixation – toward improved reliability. Orthop. Clin. N. Am. 36:, 113–119. Ratner, B.D. and Hoffman, A.S. (2004). Physicochemical surface modification of materials used in medicine. In: Ratner, B.D., Hoffman, A.S., Schoen, F.J. and Lemons, J.E. (eds.), “Biomaterials Science: An Introduction to Materials in Medicine.” San Diego: Academic Press, pp. 201–218. Ratner, B.D., Hoffman, A.S., Schoen, F.J. and Lemons, J.E. (2004). “Biomaterials Science: An Introduction to Materials in Medicine.” San Diego: Elsevier Academic Press. Sioshansi, P. (1987). Surface modification of industrial components by ion implantation. Mater. Sci. Eng. 90: 373–383. Ulman, A. (1991). “An Introduction to Ultrathin Organic Films: From Langmuir-Blodgett to Self-Assembly.” San Diego: Academic Press. Weetall, H.H. (1976). Covalent coupling methods for inorganic support materials. Meth. Enzymol. 44: 134–148. Whitesides, G.M., Ostuni, E., Takayama, S., Jiang, X. and Ingber, D.E. (2001). Soft lithography in biology and biochemistry. Annu. Rev. Biomed. Eng. 3: 335–373. Woodruff, D.P. and Delchar, T.A. (1994). “Modern Techniques of Surface Science.” Cambridge: Cambridge University Press. Yamamoto, M., Tabata, Y., Kawasaki, H., and Ikada, Y. (2000). Promotion of fibrovascular tissue ingrowth into porous sponges by basic fibroblast growth factor. J Mater Sci. Mater Med. 11: 213–218.

38 Cell–Substrate Interactions Aparna Nori, Evelyn K.F. Yim, Sulin Chen, and Kam W. Leong

INTRODUCTION Engineering of tissues in vitro or in vivo in many cases require a scaffold to provide the optimal microenvironment for the seeded cells. There has been a growing trend toward the use of synthetic substrates to mimic the natural, physiological system for the purpose of tissue engineering or regenerative medicine. Cell–substrate interaction is of fundamental importance to studies geared toward designing biomimetic substrates that may replace damaged, vital organs or tissues, or assist in the natural healing processes of the body. This chapter first reviews cell interactions with the extracellular matrix (ECM). This is followed by sections detailing the modification of cell behavior by different aspects of a biomimetic substrate, such as its physical, chemical, and biological properties. The role of surface topography in modulating cell interactions is also discussed. As several studies have underscored the necessity of a three-dimensional (3D) environment to yield physiologically relevant data, a section is dedicated to the effect of dimensionality on cell behavior. The chapter concludes with a discussion on the importance of mechanical stress in tissue development, both at the cellular as well as tissue level. CELL–ECM INTERACTIONS The principal building blocks of organs are cells. These functional units are held together to give rise to structures such as tissues, by a hydrated gel-like material known as the ECM. The ECM provides spatial organization, anchorage, and mechanical strength for different cells within a tissue. In addition, it is also responsible for the control and regulation of cell functions such as adhesion, spreading, proliferation, migration, differentiation, and apoptosis by providing mechanical as well as biochemical stimuli. The functions of the ECM are carried out by its key components such as fibrillar and non-fibrillar glycoproteins, hydrated proteoglycans (insoluble macromolecules), and soluble molecules such as growth factors and cytokines (Lutolf and Hubbell, 2005). The most important fibrous proteins that crosslink the matrix are collagen and elastin, which contribute to the tensile and contractile strength of the tissue, respectively. For instance, elastin bears the recoil after the transient stretch (Rosenbloom et al., 1993) such as contraction of the heart tissue. The non-fibrous proteins include fibronectin (FN), vitronectin, and laminin, which act as anchors that initiate cell binding via cell surface adhesion receptors such as integrins. These proteins also stimulate cell signaling pathways in a bidirectional manner between the cells and the ECM. While FN is ubiquitous, vitronectin plays a greater role in adhesion involving endothelial cells. Laminin is a vital protein secreted by epithelial cells and forms an important constituent of the basal lamina (Kleinman et al., 1985). The integrin family of cell adhesion receptors consists of heterodimeric glycoproteins (comprised of α and β subunits) that show specificity for different cell adhesion proteins (Hynes, 1992) as well as collagen. For example, α1β1 binds collagen whereas α5β1 and αvβ3 have affinity for FN and vitronectin specifically.

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The complete integrin receptor comprises of an extracellular domain that binds the ECM in a cation-dependent manner, and short cytoplasmic tails that lack intrinsic kinase activity (Vuori, 1998; Humphries, 2000; Leitinger et al., 2000; Plow et al., 2000; Schwartz, 2001; Xiong et al., 2001). These proteins mediate cell adhesion via cell surface receptor–ligand binding, leading to the clustering of the integrin receptors and formation of transient cell–ECM contacts known as focal contacts. Clustering (or mechanical force and presence of growth factors) can bring the cytoplasmic segments of the integrins in close proximity, possibly resulting in dimerization or autophosphorylation of tyrosine kinase proteins such as focal adhesion kinase (FAK). Focal contacts stabilize into focal adhesions, which are molecularly complex structures, containing proteins including α-actinin and talin (which connect integrins to the actin cytoskeleton), paxillin and signaling molecules such as FAK (bound to integrins directly, but not to the cytoskeleton). Further, focal adhesions also contain vital molecules such as vinculin, which link adhesion molecules to actin and other adaptor molecules. All of these molecules transmit signals for various regulatory pathways between the cell and the ECM (Geiger et al., 2001). These complexes also serve as the termination points of actin filament bundles known as stress fibers. Recruitment of Src homology 2 (SH2) domain containing proteins such as Src kinase and p130 by phosphorylated FAK to these complexes may induce their subsequent phosphorylation. The phosphorylation in turn activates downstream pathways and alter gene expression that is ultimately translated into specific processes such as migration, proliferation, and differentiation. For instance, activation of small GTPase Rac or Erk and JNK pathways results in cell migration and proliferation respectively (Schwartz, 2001). Glycosaminoglycans such as heparan sulfate, hyaluronic acid, and chondroitin sulfate in the glycosylated proteins are also involved in cell adhesion, cell signaling, and communication. Additionally, growth factors and cytokines such as IL-2, transforming growth factor β (TGF-β), and platelet derived growth factor (PDGF), which may be either immobilized or solubilized in the ECM, are also responsible for cell proliferation and differentiation. Cell interaction with the ECM has been demonstrated in the adhesion of cells to substrates coated with ECM molecules, change of cell shape, migration of cells along a concentration gradient of ECM ligands, and demonstration of differentiated phenotype (development of neurites) in response to the ECM (Ruoslahti and Pierschbacher, 1987).

CELL–SUBSTRATE INTERACTIONS Importance of Substrate Though it may appear that the interactions between the ECM and the cells control cellular functions only, the ECM–cell communication is in reality, bidirectional. The cells can regulate how much ECM is synthesized or the extent to which it is degraded by controlling the amount of ECM degrading enzymes produced (proteases such as matrix metalloproteases, collagenase, or plasmin). This is an integral part of ECM remodeling. Further, the extent and specific function of the ECM varies from tissue to tissue and is governed by the need and function of the tissue itself (e.g. connective tissue versus epithelial cells). As the cells are in constant close contact with the ECM, these two components exert a considerable degree of influence over each other, which is known as “ECM–cell dynamic reciprocity” (Lutolf and Hubbell, 2005). The realm of tissue engineering ranges from basic studies elucidating the mechanisms underlying cell behavior to applications for the purpose of tissue regeneration or organ replacement. This involves the use of biomaterials that can mimic the natural environment spatially and temporally so as to facilitate cell adhesion, proliferation, or differentiation according to the tissue-specific requirements. Within the context of cell–ECM interactions, the main aspects that require consideration in the design of a potential biomaterial are discussed below. The influence of these factors on cell behavior is further illustrated with pertinent examples from the literature.

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Effect of Physical Properties Although metals and ceramics are important biomaterials, polymers are much more commonly used in tissue engineering and regenerative medicine applications. The focus of this chapter is therefore on biomedical polymers. Understanding cell interactions with polymers is important for designing substrata for in vitro cell culture or in vivo implantation. As cells are constantly interacting with the extracellular environment, they are sensitive to changes in the surface and bulk properties. In addition, for in vivo application, the mechanical property of the substrate or scaffold may need to match the mechanical requirement of the implantation site so as not to adversely affect the biomechanical stimulus provided to the seeded cells in situ. Physiochemical properties such as crystallinity, morphology, and stiffness/compliance of materials affect cell attachment and cellular behavior. Crystallinity Crystallinity in a polymer represents a state in which a periodic and repeating atomic arrangement is achieved by molecular chain alignment (Callister, 1997). Due to their size and complexity, polymer molecules are often only partially crystalline, having crystalline regions dispersed within the amorphous material. The degree of crystallinity of a polymer depends on the rate of cooling during solidification as well as on chain configuration. Upon cooling through the melting temperature, sufficient time must be allowed during crystallization for the chains to move and align. Crystallization is favored in polymers with a chemically simple structure. Some of the common crystalline polymers used in biomedical applications include polyethylene, polypropylene, polytetrafluoroethylene, poly(vinyl chloride), poly(glycolide) (PGA), poly(L-lactide) (PLLA), and poly(-caprolactone) (PCL). Crystalline polymers are usually stronger and more resistant to dissolution and softening by heat. Crystallinity not only affects the mechanical properties such as strength and fatigue resistance of the polymer, but also plays an important role in determination of the surface physiochemical properties including surface free energy, chemical states, polarity, surface roughness, and wettability, which influence cellular response. When blood compatibility was tested on polypropylene surfaces with different crystalline states, an increase in platelet adhesiveness was observed with decreasing surface crystallinity and interlamellar spacing (Kawamoto et al., 1997). A decrease in interlamellar spacing resulted in enhancing albumin adsorption and diminishing fibrinogen adsorption. Therefore, controlling the crystalline–amorphous microstructure at the surface layer may improve the blood compatibility of polypropylene surfaces. When designing scaffolds for implantation, crystallinity can influence the biodegradability and cellular responses of the scaffold. Crystalline region is more resistant to water penetration and hence retards biodegradation. The adhesion, proliferation, and morphology of human articular cartilage chondrocytes were different when cultured on various degradable polymers with various crystallinity, such as PGA, PLLA, poly (D,L-lactide) (PDLA), different ratios of poly(D,L-lactide-co-glycolide)s, PCL, poly(glycolide-co-trimethylene carbonate) (PCTMC), and poly(dioxanon) (PDO) films (Ishaug-Riley et al., 1999). Significantly higher number of chondrocytes are attached to PGA and 67:33 PCTMC polymer films than on PCL and PLLA films. The total cell numbers and hence fold expansion were significantly higher than the fold expansion on tissue culture polystyrene (TCPS), although the greater fold expansion may be attributed to the lower initial cell attachment. Park and Griffith performed a study of spheroid formation by hepatocytes and proliferation of fibroblasts on PLLA substrates (Park and Cima, 1996). Their results suggest that cells proliferate more slowly on crystalline versus amorphous PLLA and faster spheroid formation on crystalline substrates. This highlights the interesting dynamics between cell–substrate and cell–cell interactions in dictating cell aggregation. Mikos et al. investigated tissue in growth through porous scaffolds composed of semicrystalline or amorphous PLLA that were implanted in rat mesentery (Mikos et al., 1993). There was a two-fold reduction in percent tissue in-growth through the crystalline scaffolds after 10 days as compared to the amorphous scaffolds.

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Variations in crystallinity can also lead to changes in surface roughness on nanometer length scales (Washburn et al., 2004). MC3T3-E1 osteoblastic cells show a higher rate of proliferation on smooth region with a monotonic variation in rate as a function of roughness. Morphology Morphology of the substrate can affect cell attachment by influencing the ability of the substratum to adsorb protein and/or by altering the conformation of the adsorbed protein. ECM proteins are present in serum which is used in most cell cultures. Cell attachment to the substratum is almost always mediated by these ECM proteins adsorbed on the culture surface. Rough and porous surfaces are routinely used in clinical applications such as orthopedic, dental, and cardiovascular prosthesis (Clark et al., 1974; Haddad et al., 1987; Chehroudi et al., 1990; Singhvi et al., 1994). For example, numerous studies have suggested that implants with a porous surface can form better tissue-implant seals to enhance tissue integration (Haddad et al., 1987). Roughness has been shown to alter adhesiveness of platelets to hydrophobic and hydrophilic surfaces (Zingg et al., 1982). The details of surface topography and surface chemistry will be discussed in later sections. Stiffness and Compliance Stiffness of a material is measured by modulus of elasticity or Young’s modulus, while compliance is the inverse of stiffness. Sufficient substrate stiffness is important for anchorage-dependent cells, which often rely on finite resistance to cell-generated forces in order to induce outside-in mechanical signals. Such signals feed back into cell tension (Wang et al., 2002), cell adhesion (Choquet et al., 1997), protein expression and cytoskeletal organization (Cukierman et al., 2001), and cell viability (Wang et al., 2000). Stiffness and compliance encountered during cell–cell adhesion and cell–substrate adhesion are important interactions that modulate intracellular signaling pathways and cellular events from gene expression to cell locomotion. When NRK epithelial cells and 3T3 fibroblasts were cultured on collagen I substrates with varying Young’s modulus, they exhibited different motility and cytoskeletal organization (Pelham and Wang, 1998). Both cells were well spread on rigid substrates. NRK cells became less well spread and irregular shaped, while 3T3 cells lost most of their stress fibers with an increase in locomotion rate when they were cultured on increasingly compliant substrates. Cell movement can also be guided by the manipulation of flexible substrates to produce mechanical strains in polarized cells. When NIH 3T3 fibroblasts were cultured on flexible polyacrylamide sheets with type I collagen coating and transition in rigidity (Lo et al., 2000), cells approaching the transition region from the soft side could easily migrate across the boundary, while cells migrating from the stiff side turned around or retracted as they reached the boundary. Cell also spread to a greater extent on stiff substrates compared with more compliant counterparts (Engler et al., 2004). Contractile myocytes sense the mechanical as well as molecular microenvironment. Myoblast culture has been studied on collagen strips attached to glass or polymer gels of varied elasticity (Engler et al., 2004). Myosin/actin striation emerges only on gels with stiffness typical of normal muscle. Adhesion strength increases monotonically with increasing substrate stiffness. Effect of Chemical Properties The chemical properties of a polymer play an important role in its surface functionality and consequently, cell behavior. When cells are exposed to a polymeric surface, a layer of protein adsorbs onto the surface within milliseconds. Thus, cells “see” the adsorbed protein layer rather than the actual polymer surface. The surface chemistry of a polymer may be fine-tuned to control protein adsorption, which in turn controls cell adhesion. Depending on the desired outcome, the surface chemistry of a polymer can be modified to

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modulate the interactions of the adherent cells, such as morphology, migration, differentiation, proliferation, and apoptosis. In the context of biointeractions, the important chemical properties of a polymeric surface may be categorized into surface wettability and charge. Surface Wettability The wettability of a polymer surface is a measure of its hydrophobicity or hydrophilicity, or its surface energy. Water molecules at a polymeric surface rearrange around proteins, causing the native protein to unfold and adsorb irreversibly to the surface. Water molecules are unable to form hydrogen bonds with hydrophobic substrates. Hence, they form hydrogen bonds within themselves leading to a more ordered structure with lower entropy. Proteins present in the serum can act as surfactants where their hydrophobic domains interact with the substrate, and their hydrophilic domains form hydrogen bonds with the water molecules. This results in the release of the ordered water molecules which is energetically favorable due to the increase in entropy. This is known as the hydrophobic effect (Tanford, 1978). In general, proteins preferentially adsorb onto a hydrophobic surface, as mediated by their hydrophobic domains. The adsorbed protein monolayer is seen by the cells instead of the underlying surface, modulating adhesion to a great extent. For example, fibrinogen adsorbed onto a polymer surface greatly increases platelet adhesion. Functionalization of polymer surfaces with poly(ethylene oxide) (PEO) creates a hydrophilic surface that becomes easily hydrated. Currently, PEO is the gold standard for creating a hydrated hydrophilic surface and is commonly employed to reduce uncontrolled protein adsorption or biofouling onto an implant device. Surface Charge The surface charge of a polymer affects protein adsorption and unfolding on its surface. Unlike surface wettability, the driving force for protein unfolding onto a charged surface is ionic interactions, and not hydrophobic interactions. Protein unfolding depends on the net charge that proteins and cells encounter on the surface. For example, PEO is hydrophilic but has a net neutral charge. In contrast, NH2 and COOH groups become ionized in solution giving rise to a net positive and negative charge, respectively. Many proteins have a net negative surface charge, which promotes their adsorption to a positively charged surface. In addition, the glycocalyx on a cell surface (the polysaccharide mucosal layer) has a largely negative charge, adhering to positively charged surfaces via non-specific interactions. Cellular Response The effect of surface chemistry on cellular behavior begins at the point of interaction. The surface chemistry influences the pattern of protein immobilization, absorptive or ionic, on the surface. For example, polymers with higher hydrophobicity are demonstrated to promote greater osteogenesis (bone regeneration) in vivo (Jansen et al., 2005). This effect has been attributed to a more favorable balance between hydrophobic and hydrophilic properties, which promotes greater protein adsorption onto its surfaces as well as enhanced cell adhesion. Hydrophilic surfaces appear to inhibit leukocyte adhesion and the attached cells exhibit a decreased cytokine response. This results in an attenuated inflammatory reaction and decreased macrophage fusion (Brodbeck et al., 2003). Thus, hydrophilic polymer surfaces may offer an approach for limiting leukocyte adhesion and consequently improving the biocompatibility of an implant. Surface chemistries also regulate the conformational changes in binding domains of FN, directing integrin binding affinity and specificity of cell adhesion. In turn, this provides a greater degree of control over cellular behavior. For example, the α5β1 integrin was shown to bind with greatest affinity to hydrophilic, non-charged surfaces (OH9 functionalized end groups), intermediate affinity to hydrophilic, charged surfaces (NH29 and COOH9 end groups), and least affinity to hydrophobic surfaces (CH39 end groups).

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In contrast, αVβ3 integrins bind with highest affinity to COOH surfaces, intermediate affinity to NH2, and negligible affinity to OH and CH3 modified surfaces. These differences in binding and adhesion were reflected in the varying degrees of mineral matrix deposition and osteoblast differentiation of MC3T3-E1 cells (Keselowsky et al., 2004). Methods of Altering Surface Chemistry The modifications of surface chemistry of a biomaterial allow for the selective treatment of the superficial layer without changing its bulk property. This is achieved mainly through coating of a top layer or by plasma treatment. Coating/deposition of a top layer includes several methods. Solvent coating or casting is a method where a polymer is dissolved in a solvent (usually an organic solvent), which in turn is then soaked, brushed, or sprayed onto a surface. Polyelectrolyte multilayers (PEMs) (Dubas and Schlenoff, 1999) are generated by the deposition of alternating layers of polycationic and polyanionic monolayers onto a surface. Self-assembled monolayers (SAMs) consist of chemisorbed monolayers of closely packed alkanethiols onto surfaces like gold, silver, or mercury. The head groups with a hydrocarbon tail may be functionalized with different end groups such as hydrophobic CH3, hydrophilic OH, or charged COO end groups (Whitesides et al., 2005). Another method of altering polymer chemistry is via plasma treatment of the surfaces of interest. Plasma treatment creates ionized gases such as ions, free radicals, and electrons, from electron and ion impact in an electric field. These ionized partices create oxidized and groups on the polymer surface from the breaking of chemical bonds. In surface etching, inert gases such as argon are employed to remove impurities and increase surface roughness. In addition, plasma etching also allows the alteration of surface reactivity by crosslinking polymer chains. Effect of Biological Properties While the physical and chemical properties of the biomaterial play an integral role in modulating cell behavior, biological features may be equally, if not more, important as they represent the natural ECM. Thus, bioactive molecules such as adhesion ligands, growth factors, or enzymes may be physically entrapped, surface-immobilized, or covalently conjugated onto the substrate to resemble the ECM and control cell behavior. Commonly Used Ligands Bioactive substrates can be produced by the surface immobilization of a vast variety of ECM cell adhesion proteins such as FN, vitronectin, laminin, and collagen. With the advent of molecular biology tools, the amino acid sequences of the cell-binding sites of these proteins have been identified. The RGD tripeptide is a commonly occurring motif present in several cell adhesion glycoproteins and mediates binding to specific members of the integrin family (FN and vitronectin bind via RGD to α5β1 and αvβ3 integrins respectively) (Pierschbacher and Ruoslahti, 1984). The YGISR peptide derived from laminin binds to a family of non-integrin cell adhesion receptors and can elicit cell adhesion and motility (Graf et al., 1987). Substrates coated with short, cell adhesion peptides are being increasingly employed to develop biomimetic substrates. In addition, cell surface proteoglycans have been known to bind to proteins containing a large number of cationic amino acids. This finding has led to the generation of substrates modified with a positively charged surface. Development of Bioactive Surfaces Based on naturally occurring hydrophibic effect of proteins in solution, several strategies have been developed to immobilize either naturally occurring ECM proteins or ECM-derived peptides (cell adhesion ligands) by

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directly adsorbing them onto a substrate. Covalent conjugates of non-adhesive bovine serum albumin (BSA) with RGD peptide adsorbed to tissue culture plates via BSA adsorption to the surface have supported cell adhesion (Danilov and Juliano, 1989). The disadvantages of such methods include easy desorption of the bioactive molecules by other proteins such as antibodies that may be routinely employed in assay procedures. More importantly, protein adsorption may cause burial of the active site of the adhesion ligand, thus rendering it inaccessible for cell binding. Furthermore, protein adsorption may constrain the peptide in a conformation that reduces its binding affinity (Massia et al., 1993). To overcome these problems, new methods for protein/peptide immobilization have emerged where peptides are covalently grafted onto surfaces that do not support cell adhesion. The inert nature of these substrates ensures that any cell adhesion observed can be solely attributed to the bound peptide. For instance, human foreskin fibroblasts seeded on glycophase glass-modified substrates exhibit increasingly spread out morphologies with higher ligand concentrations. Higher ligand concentrations also generate more focal contacts and stress fibers. This technique allows precise control and quantification of ligand density. Thus, the researchers were able to conclude that an RGD spacing of 440 and 140 nm is sufficient to promote cell spreading and cell adhesion respectively (Massia and Hubbell, 1991). Glycophase glass-modified substrates specific for laminin demonstrated both cell attachment and spreading of different cell types via non-integrin cell adhesion receptors. In contrast, surfaces containing adsorbed laminin supported only cell attachment (Massia et al., 1993). Another approach to generate non-adhesive substrates derivatized with cell specific ligands is to employ the non-adhesive polymer poly(ethylene glycol) (PEG). Graft copolymers of RGD modified-PEG and poly-L-lysine were found to undergo surface immobilization onto negatively charged surfaces via electrostatic interactions. These surfaces supported the binding and spreading of human dermal fibroblasts (VandeVondele et al., 2003). Alkanethiolates bind to gold and silver monolayer surfaces and form SAMs. Alkanethiolates modified terminally with the biologically inactive oligo(ethylene glycol) (OEG) have been extensively utilized in studying the modulation of cell behavior by substrates based on mixed SAMs. Mixed SAMs consisting of alkanethiolates modified terminally with either RGD or OEG promoted cell attachment via the cell adhesive RGD or cell repulsion by virtue of the OEG regions (Roberts et al., 1998). Cell–Bioactive Surface Interactions The property of the underlying ECM and its corresponding ligands determine cell behavior and response. Thus, the desired application (to evoke cell adhesion and proliferation versus cell migration and differentiation) may be tailored by choosing and designing the appropriate ligand, allowing an engineered cell response. Thus, substrates can be tailored to evoke specific cell responses which in turn will affect the substrate itself, similar to the cell–ECM adaptive behavior inherent in vivo. The effects of some ligands on various aspects of cell behavior are presented below. Cell Adhesion

Cell adhesion and its natural outcome of cell spreading are one of the first interactions that occur between the cell and the ECM. Cells adhere to the underlying ECM via cell–substratum bonds, which are typically receptor–ligand complexes formed between adhesion receptors (such as integrins) and their ligands (such as FN). Features of the substratum such as ligand density have been shown to affect cell adhesion and spreading via intracellular mechanisms that have not been completely delineated. Hepatocytes cultured on surfaces expressing interstitial ligands (laminin, collagen type I) or basement membrane (FN, collagen type IV) demonstrate a greater rate and degree of spreading at higher ligand densities as compared to lower coating densities. This direct dependence on ligand density was attributed to the increased number and density of cell adhesion receptor–ligand bonds which may generate forces that

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overcome cellular traction (Ingber, 1997). Alternately, formation of focal contacts could be enhanced, leading to reorganization of the cytoskeleton. Initial cell adhesion was found to occur with an increase in actin microfilamant (MF) mass, and concurrent rapid cell spreading phase. This is later followed by a slow rate of ECM-independent spreading. It was hypothesized that the MFs may have formed focal contacts with the ECM which assisted in providing force for cell spreading. The fact that only a combination of cytoskeleton disrupting drugs (against actin and microtubules) could inhibit cell spreading implies cytoskeletal redundancy, that is, one structural component can bear the supporting role of the other. Further, these intracellular forces generated may be transmitted to the ECM as well (Mooney et al., 1995). Cell Motility

Migration of cells is an important aspect of cell behavior, especially during times of development, regeneration of organs, and other processes of repair such as wound healing and angiogenesis. One of the main requirements to promote cell migration is the breaking of existing cell–ECM bonds and the formation of new ones in the desired direction and at the next site of attachment on the substratum. Studies on factors affecting cell motility such as ligand concentration have shown that the migration speed of human mesenchymal stem cells on FN or collagen IV coated surfaces has a biphasic dependence on ligand concentration. The maximal speed is attained at intermediate ligand density. This suggests that very low densities do not afford the traction needed for movement whereas extremely high densities confer strong cell adhesiveness that deters cell detachment from the substrate. As the shear forces for cell detachment for both FN and collagen are similar, it was also concluded that the strength or adhesiveness of the initial cell–substratum bond governed cell migration speeds (DiMilla et al., 1993). The initial bonds formed between the cell and ECM can also be affected by the affinity of the receptor to the ligand as well as the expression levels of the receptor itself. In one study, cell migration speed was inversely dependent on the expression and affinity of the integrin receptor. Maximal cell speed was achieved with intermediate ligand densities, integrin levels, and affinity. All these parameters represent an optimal initial adhesiveness or an adequate number of cell–substratum bonds to support initial attachment followed by migration. At conditions where very few or too many cell–substratum bonds were formed, cells were observed to form short unstable lamellipodia or were too spread out to support movement respectively. Interestingly, the maximal speed attained was found to be independent of ligand concentration, receptor expression levels, or affinity (Palecek et al., 1997). This suggests the involvement of other factors such as intracellular contractile force or the induction of intracellular signaling. As integrins have been shown to cluster and mediate downstream cytoplasmic signaling, the spatial arrangement of the ligand also influences motility. Murine fibroblasts were grown on surfaces that presented the YGRGD ligand either singly or in clusters while maintaining the same average ligand density for the entire surface. Cells showed higher adhesion strength and migration speeds when the ligand was presented in increasing cluster numbers as compared to single ligands. Furthermore, the maximum speed attained by the clustered ligands was achieved with lower overall ligand densities compared to the higher ligand densities required when ligands were expressed singly. Lastly, the presence of stress fibers and distinct focal adhesions in cells exposed to the clustered ligand confirmed the importance of the orientation of the ligands in cell adhesion and motility (Maheshwari et al., 2000). Cell Proliferation and Differentiation

Cell–ECM interactions can also modulate cell function, although the underlying mechanism remains unclear. Identification of critical parameters such as cell–cell contacts, key proteins, and intracellular tension will allow their incorporation into an artificial ECM which can stimulate desired cell function when necessary.

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Cell shape has been long proved to be a vital regulator of cell function. Bioactive surfaces can be tailored to modulate cell shape by controlling the ligand density or arrangement (to create patterns) on the surface. Hepatocytes cultured at very low initial densities (to minimize cell–cell interactions) on substrates coated with different densities of either FN, collagen type I, laminin, or collagen type IV would exhibit a densitydependent alteration of cell shape. The lowest densities were sufficient to support cell attachment but not promote cell spreading as was evident from the rounded cell morphology. In contrast, higher densities mediated a shift to extensive cell spreading. These changes in cell shape were accompanied by corresponding changes in cell function. Rounded cells observed on low density surfaces exhibited a lower degree of proliferation and bore signs of differentiation such as the increased production of liver specific proteins such as albumin. On the other hand, highly spread out cells on high density surfaces tended toward proliferation rather than differentiation (Mooney et al., 1992). Density of ligands influences the number of cell–ECM bonds formed and consequently cell shape. This in turn may regulate the switch between proliferation and differentiation via several possible mechanisms such as the clustering of the integrin receptors leading to cytoskeletal reorganization, upregulation of downstream signaling pathways, or the distribution of intracellular forces. The effect of cell shape on function was also convincingly demonstrated when hepatocytes were grown on laminin-coated adhesive islands (2–80 μM) surrounded by non-adhesive PEG areas. Primary rat hepatocytes were restricted to these adhesive islands and adopted the underlying square or rectangular island morphology as opposed to cells showing pleiomorphic forms when grown on non-patterned adhesive surfaces. This confinement of cell shape led to increased cell differentiation (albumin secretion) with concomitant reduction in growth. As the ligand density was maintained over the various islands, this regulation of cell function could not be attributed to a lack of cell–ECM contacts but primarily to cell shape and consequently triggered molecular pathways (Singhvi et al., 1994). Effect of Topography Cells in tissues or organs respond to organized spatial and temporal stimuli. In the development of an embryo, the surrounding ECM provides a hierarchical organization of topography that ranges from meso to molecular scales. Topography, coupled with biochemical and physical cues, regulates cellular functions such as migration, adhesion, morphogenesis, differentiation, and apoptosis in a developing embryo (Zagris, 2001). Defined topographical cues not only allow the systematic study of cell–substrate interactions (termed contact guidance), but can also control cellular orientation and morphology which may be extended in turn, to control other cellular responses. Fabrication Techniques To study the effect of topography on cellular behavior, patterning techniques have been developed to create defined substrate topography at the micron scale. With further advancements in patterning techniques, topographical structures may now be fabricated at the nanometer scale over larger areas and with greater ease (Odom et al., 2002). This is important as most in vivo structures are found at the nanometer scale, such as the 40–120 nm collagen fibrils of the basement membrane. Several reviews have been published on the common techniques used to generate topographically modified surfaces to study cellular behavior (Flemming et al., 1999; Curtis, 2004). Many of the techniques used today have been developed from photolithography. One of the first studies using photolithography to fabricate patterned structures involves coating a resist onto a surface and subsequently exposing it to UV light through a patterned mask (Brunette et al., 1983). Thus, a photochemical reaction occurs only at the exposed areas of the resist. These reacted areas of the resist are either retained or dissolved away when soaked in a developer solvent, depending on whether a positive or negative resist is used. Photolithography allows the

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facile fabrication of sub-micron sized topography with great reproducibility over large areas. However, the size of the feature achieved is curtailed by the wavelength of light diffraction. This limitation may be overcome by the process of electron beam lithography (EBL) to fabricate fine features at the sub-micron level. However, this method is both expensive and time-intensive. The patterns obtained by photolithography may be transferred to elastomeric molds such as poly(dimethyl siloxane) (PDMS). This technique known as soft lithography allows structures as small as 30–50 nm to be fabricated (Odom et al., 2002). Soft lithography can also be utilized to “stamp” patterns onto a surface and this is known as microcontact printing (Quist et al., 2005). Polymer demixing, which is based on phase separation, has also been shown to generate nano-scale structures. For instance, polystyrene and poly(4-bromostyrene) were shown to spontaneously demix, producing islands that varied from 13 to 95 nm in height, depending on the ratio of polymer to solvent mixture (Dalby et al., 2002). Though this technique is simple and inexpensive, its main disadvantage is a compromise in precision. Electrospinning (Ma et al., 2005) has recently emerged as a simple, efficient method to produce polymer fibers as scaffolds for cell and tissue engineering. Briefly, electrospinning uses a high voltage field to overcome the surface tension of a polymer solution, to form fibers that are deposited onto a grounded surface. This yields a non-woven mesh, which if collected over a longer time period can form a non-woven 3D scaffold. Alignment of the fibers can be accomplished by spinning the grounded surface at the same speed as the rate of fiber deposition. Cellular Responses to Topographical Cues The complex crosstalk between cell adhesion molecules and the ECM, intracellular, and ECM-generated mechanical forces, and biochemical signaling molecules, elucidates a correlation between cell shape and function (Schwartz and Ginsberg, 2002). Control of cell morphology dictates cell behavior such as growth, differentiation, and survival. Topographical cues can induce changes in cell morphology, thus affecting cellular responses such as proliferation, gene expression, cytokine production, and cellular function. These responses vary depending on cell type and the geometry and size of the topographical features and have been reviewed elsewhere (Yim and Leong, 2005b). The importance of the feature size was illustrated when smooth muscle cells were cultured on PDMS and poly(methyl methacrylate) surfaces presenting a range of different sized grooves. Cells exhibited superior alignment along the grooves of surfaces with the smallest topographical features. Epithelial cells aligned on uniform grooves and ridges showed greater adhesion strength when the groove size was reduced from 4,000 to 400 nm by exposure to fluid shear stress (Karuri et al., 2004). Illustrating the effect of geometry on cellular behavior, it was shown that cells grown on adhesive, patterned islands of varying sizes and shapes conformed to the underlying substrate geometry. Further, cells stretched by the underlying substrate were observed to switch from an apoptotic mode to growth (Chen et al., 1997). The geometry of the underlying substrate was also found to affect fibroblast attachment, where the greatest adhesion occurred at ridges, but diminished on nanosized pits and columns (Curtis et al., 2001). Genetic expression can be influenced by the surface topography. Fibroblasts aligned along 3 μm V-shape grooves demonstrated increased mRNA production of FN as compared to non-aligned controls (Chou et al., 1995). At the nanometer scale, uroepithelial cells cultured on hemispherical pillars or step edges demonstrated a stellate morphology as compared to the spread out morphology adopted when grown on smooth surfaces. The less spread morphology also correlated with a decrease in cytokine production. Interestingly, uroepithelial cells cultured on parallel grooves and ridges showed only differences in morphology while cytokine production remained unaffected (Andersson et al., 2003). Furthermore, smooth muscle cells showed a decreased

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rate of proliferation when cultured on nanoimprinted gratings compared to those cultured on non-patterned surfaces (Yim et al., 2005). Effect of Dimensionality Research aimed at either gaining a fundamental understanding of tissue development or designing a new biomaterial has always employed substrates that mimic the ECM as closely as possible. However, one aspect that has been more difficult to imitate is the 3D nature of the tissue. Here, cells are not only surrounded by other cells but also encompassed by the ECM (between the cells or as basement membrane) and its various components (growth factors, proteins). Importance of 3D Culturing cells in a 3D environment promotes normal cell polarity as opposed to two-dimensional (2D) culture where cells are exposed to different upper and lower microenvironments, thus resulting in an artificial cell polarity (Cukierman et al., 2001). Growth on 3D substrates is not curtailed to a single plane. Hence the surface area over which cells can adhere and cell–cell communication can take place is enhanced. Due to the proximity of the cells to the ECM, the local concentration of important cytokines and enzymes may be greater as compared to that distributed over a cell monolayer. Moreover, the presentation of ECM ligands to spatially oriented cells may afford the simultaneous stimulation of several signal transduction pathways (Cukierman, 2002),leading to a greater control over cell behavior and function. Thus, 3D systems not only provide spatial regulation of the cells but may also affect cellular responses to the physical and biochemical cues provided (Schmeichel and Bissell, 2003). This may in turn lead to the remodeling of the ECM itself, thus modulating the dynamic reciprocity of the cell–ECM system more effectively. Substrates for 3D Culture It has been demonstrated that experiments carried out on a planar, rigid substrate elicit results that may not be comparable to those obtained under in vivo conditions (Cukierman et al., 2002). To overcome this hurdle, sufficiently porous substrates are being increasingly developed that provide spatial freedom to allow both the movement of cells as well as the transport of nutrients. In addition, cell adhesion ligands and growth factors are being incorporated to bestow adhesive and proliferative properties to these substrates to recreate the natural environment. Substrates that serve as 3D environments include polymer scaffolds, hydrogels, porous cellulose beads, and non-woven polyester disks (Yim and Leong, 2005a). This section discusses hydrogels as example of 3D substrates that are being currently developed to mimic the biochemical features of the natural ECM. Migration of cells through the 3D matrix during angiogenesis or wound healing requires its degradation by proteolytic enzymes such as matrix metalloproteases followed by cell adhesion in order to promote ECM remodeling (Friedl and Brocker, 2000). To meet this requirement, hydrogels are prepared by the copolymerization of macromonomers containing the non-adhesive polymer PEG, flanked by oligopeptide sequences that serve as substrates for proteolytic enzymes such as collagenase or plasmin. These oligopeptides crosslinked the hydrogel and are susceptible to specific enzymatic degradation (West and Hubbell, 1999). As these hydrogels lacked sufficient adhesiveness to promote cell traction, pendant cell adhesive ligands such as RGD were introduced. This enzymatically degradable hydrogel now successfully promoted the attachment of ensconced human dermal fibroblasts. Upon stimulation to secrete MMP, these cells were found to selectively degrade the oligopeptide crosslinks and migrate, a phenomenon evident only in these cell-adhesive and responsive hydrogels. In addition, evidence of ECM remodeling was visible in the production of a continuous cellular meshwork with no loss of traction. In vivo, these gels containing vascular endothelial growth

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factor were found to generate new connective tissue with a vascular network (Pratt et al., 2004). Similarly, PEG-based hydrogels crosslinked with peptides sensitive to plasmin and bearing pendant RGD peptides and heparin binding sites for immobilizing the growth factor bone morphogenetic protein-2 were prepared. These cell-responsive hydrogels were found to effectively regenerate bone in rats (Lutolf et al., 2003). Cellular Responses to 3D Substrates Cell Adhesion Fibroblasts grown in 3D cell-derived matrices exhibited increased cell adhesion, migration, rapid stabilization of cell shape (approximating that observed in vivo), and higher proliferation compared to fibroblasts grown on 2D substrates. Cells were found to attach to the 3D matrix via “3D-matrix adhesions” similar to those observed in vivo, but structurally and molecularly distinct from the focal and fibrillar adhesions seen in vitro. These 3D adhesions lacked phosphorylated FAK (responsible for the activation/phosphorylation of proteins involved in different pathways), suggesting the involvement of other pathways such as the MAP kinase pathway (Cukierman et al., 2001). Cell Proliferation and Differentiation

Rat aortic smooth muscle cells showed reduced proliferation and lower expression of the contractile smooth muscle α-actin protein when embedded in collagen type I containing 3D gels as compared to collagen coated 2D substrates. These effects were attributed to gel compaction leading to contact inhibition (Stegemann and Nerem, 2003). The expansion potential of human embryonic germ cell derivatives grown on a fibrous cellulose acetate scaffold was found to be superior to that observed on 2D controls. Further, these cells maintained their potential to differentiate into various lineages (Yim and Leong, 2005a). Cell–ECM Reciprocity

3D culture also governs the dynamic cell–ECM reciprocity as seen in the response of cells to mechanical stress during matrix remodeling. In the case of fibroblasts grown on collagen matrices, this involves an increased density of collagen fibrils. As the fibroblasts exert a mechanical force on the underlying matrix, these collagen fibrils are either aligned or randomly oriented depending on whether the matrix is restrained to the dish (scenario mimicking 2D culture) or free floating, respectively. Alignment of the fibrils in turn affects cell morphology as stimulation with PDGF yields fibroblasts with a stellate morphology and isometric tension. In contrast, fibroblasts grown on the floating matrices (lack of tension in the matrix) adopt neuronal extensions and a dendritic network (Grinnell, 2003). Effect of Mechanical Stimuli Mechanical stimuli are particularly important in several areas of tissue engineering, such as cardiovascular grafts, bone, cartilage, and tendon/ligament engineering. The polymer substrate should possess physical properties that match the mechanical properties of the implant site. It should also have the ability to support the mechanical force exerted on the implanted graft at the site, such as pulsating blood flow though arteries and weight-bearing bone grafts. Different types of mechanical stresses are experienced by various tissues. In vitro systems have been developed to model the effect of mechanical stimuli, such as compressive stress on chondrocytes (Wong et al., 1997) or fluid shear stress on endothelial cells (Davies et al., 1997). One type of experimental setup involves seeding cells on a flexible substrate such as an elastic membrane, and to apply defined, stepwise, or cyclic strain to the substrate (Leung et al., 1976; Sadoshima and Izumo, 1997; Chiquet, 1999).

Compression Shear stress Tension ECM/polymer scaffold

Changes in cellular response

Intracellular changes

Mechanical sensor

Mechanical stimuli

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Mechanical sensing receptor (e.g. integrin)

Tyrosine phosphorylation Intracellular secondary mediators (e.g. FAK, MAPK, Rho) Ion channel modulation (e.g. Ca2+,K+) Intracellular calcium modulation

Actin polymerization/ depolymerization Integrin activation, clustering

Cell differentiation Extracellular matrix modification Phenotypic change Changes in physiological function (e.g. myogenic response)

Cytoskeleton reorganization Nuclear elongation MTOC polarization

Gene expression (e.g. c-fos, c-jun) DNA synthesis Protein synthesis

Cell orientation Tissue development (e.g. angiogenesis) Cell migration, proliferation, apoptosis, cytoskeleton remodeling

Figure 38.1 Effects of mechanical stress on cellular behavior.

Cells reside in a dynamic environment in the body and are sensitive to changes in the microenvironment. Mechanical forces applied on the cell–polymer construct will often change the cellular response, thus rendering cell–polymer interactions in the presence of mechanical stimuli an important area of study (Figure 38.1). Mechanotransduction When a force is applied to cells growing on substrates, the cells sense the changes in the physical environment and transmit the mechanical signal to intracellular biochemical signals via signal transduction. This mechanism is called mechanotransduction. One of the cellular mechanosensors for mechanotransduction is the integrin class of adhesion receptors (Martinez-Lemus et al., 2005). As integrins physically link the ECM to the cytoskeleton, they allow for a direct mechanical connection between the two. Hence, they are responsible for establishing a mechanical continuum by which forces are transmitted between the outside and the inside of cells in a bidirectional manner (Ingber et al., 1994). Application of mechanical stress to integrin adhesion sites in a variety of cells induces numerous cellular responses. Cellular Responses in Modifying ECM The ECM forms the underlying substrate for cell adhesion, growth, differentiation, and mechanical support. It is known that connective tissue cells adapt their ECM to changes in mechanical load, such as in bone remodeling

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or wound healing. In turn, changes in the ECM content can influence the performance of the tissue, such as the stiffness of heart and vasculature and the strength of bone and cartilage. Experimental evidence confirms that specific ECM proteins are regulated by mechanical stimuli in vivo. For example, tenascin-C and collagen XII are two ECM components associated with fibrillar collagen in tissues bearing high tensile stress such as tendons, ligament, periosteum, arterial smooth muscle, and heart valves (Chiquet and Fambrough, 1984; Walchli et al., 1994; Koch et al., 1995). In cardiac tissue, fibroblasts are the principal cell type responsible for secreting components of the ECM. In vitro studies using rat cardiac fibroblasts have shown long-term ECM component changes in response to variations in mechanical load. In response to both cyclic and static uniaxial stretch, an increase in both collagen I and collagen II mRNA expression was obtained (Carver et al., 1991). When the rat cardiac fibroblasts were cyclically stretched for various durations, mitogen-activated protein kinase was most rapidly activated, and collagen I expression became most abundant (Atance et al., 2004). Vascular Grafts Endothelial cells act as primary transducers of local hemodynamic forces into signals that maintain physiological function or initiate pathological processes in vessel walls (Helmke, 2005). When subjected to sustained fluid shear stress, cultured vascular endothelial cells undergo significant morphological changes including elongation and cell alignment in the direction of the applied flow (Suciu et al., 1997). In another study, human aortic endothelial cells were seeded on deformable silicone membranes and subjected to various magnitudes and rates of stretching or compression (Wille et al., 2004). Both stretching and compression resulted in magnitude-dependent orientation responses away from the deforming direction. Compression produced a slower temporal response than stretching. Vascular smooth muscle cells are the major source of ECM protein production within the vessel wall. The accumulation of rigid ECM proteins such as collagen can affect arterial wall stiffness and therefore arterial compliance, pulse wave propagation, and pulse pressure. When chronic cyclical mechanical strain was applied to human vascular smooth muscle cells, the concentration of FN and collagen increased (O’Callaghan and Williams, 2000). An increase in MMP-2 activity showed that ECM accumulation was not due to inhibition of ECM protein degradation. TGF-β1 expression was also higher and this induced TGF-β1 production may be a mechanism for increased vascular ECM deposition in hypertension. Mechanical stimuli can also be harnessed to ameliorate tissue formation. “Functional tissue engineering” involves the growth of tissues that normally experience mechanical loading in vivo (Butler et al., 2000). The application of in vitro physical loading mimics the physiological environment and accelerates the development of tissue constructs that can meet the functional and mechanical demands in vivo. Mechanical preconditioning of tissue-engineered constructs in vitro can also improve its post-transplantation survival and performance. Dynamic mechanical conditioning has been applied to tissue engineering blood vessel constructs composed of smooth muscle cells embedded in collagen–gel constructs. This resulted in improved contraction and mechanical strength (reflected by ultimate stress and material modulus) as compared to statically cultured controls (Seliktar et al., 2000). The dynamic mechanical conditioning also led to an improvement of tissueengineered blood vessel constructs in terms of histological organization, where circumferential orientation was increased. In another study, it was demonstrated that smooth muscle cells and fibroblast-seeded collagen constructs exposed to 10% cyclic strain showed increased MMP-2, elastin and collagen gene expression (Seliktar et al., 2003). Strain-stimulated MMP2-activity can have a favorable impact on the structural development of the constructs, but overexpression of MMPs can have adverse consequences on the structural integrity. Shear stress also plays an important role in enhancing angiogenesis. The effect of shear stress stimulus on 3D microvessel formation has been investigated in vitro (Ueda et al., 2004). Bovine pulmonary microvascular

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endothelial cells were seeded onto collagen gels with basic fibroblast growth factor to model microvessel formation. The model was placed in a parallel-plate flow chamber. Shear stress applied to the surfaces of endothelial cells on the collagen gel promoted microvessel network formation and expansion in the gel, with increased bifurcations and endpoints observed. The role of fluid shear stress in collateral vessel growth has also been investigated in pig and rabbit hind limbs (Pipp et al., 2004). A side-to-side anastomosis was created between the distal stumps of one of the bilaterally occluded femoral arteries with the accompanying vein. The increased fluid shear stress increased capillary density in the lower leg muscles and augmented proliferative activity of endothelial and smooth muscle cells. High levels of fluid shear stress caused a strong arteriogenic response, reinstated cellular proliferation, stimulated cytoskeletal rearrangement, and normalized maximal conductance. Cartilage/Bone Engineering Differentiation of chondrocytes to osteoblastic phenotype occurs during an interim period of bone development, fracture repair, and distraction osteogenesis. Uniaxial strains were applied in a rabbit model of mandibular distraction osteogenesis (Meyer et al., 2001). Cell differentiation, apoptosis, and tissue development in the newly formed gap tissue showed a correlation to the magnitude of the applied strain. Specimens exposed to 20,000 microstrain of cyclic loading resulted in a statistically significant formation of cartilage struts with embedded chondrocyte-like cells. Chondrocytes cultured in agarose hydrogels develop a functional ECM. Application of dynamic strain at physiological levels to these constructs over time can increase their mechanical properties. The application of daily, dynamic deformational loading to constructs over a long term period (more than a week) reult in enhanced biochemical content and mechanical properties (Mauck et al., 2002). Tendon/Ligament Engineering Mechanical stimuli have been shown to enhance proliferation of human anterior cruciate ligament and medial collateral ligament seeded on biodegradable polymer fiber scaffolds (Lin et al., 1999). Mechanical stimulation in vitro, without ligament-selective exogenous growth and differentiation factors, induced the differentiation of mesenchymal progenitor cells from the bone marrow into a ligament cell lineage in preference to alternative lineages such as bone or cartilage cells (Altman et al., 2002). The application of mechanical stress yielded features characteristic of ligament cells. These included upregulated ligament fibroblast markers such as collagen types I and III and tenascin-C, statistically significant cell alignment and density and the formation of oriented collagen fibers.

CONCLUSION Cell–substrate interaction is of central importance to many biological processes and has been investigated extensively from various angles. Discussed in the context of tissue engineering and regenerative medicine, this chapter covers the basics and general principles of this tremendously complex phenomenon. The examples cited above do confirm the importance and relevance of elucidating the interactions of cells with substrates (Figure 38.2). Hopefully this review would provide a starting point for the readers to design the optimal substrates for specific tissue development. Many challenges remain for a deeper understanding of the cell–substrate phenomenon. Primary among them is the heterogeneous nature of any cell type. Effort to start with a more homogeneous cell population, preferably at the same cell cycle, should yield more reproducible results. The ability to analyze cells at the single cell level in situ will also provide important insight, for instance, in uncoupling the effects of cell density and cell–cell communication from cell–substrate interaction. This may become feasible in the near future as optical techniques to image gene expression at the single cell level materialize. There is also in general a lack of quantitative approach to the studies. As the quality and quantity of the

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Figure 38.2 Schematic depicting the different properties of a substrate that influence cellular behavior. data improve with the characterization techniques, it may become possible to develop a meaningful theoretical framework to describe and predict these cell–substrate interaction phenomena. Finally, as the field of regenerative medicine continues to be fueled by advances in stem cell biology, studies of cell–substrate interactions, particularly with stem cells, will become increasingly more important and rewarding.

ACKNOWLEDGMENT The authors would like to acknowledge the partial support of NIH (EB003447) to this work.

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39 Histogenesis in ThreeDimensional Scaffolds Nicole M. Bergmann and Jennifer L. West

REPLACING DISEASED TISSUES Regeneration or replacement of diseased tissues and organs is one of the biggest problems facing the medical industry. According to the US Scientific Registry of Transplant Recipients, in 2003, patients on the National Organ Registry numbered over 80,000 and that number is expected to rise drastically over the next 10 years (OPTN/SRTR, 2003). Of the patients awaiting life-saving transplants, approximately 10% died before they could receive donor organs. In addition, tissue disease and organ failure lead to an estimated 8 million surgical procedures annually in the United States (Angelos et al., 2003). From these statistics, one fact is clear – the need for replacement organs far outweighs the supply, and the discrepancy is only expected to increase as populations rise and life expectancies increase (Smith-Brew and Yanai, 1996). Thus, the regeneration of lost tissue (i.e. by tissue engineering) is seen as the solution to the lack of suitable replacement organs. Since synthetic therapies to replace or regenerate damaged tissues are limited, formulating replacement methodologies that allow the patient to self-heal would allow physicians to begin treatment before patients are critically ill leading to a decrease in mortality rates. The idea that tissue function can be restored is as old as the medical profession. Many cultures have been performing successful nose operations for thousands of years (Ang, 2005). One of the most popular sites of implant was the nose which other than the hands was the body part most likely to be injured in battle (Wallace, 1978). In 1596, a pioneering technique for regeneration of nasal tissue was developed that involved connecting a flap of skin and underlying vessels from the arm to the nose. In many cases, tissues were successfully regenerated due to the blood supplied from the body. In the 19th century, many surgeons successfully transplanted skin between individuals (Hauben, 1985; Ang, 2005). The ancestry of modern tissue engineering can be traced to World War II. Due to the high number of battlefield casualties in this war, many surgeons began experimenting with replacing native tissues with artificial materials in order to attempt to reduce battlefield mortality (DeBakey, 1946). Although most of these attempts were unsuccessful for various reasons, the idea of replacing a lost tissue with synthetic materials endured. Romanced by the thought of creating new organs along with the observance of tissue growth around a silk suture, Arthur B. Voorhees from Columbia University surgically replaced canine blood vessels with parachute nylon (Voorhees et al., 1952). Voorhees believed that the material possessed suitable properties that mimicked the properties of blood vessels including an elastic nature. However, this material proved unsuitable as a replacement as thrombi formed in the grafts leading to death. Nonetheless, Voorhees hypothesized that with a better understanding of materials and the human body, organ replacement could be possible. Although most mammals have some form of limited regeneration capacity, the amount of tissue that can be spontaneously formed is small compared to the size of whole organ. Although some types of organ pathologies

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such as diabetes or kidney disease can be treated pharmacologically or mechanically to restore lost tissue function, most organ failures are catastrophic to the host. Drugs cannot be used to treat these disease states because the defect occurs to the whole tissue. Cells, proteins, and other organotypic molecules are lost at the macromolecular level, and no one drug is all-encompassing. Machines can replace lost function of the kidneys for a limited time, but mechanical means cannot replace vital organs such as the heart, liver, stomach, and lungs. In addition, many non-biological artificial replacements can induce chronic inflammation and immune response (Tomazic et al., 1991). Therefore, biological replacements have been sought through the use of tissue engineering. Histogenesis Biological tissues are comprised of three components: cells, the extracellular matrix (ECM), and the signaling systems that are encoded by genes in the nuclei of the cells and are activated through cues from the ECM or from other cells (Shin et al., 2003) (Figure 39.1). Together, the three components interact in balance to form tissues and organs, and it is mimicking these interactions that are the focus of tissue engineering. A greater understanding of these interactions will lead to better biologic materials. The ECM can be widely viewed as the natural scaffolding that supports tissues and organs, but it should not be viewed as merely providing strength and physical support. The ECM is now believed to be intricately involved in the events that lead to the eventual formation of tissue. The ECM is composed of a fibrillar and an amorphous component (Reid and Zern, 1993). These two broad components interact with cells via cell surface receptors and other membrane proteins. Cell–ECM interactions can determine everything from cell differentiation and cell growth to cell orientation and the secretion of other molecules by the cell (Reid and Zern, 1993). For instance, ECM interactions such as stress forces due to injury or disease can cause cytoskeletal rearrangement in cells leading to changes in protein expression and nuclear events. The cell may divide to produce more cells for tissue formation, and the cell may also differentiate into a more specific cell phenotype. Concurrently, ECM cues can lead to a production and release of matrix molecules in a dynamic loop. These matrix molecules form a place for anchorage for newly divided cells which will act with the ECM in this loop to lead to the formation of natural tissue. From an engineering design standpoint, the tissues of the human body are infinitely complex, and the dynamic processes the comprise organs are not confined to just one dimension. At the most basic level, cells attach to the ECM and spread secrete growth factors, proteases, matrix proteins, and other chemicals, and proliferate in response to specific signals and the availability of oxygen and nutrients. Although molecular

Figure 39.1 The three components of mammalian tissue.

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biology and other fields have come up with techniques for characterizing many of these traits, most techniques are limited to two-dimensional cell culture. However, tissues are three-dimensional, and as such, the means of understanding and mimicking these processes artificially is immensely difficult. When a third dimension of tissue is added, cell migration to form distinct tissue layers is often seen. Cell-secreted factors become chemical gradients that can attract other cells in a concentration dependent manner. Thus, in threedimensions, tissues exhibit a high level of heterogeneity that becomes very difficult to both characterize and mimic artificially. Engineers have therefore been attempting to identify the most important factors needed for new tissue formation, and the materials and cells needed to begin regeneration. Regeneration of Diseased Tissues Regeneration is defined as the synthesis of physiologic tissues, and the objective is to restore lost function. This is in direct contrast with repair, which is defined as wound closure (Yannas, 2005a). The purpose of repair is to close a wound and return homeostatic function to a site of injury. Repair results in contraction of the organ and formation of a scar consisting of epithelial tissue that has not differentiated into the proper tissue phenotype. Scar tissue usually exhibits reduced or no physiological function. Therefore, scar formation is seen as an adverse event in tissue engineering which directly inhibits the regeneration of new tissues (Yannas, 2005b). While a fetus has a large capacity for regeneration, the adult human has only limited capability, in the case of some tissues, essentially no regeneration is observed. It is not widely understood what chemical changes occur that suppress regeneration, but studies have shown that the removal of the ECM through disease or injury disrupts normal tissue regeneration, and directly leads to the repair state (Yannas, 2005a). Indeed, studies have shown that cells cannot be placed directly into large tissue defects. Upon implantation, the cells begin attempting to reform the structure from which they were isolated. However, due to the size of the defect, the ability of these cells to form whole structures is limited by a lack of physical and chemical cues found in the ECM (Yannas, 2005b). In the absence of structural support, cells cannot reorganize and differentiate into the higher ordered structures of organs. Three interconnected layers of tissue derive organs – epithelium, basal lamina, and stroma. Epithelium is composed entirely of cells with little or no ECM or vasculature. The basement membrane contains only ECM, while the stroma contains cells, ECM, and blood vessels (Wetzels et al., 1991) (Figure 39.2). Both the epithelia and basal lamina can regenerate without formation of a scar, but the stroma is usually non-regenerative, and injury to this layer directly leads to the repair state (Yannas, 2004). When injury occurs in a tissue the inflammatory response is swift, and the primary cell types that are recruited to the wound bed fibroblasts and myofibroblasts (Yannas, 2004). In normal tissue formation, fibroblasts secrete collagen in a randomly oriented manner in three-dimensions (Yannas, 2005a). When stromal tissue is destroyed, the myofibroblasts will secrete collagen fibrils to attempt to close the wound. Since myofibroblasts are contractile cells, planes of stress are placed on the fibers due to contraction by these cells. The stress forces are usually in one-dimension and cause both the myofibroblasts and fibers to align parallel to this force. Thus, ECM secretion occurs along the plane of stress from tissue contraction (Ng et al., 2005). The highly aligned scar fibrils do not possess the randomly oriented cells and fibers of normal tissue and directly inhibit further infiltration by other cell types. It is believed the main purpose of contraction is to speed healing events by bringing healthy tissue closer together and thus reducing the size of the wound to be healed. Because the absence of an ECM induces contraction and indirectly scar formation, it is believed that directly blocking contraction will help induce regeneration. Indeed many groups have observed that placing a biomaterial scaffold into the site will inhibit contraction and allow regeneration to occur. When Yannas and colleagues placed a collagen scaffold into a full-thickness skin wound, contraction was blocked and scar formation was greatly reduced (Yannas, 2004). In addition, skin was partially regenerated. The collagen construct was

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Epithelia Basal lamina

Stroma

Figure 39.2 The three layers of tissue. The epithelia is the outermost layer of tissue and is composed mainly of cells. The basal lamina is a thin ECM membrane that separates the epithelia from the stroma. The stroma is mainly composed of ECM fibers and usually cannot spontaneously regenerate. shown to greatly reduce the inflammatory response and consequently, the number of myofibroblasts recruited to the wound bed. The suppression of contraction and subsequent regeneration was not shown in defects treated with cytokines, cell suspensions, or scaffolds that induced a high inflammatory response (Yannas, 2005b). In addition, inhibition of contraction by steroids or other chemicals (Ehrlich and Hunt, 1968) will not induce organ regrowth in the absence of a template. From the above studies, it can be concluded that a successful ECM for regeneration must be one that suppresses contraction by eliminating the inflammatory response. However, the scaffold must contain other design parameters that aid in physiologic synthesis of new tissues and organs. Important parameters include cell sources and seeding into scaffolds, microvasculature, scaffold material, porosity, degradation characteristics, and biomolecular design. Design Parameters for Histogenesis Because the presence of an ECM template has been shown to suppress scar formation, tissue engineers have sought to use an artificial scaffold composed of biocompatible, biodegradable components with characteristics similar to the natural ECM of native tissues. This scaffold acts as a temporary replacement for the ECM, and the primary function to serve as a template for cell and protein attachment while acting as a degradable support structure for new tissue ingrowth. Many design strategies have been investigated to mimic the behavior of natural biologic tissue, and in turn drive new tissue development. Cell Sources As has been mentioned previously, a biomaterial scaffold alone is generally not sufficient to induce regeneration of new organs. Cells can infiltrate the scaffold from the surrounding tissue but the distance of invasion is limited to a few microns. Thus, in conjunction with an artificial scaffold, isolated cells have been used to replace lost or damaged tissue function. Cells are grown in vitro, and then seeded onto the ECM construct at a known cell density. The cells are allowed to proliferate on the substrate under in vitro culture conditions, and then the cell-biomaterial hybrid is implanted into the site of the defect. Many different cell lines have been attempted, but generally cells are of three different types: mature, differentiated cells, adult-derived stem cells, and embryonic stem cells (Hedrick and Daniels, 2003). By implanting cells into the biomaterial scaffold, it is believed that the time for tissue infiltration by the host will be minimized and the cells can secrete bioactive factors in vitro that will encourage autologous ECM formation.

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Mature cells have historically come from three sources: autologous, allogeneic, or xenogeneic. Xenogeneic cell transplantations have generally been abandoned as a cell source due to concerns over immune rejection and cross-species disease transmission. Allogeneic cells are harvested from healthy adult donor organs and then expanded in vivo. Scaffolds with allogeneic cells are also subject to immune rejection but this method has seen success in skin regeneration of burn patients (Horch et al., 2005). Autologous cells biopsied from a patient, expanded in vitro, and then seed onto a tissue scaffold are generally viewed as the ideal replacement because of no immune rejection. However, in many disease states and/or tissue wounds, enough healthy cells for a suitable cell line are not present (Hedrick and Daniels, 2003). More recently, researchers have been using progenitor stem cells derived from the patients bloodstream, or have transplanted stem cells from other human sources (embryonic or adult-derived) (Vats et al., 2002). A stem cell is an undifferentiated cell that can produce an identical daughter cell in addition to differentiated cells. Adult-derived stem cells from the bloodstream are normally multipotent and the cell lineages that they can originate are usually restricted to the germ layer of origin. Embryonic stem cells are pluripotent and can differentiate into almost any cell type. Indeed embryonic stem cells from mice have been differentiated in vitro into neural cells, muscle cells, chondrocytes, and others (Vats et al., 2002). Differentiation was initiated in vitro through the use of media containing cytokines and growth factors specific to the cell lineage of interest. One group even has succeeded in causing embryonic stem cell differentiation to osteoblasts by culturing stem cells in media that had formally contained osteoblasts (Buttery et al., 2001). It was found that the culture medium contained growth factors secreted specifically by osteoblasts and it is believed these molecules triggered differentiation into the osteoblast lineage. Size Limitation Due to Diffusion: Importance of Microvasculature One of the biggest limitations to histogenesis of new organs is the availability of nutrients and oxygen available to the cells contained in the biomaterial construct. Most tissue engineered constructs do not contain the intricate vasculature of native tissue, and thus cells contained in scaffolds rely on oxygenation from simple diffusion (Soker et al., 2000). Since diffusion can be limited by the construct, cells in the interior of the scaffold can become anoxic and quickly die. Many experiments have shown that the critical distance that a cell can live from a capillary bed is at best a few hundred microns (Griffith et al., 2005). In an ideal situation, the implant will become vascularized from infiltration and extension of host capillaries. However, the growth and reorganization of tiny blood vessels usually takes much longer than the division of cells already seeded inside the construct (Soker et al., 2000). As a result, the developing vasculature cannot meet the demands of the rapidly increasing cell population. Some groups have successfully implanted a polymer construct close to a capillary bed and induced the construct to become vascularized (Cheng et al., 2005). The vascularized construct can then be removed from the host and seeded with cells for a particular application. Mikos and colleagues have used microporous polylactic-co-glycolic acid (PLGA) sponges implanted by the capillary bed of a sheep. The sponges were surrounded by a chamber that created a tiny bioreactor environment within the sheep. The implant showed new vascular ingrowth throughout the PLGA sponges (Cheng et al., 2005). Another method to encourage vascularization without multiple implant removals is to seed the construct with autologous endothelial cells in the hopes that the cells will merge with the existing vascular to accelerate the vascularization process (Marler et al., 1998). Constructs that have been seeded with endothelial cells and then cultured in pulsed-flow bioreactor systems have also shown promise as means of forming microvascularized constructs (Burg et al., 2000). The microenvironment is believed to better simulate the mechanical environment of native vascular than static culture, which in turn provides appropriate physiomechanical signal to cells. In this simulated environment, endothelial cells secrete growth factors and ECM molecules to form microvessels. Also, the addition of angiogenic growth factors to the matrix, either chemically immobilized or physically entrapped,

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can also greatly encourage vasculature formation. Mooney and colleagues have incorporated basic fibroblast growth factor (bFGF), vascular endothelial growth factor (VEGF), and other angiogenic factors into scaffolds to encourage vasculature from the host (Sun, C. et al., 2005; Sun, Q.H. et al., 2005). In addition, the presence of these growth factors can cause mesenchymal stem cells recruited from the bloodstream to differentiate into endothelial cells and eventually microvasculature. Porosity Tissue ingrowth into biomaterials scaffolds is absolutely necessary for successful histogenesis. In most types of synthetic materials, this requires a porous microstructure. Scaffold porosity, pore size, and the overall pore structure all have important effects upon tissue formation and infiltration into the construct. It also allows diffusion of metabolites, oxygen and growth factors into and out of the material. A porous structure therefore enables cell seeding, attachment, and proliferation while allowing vascularization from the host. Ways of creating pores or increasing the porosity of biomaterials include gas foaming (Montjovent et al., 2005), salt leaching (Gross and Rodriguez-Lorenzo, 2004), freeze drying (Moscato et al., 2006), and electrospinning (He et al., 2005). A fabrication method for a microporous polyurethaneurea has been developed using a gas foaming method. In these studies, sodium bicarbonate is added to a mixture of N,N-dimethylformamide (DMF) containing a polyurethaneurea copolymer. The polymer is then casted in a mold and the DMF evaporated. The cast polymer is placed into an acidic solution to react with the salt. Both the removal of the salt and bubbles formed during this process contribute to the porosity of the scaffold (Jun and West, 2005a). In addition, both the surface and the bulk exhibited the same pore morphology and volume fraction (Figure 39.3). The addition of the cell-specific peptide, YIGSR, also did not change the overall porosity of the scaffold (Figure 39.4). Through experimental studies, there is a characteristic pore diameter that will allow the greatest amount of cell infiltration and attachment. The ability of scaffold to bind cells can be approximated by the following equation: c 

Nc A

,

where Nc is the number of cells bound to the available surface area, A, of the template material (Yannas, 2005a). By careful analysis, it can be shown that two matrices with identical chemistries, but differing pore diameters have vastly different abilities to allow cellular infiltration. Scaffolds containing pores of 300 μm have a much lower surface area to allow cellular attachment than matrices with pore diameters of less than 50 μm and as such,

(a)

(b)

Figure 39.3 Microporous polyurethaneurea copolymer scaffold. Scaffold shows the same morphology on the surface (a) and bulk (b).

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(a)

(b)

Figure 39.4 Microporous polyurethaneurea scaffold containing cell-adhesion peptides. The presence of YIGSR (b) did not change the overall porosity of the scaffold.

a much lower number of cells contained in the matrix. Because of the low cellular infiltration, tissue formation in these constructs is slow and in certain cases inhibited due to lack of interactions between cells (Yannas, 2005a). From this calculation, it can be assumed that there is a maximum pore diameter that will allow maximum cellular infiltration but still allow the greatest number of cells to attach, divide, and maintain the cellular interplay that leads to histogenesis, a number that can be specific to cell type and scaffold material. In addition, the lower pore size is limited by the size of the cell which measures roughly 10 μm, and in general, research supports that pores larger than 10 μm will allow cell infiltration (Agrawal and Ray, 2001). The amount of porosity needed for tissue formation along with pore size is tissue and material specific. For example, osteogenesis in vivo will occur in biomaterials that contain a high porosity (70%) with average pore sizes 300 μm (Karageorgiou and Kaplan, 2005). However, in skin regeneration, successful scaffolds exhibit 20–124 μm pore sizes (Yannas et al., 1989). This is believed to be due to the low vascular needs of skin and its usual contact with the atmosphere to supply oxygen. There is, however, an upper limit in porosity and pore size set by constraints associated with mechanical properties. An increase in the void volume results in a reduction in mechanical strength of the scaffold, which can be critical for regeneration in organs requiring significant mechanical strength such as longbones, heart valves, and vasculature (Yannas, 2004). The extent to which pore size can be increased while maintaining mechanical requirements is dependent on many factors including the nature of the biomaterial and the processing conditions used in fabrication (Karageorgiou and Kaplan, 2005). In addition, the formation of an interconnected pore network has been shown to enhance the diffusion of metabolites to the center of the scaffold and has been shown to enhance vascularization (Zhang et al., 2004; O’Brien et al., 2005). Degradation Though non-degradable biomaterials have had success in many medical devices, many complications remain unsolved, mainly due to chronic foreign body responses. The most successful biomaterial will be the one that can eventually replaced by native tissues. The degradation rate of the construct is intrinsic to the success of the implant. Degradation of the material should occur at the same rate as tissue synthesis in order to insure suitable mechanical stability to allow native matrix deposition by host cells (Yannas, 2005a). The biomaterial scaffold must be presented long enough to allow cellular recruitment, attachment, and proliferation along with secretion and stabilization of ECM. The residence time of the scaffold is also tissue specific and depends upon the cell phenotype proliferation rate and ECM deposition (Yannas, 2004). If biomaterial scaffold degrades before sufficient

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Intramolecularly quenched substrate

Fluorescent cleavage product

Figure 39.5 Fluorogenic degradable substrate upon fabrication, fluorescent molecules are self-quenched and no fluorescence is detected. Upon proteolysis, fluorophores are no longer quenched.

ECM deposition has occurred, cells will lose important physiochemical factors for tissue regeneration and repair is likely to occur resulting in scar formation. However, if the scaffold residence time is too long, ECM deposition and cell proliferation will be suppressed. Additionally, the degradation products of the scaffold can be toxic not only to cells of the surrounding tissue, but also to the vital organs of the lymphatic system. Poly(lactic) acid (PLA) and poly(glycolic) acid (PGA) scaffolds upon degradation show a marked localized pH drop in the area around the template due to acidic degradation products (Martin et al., 1996; Lu et al., 2000). The pH decrease can be detrimental to cells and surrounding organs and over time can lead to an inflammatory response with possible capsule formation and even necrosis of surrounding tissue (Sung et al., 2004). In most scaffold materials in current use, degradation occurs via hydrolysis of chemical bonds in the polymer backbone from the aqueous environment in vivo. Chemical functionalities, percentage of cross-linking, and molecular determine the degradation characteristics. Higher molecular weight polymers tend to degrade more slowly over time as do polymers with a higher hydrophobicity and crystallinity. Using a combination of these factors, predictable degradation profiles can be utilized to match expected tissue formation rates. However, polymers that undergo bulk erosion can become rapidly unstable due to formation of large pores with low mechanical stability (Lu et al., 2000). Instead of utilizing hydrolysis for polymer degradation, chemical sequences have been introduced into the backbone of the polymer that can be degraded specifically by cells. Natural ECM proteins are degraded by matrix metalloproteinases (MMPs) and serine proteases that are either secreted or activated by local cells. Since proteolysis induced degradation is required for cell migration and invasion, researchers have had success in introducing synthetic hydrogels that are sensitive to cell proteases. Hydrogels containing amino acid sequences that can be degraded by plasmin (Halstenberg et al., 2002), MMPs (Kim et al., 2005), or both of these protease families (Mann et al., 2001a; Raeber et al., 2005) all show sustained degradation upon cellular infiltration. West and colleagues have fabricated MMP-degradable hydrogels that become fluorogenic when degraded by cell proteases (Lee et al., 2005). These polyethylene glycol (PEG)-based hydrogels incorporate MMP-degradable biomolecules into the polymer backbone. The biomolecules are labeled with fluorescent molecules that selfquench. Thus, quenched substrates show no fluorescence, but upon degradation by cell proteases, the fluorophores are no longer quenched and fluorescence can be measured (Figure 39.5). Cells seeded upon these fluorogenic substrates showed marked increase in fluorescence in the areas immediately around the cell, and cell remained viable after 7 days (Figure 39.6). In addition, cell migration trails could be seen in the hydrogels. It is believed these gels will contribute to the understanding of cell migration and degradation of material in three dimensions.

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10 µm (a)

(b)

Figure 39.6 Fibroblast encapsulated within fluorogenic substrate DIC image (a) and fluorescent image (b) showing fluorescence around cell.

Biomolecular Factors In many cases, seeding cells inside a porous scaffold is not enough for induced tissue regeneration because the material does not contain chemical cues that encourage cellular remodeling events. Thus, researchers attempt to actively modify biomaterials at the molecular by incorporating cell-specific biomolecules. One method is to make the material bioactive by incorporating relevant tissue engineering molecules such as peptides, growth factors, and other relevant tissue molecules into biomaterial carriers so that these molecules can be released from the material and trigger or modulate new tissue formation (Shin et al., 2003). One approach toward biomolecular recognition involves physically or chemically modifying biomaterials to incorporate specific cell-binding peptides. Cell-binding peptides are short amino acid sequences derived from much longer native ECM proteins that have been identified as able to incur specific, predictable interactions with cell receptors. Since most synthetic hydrogels materials are not adhesive to cells, introduction of adhesive sequences will attract and bind cells if signaling peptides are incorporated on the surface (Mann et al., 1999). Thus, incorporating peptides into these materials can potentially mimic the signaling dynamic between ECM and cells in tissues. The most studied celladhesion peptide, arginine–glycine–aspartic acid–serine (RGDS) has been widely used to encourage fibroblasts and other cells to adhere to polymer matrices to encourage tissue formation (Hern and Hubbell, 1998). The presence of this short peptide encourages adherence of specific cells on the surface of substrates that are normally non-adherent (Figure 39.7). Other amino acid sequences have been found that promote adhesion in specific cell phenotypes including endothelial cells (Gobin and West, 2003b; Heilshorn et al., 2003; Jun and West, 2005a, b), smooth muscle cells (Gobin and West, 2003b), neural cells (Adams et al., 2005), and osteoblasts (Benoit and Anseth, 2005). Various growth factors have also been studied for use as a cellular chemoattractant. Epidermal growth factor (EGF), platelet-derived growth factor (PDGF), and insulin-like growth factor (IGF) have been shown to induce both mitogenic and motogenic responses in various cell types. Griffith and colleagues showed that the presence of both fibronectin and EGF will cause cell motility in scaffolds in a co-dependent manner (Maheshwari et al., 1999). Because growth factors play a key role in tissue differentiation and repair, immobilization of growth factors

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(a)

RDGS

(b)

RGDS

Figure 39.7 Cell binding on RGDS substrate. Cells adhere and spread on PEG-hydrogel substrates containing RGDS compared to negative control RDGS (a).

High concentration

Axis of gradient

Location of initial cell seeding Low concentration

Figure 39.8 bFGF immobilized gradient scaffold. Cells seeded on immobilized bFGF gradient aligned along the axis of growth factor immobilization.

into biomaterials has been studied. To mimic this behavior in materials for tissue engineering, growth factors have been covalently coupled to PEG diacrylate (PEGDA) materials (Gobin and West, 2003a; DeLong et al., 2005b). These polymers are rendered chemoattractant to cells and in turn drive the secretion of native ECM. Further, the growth factors could be immobilized as a gradient to guide and direct tissue formation. Delong and West have formed gradients of bFGF. Cells were shown to preferentially align and migrate differentially along the bFGF gradient (Figure 39.8). In addition, bFGF and nerve growth factor have been immobilized into fibrin scaffolds in order to facilitate cellular recruitment and differentiation (Sakiyama-Elbert and Hubbell, 2000). Synthetic Materials for Histogenesis of New Organs Hydrolytically Degradable Polymers Synthetic polymers are viewed by many researchers as having the most promise as a biomaterial because they can be physically or chemically tailored to induce specific interaction with host cells or proteins. In addition, they can be structurally molded to mimic native biomechanics, they can be tailored to degrade for eventual replacement by host tissue, and they are generally less expensive to mass-produce than natural materials. The most

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widely used polymers for cellular scaffold materials are PLA, PGA, or a combination of these two polymers (PLGA). PLA, PGA, or PLGA are aliphatic esters that possess good biocompatibility (Li, 1999) and can be used as drug delivery materials to deliver biomolecules during tissue regeneration (Brannon-Peppas and Vert, 2000; Whang et al., 2000). These polymers are also among the few synthetic polymers approved by the US Food and Drug Administration (FDA) for certain human clinical applications. PGA is extremely hydrophilic in nature and, consequently, will lose its mechanical strength within 2–4 weeks of implantation (Reed and Gilding, 1981). PLA, however, contains one methyl group more than PGA and as a result it is more hydrophobic. Degradation rates for PLA scaffolds have been measured up to months and even years (Pitt et al., 1981; Brannon-Peppas and Vert, 2000). The degradation rates of these polymers can be tailored by using copolymer blends (PLGA) to give distinct degradation profiles (Brannon-Peppas and Vert, 2000; Ma, 2004). However, these polymers undergo acid-catalyzed hydrolysis and bulk erosion, which can cause the polymer to suddenly lose structural integrity before complete cellular incorporation into these ECM constructs. This lack of long-term mechanical stability could inhibit formation of new tissue (Moran and Bonassar, 1998). In addition, polyanhydrides have been synthesized for a number of biomedical applications including tissue engineering and drug delivery (Burkoth and Anseth, 2000). Polyanhydride networks exhibit excellent biocompatibility and contain a large aliphatic component possessing an ester group that undergoes surface erosion (Davis et al., 2003). The deliberate surface erosion is different from the bulk hydrolysis that is undergone by PLA or PGA and can allow biomaterials scaffolds to be made that have very predictable degradation profiles. In addition, the erosion of only the surface of the material allows anhydrides to maintain structural integrity to allow for support of cellular integration. Anhydrides have been widely studied as a scaffold for bone regeneration in vivo (Muggli et al., 1999; Burkoth and Anseth, 2000). Anhydrides exhibit mechanical properties similar to bone, and thus are ideal scaffolds for tissue infiltration. In addition, in the aqueous environment of the body, the wafers undergo slow surface erosion to allow maximum cellular migration, and the degradation products show minimal toxicity in vivo (Anseth et al., 1999). Polyanhydride networks can also be combined with other polymers to change their degradation and structural characteristics. Jiang and Zhu (1999) showed that anhydride polymers could be polymerized in the presence of PEG to form cross-linked networks with both hydrophobic and hydrophilic components. The hydrophilic PEG chains increase uptake of water to in turn drive the hydrolysis of the ester bond in the hydrophobic anhydride, and the degradation properties can be tailored by altering the amount of PEG in the polymer. Hydrogels As an alternative to aliphatic polymers, a class of polymers termed hydrogels are being studied for many tissue engineering applications. These polymers are termed hydrogels because the materials can absorb greater than 90% of the initial dry weight in water. These materials are appealing because the polymer properties are controllable and reproducible (Peppas, 2004) and the large water uptake promotes excellent biocompatibility due to low protein adsorption. In addition, the mechanical properties and hydrophilicity resemble the properties of native tissue. Many hydrogel monomers contain vinyl moieties, and as a result, many means of free radical initiated polymerizations as fabrication vehicles are possible. Photoinitiation, one such method, allows for polymers to be formed using specific wavelengths of light. Using this method, many researchers have had success forming complex three-dimensional structures with varying mechanical properties. Polyacrylamides are useful hydrogels that have induced regeneration of soft tissue in facial defects (von Buelow et al., 2005), and 2-hydroxyethyl methacrylate recently has been used as a fibrillar support for nerve regeneration (Flynn et al., 2003). Among the most studied hydrogel material is cross-linked PEG, which has been approved by the FDA for use in certain medical applications (Drury and Mooney, 2003). Like many hydrogels, the high water content of PEG causes low cellular and protein adherence and therefore a low immunorejection by the host. By changing

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the chain length, adding biological molecules or moieties, or utilizing copolymers, researcher have a large toolbox to use PEG polymers in a wide array of tissue engineering applications. For example, to gain cellular specificity, researchers have immobilized growth factors to the surfaces of biomaterials (Mann et al., 2001b; Gonzalez et al., 2004; DeLong et al., 2005a, b). The cellular peptides YIGSR and RGDS have been incorporated into PEG derivatives to encourage the formation of tissue. YIGSR is an ECM analog that binds endothelial cells to encourage intima layers to form in the artificial construct. Because of low protein adherence, PEG polymers have shown promise in the formation of small-diameter vascular grafts (Tulis et al., 2002a, b; Lipke et al., 2003; Lipke and West, 2005; Masters et al., 2005). In addition, PEG materials have been used to encapsulate cells in an attempt to encourage the cells to begin to secret native ECM molecules (Elisseeff et al., 1999; Burdick and Anseth, 2002). Groups have successfully used this technique toward the formation of new biomimetic constructs. However, initial strategies utilizing hydrogels only achieve limited success because the highly cross-linking hinders degradation as consequently, tissue induction. Consequently, groups have successfully incorporated degradable structures such as PLGA into hydrogel materials to produce a degradable structure while still maintaining the high water content of the hydrogel (Hubbell et al., 2001). Scaffolds in vivo: The Human Bioreactor Since the goal of tissue engineering is the complete regeneration of an organ, the human body can be considered a scaffold bioreactor. Because a material scaffold must eventually be placed into the defect, the use of the wound bed itself as a bioreactor has been investigated. Whether a material should be conditioned in vitro before implantation in vivo or if the implant should be directly implanted depends upon a number of factors. These factors include the size of the defect, the patient’s immune response, the health of tissue surrounding the wound, the consequences of tissue failure, and type of tissue to be replaced. When a biomaterial scaffold is placed into a tissue defect, the patient’s body immediately becomes a bioreactor for regeneration. Unlike bioreactors on the benchtop, homeostatic control of the wound bed is infinitely monitored and maintained because of the constant regulatory systems of the human body. Temperature, pH, and dissolved oxygen content in the blood are intrinsically controlled, and the removal of metabolic wastes and cells are handled by the lymphatic system. The blood that flows into the wound bed contains nutrients, cytokines, soluble proteins, and dissolved growth factors that encourage tissue formation and growth. The implanted scaffold directly interacts with the cells via biochemical and mechanical factors and the cells in turn use soluble factors in the blood to initiate repair. The interaction of the cells with the template causes a drastic change in the cascade of events that initiate the repair of the defect. The scaffold presence suppresses the rapid secretion of ECM molecules normally found in repair, and instead cells begin to systematically secrete proteases to break down the artificial ECM while also initiating the organized cellular events present in histogenesis such as division, differentiation, and apoptosis. Differentiation can be induced by the combination of a mechanical anchorage point, and biomolecules supplied by the blood and surrounding healthy tissue (Yannas, 2005b). Although the human body can be viewed as an ideal bioreactor due to the factors described above, the implantation of materials directly into the wound bed is not always ideal. When a tissue is damaged, the inflammatory response is initiated followed and by the cascade of wound healing. The inflammatory response can bring macrophages, neutrophils, and other inflammatory cells to the wound bed. These cells can secrete inhibitory molecules that block activation of growth factors bound on the scaffold. In addition, these cells can secrete proteolytic enzymes and other molecules that can prematurely degrade the scaffold material. In addition, if the tissue to be regenerated is diseased, the aspects of remodeling in the surrounding tissue can be greatly altered.

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Future Directions in Three-Dimensional Constructs: Three-Dimensional Microfabrication Since human tissues are three-dimensional entities, a way of reproducing the complex in vivo complexities of these systems is through the fabrication of three-dimensional scaffolds. One of the earliest examples of three-dimensional scaffold architecture was developed by Griffith and colleagues for hepatocyte culture and liver regeneration. Using rapid printing technique, microporous PLGA scaffolds were fabricated by directing solvent streams onto polymer granules in a controlled manner (Kim et al., 1998). The hepatocytes seeded upon these constructs exhibited increased metabolic rates that more closely mimicked hepatocytes in vivo. In addition, three-dimensional, microporous PLGA foams have been shown to successfully regenerate bone in animal models (Karp et al., 2003, 2004). In these studies, cylinders of PLGA were prepared using a drilling technique utilizing dies of a specific size. The size of the cylinders was reproducible to the millimeter scale and when placed in vivo, bone formation was seen in non-healing defects. Photopolymerizable hydrogels show promise as materials for three-dimensional fabrication due to the ease of fabricating these materials. Cells can easily be encapsulated within the gels during polymerization, thus reducing problems with seeding cells in the center of the construct. In addition, the presence of acrylate groups in PEG derivatives allows for rapid formation of many shapes and patterns of substrates via free-radical polymerization. Peppas and Ward (2004) have micropatterned hydrogels using UV polymerization on PEG hydrogels. Many different substrate morphologies were patterned with precise morphologies of less than 100 μm. It is believed that these three-dimensionally patterned substrates could be used for sensor applications and for biomaterials patterned on the microscale. Hahn et al. (2006) used photolithography to pattern cell-adhesive peptides onto the surface of PEGDA hydrogels. In this technique, an acrylated RGDS derivative was spread on the surface of a PEGDA hydrogel and a patterned transparency was placed on the surface of this solution (Figure 39.9). When exposed to UV light, the dark regions of the transparency were not photoinitiated. Thus, patterned surface was formed. Cells seeded upon the surface were only selectively bound to patterned regions of the hydrogel. In addition, Liu and Bhatia (2002) have photopatterned PEG hydrogels using a layer-by-layer method containing encapsulated cells. Cells remain viable in these scaffolds and these scaffolds can be patterned with features as small as 50 μm.

UV light

Patterned transparency

Coverslip PEGhydrogel hydrogel

Peptide solution

Figure 39.9 Photolithographic patterning of hydrogels. A peptide solution is spread on the surface of a PEG hydrogel and a patterned transparency is placed on top. UV exposure causes a patterned substrate to form.

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The use of laser patterning of hydrogels has been used to make advanced three-dimensional architecture inside hydrogel materials and natural constructs. Liu et al., used a laser ablation technique to form lines, holes, and interconnected grids in collagen matrices (Liu et al., 2005). Growth factors and peptides were patterned by Roy and colleagues using laser-based stereolithography (Mapili et al., 2005). Using this technique, layers of growth factors were patterned in different layers of a PEG-derived hydrogel. Luo and Shoichet have used a focused laser to pattern biomolecules inside agarose hydrogels (Luo and Shoichet, 2004). RGDS peptides were successfully patterned into cylinders in the hydrogel and surface seeded with primary dorsal root ganglia cells. After 3 days, neuronal cells showed migration into the hydrogel only at the RGDS patterned sections. Additionally, Hahn and colleagues have used laser scanning soft lithography of PEGDA hydrogels to successfully pattern complex three-dimensional geometries that could be used to pattern complex growth factor gradients to study cell migration (Hahn et al., 2005). In these studies, cells again only showed adherence and migration on patterned regions.

CONCLUSIONS Fabrication of functional three-dimensional tissues is the ultimate goal of tissue engineering. Many biomaterials and techniques have been investigated as tissue engineering scaffolds. Although many designs considerations still need to be investigated and certain challenges still exist, past experiments have revealed design parameters that are critical to the fabrication of replacement scaffolds. In addition, the use of cells in these scaffolds along with the use of biomolecules contained within these constructs is critical to the success of the implant. Histogenesis has been shown in vivo in many constructs, but the number of tissues that have been regenerated remain limited. New techniques of three-dimensional micropatterning have been developed that can allow precise structures to be patterned down to the micron scale. Many of these techniques show functional cells with cellular events that more greatly mimic those found in vivo. These techniques are expected to be used to fabricated material that more greatly mimic the complex organization of tissues. Use of these materials is expected to lead to greater insight into cell behavior and cell–biomaterial interactions while accelerating the field of tissue engineering.

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40 Biocompatibility and Bioresponse to Biomaterials James M. Anderson

INTRODUCTION Biocompatibility is generally defined as the ability of a biomaterial or medical device to perform with an appropriate host response in a specific application. Bioresponse or biocompatibility assessment (i.e. evaluation of biological responses) is considered to be a measure of the magnitude and duration of the adverse alterations in homeostatic mechanisms that determine the host response. From a practical view, the evaluation of biological responses to a medical device is carried out to determine that the medical device performs as intended and presents no significant harm to the patient. The goal of bioresponse evaluation is to predict whether a biomaterial or medical device presents potential harm to the patient. In regenerative medicine, biomaterials are utilized in a wide variety of ways ranging from carriers of genetic material to tissue-engineered implants that may contain autologous, allogeneic, or xenogeneic genetic materials, cells, and scaffold materials. Scaffolds may be composed of synthetic or modified-natural materials. A tissueengineered implant is a biologic–biomaterial combination in which some component of tissue has been combined with a biomaterial to create a device for the restoration or modification of tissue or organ function. Thus, tissue-engineered devices having a biologic component(s) require an expanded perspective and understanding of biocompatibility and biological response evaluation. The purpose of this chapter is to provide an overview of this expanded perspective. It must be understood that each unique tissue-engineered device requires a unique set of experiments to determine its biological responses and biocompatibility. This chapter presents an overview of host responses that must be considered in determining the biocompatibility of tissue-engineered devices that utilize biomaterials. The three major responses that must be considered for biocompatibility assessment are: (1) inflammation, (2) wound healing, and (3) immunological reactions or immunity. For the purposes of biological response evaluation, the immunological reactions or immunity are considered to be immunotoxicity. Pathologists use the terminology of inflammation and immunity to describe adverse tissue reactions whereas immunologists commonly refer to inflammation as innate immunity and activation of the immune system as being acquired immunity. Tissue/material interactions are a series of responses that are initiated by the implantation procedure, as well as by the presence of the biomaterial, medical device, or tissue-engineered device. In this chapter, we divide the series of tissue/material responses into inflammation (innate immunity) and wound healing, and immunotoxicity. Following implantation, early, transient tissue/material responses include injury (implantation), blood–materials interactions, provisional matrix formation, and the temporal sequence of inflammation and wound healing including acute inflammation, chronic inflammation, granulation tissue development, foreign body reaction, and ultimately fibrosis/fibrous capsule (scar) development. Immunotoxicity is any adverse effect on the function or structure of the immune system or other systems as

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a result of an immune system dysfunction. Two significant failure mechanisms of tissue-engineered devices are fibrosis/fibrous capsule (scar) development surrounding and infiltrating the tissue-engineered device, or the initiation of acquired or cellular immunity by the biological component of the tissue-engineered device. It must also be considered that the biological component and the biomaterial component in a tissueengineered device may act in concert or synergistically to facilitate either of these failure mechanisms.

INFLAMMATION (INNATE IMMUNITY) AND WOUND HEALING The process of implantation of a biomaterial or tissue-engineered device results in injury to tissues or organs (Anderson, 1988, 1993, 2001; Cotran et al., 1999; Gallin and Synderman, 1999). It is this injury and the subsequent perturbation of homeostatic mechanisms that lead to the inflammatory responses, foreign body reaction, and wound healing. The response to injury is dependent on multiple factors that include the extent of injury, loss of basement membrane structures, blood–material interactions, provisional matrix formation, extent or degree of cellular necrosis, and extent of the inflammatory response. The organ or tissue undergoing implantation may play a significant role in the response. These events, in turn, may affect the extent or degree of granulation tissue formation, foreign body reaction, and fibrosis or fibrous capsule (scar) development. These events are summarized in Table 40.1. These host reactions for biocompatible biomaterials are considered to be normal. It is noteworthy that these host reactions are also tissue-dependent, organ-dependent, and species-dependent. These dependencies thus provide perspectives on the biological response evaluation and the ultimate determination of biocompatibility. It is important to recognize that these reactions occur or are initiated early, that is, within 2–3 weeks of the time of implantation and undergo resolution rather quickly leading to fibrosis or fibrous capsule formation. Blood–Material Interactions and Initiation of the Inflammatory Response Blood–material interactions and the inflammatory response are intimately linked, and in fact, early responses to injury involve mainly blood and the vasculature (Anderson, 1988, 1993, 2001; Cotran et al., 1999; Gallin and Synderman, 1999). Regardless of the tissue into which a biomaterial is implanted, the initial inflammatory response is activated by injury to vascularized connective tissue. Because blood and its components are involved in the initial inflammatory responses, thrombus and/or blood clot also form. Thrombus formation involves activation of the extrinsic and intrinsic coagulation systems, the complement system, the fibrinolytic system, the kinin-generating system, and platelets. Thrombus or blood clot formation on the surface of a biomaterial is related to the well-known Vroman effect of protein adsorption. From a wound healing perspective, blood protein deposition on a biomaterial surface is described as provisional matrix formation. Although injury initiates the inflammatory response, released chemicals from plasma, cells, and injured tissue mediate the response (Salthouse, 1976; Cotran et al., 1999; Gallin and Synderman, 1999; Weisman et al., 1980). Important classes of chemical mediators of inflammation are presented in Table 40.2. Several important

Table 40.1 Sequence of host reactions Injury Blood–material interactions Provisional matrix formation Acute inflammation Granulation tissue Foreign body reaction Fibrosis/fibrous capsule development

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Table 40.2 Important chemical mediators of inflammation derived from plasma, cells, or injured tissue Mediators

Examples

Vasoactive agents

Histamine, serotonin, adenosine, endothelial derived relaxing factor (EDRF), prostacyclin, endothelin, thromboxane a2

Plasma proteases Kinin system Complement system Coagulation/fibrinolytic system Leukotrienes Lysosomal proteases Oxygen-derived free radicals Platelet activating factors Cytokines Growth factors

Bradykinin, kallikrein C3a, C5a, C3b, C5b–C9 Fibrin degradation products, activated Hageman factor (FXIIA), tissue plasminogen activator (tPA) Leukotriene B4 (LTB4), hydroxyeicosatetranoic acid (HETE) Collagenase, elastase H2O2, superoxide anion, nitric oxide Cell membrane lipids Interleukin-1 (IL-1), TNF PDGF, fibroblast growth factor (FGF), transforming growth factor (TGF-α or TGF-β), epithelial growth factor (EGF)

points must be noted in order to understand the inflammatory response and how it relates to biomaterials. First, although chemical mediators are classified on a structural or functional basis, different mediator systems interact and provide a system of checks and balances regarding their respective activities and functions. Second, chemical mediators are quickly inactivated or destroyed, suggesting that their action is predominantly local (i.e. at the implant site). Third, generally acid, lyosomal proteases and oxygen-derived free radicals produce the most significant damage or injury. These chemical mediators are also important in the degradation of biomaterials. The predominant cell type present in the inflammatory response varies with the age of the injury. In general, neutrophils, commonly called polymorphonuclear leukocytes or polys, predominate during the first several days following injury and then are replaced by monocytes as the predominant cell type. Three factors account for this change in cell type: (i) Neutrophils are short-lived and disintegrate and disappear after 24–48 h; neutrophil emigration is of short duration because chemotactic factors for neutrophil migration are activated early in the inflammatory response. (ii) Following emigration from the vasculature, monocytes differentiate into macrophages, and these cells are very long-lived (up to months). (iii) Monocyte emigration may continue for days to weeks, depending on the injury and implanted biomaterial, and chemotactic factors for monocytes are activated over longer periods of time. Provisional Matrix Formation Injury to vascularized tissue in the implantation procedure leads to immediate development of the provisional matrix at the implant site. This provisional matrix consists of fibrin, produced by activation of the coagulative and thrombosis systems, and inflammatory products released by the complement system, activated platelets, inflammatory cells, and endothelial cells (Clark et al., 1982; Tang et al., 1993; Tang, 1998). These events occur early, within minutes to hours following implantation of a medical device. Components within or released from the provisional matrix, that is, fibrin network (thrombosis or clot), initiate the resolution, reorganization, and repair processes such as inflammatory cell and fibroblast recruitment. Platelets, activated during the fibrin network formation, release platelet factor 4, platelet-derived growth factor (PDGF), and transforming growth factor β (TGF-β), which contribute to fibroblast recruitment (Wahl et al., 1989; Riches, 1998). Monocytes and lymphocytes, upon activation, generate additional chemotactic factors including LTB4, PDGF, and TGF-β to recruit fibroblasts.

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The provisional matrix is composed of adhesive molecules such as fibronectin and thrombospondin bound to fibrin as well as platelet granule components released during platelet aggregation. Platelet granule components include thrombospondin, released from the platelet α-granule, and cytokines including TGF-α, TGF-β, PDGF, platelet factor 4, and platelet-derived endothelial cell growth factor. The provisional matrix is stabilized by the cross-linking of fibrin by factor XIIIa. The provisional matrix appears to provide both structural and biochemical components to the process of wound healing. The complex three-dimensional structure of the fibrin network with attached adhesive proteins provides a substrate for cell adhesion and migration. The presence of mitogens, chemoattractants, cytokines, and growth factors within the provisional matrix provide for a rich milieu of activating and inhibiting substances for various cellular proliferative and synthetic processes. The provisional matrix may be viewed as a naturally derived, biodegradable, sustained release system in which mitogens, chemoattractants, cytokines, and growth factors are released to control subsequent wound healing processes (Dvorak et al., 1987; Ignotz et al., 1987; Muller et al., 1987; Wahl et al., 1987; Madri et al., 1988; Sporn and Roberts, 1988; Broadley et al., 1989). In spite of the rapid increase in our knowledge of the provisional matrix and its capabilities, our knowledge of the control of the formation of the provisional matrix and its effect on subsequent wound healing events is poor.

Temporal Sequence of Inflammation and Wound Healing Inflammation is generally defined as the reaction of vascularized living tissue to local injury. Inflammation serves to contain, neutralize, dilute, or wall off the injurious agent or process. In addition, it sets into motion a series of events that may heal and reconstitute the implant site through replacement of the injured tissue by regeneration of native parenchymal cells, formation of fibroblastic scar tissue, or a combination of these two processes (Cotran et al., 1999; Gallin and Synderman, 1999). The sequence of events following implantation of a biomaterial is illustrated in Figure 40.1. The size, shape, and chemical and physical properties of the biomaterial and the physical dimensions and properties of the prosthesis or device may be responsible for variations in the intensity and time duration of the inflammatory and wound healing processes. Thus, intensity and/or time duration of inflammatory reaction may characterize the biocompatibility of a biomaterial, or device. Classically, the biocompatibility of an implanted material has been described in terms of the morphological appearance of the inflammatory reaction to the material; however, the inflammatory response is a series of complex reactions involving various types of cells, the densities, activities, and functions of which are controlled by various endogenous and autocoid mediators. The simplistic view of the acute inflammatory response progressing to the chronic inflammatory response may be misleading with respect to biocompatibility studies and the inflammatory response to implants. In vivo studies using the cage implant system show that monocytes and macrophages are present in highest concentrations when neutrophils are also at their highest concentrations, that is, the acute inflammatory response (Marchant et al., 1983; Spilizewski et al., 1985). Neutrophils have short lifetimes – hours to days – and disappear from the exudates more rapidly than do macrophages, which have lifetimes of days to weeks to months. Eventually macrophages become the predominant cell type in the exudates, resulting in a chronic inflammatory response. Monocytes rapidly differentiate into macrophages, the cells principally responsible for normal wound healing in the foreign body reaction. Classically, the development of granulation tissue has been considered to be part of chronic inflammation, but because of unique tissue–material interactions, it is preferable to differentiate the foreign body reaction – with its varying degree of granulation tissue development, including macrophages, fibroblasts, and capillary formation – from chronic inflammation.

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Injury, Implantation Inflammatory Cell Infiltration PMNs, Monocytes, Lymphocytes

Exudate/Tissue

Biomaterial

Acute Inflammation PMNs Monocyte adhesion Macrophage differentiation Macrophage mannose Receptor upregulation

Chronic Inflammation Monocytes Lymphocytes

Th2: IL-4, IL-13

Macrophage fusion

Granulation Tissue Fibroblast proliferation and migration Capillary formation

Fibrous Capsule Formation

Foreign Body Giant Cell Formation

Figure 40.1 Sequence of events involved in inflammatory and wound healing responses leading to FBGC formation. This shows the importance of Th2 lymphocytes in the transient chronic inflammatory phase with the production of IL-4 and IL-3 that can induce monocyte/macrophage fusion to form FBGCs.

Acute Inflammation Acute inflammation is of relatively short duration, lasting from minutes to days, depending on the extent of injury. The main characteristics of acute inflammation are the exudation of fluid and plasma proteins (edema) and the emigration of leukocytes (predominantly neutrophils). Neutrophils and other motile white cells emigrate or move from the blood vessels to the perivascular tissues and the injury (implant) site (Henson et al., 1987; Malech et al., 1987; Ganz, 1988). The accumulation of leukocytes, in particular neutrophils and monocytes, is the most important feature of the inflammatory reaction. Leukocytes accumulate through a series of processes including margination, adhesion, emigration, phagocytosis, and extracellular release of leukocyte products (Jutila, 1990). Increased leukocytic adhesion in inflammation involves specific interactions between complementary “adhesion molecules” present on the leukocyte and endothelial surfaces (Cotran and Pober, 1990; Pober and Cotran, 1990). The surface expression of these adhesion molecules is modulated by inflammatory agents; mechanisms of interaction include stimulation of leukocyte adhesion molecules (C5a, LTB4), stimulation of endothelial adhesion molecules (IL-1), or both effects tumor necrosis factor-α (TNF-α). Integrins comprise a family of transmembrane glycoproteins that modulate cell–matrix and cell–cell relationships by acting as receptors to extracellular protein ligands and also as direct adhesion molecules (Hynes, 1992). An important group of integrins (adhesion molecules) on leukocytes include the CD11/CD18 family of adhesion molecules.

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Inflammatory mediators (i.e. cytokines) stimulate a rapid increase in these adhesion molecules on the leukocyte surface as well as increased leukocyte adhesion to endothelium. Leukocyte–endothelial cell interactions are also controlled by endothelial–leukocyte adhesion molecules (ELAMs, E-selectins) or intracellular adhesion molecules (ICAM-1, ICAM-2, and vascular cell adhesion molecules (VCAMs)) on endothelial cells (Butcher, 1991). Inflammatory cell emigration is controlled in part by chemotaxis, which is the unidirectional migration of cells along a chemical gradient. A wide variety of exogenous and endogenous substances have been identified as chemotactic agents (Henson, 1971, 1980; Weisman et al., 1980; Henson et al., 1987; Malech and Gallin, 1987; Ganz, 1988; Weiss, 1989; Cotran and Pober, 1990; Jutila, 1990; Paty et al., 1990; Pober and Cotran, 1990; Butcher, 1991; Hynes, 1992). Important to the emigration or movement of leukocytes is the presence of specific receptors for chemotactic agents on the cell membranes of leukocytes. These and other receptors may also play a role in the activation of leukocytes. Following localization of leukocytes at the injury (implant) site, phagocytosis and the release of enzymes occur following activation of neutrophils and macrophages. The major role of the neutrophils in acute inflammation is to phagocytose microorganisms and foreign materials. Phagocytosis is seen as a three-step process in which the injurious agent undergoes recognition and neutrophil attachment, engulfment, and killing or degradation. With regard to biomaterials, engulfment and degradation may or may not occur depending on the properties of the biomaterial. Although biomaterials are not generally phagocytosed by neutrophils or macrophages because of the size disparity (i.e. the surface of the biomaterial is greater than the size of the cell), certain events in phagocytosis may occur. The process of recognition and attachment is expedited when the injurious agent is coated by naturally occurring serum factors called opsonins. The two major opsonins are IgG and the complement-activated fragment, C3b. Both of these plasma-derived proteins are known to adsorb to biomaterials, and neutrophils and macrophages have corresponding cell membrane receptors for these opsonization proteins. These receptors may also play a role in the activation of the attached neutrophil or macrophage. Because of the size disparity between the biomaterial surface and the attached cell, “frustrated phagocytosis” may occur (Henson, 1971, 1980). This process does not involve engulfment of the biomaterial but does cause the extracellular release of leukocyte products in an attempt to degrade the biomaterial. Neutrophils adherent to complement-coated and immunoglobulin-coated non-phagocytosable surfaces may release enzymes by direct extrusion or exocytosis from the cell (Henson, 1971, 1980). The amount of enzyme released during this process depends on the size of the polymer particle, with larger particles inducing greater amounts of enzyme release. This suggests that the specific mode of cell activation in the inflammatory response in tissue is dependent upon the size of the implant and that a material in a phagocytosable form (e.g. powder or particulate) may provoke a degree of inflammatory response different from that of the same material in a non-phagocytosable form (e.g. film). Tissue-engineered constructs containing biomaterial scaffolds alone, or with cells and/or chemokines, growth factors, or other biological components are thus subjected to an aggressive microenvironment that may quickly compromise the intended function of the construct (Babensee et al., 1998). Chronic Inflammation Chronic inflammation is less uniform histologically than is acute inflammation. In general, chronic inflammation is characterized by the presence of monocytes and lymphocytes with the early proliferation of blood vessels and connective tissue (Williams et al., 1983; Johnston, 1988; Cotran et al., 1999; Gallin and Synderman, 1999). It must be noted that many factors modify the course and histological appearance of chronic inflammation. Persistent inflammatory stimuli lead to chronic inflammation. Although the chemical and physical properties of the biomaterial may lead to chronic inflammation, motion in the implant site by the biomaterial may

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also produce chronic inflammation. The chronic inflammatory response to biomaterials is confined to the implant site. Inflammation with the presence of mononuclear cells, including lymphocytes and plasma cells, is given the designation chronic inflammation, whereas the foreign body reaction with granulation tissue development is considered the normal wound healing response to implanted biomaterials (i.e. the normal foreign body reaction). Chronic inflammation with biocompatible materials is usually of very short duration (i.e. a few days). Lymphocytes and plasma cells are involved principally in immune reactions and are key mediators of antibody production and delayed hypersensitivity responses. Their roles in non-immunological injuries and inflammation are largely unknown. Little is known regarding immune responses and cell-mediated immunity to synthetic biomaterials. The role of macrophages must be considered in the possible development of immune responses to synthetic biomaterials. Macrophages process and present the antigen to immunocompetent cells and thus are key mediators in the development of immune reactions. The macrophage is probably the most important cell in chronic inflammation because of the great number of biologically active products its produces (Johnston, 1988). Important classes of products produced and secreted by macrophages include neutral proteases, chemotactic factors, arachidonic acid metabolites, reactive oxygen metabolites, complement components, coagulation factors, growth-promoting factors, and cytokines. Growth factors such as PDGF, FGF, TGF-β, TGF-α/EGF, and IL-1 or TNF are important to the growth of fibroblasts and blood vessels and the regeneration of epithelial cells. Growth factors, released by activated cells, stimulated production of a wide variety of cells; initiate cell migration, differentiation, and tissue remodeling; and may be involved in various stages of wound healing (Mustoe et al., 1987; Wahl et al., 1989; Fong et al., 1990; Sporn and Roberts, 1990; Golden et al., 1991; Kovacs, 1991). It is clear that there is a lack of information regarding interaction and synergy among various cytokines and growth factors and their abilities to exhibit chemotactic, mitogenic, and angiogenic properties. Granulation Tissue Within 1 day following implantation of a biomaterial (i.e. injury), the healing response is initiated by the action of monocytes and macrophages, followed by proliferation of fibroblasts and vascular endothelial cells at the implant site, leading to the formation of granulation tissue, the hallmark of healing inflammation. Granulation tissue derives its name from the pink, soft granular appearance on the surface of healing wounds, and its characteristic histological features include the proliferation of new small blood vessels and fibroblasts. Depending on the extent of injury, granulation tissue may be seen as early as 3–5 days following implantation of a biomaterial. The new small blood vessels are formed by budding or sprouting of pre-existing vessels in a process known as neovascularization or angiogenesis (Ziats et al., 1985; Thompson et al., 1988; Maciag, 1990). This process involves proliferation, maturation, and organization of endothelial cells into capillary tubes. Fibroblasts also proliferate in developing granulation tissue and are active in synthesizing collagen and proteoglycans. In the early stages of granulation tissue development, proteoglycans predominate; later, however, collagen – especially type I collagen – predominates and forms the fibrous capsule. Some fibroblasts in developing granulation tissue may have features of smooth muscle cells. These cells are called myofibroblasts and are considered to be responsible for the wound contraction seen during the development of granulation tissue. Macrophage Interactions The inflammatory and immune systems overlap considerably through the activity and phenotypic expression of macrophages that are derived from blood-borne monocytes. Monocytes and macrophages belong to the

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mononuclear phagocytic system (MPS), Table 40.3. Cells in the MPS may be considered as resident macrophages in the respective tissues that take on specialized functions that are dependent on their tissue environment. From this perspective, the host defense system may be seen as blood-borne or circulating inflammatory and immune cells as well as mononuclear phagocytic cells that reside in specific tissues with specialized functions. In the inflammatory and immune responses, the macrophage plays a pivotal role in both the induction and effector phases of these responses. Two factors that play a role in monocyte/macrophage adhesion and activation and foreign body giant cell (FBGC) formation are the surface chemistry of the substrate onto which the cells adhere and the protein adsorption that occurs before cell adhesion. These two factors have been hypothesized to play significant roles in the inflammatory and wound healing responses to biomaterials and medical devices in vivo. Macrophage interactions with biomaterials are initiated when blood-borne monocytes in the early, transient responses migrate to the implant site and adhere to the blood protein adsorbed biomaterial through monocyte–integrin interactions. Following adhesion, adherent monocytes differentiate into macrophages that may then fuse to form FBGCs. Figure 40.2 demonstrates the progression from circulating blood monocyte to tissue macrophage to FBGC development that is most commonly observed. Because of the progression of monocytes to macrophages to FBGCs (Figure 40.2), the following discussion of macrophage interactions also includes perspectives on how macrophages are formed (i.e. monocyte adhesion) and what happens to macrophages on biomaterial surfaces (i.e. FBGC formation) (McNally et al., 1994; McNally et al., 1995).

Table 40.3 The mononuclear phagocytic system Tissues

Cells

Implant sites Liver Lung Connective tissue Bone marrow Spleen and lymph nodes Serous cavities Nervous system Bone Skin Lymphoid tissue

Inflammatory macrophages, FBGCs Kupffer cells Alveolar macrophages Histiocytes Macrophages Fixed and free macrophages Pleural and peritoneal macrophages Microglial cells Osteoclasts Langerhans’ cells, dendritic cells Dendritic cells

Macrophage

Foreign Body Giant Cell

Tissue/Biomaterial

Biomaterial

Monocyte Blood

Tissue

Chemotaxis Migration

Chemotaxis Migration Adhesion Differentiation

Adhesion Differentiation Signal Transduction Activation

Activity Phenotypic Expression

Figure 40.2 In vivo transition from blood-borne monocyte to biomaterial adherent monocyte/macrophage to FBGC at the tissue/biomaterial interface. Little is known regarding the indicated biological responses that are considered to play important roles in the transition to FBGC development.

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Material surface property-dependent blood protein adsorption occurs immediately upon surgical implantation of a biomaterial and it is the protein-modified biomaterial that inflammatory cells subsequently encounter. Monocytes express receptors for various blood components, but they recognize naturally occurring foreign surfaces by receptors for opsonins such as fragments of complement component C3. Complement activation by biomaterials has been well documented. Exposure to blood during biomaterial implantation may permit extensive opsonization with the labile fragment C3b and the rapid conversion of C3b to its hemolytically inactive but nevertheless opsonic and more stable form, C3bi. C3b is bound by the CD35 receptor, but C3bi is recognized by distinct receptors, CD11b/CD18 and CD11c/CD18 on monocytes (McNally et al., 1994). Fibrinogen, a major plasma protein that adsorbs to biomaterials, is another ligand for these receptors that together with CD11a/CD18 constitutes a subfamily of integrins that is restricted to leukocytes (McNally et al., 1994, 1995). Studies with monoclonal antibodies to their common β2 subunit (CD 18) and distinct α chains have implicated CD11b/CD18 and CD11c/CD18 in monocyte/macrophage responses. Other potential adhesion-mediating proteins that adsorb to biomaterials include IgG, which may interact with monocytes via various receptors and fibronectin, for which monocytes also express multiple types of receptors (Jenney and Anderson, 2000; McNally and Anderson, 2002). FBGC Formation and Interactions The foreign body reaction is composed of FBGCs and the components of granulation tissue, which consist of macrophages, fibroblasts, and capillaries in varying amounts, depending upon the form and topography of the implanted material. Relatively flat and smooth surfaces, such as those found on breast prostheses, have a foreign body reaction that is composed of a layer of macrophages one to two cells in thickness. Relatively rough surfaces, such as those found on the outer surfaces of expanded poly(tetrafluroethylene) (ePTFE) vascular prostheses or poly(methyl methacrylate) (PMMA) bone cement, have a foreign body reaction composed of several layers of macrophages and FBGCs at the surface. Fabric materials generally have a surface response composed of macrophages and FBGCs with varying degrees of granulation tissue subjacent to the surface response. As previously discussed, the form and topography of the surface of the biomaterial determines the composition of the foreign body reaction. With biocompatible materials, the composition of the foreign body reaction in the implant site may be controlled by the surface properties of the biomaterial, the form of the implant, and the relationship between the surface area of the biomaterial and the volume of the implant. For example, high surface-to-volume implants such as fabrics or porous materials will have higher ratios of macrophages and FBGCs in the implant site than will smooth-surface implants, which will have fibrosis as a significant component of the implant site. The foreign body reaction consisting mainly of macrophages and/or FBGCs may persist at the tissue– implant interface for the lifetime of the implant (Chambers and Spector; 1982; Rae, 1986; Anderson, 1988, 1993, 2000; Greisler, 1988). Generally, fibrosis (i.e. fibrous encapsulation) surrounds the biomaterial or implant with its interfacial foreign body reaction, isolating the implant and foreign body reaction from the local tissue environment. Early in the inflammatory and wound healing response, the macrophages are activated upon adherence to the material surface. Although it is generally considered that the chemical and physical properties of the biomaterial are responsible for macrophage activation, the nature of the subsequent events regarding the activity of macrophages at the surface is not clear. Tissue macrophages, derived from circulating blood monocytes, may coalesce to form multinucleated FBGCs. FBGCs containing large numbers of nuclei are typically present on the surface of biomaterials. Although these FBGCs may persist for the lifetime of the implant, it is not known if they remain activated, releasing their lysosomal constituents, or become quiescent. FBGCs have been implicated in the biodegradation of polymeric medical devices (Zhao et al., 1990, 1991; Wiggins et al., 2001).

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Figure 40.1 demonstrates the sequence of events involved in inflammation and wound healing when medical devices are implanted. In general, the neutrophil (PMN) predominant acute inflammatory response and the lymphocyte/monocyte predominant chronic inflammatory response resolve quickly (i.e. within 2 weeks) depending on the type and location of implant. Studies utilizing IL-4 demonstrate the role for Th2 helper lymphocytes in the development of the foreign body reaction at the tissue/material interface. Th2 helper lymphocytes have been described as “anti-inflammatory” based on their cytokine profile of which IL4 is a significant component. Th2 helper lymphocytes also produce IL-13 that has a similar effect to IL-4 on FBGC formation. In this regard, it is noteworthy that anti-IL-4 antibody does not inhibit IL-13 induced FBGC formation nor does anti-IL-13 antibody inhibit IL-4 induced FBGC formation. In IL-4 and IL-13 FBGC culture systems, the macrophage mannose receptor (MMR) has been identified as critical to the fusion of macrophages in the formation of FBGC (McNally et al., 1996; DeFife et al., 1997). FBGC formation can be prevented by competitive inhibitors of MMR activity (i.e. α-mannan) or inhibitors of glycoprotein processing that restrict MMR surface expression.

FIBROSIS AND FIBROUS ENCAPSULATION The end-stage healing response to biomaterials is generally fibrosis or fibrous encapsulation. However, tissueengineered devices may be exceptions to this general statement (e.g. porous materials inoculated with parenchymal cells or porous materials implanted into bone). Repair of implant sites involves two distinct processes: regeneration, which is the replacement of injury tissue by parenchymal cells of the same type, or replacement by connective tissue that constitutes the fibrous capsule. These processes are generally controlled by either (i) the proliferative capacity of the cells in the tissue receiving the implant and the extent of injury as it relates to the destruction or (ii) persistence of the tissue framework of the implant site. The regenerative capacity of cells permits classification into three groups: labile, stable (or expanding), and permanent (or static) cells. Labile cells continue to proliferate throughout life, stable cells retain this capacity but do not normally replicate, and permanent cells cannot reproduce themselves after birth. Perfect repair with restitution of normal structure theoretically occurs only in tissue consisting of stable and labile cells, whereas all injuries to tissues composed of permanent cells may give rise to fibrosis and fibrous capsule formation with very little restitution of the normal tissue or organ structure. Tissues composed of permanent cells (e.g. nerve cells, skeletal muscle cells, and cardiac muscle cells) most commonly undergo an organization of the inflammatory exudates, leading to fibrosis. Tissues composed of stable cells (e.g. parenchymal cells of the liver, kidney, and pancreas), mesenchymal cells (e.g. fibroblasts, smooth muscle cells, osteoblasts, and chondroblasts), and vascular endothelial and labile cells (e.g. epithelial cells and lymphoid and hematopoietic cells) may also follow this pathway to fibrosis or may undergo resolution of the inflammatory exudates, leading to restitution of the normal tissue structure. The condition of the underlying framework or supporting stroma of the parenchymal cells following an injury plays an important role in the restoration of normal tissue structure. Retention of the framework may lead to restitution of the normal tissue structure, whereas destruction of the framework most commonly leads to fibrosis. It is important to consider the species-dependent nature of the regenerative capacity of cells. For example, cells from the same organ or tissue but from different species may exhibit different regenerative capacities and/or connective tissue repair. The extent of provisional matrix formation is an important factor as it is related to wound healing by first or second intention. First intention (primary union) wound healing occurs when there is minimal to no space between the tissue and device whereas second intention (secondary union) wound healing occurs when a large space, providing for extensive provisional matrix formation, is present. Obviously, inappropriate or inadequate

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Table 40.4 Common components in the inflammatory (innate) and immune (adaptive) responses Components Complement cascade components Immunoglobulins Cellular components Macrophages NK (natural killer) cells Dendritic cells Cells with dual phagocytic and antigen presenting capabilities

preparation of the implant site leading to extensive provisional matrix formation may predispose the implant to failure through mechanisms related to fibrous capsule formation. The inflammatory (innate) and immune (adaptive) responses have common components. It is possible to have inflammatory responses only with no adaptive immune response. In this situation, both humoral and cellular components that are shared by both types of responses may only participate in the inflammatory response. Table 40.4 indicates the common components to the inflammatory (innate) and immune (adaptive) responses. Macrophages and dendritic cells are known as professional antigen-presenting cells responsible for the initiation of the adaptive immune response.

IMMUNOTOXICITY (ACQUIRED IMMUNITY) The acquired or adaptive immune system acts to protect the host from foreign agents or materials and is usually initiated through specific recognition mechanisms and the ability of humoral and cellular components to recognize the foreign agent or material as being “non-self” (Coligan et al., 1992; Burleson et al., 1995; Smialowicz and Holsapple, 1996; Janeway and Travers, 1997; Rose et al., 1997). Generally, the adaptive immune system may be considered as having two components: humoral or cellular. Humoral components include antibodies, complement components, cytokines, chemokines, growth factors, and other soluble mediators. These components are synthesized by cells of the immune response and, in turn, function to regulate the activity of these same cells and provide for communication between different cells in the cellular component of the adaptive immune response. Cells of the immune system arise from stem cells in the bone marrow (B lymphocytes) or the thymus (T lymphocytes) and differ from each other in morphology, function, and the expression of cell-surface antigens. They share the common features of maintaining cell-surface receptors that assist in the recognition and/or elimination of foreign materials. Regarding tissue-engineered devices, the adaptive immune response may recognize the biological components, modifications of the biological components, or degradation products of the biological components, commonly known as antigens, and initiate immune response through humoral or cellular mechanisms. Components of the humoral immune system play important roles in the inflammatory responses to foreign materials. Antibodies and complement components C3b and C3bi adhere to foreign materials, act as opsonins and facilitate phagocytosis of the foreign materials by neutrophils and macrophages that have cell-surface receptors for C3b. Complement component C5a is a chemotactic agent for neutrophils, monocytes, and other inflammatory cells and facilitate the immigration of these cells to the implant site. The complement system is composed of classic and alternative pathways that eventuate in a common pathway to produce the membrane attack complex (MAC), which is capable of lysing microbial agents. The complement system (i.e.complement cascade) is closely controlled by protein inhibitors in the host cell membrane that may prevent damage to host cells. This inhibitory mechanism may not function when non-host cells are used in tissue-engineered devices.

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T (thymus-derived) lymphocytes are significant cells in the cell-mediated adaptive immune response and their cell-adhesion molecules play a significant role in lymphocyte migration, activation, and effector function. The specific interaction of cell membrane adhesion molecules, sometimes also called ligands or antigens, with antigen-presenting cells (APCs) produce specific types of lymphocytes with specific functions. Table 40.5 indicates cell types and function in the adaptive immune response. Obviously, the functions of these cells are more numerous than that indicated in Table 40.5 but the major function of these cells is provided to indicate similarities and differences in the interaction and responsiveness of these cells. Effector T-cells (Table 40.6) are produced when their antigen-specific receptors and either the CD4 or the CD8 co-receptors bind to peptide-MHC (major histocompatibility complex) complexes. A second, co-stimulatory signal is also required and this is provided by the interaction of the CD28 receptor on the T-cell and the B7.1 and B7.2 glycoproteins of the immunoglobulin superfamily present on APCs. B lymphocytes bind soluble antigens through their cell-surface immunoglobulin and thus can function as professional APCs by internalizing the soluble antigens and presenting peptide fragments of these antigens as MHC: peptide complexes. Once activated, T-cells can synthesize the

Table 40.5 Cell types and function in the adaptive immune system Cell type

Motor function

Macrophages (APC)

Process and present antigen to immunocompetent T-cells Phagocytosis Activated by cytokines (i.e. IFN-γ) from other immune cells

T-cells

Interact with APCs and are activated through two required cell membrane interactions Facilitate target cell apoptosis Participate in transplant rejection (type IV hypersensitivity)

B-cells

Form plasma cells that secrete immunoglobulins (IgG, IgA, and IgE) Participate in antigen–antibody complex mediated tissue damage (type III hypersensitivity)

Dendritic cells (APC)

Process and present antigen to immunocompetent T-cells Utilize Fc receptors for IgG to trap antigen–antibody complexes

NK cells (non-T, non-B lymphocytes)

Innate ability to lyse tumor, virus infected, and other cells without previous sensitization Mediates T- and B-cell function by secretion of IFN-γ

Table 40.6 Effector T lymphocytes in adaptive immunity Th1 helper cells

CD4 Pro-inflammatory Activation of macrophages Produces IL-2, interferon-γ (IFN-γ), IL-3, TNF-α, GM-CSF, macrophage chemotactic factor (MCF), migration inhibitor factor (MIF) Induce IgG2a

Th2 helper cells

CD4 Anti-inflammatory Activation of B-cells to make antibodies Produces IL-4, IL-5, IL-6, IL-10, IL-3, GM-CSF, and IL-13 Induce IgG1

Cytotoxic T-cells (CTL)

CD8 Induce apoptosis of target cells Produce IFN-γ, TNF-β, and TNF-α Release cytotoxic proteins

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T-cell growth factor IL-2 and its receptor. Thus, activated T-cells secrete and respond to IL-2 to promote T-cell growth in an autocrine fashion. Cytokines are the messenger molecules of the immune system. Most cytokines have a wide spectrum of effects, reacting with many different cell types, and some are produced by several different cell types. Table 40.7 presents common categories of cytokines and lists some of their general properties. It should be noted that while cytokines can be subdivided into functional groups, many cytokines such as IL-1, TNF-α, and IFN-γ are pleotropic in their effects and regulate, mediate, and activate numerous responses by various cells. Immunotoxicity is any adverse effect on the function or structure of the immune system or other systems as a result of an immune system dysfunction (Langone, 1998). Adverse or immunotoxic effects occur when humoral or cellular immunity needed by the host to defend itself against infections or neoplastic disease (immunosuppression) or unnecessary tissue damage (chronic inflammation, hypersensitivity, or autoimmunity) is compromised. Potential immunological effects and responses that may be associated with one or more of these effects are presented in Table 40.8. Hypersensitivity responses are classified on the basis of the immunological mechanism that mediates the response. There are four types: type I (anaphylactic), type II (cytotoxic), type III (immune complex), and type IV (cell-mediated delayed hypersensitivity). Hypersensitivity is considered to be increased reactivity to an antigen to which a human or animal has been previously exposed, with an adverse rather than a protective effect. Hypersensitivity is a synonym for allergy. Type I (anaphylactic) reactions and type IV (cell-mediated delayed hypersensitivity) reactions are the most common. Types II and III reactions are relatively rare and are less likely to occur with medical devices and biomaterials, however, with tissue-engineered

Table 40.7 Selected cytokines and their effects Cytokine

Effect

IL-1, TNF-α, INF-γ, IL-6 IL-1, TNF-α, IL-6 IL-2, IL-4, IL-5, IL-12, IL-15 and TGF-β IL-2 and IL-4 IL-10 and TGF-β IL-1, INF-γ, TNF-α, and MIF IL-8

Mediate natural immunity Initiate non-specific inflammatory responses Regulate lymphocyte growth, activation, and differentiation Promote lymphocyte growth and differentiation Down-regulate immune responses Activate inflammatory cells Produced by activated macrophages and endothelial cells Chemoattractant for neutrophils Chemoattractant for monocytes and lymphocytes Stimulate hematopoiesis Promote macrophage fusion and foreign body giant cell formation

MCP-1, MIP-α, and RANTES GM-CSF and G-CSF IL-4 and IL-13

Table 40.8 Potential immunological effects and responses Effects

Responses

Hypersensitivity Type I – anaphylactic Type II – cytotoxic Type III – immune complex Type IV – cell-mediated (delayed) Chronic inflammation Immunosuppression Immunostimulation Autoimmunity

Histopathological changes Humoral responses Host resistance Clinical symptoms Cellular responses T-cells NK cells Macrophages Granulocytes

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devices containing potential antigens (i.e. proteins), extracellular matrix (ECM) components, and/or cells, types II and III reactions must be considered in biological response evaluations. Type I (anaphylactic) hypersensitivity reactions are mediated by IgE antibodies which are cytotropic and affect the immediate release of basoactive amines and other mediators from basophils and mast cells followed by recruitment of other inflammatory cells. Type IV cell-mediated (delayed) hypersensitivity responses involve sensitized T lymphocytes that release cytokines and other mediators that lead to cellular and tissue injury. Type IV hypersensitivity (cell-mediated) reactions are initiated by specifically sensitized T lymphocytes. This reaction includes the classic delayed-type hypersensitivity reaction initiated by CD4 T-cells and direct cell cytotoxicity mediated by CD8 T-cells. The less common type II (cytotoxic) hypersensitivity involves the formation and binding IgG and/or IgM to antigens on target cell surfaces that facilitate phagocytosis of the target cell or lysis of the target cell by activated complement components. Type II hypersensitivity (cytotoxic) is mediated by antibodies directed toward antigens present on the surface of cells or other tissue components. Three different antibody-dependent mechanisms may be involved in this type of reaction: complement-dependent reactions, antibody-dependent, cell-mediated cytotoxicity, or antibody-mediated cellular dysfunction. Type III immune complex hypersensitivity is present when circulating antigen–antibody complexes activate complement whose components are chemotactic for neutrophils that release enzymes and other toxic moieties and mediators leading to cellular and tissue injury. Immunological reactions that occur with organ transplant rejection also offer insight into potential immune responses to tissue-engineered devices. Mechanisms involved in organ transplant rejection include T-cell-mediated reactions by direct and indirect pathways and antibody-mediated reactions. Immune responses may be avoided or diminished by using autologous or isogeneic cells in cell/polymer scaffold constructs. The use of allogeneic or xenogenic cells incorporated into the device requires prevention of immune rejection by immune suppression of the host, induction of tolerance in the host, or immunomodulation of the tissue-engineered construct. The development of tissue-engineered constructs by immunoisolation using polymer membranes and the use of non-host cells have been compromised by immune responses. In this concept, a polymer membrane is used to encapsulate non-host cells or tissues thus separating them from the host immune system. However, antigens shed by encapsulated cells were released from the device and initiated immune responses (Brauker, 1992; Brauker et al., 1995; Babensee et al., 1998). Although exceptionally minimal and superficial in its presentation, the previously discussed humoral and cell-mediated immune responses demonstrate the possibility that any known tissue-engineered construct may undergo immunological tissue injury. To date, our understanding of immune mechanisms and their interactions with tissue-engineered constructs is markedly limited. One of the obvious problems is that preliminary studies are generally carried out with non-human tissues and immune reactions result when tissue-engineered constructs from one species are used in testing the device in another species. Ideally, tissue-engineered constructs would be prepared from cells and tissues of a given species and subsequently tested in that species. While this approach does not guarantee that immune responses will not be present, the probability of immune responses in this type of situation is markedly decreased. The following examples provide perspective to these issues. They further demonstrate the detailed and in-depth approach that must be taken to appropriately and adequately evaluate tissue-engineered constructs or devices and their potential adverse responses. The inflammatory response considered to be immunotoxic is persistent chronic inflammation. With biomaterials, controlled release systems and tissue-engineered devices, potential antigens capable of stimulating the immune response may be present and these agents may facilitate a chronic inflammatory response that is of extended duration (weeks, months). Regarding immunotoxicity, it is this persistent chronic inflammation that is of concern as immune granuloma formation and other serious immunological reactions such as autoimmune

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disease may occur. Thus, in biological response evaluation, it is important to discriminate between the shortlived chronic inflammation that is a component of the normal inflammatory and healing responses versus longterm, persistent chronic inflammation that may indicate an adverse immunological response. Immunosuppression may occur when antibody and T-cell responses (adaptive immune response) are inhibited. Potentially significant consequences of this type of response are frequent and serious infections resulting from reduced host defense. Immunostimulation may occur when unintended or inappropriate antigen-specific or non-specific activation of the immune system is present. From a biomaterial and controlled release system perspective, antibody and/or cellular immune responses to a foreign protein may lead to unintended immunogenicity. Enhancement of the immune response to an antigen by a biomaterial with which it is mixed ex vivo or in situ may lead to adjuvancy, which is a form of immunostimulation. This effect must be considered when biodegradable controlled release systems are designed and developed for use as vaccines. Autoimmunity is the immune response to the body’s own constituents, which are considered in this response to be autoantigens. An autoimmune response, indicated by the presence of autoantibodies or T lymphocytes that are reactive with host tissue or cellular antigens may, but not necessarily, result in autoimmune disease with chronic, debilitating and sometimes life-threatening tissue and organ injury. Representative tests for the evaluation of immune responses are given in Table 40.9. Table 40.9 is not all-inclusive and other tests may be applicable. The examples presented in Table 40.9 are only representative of the large number of tests that are currently available (Coligan et al., 1992; Burleson et al., 1995; Smialowicz and Holsapple, 1996; Rose et al., 1997). Table 40.9 is informative but incomplete as in the future direct and indirect markers of immune response may be validated and their predictive value documented thus providing new tests for immunotoxicity. Direct measures of immune system activity by functional assays are the most important types of test for immunotoxicity. Functional assays are generally more important than tests for soluble mediators, which are more important than phenotyping. Signs of illness may be important in in vivo experiments but symptoms may also have a significant role in studies of immune function in clinical trials and postmarket studies. As with any type of test for biological response evaluation, immunotoxicity tests should be valid and have been shown to provide accurate, reproducible results that are indicative of the effect being studied and are useful in a statistical analysis. This implies that appropriate control groups are also included in the study design.

Table 40.9 Representative tests for the evaluation of immune responses Functional assays

Soluble mediators

Skin testing Immunoassays (e.g. ELISA) Lymphocyte proliferation Plaque-forming cells Local lymph node assay Mixed lymphocyte reaction Tumor cytotoxicity Antigen presentation Phagocytosis Degranulation Resistance to bacteria, viruses, and tumors

Antibodies Complement Immune complexes Cytokine patterns (T-cell subsets) Cytokines (IL-1, IL-1ra, TNF-α, IL-6, TGF-β, IL-4, IL-13) Chemokines Basoactive amines

Phenotyping Cell-surface markers MHC markers

Signs of illness Allergy Skin rash Urticaria Edema Lymphadenopathy

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Immunogenicity involving a specific immune response to a biomaterial is an important consideration as it may lead to serious adverse effects. For example, a foreign, non-human, protein may induce IgE antibodies that cause an anaphylactic (type I) hypersensitivity reaction. An example of this type of response is latex protein found in latex gloves. Low molecular weight compounds such as chemical accelerators used in the manufacture of latex gloves may also induce a T-cell-mediated (type IV) reaction resulting in contact dermatitis. Tests for type I (e.g. antigen-specific IgE) and type IV (e.g. guinea pig) maximization tests, hypersensitivity should be considered for materials with the potential to cause these allergic reactions. In addition to hypersensitivity reactions, a device may elicit autoimmune responses (i.e. antibodies or T-cells) that react with the body’s own constituents. An autoimmune response may lead to the pathological consequences of an autoimmune disease. For example, a foreign protein may induce IgG or IgM antibodies that cross-react with a human protein and cause tissue damage by activating the complement system. In a similar fashion, a biomaterial or controlled release system which has a gel or oil constituent may act as an adjuvant leading to the induction of an autoimmune response. Even if an autoimmune response (autoantibodies and/or autoreactive T lymphocytes) is suggested in preclinical testing, it is difficult to obtain convincing evidence that a biomaterial or controlled release system causes autoimmune disease in animals. Therefore, routine testing for induction of autoimmune disease in animal models is not recommended. Babensee and co-workers have tested the hypothesis that the biomaterial component of a medical device, by promoting an inflammatory response can recruit APCs (e.g. macrophages and dendritic cells) and induce their activation, thus acting as an adjuvant in the immune response to foreign antigens originating from the histological component of the device (Babensee et al., 2002; Matzell and Babensee, 2004). Utilizing polystyrene and polylactic-glycolic acid microparticles and polylactic-glycolic scaffolds together with their model antigen, ovalbumin, in a mouse model for 18 weeks, Babensee et al. demonstrated that a persistent humoral immune response that was Th2 helper T-cell dependent, as determined by the IgG1, was present. These findings indicated that activation of CD4 T-cells and the proliferation and isotype switching of B-cells had occurred. A Th1 immune response characterized by the presence of IgG2a was not identified. Moreover, the humoral immune responses for all three types of microparticles were similar indicating that the production of antigen-specific antibodies was not material chemistry-dependent in this model. Babensee suggests that the presence of the biomaterial functions as an adjuvant for initiation and promotion of the immune response and augments the phagocytosis of the antigen with expression of MHC class II and co-stimulatory molecules on APCs with the presentation of antigen to CD4 T-cells. Babensee and co-workers have identified differential levels of dendritic cell maturation on different biomaterials used in combination products (Babensee and Paranjpe, 2005; Bennewitz and Babensee, 2005). The effect of biomaterials on dendritic cell maturation, and the associated adjuvant effect, is a novel biocompatibility selection and design criteria for biomaterials to be used in combination products in which immune consequences are potential complications or outcomes. Badylak and colleagues have carried out extensive studies on the utilization of xenogeneic ECM as a scaffold for tissue reconstruction (Allman et al., 2002; Badylak, 2004). Use of the small intestinal submucosa (SIS) ECM in animals has indicated a restricted Th2-type immune response. The presence of natural antibodies to the terminal galactose-α1,3-galactose (α-gal) epitope is considered to be a major barrier to xenotransplantation in humans. Cell membranes of all animals except those of the humans express this epitope and naturally occurring antibodies mediate hyperacute or delayed rejection of transplanted organs through complement fixation or antibody dependence cell-mediated cytotoxicity. While ECM derived from porcine tissues, SIS, contain small amounts of the gal epitope, it appears that the quantity or distribution of this epitope and/or the subtype of immunoglobulin response to the epitope is such that complement activation does not occur (McPherson et al., 2000). In addition, the resorbable characteristics of this non-chemically cross-linked ECM

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scaffold demonstrate constructive tissue remodeling and deposition of new matrix whereas chemically crosslinked ECM leads to active inflammation and eventually scar formation. The role of Th1 and Th2 lymphocytes in cell-mediated immune responses to xenografts has been examined. Activation of the Th1 pathway leads to macrophage activation, stimulation of complement fixing antibody isotypes, and differentiation of CD8 cells to a cytotoxic type phenotype that is associated with both allogeneic and xenogeneic transplant rejection. The Th2 lymphocyte response does not activate macrophages and leads to production of non-complement fixing antibody isotypes and usually is associated with transplant acceptance. The use of appropriate animal models is an important consideration in the safety evaluation of controlled release systems that may contain potential immunoreactive materials (Greenwald and Diamond, 1988; Cohen and Miller, 1994; Rose, 1997). A recently published study involving the in vivo evaluation of recombinant human growth hormone in poly(lactic-co-glycolic acid) (PLGA) microspheres demonstrates the appropriate use of various animal models to evaluate biological responses and the potential for immunotoxicity. Utilizing biodegradable PLGA microspheres containing recombinant human growth hormone (rhGH), Cleland et al. used rhesus monkeys, transgenic mice expression rhGH and normal control (Balb/C) mice in their in vivo studies (Cleland et al., 1997). Rhesus monkeys were utilized for serum assays in the pharmacokinetic study of rhGH release as well as tissue responses to the injected microcapsule formulation. Placebo injection sites were also utilized and a comparison of the injection sites from rhGH PLGA microspheres and placebo PLGA microspheres demonstrated a normal inflammatory and wound healing response with a normal focal foreign body reaction. To further examine the tissue response, transgenic mice were utilized to assess the immunogenicity of the rhGH PLGA formulation. Transgenic mice expressing a heterologous protein have been previously used for assessing the immunogenicity of sequence or structural mutant proteins (Stewart et al., 1989; Stewart, 1993). With the transgenic animals, no detectable antibody response to rhGH was found. In contrast, the Balb/C control mice had a rapid onset of high titer antibody response to the rhGH PLGA formulation. This study points out the appropriate utilization of animal models to not only evaluate biological responses but also one type of immunotoxicity (immunogenicity) of controlled release systems.

SUMMARY Tissue-engineered devices are biologic–biomaterial combinations in which some component of tissue has been combined with a biomaterial to create a device for the restoration or modification of tissue or organ function. The biocompatibility and bioresponse requires the ultimate achievement of four significant goals if these devices are to function adequately and appropriately in the host environment. These goals are: (1) restoration of the target tissue with its appropriate function and cellular phenotypic expression; (2) inhibition of the macrophage and FBGC foreign body response that may degrade or adversely modify device function; (3) inhibition of scar and fibrous capsule formation that may be deleterious to the function of the device; and (4) inhibition of immune responses that may inhibit the proposed function of the device and ultimately lead to the destruction of the tissue component of the tissue-engineered device. This chapter has presented a brief and limited overview of mechanisms and biological responses that determine biocompatibility: inflammation, wound healing, and immunotoxicity. Given the unique nature of the combination of tissue component and biomaterial in tissue-engineered devices, coupled with the species differences in biological responses, a significant future challenge in the development of tissue-engineered devices is the construction and utilization of a unique set of tests that will ensure that the four goals indicated above are achieved for the lifetime of the device in its in vivo environment in humans.

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41 Essential Elements of Wound Healing William J. Lindblad INTRODUCTION Wound healing represents a primary survival process for all multi-cellular organisms, and involves the replacement of damaged tissue with connective tissue. The process of repairing damaged tissue by deposition of connective tissue provides for a rapid repair, but in general it fails to return the original function to the tissue. It is interesting to consider that as organisms of greater complexity evolved, the mechanism of wound healing changed from one of regeneration to that of repair. Thus, fitness of the organism appeared to align with rapidity of repair rather than the benefits of a reconstituted tissue. Unfortunately, this emphasis on repair by scar formation can lead to deleterious outcomes, such as adhesions, keloids, and hypertrophic scars that are not a concern for organisms that have not evolved to the level of complexity of human beings. The process of tissue regeneration in many respects would represent a preferred outcome to injury and it is possible in many invertebrates and some tissues, even in mammals. For example, in humans the liver undergoes significant regeneration following acute injury, but even here, the organ will undergo fibrosis given chronic injury. The classic pictures of salamanders regenerating entire limbs following amputation (so-called epimorphic regeneration) have kindled the imagination of researchers interested in the translation of those results to humans. Clearly, if one could restore not only the volume of damaged tissue, but also its original function, the result of traumatic injury would be inconsequential. Thus, much research in the field has recently focused on attempting to understand why human beings do not invoke a regenerative response and rather activate a repair response. Much of this work has been performed in fetal wound healing models which will briefly be discussed. Our understanding of repair has undergone major advances over the past 20 years, particularly with the realization that stem cells may actively participate in the process to a far greater extent than earlier appreciated. This chapter will review our current knowledge of tissue repair by focusing on the repair of dermal lesions. By necessity, this will highlight specific topics rather than provide in-depth coverage, but hopefully the reader will obtain insights into the essential biological processes required to repair damage at the tissue level. REGULATION OF TISSUE HEALING One area of understanding wound healing that has undergone extensive growth over the past 25–30 years is that of control/regulation of the individual cellular and biochemical events. We now know that an impressive array of soluble, insoluble, and gaseous mediators is able to control cell behavior and thereby respond to injury in a rapid, concerted fashion (Table 41.1). These multiple interacting factors, sometimes with apparent overlapping functions, appear to form a network of redundant mechanisms that ensures the process will go forward to completion despite loss of function in any one system.

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Table 41.1 Partial list of well-studied regulatory signals for controlling wound healing. Note the wide variety of mediator types and biochemical forms Physical state

Biochemical class

Examples

Soluble mediators

Growth factors Prostanoids Peptides Cytokines Metabolic products Redox potential Matrix proteins

CTGF, EGF, FGF, PDGF, TGF-β, VEGF TxA2, LTB4 Collagen fragments MIP-1 Lactate

Insoluble signals Gaseous mediators

Thrombospondin, tenascin, laminin O2, NO

It is clear that regulation of healing is based on a series of sequentially triggered responses following injury to the tissue. For example, upon damage to the blood compartment, platelets are activated and release a series of bioactive mediators either from preformed cytoplasmic granules (e.g. platelet-derived growth factor (PDGF) and transforming growth factor-β (TGF-β)) or by de novo synthesis at the lipid cell membrane (e.g. thromboxane A2 (TxA2) and leukotriene B4 (LTB4)). These mediators are then used as stimuli to promote the influx of cell populations and/or activation of resident cells (Diegelmann and Evans, 2004). Less clear, are the signals to terminate the various repair processes, which may be as critical to obtaining an optimally repaired tissue as those for initial activation and recruitment. It is now well established that the formation of dermal ulcers includes the prolonged entry and residence of inflammatory cells (Mast and Schultz, 1996). The question then becomes is the abnormal prolongation of the inflammatory response one of continual recruitment or one of impaired termination of the inflammatory process. The signals for termination appear to be one of triggering neutrophil apoptosis which are clearly different from those of recruitment (Brown et al., 1997). Similar arguments can be invoked for explaining other pathological healing responses resulting from excessive deposition of connective tissue (keloids) and excessive wound contraction (contractures). Therefore, the challenge in the future is to understand not just those mediators that activate the repair pathways, but to understand how the termination signals interplay with the activation signals. It is hoped that studies employing various microarray techniques will provide insights into these regulatory processes. In addition to the interplay of soluble mediators with cells that re-populate a damaged tissue, there are significant tissue–tissue interactions, of particular note – mesenchymal–epithelial interactions (Yamaguchi et al., 2005). It is clear that the cells of newly reforming epithelium express regulatory proteins that influence the behavior of mesenchymal cells in the subjacent dermis. For example, fibroblasts are able to express hepatocyte growth factor/scatter factor that is able to influence the behavior of melanocytes and keratinocytes in the overlying epithelium (Imokawa, 2004). Therefore, regulation of healing should not be considered simply an interplay between transient cell populations and mediators, or resident cells and transient cells, but also between cells that are resident in the final reconstituted tissue.

INFLAMMATION Tissue damage is a potent stimulus for the inflammatory system, resulting in an initial vascular response that also includes initiating events for the multiple interacting events of inflammation. Damage to vascular structures exposes the blood compartment to sub-endothelial collagen which serves as a binding ligand for the circulating protein von Willebrand factor, which also possesses a binding site for the GpIb expressed on the surface of

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platelets (Nieswandt and Watson, 2003). Binding of the platelet through von Willebrand factor to the vessel wall serves as an activation signal, leading to multiple changes in the platelet including cell surface ruffling; de novo synthesis and release of platelet activating factor (PAF); the prostanoids, TxA2 and LTB4; and release of the growth factors PDGF and TGF-β1 from α-granules (Gosain and Gamelli, 2005; Tettamanti et al., 2006). All of these mediators perform specific functions that promulgate the inflammatory response to the initial tissue damage. Of particular note are the mediators LTB4, PDGF, and TGF-β1 that all serve as chemotactic mediators for inflammatory cells and connective tissue cells (Gillitzer and Goebeler, 2001; Tettamanti et al., 2006). Upon the activation of platelets and release of mediators, the vascular endothelium is activated by among other mediators PAF and tumor necrosis factor-α (TNF-α) with the synthesis and cell surface expression of selectins and other adhesion molecules that enable circulating neutrophils and monocytes to localize to the site of tissue injury (Muller, 2003). The translocation of P-selectin to the endothelial cell luminal surface and expression of E-selectin by the endothelial cells allows circulating neutrophils to stick and roll along the activated endothelium. Once these cells have bound to the endothelial cell surface by the moderately strong selectin interactions, the cells become more tightly affixed to the endothelium by interaction with other classes of adhesion molecules (Muller, 2003). The cells access the extravascular compartment by diapodesis, a process that involves not only the active participation of the inflammatory cells, but also the activated endothelial cells because the latter cells need to loosen their intercellular junctions to allow neutrophil passage (Martin and Leibovich, 2005). Once in the tissue, the neutrophils are able to phagocytize devitalized tissue and any infectious agent introduced into the wound. Although this movement of neutrophils from the vascular compartment into the extracellular space is essential for proper healing, a prolonged or excessive accumulation of neutrophils may lead to extensive breakdown of the tissue potentially leading to the formation of a chronic non-healing wound (Mast and Schultz, 1996). The neutrophil uses the generation of reactive oxygen species (ROS) to oxidize biological membranes and promote the clearance of damaged tissues (Sen, 2003). Because of this generation of ROS, the neutrophil has a large requirement for O2 which largely determines the need for O2 during the early phase of wound healing (Albina and Reichner, 2003). Subsequent to the active recruitment of neutrophils to a wound site, monocytes are actively recruited to the wound site by a number of chemotactic proteins including TGF-β and monocyte chemotactic protein-1 (MCP-1; Wahl et al., 1987). Once at the wound site these cells become activated to highly synthetic macrophages, cells that are able to express a large array of proteins that modulate the repair response and start the process of reconstitution of the tissues. Early studies strongly suggested that the monocyte/macrophage was central to wound healing (Leibovich and Ross, 1975) however, recently studies using knock-out mice suggest that the absolute need of macrophages for wound healing may have been overstated (Martin et al., 2003). In addition to the generation of ROS, the monocyte/macrophage is able to generate large amounts of NO at levels sufficient to be cytotoxic (Schwentker and Billiar, 2003). This level of NO is the result of enhanced expression of the inducible nitric oxide synthetase (iNOS) gene which has been shown to be essential for normal healing through the use of iNOS knock-out mice which show an abnormal healing phenotype (Yamasaki et al., 1998). As with many other facets of wound healing, the production of NO is under the control of inflammatory mediators, such that the levels of NO are dependent on the overall inflammatory response as noted in tumor necrosis-α (TNF-α) and interferon-γ (IFN-γ) levels (Schaffer et al., 2006). This also suggests that NO levels may contribute to the dysregulation of healing seen when TNF-α levels remain high after injury (Mast and Schultz, 1996).

FIBROPLASIA Following the influx and subsequent demise of neutrophils by apoptosis, and the arrival of the biosynthetically active monocyte/macrophages, the wounded tissue undergoes a conversion of the damaged area from

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one of acute inflammation and tissue destruction, into one of nascent tissue synthesis and deposition. In addition to the proteins expressed above, macrophages are activated to express a number of genes encoding for extracellular matrix (ECM) proteins, for example, fibronectin (Liptay et al., 1993). Additionally, a selected subset of the cells are able to develop a type I collagen-synthetic phenotype that may help to assemble the initial scaffold onto which mesenchymal cells migrate and vascular endothelial cells use to begin the neovascularization process (Lindblad, 2004). The ultimate ECM formed to replace damaged tissue represents the end stage of an orchestrated series of transient ECMs that finally develop into the extensively cross-linked type I collagen rich ECM of a mature scar. The process is initiated with the deposition of the polymerized fibrin blood clot used to prevent blood loss during coagulation (Laurens et al., 2006). Physical (fiber thickness and matrix porosity) and biochemical characteristics of the fibrin network that forms have a pronounced effect on the ECM that is subsequently assembled upon this basic matrix. Fibronectin coats the fibrin matrix and this serves as a transient or provisional ECM that allows formed blood cells to attach, migrate, become activated and participate in the initial acute inflammatory reaction. The fibrin clot is stabilized by cross-linking of the fibrin via the action of plasma-derived transglutaminases (Inbal and Dardik, 2006). Another set of enzymes of the transglutaminases family, the tissue transglutaminases have been implicated in normal wound healing (Telci and Griffin, 2006). These enzymes, by cross-linking proteins via ε (γ-glutamyl) lysine bridges, influence the biophysical characteristics of ECM proteins, which in turn alters the ability of cells to interact with these structural matrices. Once the blood clot has been cross-linked it is slowly degraded and replaced with a collagen enriched structure incorporating cell attachment proteins such as fibronectin (Hynes, 1990), thrombospondin (Reed et al., 1993), and tenascin (Erickson and Bourdon, 1989). The incorporation of tenascin is essential for the correct attachment and activation of mesenchymal cells that are actively attracted to the area under the chemoattractant influence of PDGF and TGF-β (Whitby and Ferguson, 1991). Cells exposed to the tenascin epitopes bind through an integrin-dependent receptor (αVβ3), as well as non-integrin sites (Prieto et al., 1992), with these interactions particularly notable in fetal wounds suggesting the possible importance of the interaction for a fetal-type repair response. This provisional matrix is a relatively open molecular construct that allows for migration of cells on top (epithelial cells) and through (endothelial and various mesenchymal cells) the matrix to form the permanent ECM and vascular structures. Cells originating from the surrounding intact ECM migrate into the provisional matrix in response to multiple chemoattractant molecules produced in the wound site. These molecules include soluble growth factors (PDGF and TGF-β), collagen fragments, and lipid-based mediators (Gillitzer and Goebeler, 2001; Tettamanti et al., 2006). Therefore, multiple apparently redundant signals trigger and promote this cell movement (Figure 41.1). Deposition of the final type I collagen rich ECM occurs by fibroblasts that have re-populated the damaged and replaced matrix. Although fibroblasts microscopically have few distinguishing features; biochemically it has been shown that there are multiple populations of fibroblasts within the dermis (Bordin et al., 1984; Chang et al., 2002). It is possible that the connective tissue response is based on the population of fibroblasts present at the site and whether a particular population is selected based on this mediator milieu. It is also possible that pathologies of healing result from an abnormal expansion of one of these populations leading to an abnormal amount of ECM expressed by this population of fibroblasts. Additional research has questioned the origin of the fibroblastic cells within the granulating wound bed, that is, do these reflect resident cells or recruited cells. We have shown that circulating blood cells are able to upregulate the expression of prolyl hydroxylase and secretion of collagenous proteins within 24–48 h of isolation (Lindblad et al., 1987). Additionally, pure populations of macrophages have been shown to synthesize and deposit collagenous matrices (Vaage and Lindblad, 1990). These data suggest that circulating cells, specifically

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Collagen content

Cellularity

Initial fibrin clot

Cross-linked clot coated with attachment proteins

Dense collagenous connective tissue

Figure 41.1 Dynamic changes in wound extracellular matrix during healing. The initial provisional matrix is a well-hydrated, open structure coated with attachment proteins and containing hyaluronic acid and other glycosaminoglycans. With maturation of the granulation tissue to the final scar, the matrix becomes dense, relatively anhydrous with large bundles of type I collagen.

those of the monocyte/macrophage lineage, are able to contribute to the deposition of the neomatrix in granulation tissue. Lastly, there has been extensive work with a circulating cell population termed fibrocytes suggesting that another population of cell is also able contribute to the ECM of healing wounds (Bucala et al., 1994; Yang et al., 2002). However, this contribution may primarily contribute to pathological healing as these cells appear to be enriched in hypertrophic burn scars (Wang et al., 2007). Another difficulty with invoking these cells in normal healing is the long time delay before this phenotype of cell is found in purified whole blood cells.

NEOVASCULARIZATION Formation of a new vasculature is essential to provide a well-nourished, stable tissue after injury, and forms simultaneous with the deposition of ECM. Theoretically, this formation of a new blood supply could occur either by the process of angiogenesis or by vasculogenesis. Angiogenesis refers to the formation of new blood vessels from pre-existing ones by outgrowth of capillary buds and sprouts, whereas vasculogenesis refers to the formation of new blood vessels in the absence of pre-existing blood vessels by recruitment and differentiation of endothelial precursor cells (Hristov and Weber, 2004). It has generally been assumed that the neovascularization of damaged tissue occurred by angiogenesis alone, but recent studies suggest that neovascularization is a combination of angiogenic and vasculogenic processes (Montesinos et al., 2004). The angiogenic response reflects a concerted series of steps including the development of a highly branched network of vessels, elimination of areas of the network by closure of selected vessel lumens, and subsequent degeneration of the vessel distal to the obliterated vessel lumen (Madri et al., 1996). These processes can be further divided into the processes of endothelial cell activation (and recruitment), migration, cellular proliferation, tube formation and stabilization and tube regression, remodeling and involution. Control of this angiogenic response is multifactorial including the involvement of soluble mediators (e.g. vascular endothelial growth factor (VEGF)); matrix proteins (e.g. thrombospondin); proteases (e.g. MMPs), and levels of oxygen (e.g. HIF). These different control mechanisms while listed separately really represent overlapping, interconnected processes. For example, the matrix metalloproteinases and related ADAM (a disintegrin and metalloprotease

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domain) proteases can regulate the angiogenic process by modulating the basement membrane upon which the endothelial cells will use to form tubules, by releasing growth factors stored within the ECM and by proteolytically altering the growth factor receptors required for cellular activity (Roy et al., 2006).

RE-EPITHELIALIZATION Tissues in which an epithelium is in contact with the stroma will require the reformation of a stratified epithelia following injury. This process is of particular importance as it is the cornified epithelium of the skin that constitutes a physical barrier between the outer and inner organism’s environments. Without this barrier function, water loss increases dramatically and the underlying tissue becomes a ready nutrient source for microbial growth. Therefore, the process of reforming an intact epithelium occurs rapidly with cellular changes occurring within a few hours of injury. Stem cells have been shown to contribute to re-epithelialization and renewal of hair follicles, which was one of the first demonstrations that activation of a stem cell population was essential for tissue repair (Oshima et al., 2001). It was of note that the stem cell niche for these stem cells was in the bulge region of hair follicles and not in the basal epithelium. However, others have shown that there may be a different stem cell population located in the interfollicular region (Watt, 2001). However, the bulb-region location is consistent with the observation that re-epithelialization requires intact hair follicles, and in cases where these dermal structures are eliminated by, for example burn injury, the reformation of the epithelium is delayed and occurs strictly by in migration from the surrounding intact dermis. Following injury, the cells adjacent to the denuded area become activated and will start to migrate into the wound defect by extending long processes or lamellipodia to attach and draw the rest of the cell toward the anchored projection. The cells secrete MMP-type proteases and re-synthesize a matrix on their basal cell surface that facilitates their migration over the damaged tissue (Pilcher et al., 1997). It is clear that the cells are able to migrate only in the presence of an appropriate substratum. Therefore, conditions that alter this environment will slow or stop the migration of cells and thereby inhibit the healing process. The process of covering the denuded area initially does not involve cell proliferation at the migrating edge, but cells are required to “fill-in” behind the front. Once the wound surface is covered, the cells continue to proliferate and begin to orient and develop into multilayered structures associated with the final keratinized epithelium (Singer and McClain, 2002). Clinically, impairment of the re-epithelialization process may result from a number of conditions including infection (Singer and McClain, 2002). WOUND CONTRACTION Formation of granulation tissue represents the hallmark of new tissue growth to replace devitalized human tissue. However, in addition to the formation of granulation tissue, closure of the damaged tissue may occur by wound contraction. This process represents the active movement of the surrounding intact tissue over the area of damage to rapidly close the open defect. This observation was reported over 80 years ago by Alexis Carrel; however, the mechanism for this process is still controversial, although clarity of the process has come over the past 10–15 years. In general, and in other species of animals in which the dermis is not well affixed to the sub-cutis, wound contraction may represent over 60% of the mechanism for covering an open defect wound. However in humans, who generally have a relatively tight connection between the dermis and sub-cutis, contraction is not a primary mechanism for normal wound healing, but may be the cause of major clinical problems. Burn wounds that occur over joints of the extremities are particularly prone to contracture formation which reflects extensive migration of tissue over the affected joint limiting movement of the joint.

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The cellular component of contraction is the myofibroblast, a cell originally described by Gabbiani et al. (1971) which represents a specialized phenotype of fibroblasts. These cells are able to produce their contractile force by expression of α-smooth muscle actin (Hinz et al., 2001) a protein that is up-regulated in the fibroblasticlike cells following application of mechanical stress and inflammatory mediators (Desmoulière et al., 2005). The myofibroblast is present in many fibrotic lesions leading to the hypothesis that the (trans) differentiation of fibroblast sub-populations into myofibroblasts may represent a major transition from normal healing to fibrosis (Desmoulière et al., 2005). Currently, different compounds are being examined for their ability to alter the extent of this fibroblast differentiation as novel anti-fibrotic agents (Sheffer et al., 2007). As noted above, the mechanism of contraction has been controversial over the years, with several different competing mechanisms proposed (Rudolph et al., 1992). However, the myofibroblast-based model appears to be gaining acceptance within the wound healing research community, although the model does not explain all facets of the contracting wound state (Ehrlich et al., 2006).

REGENERATIVE DERMAL HEALING: FETAL HEALING It has been understood for many years that injuries to an embryo and early gestation fetus will result in a regenerative response, rather than a repair phenomena, if the fetus was not too advanced in gestational age (Whitby and Ferguson, 1991). Therefore, numerous investigations were conducted starting in the 1980s and 1990s in an attempt to identify those factors in the fetal environment that facilitate the regenerative phenotype of healing versus repair phenomena of healing (Ferguson and Kane, 2004). Investigators examined a number of different potential factors that could impart a regenerative response onto tissue injury in the fetus. The presence of amniotic fluid was studied extensively as it is clearly a factor present with the fetus but not with the adult. Using a variety of approaches it was shown that the ability to regenerate was inherent in the tissue and not in the amniotic environment (Longaker et al., 1994). Another area of extensive research was that of the inflammatory response. In the embryo and early gestation fetus the inflammatory response is blunted and therefore the release of the numerous regulatory mediators diminished as well. These findings, and others, lead to the examination of the multitude of growth factors and other soluble mediators present in the wound environment and particularly which factors are uniquely present in the fetal regenerative tissue versus those seen in adult repair tissue. Focusing on the TGF-β family of growth factors, Ferguson et al. demonstrated that the expression of the TGF-β3 isoform is associated with a regenerative-type response (Shah et al., 2000). The influence of the re-epithelialization process on embryonic healing highlights another difference in the healing response. Whereas the adult re-epithelializes by migration of a sheet of epithelial cells, the embryo uses an actin dependent purse-string contraction mechanism (Redd et al., 2004).

CONCLUSIONS The process of wound healing represents a sophisticated series of cellular events that encompass numerous cell types from resident tissues and cells recruited from the circulation and unique niches. We currently have a broad understanding of the manner in which these different cells react to an injury state but overall regulation of these events is not complete. In particular, we still have only a rudimentary understanding of termination signaling and how deficits in these signals could lead to pathological healing. Similarly, interaction between cells of the mesenchymal and epithelium layers is only now being elucidated. Clearly, our understanding has advanced significantly over the past 2–3 decades, but additional challenges to our understanding of normal healing remain. However, these advances now make the possibility of a regenerative response rather

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than a repair response within the realm of possibility. Continued work is needed to further elucidate how triggering a regenerative response could be accomplished.

REFERENCES Albina, J.E. and Reichner, J.S. (2003). Oxygen and the regulation of gene expression in wounds. Wound Repair Regen. 11: 445–451. Bordin, S., Page, R.C. and Narayanan, A.S. (1984). Heterogeneity of normal human diploid fibroblasts: isolation and characterization of one phenotype. Science 223: 171–173. Brown, D.L., Kao, W.W.Y. and Greenhalgh, D.G. (1997). Apoptosis down-regulates inflammation under the advancing epithelial wound edge: delayed patterns in diabetes and improvement with topical growth factors. Surgery 121: 372–380. Bucala, R., Spiegel, L.A., Chesney, J., Hogan, M. and Cerami, A. (1994). Circulating fibrocytes define a new leukocyte subpopulation that mediates tissue repair. Mol. Med. 1: 71–81. Chang, H.Y., Chi, J.T., Dudoit, S., Bondre, C., van de Rijn, M., Botstein, D. and Brown, P.O. (2002). Diversity, topographic differentiation, and positional memory in human fibroblasts. Proc. Natl Acad. Sci. USA. 99: 12877–12882. Desmoulière, A., Chaponnier, C. and Gabbiani, G. (2005). Tissue repair, contraction, and the myofibroblast. Wound Repair. Regen. 13: 7–12. Diegelmann, R.F. and Evans, M.C. (2004). Wound healing: an overview of acute, fibrotic and delayed healing. Front. Biosci. 9: 283–289. Ehrlich, H.P., Allison, G.M. and Leggett, M. (2006). The myofibroblast, cadherin, alpha smooth muscle actin and the collagen effect. Cell Biochem. Funct. 24: 63–70. Erickson, H.P. and Bourdon, M.A. (1989). Tenascin: an extracellular matrix protein prominent in specialized embryonic tissues and tumors. Annu. Rev. Cell Biol. 5: 71–92. Ferguson, M.W.J. and Kane, O. (2004). Scar-free healing: from embryonic mechanisms to adult therapeutic intervention. Phil. Trans. Roy. Soc. Lond. B 359: 839–850. Gabbiani, G., Ryan, G.B. and Majno, G. (1971). Presence of modified fibroblasts in granulation tissue and their possible role in wound contraction. Experientia 27: 549–550. Gillitzer, R. and Goebeler, M. (2001). Chemokines in cutaneous wound healing. J. Leukoc. Biol. 69: 513–521. Gosain, A. and Gamelli, R.L. (2005). A primer in cytokines. J. Burn Care Rehabil. 26: 7–12. Hinz, B., Celetta, G., Tomasek, J.J., Gabbiani, G. and Chaponnier, C. (2001). α-smooth muscle actin expression upregulates fibroblast contractile activity. Mol. Biol. Cell 12: 2730–2741. Hristov, M. and Weber, C. (2004). Endothelial progenitor cells: characterization, pathophysiology, and possible clinical relevance. J. Cell. Mol. Med. 8: 498–508. Hynes, R.O. (1990). “Fibronectins.” New York: Springer-Verlag. Imokawa, G. (2004). Autocrine and paracrine regulation of melanocytes in human skin and in pigmentary disorders. Pigm. Cell Res. 17: 96–110. Inbal, A. and Dardik, R. (2006). Pathophysiol. Haemost. Thromb. Role of coagulation factor XIII (FXIII) in angiogenesis and tissue repair. 35: 162–165. Laurens, N., Koolwijk, P. and de Maat, M.P. (2006). Fibrin structure and wound healing. J. Thromb. Haemost. 4: 932–939. Lindblad, W.J. (2004). Stem cells in dermal wound healing. In: Sell, S. (ed.), Stem Cells Handbook. Totowa, NJ: Humana Press, pp. 101–105. Lindblad, W.J., French, J.A., Redford, K.S., Buenaventura, S.K. and Cohen, I.K. (1987). Induction of prolyl hydroxylase activity in a nonadherent population of human leukocytes. Biochem. Biophys. Res. Comm. 147: 486–493. Liptay, M.J., Parks, W.C., Mecham, R.P., Roby, J., Kaiser, L.R., Cooper, J.D. and Botney, M.D. (1993). Neointimal macrophages colocalize with extracellular matrix gene expression in human atherosclerotic pulmonary arteries. J. Clin. Invest. 91: 588–594. Longaker, M.T., Whitby, D.J., Ferguson, M.W.J., Jennings, R.W., Lorenz, H.P., Harrison, M.R. and Adzick, N.S. (1994). Adult skin wounds in the fetal environment heal with scar formation. Ann. Surg. 219: 65–72. Madri, J.A., Sankar, S. and Romanic, A.M. (1996). Angiogensis. In: Clark, R.A.F. (ed.), The Molecular and Cellular Biology of Wound Repair. New York: Plenum Press, pp. 355–371.

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Martin, P. and Leibovich, S.J. (2005). Inflammatory cells during wound repair: the good, the bad and the ugly. Trends Cell Biol. 15: 599–607. Martin, P., D’Souza, D., Martin, J., Grose, R., Cooper, L., Maki, R. and McKercher, S.R. (2003). Wound healing in the PU.1 null mouse – tissue repair is not dependent on inflammatory cells. Curr. Biol. 13: 1122–1128. Mast, B.A. and Schultz, G.S. (1996). Interactions of cytokines, growth factors, and proteases in acute and chronic wounds. Wound Repair. Regen. 4: 411–420. Montesinos, M.C., Shaw, J.P., Yee, H., Shamamian, P. and Cronstein, B.N. (2004). Adenosine A2A receptor activation promotes wound neovascularization by stimulating angiogenesis and vasculogenesis. Am. J. Pathol. 164: 1887–1892. Muller, W.A. (2003). Leukocyte–endothelialcell interactions in leukocyte transmigration and the inflammatory response. Trends Immunol. 24: 327–334. Nieswandt, B. and Watson, S.P. (2003). Platelet–collagen interaction: is GPVI the central receptor? Blood 102: 449–461. Oshima, H., Rochat, A., Kedzia, C., Kobayashi, K. and Barrandon, Y. (2001). Morphogenesis and renewal of hair follicles from adult multipotent stem cells. Cell 104: 233–245. Pilcher, B.K., Dumin, J.A. Sudbeck, B.D., Krane, S.M., Welgus, H.G. and Parks, W.C. (1997). The activity of collagenase-1 is required for keratinocyte migration on a type I collagen matrix. J. Cell Biol. 137: 1445–1457. Prieto, A.L., Andersson-Fisone, C. and Crossin, K.L. (1992). Characterization of multiple adhesive and counteradhesive domains in the extracellular matrix protein cytotactin. J. Cell Biol. 119: 663–678. Reed, M.J., Puolakkainen, P., Lane, T.F., Dickerson, D., Bornstein, P. and Sage, E.H. (1993). Differential expression of SPARC and thrombospondin 1 in wound repair: Immunolocalization and in situ hybridization. J. Histochem. Cytochem. 41: 1467–1477. Redd, M.J., Cooper, L., Wood, W., Stramer, B. and Martin, P. (2004). Wound healing and inflammation: embryos reveal the way to perfect repair. Phil. Trans. Roy. Soc. Lond. B 359: 777–784. Roy, R., Zhang, B. and Moses, M.A. (2006). Making the cut: protease-mediated regulation of angiogenesis. Exp. Cell Res. 312: 608–622. Rudolph, R., Vande Berg, J. and Ehrlich, H.P. (1992). Wound contraction and scar contracture. In: Cohen, I.K., Diegelmann, R.F. and Lindblad, W.J. (eds.), Wound Healing: Biochemical and Clinical Aspects. Philadelphia: W.B. Saunders, pp. 96–114. Schaffer, M., Bongartz, M., Hoffmann, W. and Viebahn, R. (2006). Regulation of nitric oxide synthesis in wounds in IFNgamma depends on TNF-alpha. J. Investig. Surg. 19: 371–379. Schwentker, A. and Billiar, T.R. (2003). Nitric oxide and wound repair. Surg. Clin. North Am. 83: 521–530. Sen, C.K. (2003). The general case for redox control of wound repair. Wound Repair. Regen. 11: 431–438. Shah, M., Rorison, P. and Ferguson, M.W.J. (2000). The role of transforming growth factors beta in cutaneous scarring. In: Garg, H.G. and Longaker, M.T. (eds.), Scarless Wound Healing. New York: Marcel Dekker Inc., pp. 213–226. Sheffer, Y., Leon, O., Pinthus, J.H., Nagler, A., Mor, Y., Genin, O., Iluz, M., Kawada, N., Yoshizato, K. and Pines, M. (2007). Inhibition of fibroblast to myofibroblast transition by halofuginone contributes to the chemotherapy-mediated antitumoral effect. Mol. Cancer Ther. 6: 570–577. Singer, A.J. and McClain, S.A. (2002). Persistent wound infection delays epidermal maturation and increases scarring in thermal burns. Wound Repair. Regen. 10: 372–377. Telci, D. and Griffin, M. (2006). Tissue transglutaminase (TG-2) – a wound response enzyme. Front. Biosci. 11: 867–882. Tettamanti, G., Malagoli, D., Benelli, R., Albini, A., Grimaldi, A., Perletti, G., Noonan, D.M., de Eguileor, M. and Ottaviani, E. (2006). Growth factors and chemokines: a comparative functional approach between invertebrates and vertebrates. Curr. Med. Chem. 13: 2737–2750. Vaage, J., and Lindblad, W.J. (1990). Production of collagen type I by mouse peritoneal macrophages. J. Leukoc. Biol. 57: 2–10. Wahl, S.M., Hunt, D.A., Wakefield, L.M., McCartney-Francis, N., Wahl, L.M. Roberts, A.B. and Sporn, M.B. (1987). Transforming growth factor type β induces monocyte chemotaxis and growth factor production. Proc. Natl Acad. Sci. USA 84: 5788–5792. Wang, J.F., Jiao, H., Stewart, T.L., Shankowsky, H.A., Scott, P.G. and Tredget, E.E. (2007). Fibrocytes from burn patients regulate the activities of fibroblasts. Wound Repair. Regen. 15: 113–121. Watt, F.M. (2001). Stem cell fate and patterning in mammalian epidermis. Curr. Opin. Genet. Dev. 11: 410–417.

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Whitby, D.J. and Ferguson, M.W.J. (1991). The extracellular matrix of lip wounds in fetal, neonatal and adult mice. Development 112: 651–668. Yamaguchi, Y., Hearing V.J., Itami, S., Yoshikawa, K. and Katayama, I. (2005). Mesenchymal–epithelial interactions in the skin: aiming for site-specific tissue regeneration. J. Dermatol. Sci. 40: 1–9. Yamasaki, K., Edington, H.D., McClosky, C., Tzeng, E., Lizonova, A., Kovesdi, I., Steed, D.L. and Billiar, T.R. (1998). Reversal of impaired wound repair in iNOS-deficient mice by topical adenoviral-mediated iNOS gene transfer. J. Clin. Invest. 101: 967–971. Yang, L., Scott, P.G., Giuffre, J., Shankowsky, H.A., Ghahary, A. and Tredget, E.E. (2002). Peripheral blood fibrocytes from burn patients: identification and quantification of fibrocytes in adherent cells cultured from peripheral blood mononuclear cells. Lab. Invest. 82: 1183–1192.

42 Proteins Controlled with Precision at Organic, Polymeric, and Biopolymer Interfaces for Tissue Engineering and Regenerative Medicine Buddy D. Ratner

INTRODUCTION The physiologic environment is a “biological broth” comprised of thousands of different proteins, lipids, sugars, ions, and water. Normal wound healing and tissue regeneration exploit selected molecules from the biomolecule pool in this biological milieu to direct events such as cell attachment, growth, extracellular matrix (ECM) formation, and phenotypic differentiation. For tissue engineering, these same processes must be mimicked to encourage tissue development and regeneration within artificial scaffolds, gels, and other tissue engineering strategies. Nature does not offer a large number of alternative strategies to achieve the goal of reconstruction or regeneration. Evolution and the requirement for fitness in the environment have generally led to conserved pathways and these pathways, typically highly biospecific in nature, must be appropriately followed. This chapter addresses methods we might use to control proteins and other biomolecules with precision for tissue engineering scaffolds and other biomaterials used in tissue engineering. The concept of “precision control” is schematically illustrated in Figure 42.1. The active or recognition site on the protein is illustrated with a star in the figure – if this star is buried within the protein film, it cannot participate in guiding biospecific reaction. The goals of precision control of biomolecules at the biology–material interface are appropriate cell attachment, cell proliferation, reduced inflammation, angiogenesis, ECM production, and ultimately tissue regeneration. This has been referred to as an “instructive scaffold.” There is already a substantial body of literature on attaching proteins and other biomolecules to scaffolds, gels, and surfaces to direct specific aspects of tissue formation. A few examples are cited here. (Hern and Hubbell, 1998; Stile and Healy, 2001; Chua et al., 2005; Zhu et al., 2006). This chapter has four sections. First I will consider the case where proteins and biomolecules are used non-specifically – this is the case for most tissue engineering studies today. Non-specific use of proteins is never observed in normal physiology, but is widely observed in cell culture, biomaterials, and tissue engineering. Next, inhibition of protein adsorption will be addressed. Then this chapter will review a series of methods

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Native protein ( receptor site)

(a)

Scaffold fibril (b)

Scaffold fibril (c)

non-fouling region

non-fouling region

Scaffold fibril

Figure 42.1 Biomolecule signals in tissue engineering scaffolds are most often delivered in a non-specific fashion where the receptor site (star) may be buried or inaccessible (a); to better emulate the biology of normal tissue reconstruction, the hypothesis is made that there should be an advantage in delivering signals with precision (b); and possibly patterning this signal delivery to control surface density of sites (c).

that can be used to deliver biomolecule signals with precision, emulating the way these signals are delivered in vivo. Finally, implications and perspectives will be presented.

NON-SPECIFIC PROTEIN ADSORPTION Almost all synthetic materials adsorb a monolayer coating of proteins seconds after being placed in an in vivo environment. This monolayer consists of a mixture of proteins (human plasma may contain 700 or more proteins), and these proteins are oriented randomly (“up, down, sideways”) on the surface and can be in their native conformation, denatured, or partially unfolded. Non-specific protein adsorption has been widely studied and is important for phenomena such as cell attachment, cell growth, and blood compatibility (Johnston and Ratner, 1997; Jennissen, 1998; Steele et al., 1998; Horbett, 2004). Some characteristics of this non-specific adsorption are as follows: Adsorption is rapid (a monolayer forms in seconds). Proteins compete with each other for surface sites. Adsorption is pseudo-Langmuirian. Monolayer adsorption is most commonly observed (typically, a monolayer of adsorbed protein inhibits further protein adsorption). 5. Adsorption is often irreversible. 6. One protein can sometimes displace another. 7. Adsorption can lead to protein denaturation. 1. 2. 3. 4.

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8. 9. 10. 11.

The longer a protein sits on the surface, the more it unfolds. The longer a protein sits on a surface, the harder it is to desorb. High densities of proteins at surfaces stabilize conformation. Proteins have “many faces” and many binding and functional regions. In non-specific adsorption there is no control of whether the protein molecule adsorbs up, down, or sideways.

Though non-specific protein adsorption readily occurs, and is almost always seen with today’s biomaterials and scaffolds, nature never uses such a method to deliver protein signals. In fact, non-specific proteins on surfaces may be seen by the body as “foreign” leading to undesirable reactions (Ratner, 1993, 2002). This possibility leads to a materials design strategy that says either inhibit all non-specific proteins at surfaces, or control proteins at surfaces with precision.

INHIBITION OF NON-SPECIFIC PROTEIN ADSORPTION There have been numerous strategies devised to inhibit all non-specific protein adsorption, and great strides have been made in surface design and theoretical understanding in the past few years (Hoffman and Ratner, 1996; Herrwerth et al., 2003; Kane et al., 2003; Ma et al., 2004; Johnston et al., 2005). Typically, surfaces that resist non-specific protein uptake also resist the adhesion of cells. Strategies that have been used to inhibit non-specific protein adsorption at surfaces are listed in Table 42.1. The description and literature associated with all these methods is beyond the scope of this review. Also, the mechanisms by which they might function represent a huge subject for discussion, with no real resolution having been reached to date by the scientific community. However, some of the key considerations for such non-fouling surfaces are: (1) How effective are they in resisting non-specific adsorption, especially in 100% plasma situations and other “real world” applications? (2) How low is low? One theory says that 10 ng/cm2 or less of key reactive proteins must be achieved to not trigger undesirable reactions (Tsai et al., 1999). (3) How durable are they over days, months, or years to oxidation and degradation? (4) How readily are they applied to technologically useful surfaces? (5) How will they be viewed by the regulatory agencies for medical device applications?

Table 42.1 Strategies to achieve non-fouling (protein and cell-resistant) surfaces • • • • • • • • • • • • • • • • •

Poly(ethylene glycol)(PEG) surfaces (networks, brushes, etc.) PEG oligomers (as head groups on self-assembled monolayers or as plasma-deposited thin films) Phase change polymers (NIPAM, peptide) Other hydrogels Protein films (adsorbed/immobilized) Saccharides Choline headgroups Betaine, taurine H-bond acceptors Kosmotropes Ablative surfaces Controlled release surfaces Negatively charged gels and proteins Protease surface DNA surface Surface electrical potential Reverse flow of water

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CONTROLLING PROTEINS AT INTERFACES WITH PRECISION Biological systems deliver protein chemical signals in an optimal manner by orienting and organizing the proteins at biosurfaces. Since proteins are multi-faced, the correct face to accomplish the job must be involved and since the organization of amino acids on the protein face is conformation dependent, the conformation must be controlled. Methods to control proteins at interfaces can be biomimetic (copy the organization of cell surfaces, for example) or there can be novel strategies to achieve this control. A list of possible methods is presented in Table 42.2. Consistent with the theme of this chapter, it is worthwhile examining some of these methodologies in Table 42.2. Control of Packing Density Control of packing density is perhaps the simplest method to begin ordering and organizing a protein interface. The concept is straightforward. For a protein to unfold (denature), molecular volume is required since the native form is typically highly compact. By packing proteins tightly on the surface, they have insufficient room to denature and are thus conformationally stabilized. Conformational stabilization is one aspect of using proteins with precision at surfaces, but it does not address the issue of correctly orienting the protein. Preservation Agents Such As Trehalose Proteins can be conformationally stabilized at interfaces by the use of saccharide compounds. Trehalose, originally noted in the seeds of dessert plants and probably allowing the seed proteins to survive long, dry periods, has been applied to stabilize biological activity at surfaces (Xia and Castner, 2003). Such stabilization addresses the issue of conformation at surfaces but not orientation or alignment at surfaces. Histidine (HIS6) Tags Hexahistidine (HIS6) peptide chains, if site specifically incorporated into a protein molecule, can be used to consistently orient the protein molecule on the surface. HIS6 tags were originally developed for protein isolation and purification. The HIS6 binds somewhat specifically to a nickel-nitrilotriacetic acid (NTA) organic functional group that might be bound to an agarose chromatography column, pulling only labeled molecules from the solution phase. A few researchers soon realized that if the NTA groups were precision affixed to a surface (as a headgroup on a self-assembled monolayer (SAM), for example), that every HIS6- tagged protein that bound to the surface would bind with the same orientation (Frey et al., 1996; Sigal et al., 1996). This technique has been used in many situations where surface protein orientation would be desirable, but the presence of the nickel cation raises some concerns for in vivo application in tissue engineering.

Table 42.2 Some strategies to control protein orientation and conformation at interfaces • • • • • • • • • • •

Control of packing density Preservation agents such as trehalose Histidine (HIS6) tags Ionic charge (pH) control of orientation Immobilization in lipid layers, tethered lipid bilayers Streptavidin for orientation control Antibodies to control orientation/G-protein to orient IgGs and other molecules Hydroxyapatite for orientation control Collagen and extracellular matrix to control protein orientation Site specific protein modification to introduce a position-defined linking group Templating methods

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Surface Ionic Charge (pH) Control of Orientation Since proteins have different “faces,” each with unique distributions of amino acids (Kim et al., 2006), these faces would also be expected to have different isoelectric points. At physiologic pH, some protein faces will be more positively charged while other faces will have more negative charge. Thus, a positive or a negative charged surface can interact more strongly with one or the other protein face. This orientation effect was demonstrated with secondary ion mass spectrometry (SIMS) for IgG molecules that have Fab portions with an isoelectric point of 8.5 and Fc portions with an isoelectric point of 6.1 (Wang et al., 2006). Using the 7–10 type III module of fibronectin (FnIII7–10)(43 kD), this charge mechanism to orient proteins on surfaces was dramatically shown (Wang et al., 2006). FnIII7–10 was adsorbed to the surfaces of SAMs of N-alkyl thiols on gold with 9NH2 and 9COOH headgroups. Based upon I125 adsorption studies, the solution concentrations were adjusted to give the same amount of FnIII7–10 on both surfaces. When the binding of an antibody specific for the cell-binding domain of FnIII7–10 was observed by surface plasmon resonance (SPR), significant binding was shown to the FnIII7–10 on the amine surface, with little binding to the FnIII7–10 on the carboxylic acid surface (Figure 42.2). Bovine aortic endothelial cell adhesion was consistent with this SPR result with excellent adhesion on the FnIII7–10 on the 9NH2 surface and little on the FnIII7–10-coated 9COOH surface (though the same total amount of peptide was adsorbed to both surfaces.) The results strongly suggest that the RGD domain on the FnIII7–10 was oriented up (accessible) on the amine surface and down (inaccessible) on the carboxylic acid surface. Streptavidin and Avidin for Orientation Control The tetravalency and symmetry of streptavidin and avidin has suggested many strategies that can be used to orient proteins at interfaces. In particular, with the ability to site specifically position a biotin molecule on a protein, that strategy interfaced with streptavidin or avidin-coated surfaces can be used to precisely orient a protein (McLean et al., 1993). As examples, a few papers are cited (McLean et al., 1993; Muller et al., 1994; Wilson and Nock, 2001; Jung et al., 2006). The ability to add biotin to a specific location in the protein sequence

575 550

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525 NH2 NH2 NH2 NH2 NH2 NH2

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475

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S

S

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H H H H O O O O O O O O C C C C

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375 S

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Figure 42.2 The binding of an antibody specific for the cell-binding domain of FnIII7–10 was observed by SPR. Significant binding is seen to the FnIII7–10 on the amine surface, with little binding to the FnIII7–10 on the carboxylic acid surface (surface concentration of FnIII7–10 on both surfaces is approximately 150 ng/cm2). This figure is adapted from reference (Wang et al., 2006).

Proteins Controlled with Precision at Organic, Polymeric, and Biopolymer Interfaces 739

can be further generalized – if you can engineer a “handle” (cysteine, for example) into a specific location on a protein, a tether can be attached that will allow specific, oriented immobilization (Stayton et al., 1992). Immobilization in Lipid Layers and Tethered Lipid Bilayers Proteins contained within the lipid bilayers of cells are almost always oriented correctly. A hydrophobic region of the protein typically penetrates the lipid layer fixing the protein orientation normal to the membrane (in the “z” direction) while permitting mobility laterally (in the “x” and “y” directions). This general principle has been adopted to synthetic systems. Tethered lipid bilayers and aperture-suspended bilayers (black lipid bilayer membranes) can be used to orient protein molecules (Salafsky et al., 1996; Giess et al., 2004; Castellara and Cremer, 2006). Antibodies to Control Orientation/G-Protein to Orient IgGs and Other Molecules An oriented, monoclonal antibody at a surface can bind to a specific surface domain of a protein thereby leading to an oriented, immobilized protein molecule (Koyama et al., 1994; Karyakin et al., 2000; Klueh et al., 2003; Lee et al., 2006). Protein G has a specific binding site for the Fc portion of an antibody. Since one generally wants the Fab portion of the antibody oriented facing outward from the immobilization surface, protein G on a surface serves to facilitate this orientation. Both protein G (biotinylated) and streptavidin have been used together to orient antibodies (Jung et al., 2006). A variant of protein G was prepared by engineering cysteine residues into the N-terminus of the protein allowing it to assemble and order on a gold surface (Lee et al., 2007). Protein G on a surface was recently used to orient a ligand for the cell surface receptor, Notch (Beckstead et al., 2007). The Notch ligand (Jagged-1) was one component of a fusion protein with the Fc portion of an antibody. The protein G-coated surface bound the Fc portion of the fusion protein orienting the Notch ligand to be accessible in the solution. When esophageal epithelial cells were seeded on this surface, they exhibited appropriate differentiation (stratification). The Fc portion of the antibody by itself, presented on the protein G surface, did not trigger this differentiation. Hydroxyapatite for Orientation Control Specific peptide sequences can recognize faces of biominerals (Addadi and Weiner, 1985). This process is important in normal biomineralization. This interaction can be exploited by using biominerals to orient adsorbed biomolecules. For example, NMR studies demonstrated that a 15 amino acid fragment of the protein statherin was specifically immobilized on hydroxyapatite (Shaw et al., 2000). This idea was applied to enhancing cell adhesion to hydroxyapatite by fusing RGD and flanking residues from osteopontin (OPN) to the C-terminus of the statherin peptide (Gilbert et al., 2000). Solid state NMR again demonstrated the orientation by noting the high mobility of terminal RGD. Collagen and ECM to Control Protein Orientation The ECM is nature’s own scaffold for constructing tissue. The ECM is comprised on a number of structural proteins including collagens, laminins, glycosaminoglycans, and fibronectin. Collagen has been shown to specifically bind at least 50 other proteins (de Lullo et al., 2002). The discovery of the function of matricellular proteins as modulators between ECM and cells furthered our understanding of the mechanism and biological function of these interfacial proteins (Bornstein et al., 1978; Bornstein, 2000). If cells are removed from a tissue leaving behind the ECM, a distinctly porous scaffold-like structure is noted. The ECM of this decellularized tissue, when immersed in proteinaceous, biological fluid, will adsorb and retain many proteins. These proteins will modulate the healing reaction. This suggests that there may be merit to pre-exposing the ECM to specified proteins that might be important to cell interaction, inflammation, and healing. This was done in

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studies where collagen type I was immobilized to a surface and the surface was exposed to OPN, a matricellular protein (Martin et al., 2004; Liu et al., 2007). OPN contains an RGD residue and supports cell adhesion. If the RGD sequence is cell accessible (oriented outward) it will be active in supporting cell adhesion. These studies demonstrated that OPN could deliver cell-adhesive signals when bound to collagen type I much more effectively than when directly bound to non-collagenous polymeric substrates. Templating Methods The possibility of using protein molecules as templates to make imprints (“pits”) that can recognize the proteins has been demonstrated (Shi et al., 1998). Though this has not been demonstrated, it is postulated that if the templates are oriented, the imprints will be much more effective in interacting with the solution-phase template protein.

IMPLICATIONS AND PERSPECTIVES In most tissue engineering and biomaterials applications, proteins at interfaces are used with little precision. They are not specifically oriented and are conformationally uncontrolled. There is little concern for total protein concentration at the interface. For studies in complex media (serum, plasma), exactly which proteins are at the interface are rarely controlled. Engineering control requires precision and reproducibility. To achieve tissue engineering in the clinic, we must refine our control of interfacial proteins and biomolecules. Specifically, using interfacial proteins with precision we can realize appropriate cell attachment, control of inflammation, modulated cell growth, angiogenesis, and tissue regeneration. This ability to control and stabilize proteins on tissue engineering scaffolds will meet the needs of industry (reproducibility, packaging) and the demands of the regulatory agencies. But ultimately it will allow physicians to “prescribe” tissue engineered products to their patients, and have the assurance that the living constructs will develop and function as expected. ACKNOWLEDGMENTS The author has received support and intellectual input from grants NSF EEC-9529161 (UWEB Engineering Research Center), NIH R24 HL64387 (BEAT BRP), Singapore-University of Washington Alliance (A-Star) and NIBIB grant EB-002027 (NESAC/BIO) during the preparation of this review article.

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Part V Therapeutic Applications: Cell Therapy

43 Biomineralization and Bone Regeneration Jiang Hu, Xiaohua Liu, and Peter X. Ma Biomineralization is the process by which mineral crystals are deposited in the matrix of living organisms. Such process gives rise to inorganic-based skeletal structures such as bone during development, which is a complex and dynamic organ with both structural and metabolic functions. However, ectopic biomineralization often causes severe diseases, such as calcification of vascular tissues related atherosclerotic lesions (Rumberger et al., 1995). As a part of the book entitled “Principles of Regenerative Medicine,” this chapter will focus on orthotopic bone formation and bone regeneration. Bone defects, caused by tumor or trauma, are a major health problem. There is an enormous clinical need to develop safe and effective modalities to stimulate bone regeneration. Tissue engineering offers a promising new approach in facilitating bone formation by recapitulating the natural process of bone development/healing using engineering techniques. This chapter will briefly describe the biologic processes of bone development and fracture repair, summarizing the current applications of stem cells and growth/differentiation factors involved in bone regeneration, and then focus on the principles of design and fabrication of scaffolds.

DEVELOPMENT AND FRACTURE HEALING OF BONE Development of Bone Bone formation proceeds by two different ways: endochondral ossification, which is a complex, multistep process requiring the sequential formation and degradation of cartilaginous templates for the developing bones; and intramembranous ossification, which is through the direct differentiation of precursor cells into osteoblasts (Karaplis, 2002). Limb development involves a complex series of events that first define embryologic zones for future endochondral bone development, and subsequently induce cartilage and bone of precisely defined structures. These processes are regulated by a variety of signals including soluble growth/differentiation factors, cell–cell and cell–extracelluar matrix (ECM) interactions, all of which are orchestrated by an underlying genetic program. At the cellular level, the development of bone involves restrictions in lineage potential of multipotent mesenchymal precursor cells by controlling the cellular transcriptional program. This process can be broadly divided into two phases: an initial commitment phase, at which stage cells that will eventually form bone are committed in defined time and space; and the subsequent differentiation phase, when the necessary cellular phenotypes are induced to construct bone tissues. The flat bones of the skull form through an intramembranous process. Although the precursor cells in the skull are derived from the neural crest, these cells are regulated by many of the same signaling molecules found in the limb development.

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Fracture Healing of Bone Like embryological development of skeleton, fracture repair involves multiple factors and the establishment of specific morphogenetic field to drive the differentiation of precursor cells, and, in some ways, can be considered a recapitulation of bone development (Gerstenfeld et al., 2003). After a fracture happens, the initial inflammatory response recruits activated macrophages and polymorphonuclear neutrophils (PMNs) to the damaged sites. Under the control of multiple factors secreted by macrophages, an initial hematoma is formed. Then granulation tissue fibroblasts proliferate to form a blastema. Osteoprogenitors, migrating from periosteum, surrounding soft tissues, and the bone marrow space at the damaged sites, differentiate into chondrocytes and osteoblasts and form bone tissues. This process is induced and controlled by multiple soluble growth/differentiation factors. Among these, fibroblast growth factors (FGFs), insulin-like growth factors (IGFs) and platelet-derived growth factors (PDGFs), which are distributed in the soft callus early in the fracture healing, act as mitogenic factors to promote precursor cells proliferation, while other differentiation factors such as bone morphogenetic proteins (BMPs) are more responsible for differentiation of chondrocytes and osteoblasts present later in the healing tissues.

PRINCIPLE OF BONE TISSUE ENGINEERING For bone regeneration therapy to be successful, sufficient mesenchymal precursor cells must be either recruited or implanted directly to the damaged sites, and these cells must be given the appropriate signals to grow and differentiate in a controlled manner. Current clinical applicable therapies for bone defect repair include bone grafts and allogenic bone matrix implantation. Bone grafts, containing viable bone cells and osteoprogenitors, as well as growth/differentiation factors, are considered to be the “gold standard.” However, bone regeneration after bone grafting is quite variable, probably because of differences in the quality of the bone graft (Parikh, 2002). In addition, severe morbidity may occur at donor sites. Allogenic bone matrix provides a bone-like ECM and a crude source of growth/differentiation factors. These inductive factors may attract appropriate oseteoprogenitors to the regeneration site and stimulate their differentiation into osteoblast cells. However, osteoinductive activity of allogenic bone matrix is commonly inconsistent, primarily because it contains variable and often low levels of growth/differentiation factors, which are partially inactivated during processing (Iwata et al., 2002). There is also a potential risk of disease transmission if the matrix is not appropriately processed. In contrast, tissue engineering affords a new way for bone regeneration, which has the advantage to combine the use of precisely engineered scaffolds, the appropriate osteoprogenitor cells and related growth/differentiation factors (Liu and Ma, 2004). If a damaged tissue to be repaired has high activity in terms of regeneration, new tissue can form in a biodegradable scaffold directly by precursor cells infiltrating from the surrounding tissues. However, non-union or delayed union fracture sites are often too large or inflamed and associated with significant scarring that may limit the migration of osteogenic precursors. Also some bone damage sites are related to low concentrations of growth/differentiation factors. Additional components such as mesenchymal stem cells (MSCs) and BMPs are required in these cases.

STEM CELLS IN BONE TISSUE ENGINEERING Stem cells are defined as cellular population with two critical properties: self-renewal to produce daughter stem cells with identical potentialities and the ability to differentiate along one or more lineages (Wagers and Weissman, 2004). Potential sources of stem cells for bone tissue engineering include embryonic stem cells (ESCs) and adult MSCs.

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ESCs ESCs offer a potentially unlimited supply of cells that may be driven down specific lineages, giving rise to all cell types in the body (Thomson et al., 1998). ESCs can be driven to differentiate into osteoblast cells in vitro. In one method, osteogenic cells are derived from three-dimensional (3D) cell spheroids called embryoid bodies (EBs) (Bielby et al., 2004). EBs can be formed through suspension or hanging drop methods from single cell suspension. Since EBs mimic the structure of the developing embryo and recapitulate many of the stages involved during its differentiation, they create suitable conditions to drive ESCs to differentiate into precursor cells of all three germ layers. Then EBs are dispersed and committed cells are further cultured in monolayer to be induced to osteogenic cells under the presence of exogenous factors such as dexamethasone (DEX), L-ascorbic acid (AA) and sodium-β-glycerophosphate (β gP). DEX has been demonstrated to stimulate osteogenic differentiation for precursor cells derived from multiple tissues. AA is used to promote collagen secretion and deposition, and β gP is used to mineralize the deposited matrix. Alternatively, undifferentiated ESCs or dispersed EBs can be directly seeded into 3D scaffolds and driven to multiple tissues (Levenberg et al., 2003) for later implantation. MSCs MSCs are an ideal stem cell source for cell therapies because of their easy purification, amplification, multipotency, and low immunogenicity. MSCs were first identified in 1966 by Friedenstein and co-workers, who isolated bone/cartilage-forming progenitor cells from rat bone marrow cells with fibroblast-like morphology (Friedenstein et al., 1966). Although MSCs have been isolated from a number of tissues, including the fetal blood, liver, bone marrow (Campagnoli et al., 2001), and umbilical cord blood (Lee et al., 2004), the most studied and accessible source of MSCs is the bone marrow. Within the bone marrow, MSCs are estimated to comprise 0.001–0.1% of the total population of nucleated cells, which can be selected from other nucleated cells by their adherence property to plastic flasks in culture and can be expanded extensively for multiple passage numbers in vitro without loss of phenotype. Unlike the hemopoietic stem cells (HSCs), which can be defined by specific surface markers, MSCs only express a number of non-specific surface markers. MSCs express neither hemopoietic (CD34, CD45, CD14) nor endothelial cell marker (CD31), but a large number of adhesion molecules (CD44, CD29, CD90), stromal cell markers (SH-2, SH-3, SH-4), and some cytokine receptors (Pittenger et al., 1999). These MSC markers can be collectively used to identify isolated MSCs in culture. Some enrichment strategies are also developed based on selection of cells positive for STRO-1 (Simmons and Torokstorb, 1991) and SH-2 markers (Barry et al., 1999). MSCs can be driven down along mesenchymal cellular pathways, including osteogenic, chondrogenic, and adipogenic lineages, when placed in appropriate in vitro or in vivo environments (Pittenger et al., 1999). Osteogenic differentiation is stimulated under the supplement of DEX, AA, β gP. Under these culture conditions, MSCs upregulate alkaline phosphatase, osteocalcin, and osteopontin expressions, and also calcium deposition within the ECM. For bone regeneration in vivo, bone-marrow-derived MSCs have been demonstrated to facilitate bone repair when implanted locally, commonly on an artificial matrix, such as hydroxyapatite (HAP) scaffold (Kasten et al., 2005) in craniotomy and long-bone defects. In addition to multipotency, the low immunogenicity property of MSCs make the cells applicable for allogenic implantation (Barry et al., 2005). Another clinically applicable MSC source is white adipose. Like bone marrow, adipose tissue is mesodermally derived with a stromal part containing microvascular endothelial cells, smooth muscle cells, and MSCs. These cells can be enzymatically isolated from adipose tissue and separated from the buoyant adipocytes by centrifugation. A more homogeneous population can be selected and expanded under culture conditions favorable for MSC growth (Zuk et al., 2002). This population, called adipose tissue-derived stem cells (ADSCs), shares many of the characteristics of its counterpart in bone marrow, including extensive proliferative potential and multipotency (De Ugarte et al., 2003). ADSCs can be obtained in large numbers at high frequency from white tissue with minimal morbidity, representing another potential clinically useful source of MSCs for bone tissue engineering.

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GROWTH/DIFFERENTIATION FACTORS IN BONE TISSUE ENGINEERING Many growth/differentiation factors are used in bone tissue engineering. Among these, BMPs have the unique ability to stimulate the differentiation of mesenchymal precursor cells to chondrocytes and osteoblasts, and induce formation of new bone at both ectopic and orthotopic sites. BMPs It was observed that demineralized bone matrix (DBM) is able to induce ectopic bone formation in subcutaneous and intramuscular pockets in rodents (Urist, 1965). Isolation of the bone-inducing substance revealed certain proteins termed BMPs or osteogenetic proteins (OPs) (Wozney et al., 1988). BMPs belong to the transforming growth factor-β (TGFβ) superfamily, which consists of a group of related peptide growth factors. More than 40 related members of this family have been identified, including 15 BMPs (de Caestecker, 2004). They are further divided into subfamilies according to their amino acid sequence similarities. BMPs consist of dimers that are interconnected by seven disulfide bonds. This dimerization is a prerequisite for bone induction. BMPs are active both as homodimers that consist of two identical chains, and as heterodimers consisting of two different chains (Granjeiro et al., 2005). Compared to other known growth factors, BMP-2 (Boyne et al., 2005) and BMP-7 (Vaccaro et al., 2005) have the most robust osteoinductive activity as observed in both preclinical animal studies and in human trials. Growth/Differentiation Factors Delivery A simple way for bone regeneration is to supply growth factors such as BMPs to the site of defect for cell proliferation and differentiation in a controllable manner. Bone tissue regeneration is sometimes induced by use of growth/differentiation factors in soluble form, but the amount applied is much higher than that under normal physiological conditions, commonly at milligram level, which may cause adverse effects. Drug delivery systems are currently under development that allow for the controlled release of proteins, either encapsulated in poly(D,L-lactic acid-co-glycolic acid) (PLGA) microspheres (Weber et al., 2002) or incorporated into collagen carriers (Murata et al., 2000). Regional Gene Therapy Regional gene therapy offers another approach to deliver growth/differentiation factors to the healing sites. Transfected cells express growth/differentiation factors for a sustained period, thereby reducing the problem of protein degradation. Viral vectors and non-viral vectors are presently being investigated as potential gene delivery vehicles to enhance bone repair. In addition, MSCs themselves can be used as gene transfer carriers. Not only being a source of BMPs after transfection, the cells directly respond to BMPs and participate in bone formation after implantation, which may be important at some damage sites, where the supply of endogenous osteogenic precursors is limited. MSCs transfected with adenoviruses encoding BMPs have been shown to stimulate bone regeneration in several experimental models (Wang et al., 2003). Although recombinant adenovirus can be produced in high titers, and can easily infect both dividing and non-dividing cells at high efficiency (Spector et al., 2000), the immune response to the adenoviral proteins is a major obstacle to the adaptation of this approach to treat non-lethal diseases such as bone defect in humans. In contrast, non-viral vectors are easier to produce and have better chemical stability. However, the in vivo transfection efficiency of current available non-viral vectors such as liposome and poly(ethylenimine) (PEI) is low (Lollo et al., 2000). New vectors and delivery methods are being developed in this field. Combination of Growth/Differentiation Factors At anytime during bone development or fracture healing, multiple growth/differentiation factors are functioning in a coordinated manner. Therefore, combinations of bioactive factors might synergistically stimulate

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bone regeneration. Angiogenic factors and BMPs can act synergistically. To examine possible interactions between BMPs and angiogenic signals in bone regeneration, Peng and co-workers used muscle-derived stem cell (MDSC) lines genetically modified to express BMP-4 or vascular endothelial growth factor (VEGF) (Peng et al., 2002). VEGF by itself had no effect on the osteogenic activity of MDSC. However, it acted synergistically with BMP-4 to increase recruitment of mesenchymal precursor cells and to enhance cell survival, thus stimulating bone formation in a calvarial defect.

SCAFFOLDS FOR BONE TISSUE ENGINEERING Scaffolding Design Criteria for Bone Tissue Engineering In bone tissue engineering, the scaffold plays critical roles in supporting cell adhesion, migration, proliferation, differentiation, and mineralized bone tissue formation (Ma, 2003; Liu and Ma, 2004; Ma, 2004). Scaffolds for bone regeneration should meet certain criteria to serve these functions (Liu and Ma, 2004; Smith and Ma, 2004). First of all, the scaffold should have a controlled porous architecture to allow for cell growth, tissue regeneration, and vascularization. High interconnectivity between pores is desirable for uniform cell seeding and distribution, the diffusion of nutrients to and metabolites away from the cell/scaffold constructs. The scaffold should have adequate mechanical stability to provide a suitable environment for new bone tissue formation. The scaffold degradation rate must be tuned to match the rate of new bone tissue formation. Furthermore, the scaffold should be osteoconductive to enhance osteoblast attachment, migration, and differentiated function. A variety of processing technologies have been developed to fabricate porous 3D polymeric scaffolds for bone regeneration. These techniques include solvent casting/particulate-leaching (Mikos et al., 1994; Thomson et al., 1995), gas foaming (Mooney et al., 1996; Hile et al., 2000), emulsion freeze-drying (Whang et al., 1995), electrospinning (Li et al., 2002; Matthews et al., 2002), rapid prototyping (Giordano et al., 1996; Sun et al., 2004), and thermally induced phase separation (Nam and Park, 1999; Zhang and Ma, 1999a; Ma and Zhang, 2001). Several review papers have addressed the scaffolding fabrication methods, their advantages, and disadvantages (Hutmacher, 2000; Chaikof et al., 2002; Liu and Ma, 2004). This chapter is not intended to be exhaustive in detailing various processing techniques. Instead, it will focus on illustrating how to achieve the above scaffolding design goals through certain engineering methods. Important issues for scaffolding design, such as porosity, interconnectivity, mechanical strength, morphology, and surface properties, will be emphasized using examples from our group and others. Porous Scaffolds with High Interconnectivity Porosity and interconnectivity between pores are important scaffold parameters. Porous scaffolds with high interconnectivity are desirable for uniform cell seeding and distribution. Solvent casting/particulate leaching is a simple and most commonly used method to fabricate porous scaffolds for bone tissue engineering (Mikos et al., 1994). The method involves casting a mixture of polymer solution and porogen in a mold, drying the mixture, and subsequently leaching the porogen with water to obtain a porous structure. Usually, watersoluble particulates such as NaCl are used as the porogen materials. This method is simple to operate, and the pore size and porosity of the scaffold can be adequately controlled by particle size of the added salt and the salt/polymer ratio. However, the limited interpore connectivity is not desirable for uniform cell seeding and tissue growth. A new technique has been developed to fabricate scaffolds with spherical pore shape and well-controlled interpore connectivity by using paraffin spheres as pore-generating materials (Ma and Choi, 2001). The created new scaffold has a homogeneous foam skeleton and high porosity (Figure 43.1). The control of porosity and the pore size can be achieved by changing the concentration of the polymer solution, the number of the casting steps, and the size of the paraffin spheres. The degree of interconnectivity is finely tuned by the heat treatment time to bond paraffin spheres, which is critical to uniform cell seeding, tissue ingrowth, and regeneration.

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(a)

(b)

Figure 43.1 SEM micrographs of poly(α-hydroxy acids) scaffolds. (a) PLLA foams prepared with paraffin spheres with a size range of 250–420 μm (250). (b) PLGA foams prepared with paraffin spheres with a size range of 420–500 μm (50). (From Ma and Choi, copyright 2000 by Mary Ann Liebert, Inc. Reprinted with permission.)

Composite Scaffolds for Bone Tissue Engineering Although poly(α-hydroxy acids), such as poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), and PLGA, have been widely used to fabricate scaffolds for bone tissue engineering, the disadvantages of these materials are the weak mechanical properties and insufficient osteoconductivity. On the other hand, HAP, bioglass, and calcium phosphate have been demonstrated to have good osteoconductivity and bone bonding ability. They also have been shown to enhance mineralized new bone formation when implanted into bone defects (Hench, 1998; Suchanek and Yoshimura, 1998). However, the application of ceramics alone in bone tissue engineering is limited because of their fragility and low degradability in biological environment. Polymer/ceramic composite scaffolds may enhance both mechanical properties and osteoconductivity. Highly porous poly(α-hydroxy acids)/HAP scaffolds have been created through a thermally induced phase separation technique (Zhang and Ma, 1999a; Ma et al., 2001). These composite scaffolds showed significant improvement in compressive modulus and compressive yield strength over pure polymer scaffolds. Compared to pure polymer scaffolds in which cell ingrowth and tissue matrix formation were limited to the periphery of the scaffold, the composite scaffolds supported uniform cell seeding, cell ingrowth, and tissue formation throughout the scaffold (Figure 43.2). Further examination revealed that polymer/HAP scaffolds had a higher osteoblast survival rate, more uniform cell distribution and growth, enhanced bone specific gene expression, and improved new tissue formation (Ma et al., 2001). Another strategy is to prepare bone-like apatite coated composite scaffold by immersing polymeric scaffolds in a simulated body fluid (SBF) (Zhang and Ma, 1999b). In this approach, prefabricated polymeric scaffolds are incubated in SBF at 37°C to allow the in situ apatite formation on the inner pore wall surface of the 3D scaffold. After incubation, large amounts of apatite particles are formed uniformly on the scaffold pore walls (Figure 43.3). The apatite particles formed using this method are similar to the apatite of natural bone based on EDS, FTIR, and XRD analyses (Zhang and Ma, 1999b). It has been observed that the growth of apatite crystals was affected greatly by the polymer materials, porous structure, ionic concentration of SBF as well as the pH value (Zhang and Ma, 2004). Biomimetic deposition of bone-like apatite is not only of direct interest for the development of a composite scaffold but also for assessing the calcification function of existing scaffolds (Wei and Ma, 2004).

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(a)

(b)

(c)

(d)

Figure 43.2 Osteoblastic cell distribution in highly porous PLLA and PLLA/HAP composite scaffolds 1 week after cell seeding (von Kossa’s silver nitrate staining; original magnification 100): (a) The surface area of an osteoblast-PLLA construct; (b) the center of an osteoblast-PLLA construct; (c) The surface area of an steoblast-PLLA/HAP construct; and (d) The center of an osteoblast-PLLA/HAP construct. (From Ma et al., copyright 2001 by John Wiley & Sons, Inc. Reprinted with permission.)

(a)

(b)

Figure 43.3 SEM micrographs of a PLLA scaffold incubated in SBF for 30 days: original magnifications (a) 100, (b) 10,000. (From Zhang and Ma, copyright 1999 by John Wiley & Sons, Inc. Reprinted with permission.)

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Figure 43.4 SEM micrographs of a PLLA fibrous matrix prepared from 2.5% (wt/v) PLLA/THF solution at a gelation temperature of 8°C (From Ma and Zhang, copyright 1999 by John Wiley & Sons, Inc. Reprinted with permission.)

Nano-fibrous Scaffolds for Bone Tissue Engineering It is well known that the ECM environment plays an integral role in regulating cell behavior with respect to morphology, cytoskeletal structure, and functionality (Aumailley and Gayraud, 1998; Rosso et al., 2004). Thus, it is often beneficial that the scaffold replicates the cells’ natural ECM environment until the host cells can re-populate and re-synthesize a new matrix (Hubbell, 1999; Hench and Polak, 2002; Shin et al., 2003). Collagen is the main ECM component of bone, and its nano-fibrous architecture has long been known to play a role in cell adhesion, growth, and differentiated function in tissue cultures (Grinnell, 1982; Strom and Michalopoulos, 1982). To mimic the nano-fibrous architecture of collagen, a novel liquid–liquid phase separation technique has been developed to fabricate nano-fibrous PLLA (NF-PLLA) matrices (Ma and Zhang, 1999). Typically, the nano-scale fibrous matrices are fabricated with five steps: polymer dissolution, phase separation and gelation, solvent extraction, freezing, and then freeze-drying under vacuum. The fiber network formation depends on the gelling temperature and the solvent of the polymer solution. The synthetic NF-PLLA matrix is composed of interconnected fibrous network with a fiber diameter ranging from 50 to 500 nm, which is in the same range as that of collagen matrix (Figure 43.4). This NF-PLLA matrix has a much higher surface-to-volume ratio than those of fibrous non-woven fabrics fabricated with the textile technology or foams fabricated with other techniques. When combined with porogen-leaching techniques (e.g. paraffin leaching), 3D macroporous architectures can be built in the nano-fibrous matrices (Zhang and Ma, 2000; Chen and Ma, 2004). These synthetic analogs of natural ECM combine the advantages of the synthetic biodegradable polymers and the nano-scale architecture similar to the natural ECM. They have been found to selectively enhance protein adsorption and promote osteoblastic cell adhesion (Woo et al., 2003). Surface Modification of Nano-fibrous Scaffolds Surface properties as well as scaffolding architecture are important for a desirable scaffold in tissue engineering (Boyan et al., 1996; Liu et al., 2005a,b). The interactions of cells with the scaffolding materials take place

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50

*

Control Surface modified

40 DNA (μg)

* 30 20 10 0 0.5

7 Days

14

Figure 43.5 The proliferation of osteoblasts cultured on control NF-PLLA scaffolds and surface-modified NF-PLLA scaffolds (four bilayers of PDAC/gelatin). 2  106 cells were seeded on each scaffold (*p  0.05 between surface modified and control groups). (From Liu et al., copyright 2005 by American Scientific Publishers. Reprinted with permission.)

on the material surface; therefore the nature of the surface can directly affect cellular response, ultimately influencing the rate and quality of new tissue formation. Although a variety of synthetic biodegradable polymers have been used as tissue engineering scaffolding materials, one main disadvantage of these scaffolds is their lack of biological recognition on the material surface. Surface modification methods have been developed to promote cell–material interactions (Neff et al., 1998; Mann et al., 1999; Lenza et al., 2002). However, most of the surface modification methods this far are applicable to 2D films or very thin 3D constructs. A novel surface modification method based on electrostatic layer-by-layer self-assembly technique has been recently introduced to modify true 3D scaffolding (especially nano-fibrous 3D scaffolding) surface (Liu et al., 2005a). As mentioned above, NF-PLLA scaffolds fabricated by thermally induced phase separation technique mimic the physical structure of natural collagen matrix. To further mimic the chemical composition of collagen matrix, gelatin (derived from collagen by hydrolysis) is incorporated onto the surface of NF-PLLA scaffolds by the electrostatic self-assembly technique. The NF-PLLA scaffolds are first activated in an aqueous poly(diallyldimethylammonium chloride) (PDAC) solution to obtain stable positively charged surfaces. After washing the scaffolds with water, the scaffolds are dipped into gelatin solution for a designated time and then washed with water. The scaffolds are again exposed to PDAC solution. Following the same washing procedure, the scaffolds are dipped into gelatin solution and rinsed with water again. The further growth of PDAC/gelatin bilayers is accomplished by repeating the same cycle. Polyelectrolyte multilayers containing gelatin molecules are deposited on the NF-PLLA surfaces after crosslinking and drying. The amount of gelatin on the surface was controlled by the number of assembled polyelectrolyte bilayers, and increased linearly with the bilayer number after the first two bilayers. The wettability of the scaffold is controlled by varying the nature of outmost layer. The surface-modified NF-PLLA scaffolds mimick both the chemical composition and architecture of collagen matrix, and have been demonstrated to significantly improve cell adhesion and proliferation (Figure 43.5).

CONCLUSIONS Bone development and fracture healing are complex processes controlled by multiple factors. Bone tissue engineering offers a promising new approach for bone regeneration by mimicking these natural processes, which combines the stem cells, growth/differentiation factors together with supportive scaffolds in a controlled manner. Stem cells offer an ideal source for generating bone-forming cells and are especially desired for therapies to treat large defects and damaged sites with limited osteoprogenitor cells. Growth/differentiation factors can

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be used to stimulate bone regeneration by drug delivery or gene therapy approaches, and it is proposed that combinations of appropriate factors may have synergistic effects. Scaffolds play important roles in bone tissue engineering. Many characteristic parameters (e.g. porosity, interconnectivity, mechanical strength, morphology, and surface properties) should be carefully considered for the design and fabrication of scaffolds to meet the needs of a specific tissue engineering application. Mimicking the natural bone matrix structure and composition represents a new biomimetic scaffold design approach. As scientists learn more about cellular interactions with materials and growth/differentiation factors, it is likely that scaffolds will be designed to controllably manipulate stem or osteoblastic cell function to enable the development of more advanced bone regeneration therapies.

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Pittenger, M.F., Mackay, A.M., Beck, S.C., Jaiswal, R.K., Douglas, R., Mosca, J. D. Moorman, M.A., Simonetti, D.W., Craig, S. and Marshak, D.R. (1999).Multilineage potential of adult human mesenchymal stem cells. Science 284(5411): 143–147. Rosso, F., Giordano, A., Barbarisi, M. and Barbarisi, A. (2004). From cell-ECM interactions to tissue engineering. J. Cell. Physiol. 199(2): 174–180. Rumberger, J.A., Simons, D.B., Fitzpatrick, L.A., Sheedy, P.F. and Schwartz, R.S. (1995). Coronary-artery calcium area by electron-beam computed-tomography and coronary atherosclerotic plaque area – a histopathologic correlative study. Circulation 92(8): 2157–2162. Shin, H., Jo, S. and Mikos, A.G. (2003). Biomimetic materials for tissue engineering. Biomaterials 24(24): 4353–4364. Simmons, P.J. and Torokstorb, B. (1991). Identification of stromal cell precursors in human bone-marrow by a novel monoclonal-antibody, stro-1. Blood 78(1): 55–62. Smith, L.A. and Ma, P.X. (2004). Nano-fibrous scaffolds for tissue engineering. Colloid. Surf B Biointerf. 39(3): 125–131. Spector, J.A., Mehrara, B.J., Luchs, J.S., Greenwald, J.A., Fagenholz, P.J., Saadeh, P.B., Steinbrech, D.S. and Longaker, M.T. (2000). Expression of adenovirally delivered gene products in healing osseous tissues. Ann. Plastic Surg. 44(5): 522–528. Strom, S.C. and Michalopoulos, G. (1982). Collagen as a substrate for cell-growth and differentiation. Method. Enzymol. 82: 544–555. Suchanek, W. and Yoshimura, M. (1998). Processing and properties of hydroxyapatite-based biomaterials for use as hard tissue replacement implants. J. Mat. Res. 13(1): 94–117. Sun, W., Darling, A., Starly, B. and Nam, J. (2004). Computer-aided tissue engineering: overview, scope and challenges. Biotech. Appl. Biochem. 39: 29–47. Thomson, J.A., Itskovitz-Eldor, J., Shapiro, S.S., Waknitz, M.A., Swiergiel, J.J., Marshall, V.S. and Jones, J.M. (1998). Embryonic stem cell lines derived from human blastocysts. Science 282(5391): 1145–1147. Thomson, R.C., Yaszemski, M.J., Powers, J.M. and Mikos, A.G. (1995). Fabrication of biodegradable polymer scaffolds to engineer trabecular bone. J. Biomat. Sci. Polym. Edn 7(1): 23–38. Urist, M.R. (1965). Bone – Formation by autoinduction. Science 150(3698): 893–899. Vaccaro, A.R., Patel, T., Fischgrund, J., Anderson, D.G., Truumees, E., Herkowitz, H., Phillips, F., Hilibrand, A. and Albert, T.J. (2005). A 2-year follow-up pilotstudy evaluating the safety and efficacy of op-1 putty (rhbmp-7) as an adjunct to iliac crest autograft in posterolateral lumbar fusions. Eur. Spine J. 14(7): 623–629. Wagers, A.J. and Weissman, I.L. (2004). Plasticity of adult stem cells. Cell 116(5): 639–648. Wang, J.C., Kanim, L.E.A.,Yoo, S., Campbell, P.A., Berk,A.J. and Lieberman, J.R. (2003). Effect of regional gene therapy with bone morphogenetic protein-2-producing bone marrow cells on spinal fusion in rats. J. Bone Joint Surg. Am. 85A(5): 905–911. Weber, F.E., Eyrich, G., Gratz, K.W., Maly, F.E. and Sailer, H.F. (2002). Slow and continuous application of human recombinant bone morphogenetic protein via biodegradable poly(lactide-co-glycolide) foamspheres. Int. J. Oral Maxillofac. Surg. 31(1): 60–65. Wei, G.B. and Ma, P.X. (2004). Structure and properties of nano-hydroxyapatite/polymer composite scaffolds for bone tissue engineering. Biomaterials 25(19): 4749–4757. Whang, K., Thomas, C.H., Healy, K.E. and Nuber, G. (1995). A novel method to fabricate bioabsorbable scaffolds. Polymer 36(4): 837–842. Woo, K.M., Chen, V.J. and Ma, P.X. (2003). Nano-fibrous scaffolding architecture selectively enhances protein adsorption contributing to cell attachment. J. Biomed. Mat. Res. Part A 67A(2): 531–537. Wozney, J.M., Rosen, V., Celeste, A.J., Mitsock, L.M., Whitters, M.J., Kriz, R.W., Hewick. R.M. and Wang, E.A. (1988). Novel regulators of bone-formation – molecular clones and activities. Science 242(4885): 1528–1534. Zhang, R.Y. and Ma, P.X. (1999a). Poly(alpha-hydroxyl acids) hydroxyapatite porous composites for bone-tissue engineering. I. Preparation and morphology. J. Biomed. Mat. Res. 44(4): 446–455. Zhang, R.Y. and Ma, P.X. (1999b). Porous poly(L-lactic acid)/apatite composites created by biomimetic process. J. Biomed. Mat. Res. 45(4): 285–293. Zhang, R.Y. and Ma, P.X. (2000). Synthetic nano-fibrillar extracellular matrices with predesigned macroporous architectures. J. Biomed. Mat. Res. 52(2): 430–438. Zhang, R.Y. and Ma, P.X (2004). Biomimetic polymer/apatite composite scaffolds for mineralized tissue engineering. Macromol. Biosci. 4(2): 100–111. Zuk, P.A., Zhu, M., Ashjian, P., De Ugarte, D.A., Huang, J.I., Mizuno, H., Alfonso, Z.C., Fraser, J.K., Benhaim, P. and Hedrick, M.H. (2002). Human adipose tissue is a source of multipotent stem cells. Mol. Biol. Cell 13(12): 4279–4295.

44 Blood Substitutes: Reverse Evolution from Oxygen Carrying to Non-Oxygen Carrying Plasma Expanders Amy Tsai, Marcos Intaglietta, and Mark Van Dyke The scientific literature is rife with investigations of potential replacements for red blood cells (RBCs), plasma, serum, and other constituents of whole blood. The intrinsic complexity of the human cardiovascular system and arcane functionality of the many constituents of whole blood have made this a formidable task. When one considers the delicate balance of homeostasis routinely achieved in healthy individuals by the cellular and non-cellular components of whole blood, it is difficult to imagine that a man-made substitute could recapitulate this scenario. Despite this, the search for a substitute for whole blood and various blood components spans more than half a century and includes both natural and synthetic biomaterials, as well as cell-based approaches. Regardless of the approach, however, it is important to frame any discussion of blood substitutes in the context of cardiovascular physiology and fluid biomechanics, especially when considering the situations when blood substitutes would most often be required – the hypotensive patient. While many types of conditions can be treated through the use of a blood transfusion, it is typically the hemorrhagic patient that presents the most significant challenges and the greatest need.

INDICATIONS FOR BLOOD TRANSFUSION Hypotension most often arises during periods of cardiac insufficiency (e.g. ischemia such as during myocardial infarction) or hypovolemia (e.g. extreme hemorrhage). Use of a blood substitute may also be indicated in other trauma scenarios such as severe burns, or in the case of severe, chronic anemia. The average amount of blood required for several such indications is presented in Table 44.1. These data particularly emphasize the need for a blood transfusion during hemorrhagic trauma such as occurs in automobile accidents. These are the most challenging hypotensive scenarios and the focus of much of the ongoing research in blood substitutes for both military and civilian applications. Traumatic hemorrhage and ensuing shock is especially problematic due to the cascade of changes in cardiovascular biomechanics and concomitant loss of tissue perfusion. Blood substitutes are often aimed at severely hemorrhagic patients owning to an ability to reverse a downward spiral in homeostasis by restoration of mean arterial pressure (MAP) and more importantly, tissue perfusion in critical organs such as the heart and brain. In addition, fluid resuscitation can avert major organ failure that often accompanies severe hemorrhage days after the hypovolemic event. In the best case scenario, severely hemorrhagic patients can be treated with a hemostat

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Table 44.1 Average use of whole blood by indication in the United States (http://www.americasblood.org) Automobile accident Bone marrow transplant Burn Heart surgery Organ transplant

50 units 20 units 20 units 6 units 40 units

at the site of injury and concomitant fluid resuscitation. If bleeding is not brought under control, however, resuscitation with a blood substitute will accelerate blood loss and exacerbate ischemia in critical organs. Unfortunately, few resuscitation fluids can mitigate both tissue perfusion and the typical coagulopathies that accompany severe hemorrhage. At present, only fresh whole blood can accomplish both hemostasis and restoration of tissue perfusion. This fact has not stopped investigators in their quest for fluids that mimic at least some of the functionality of whole blood.

BIOMATERIAL-BASED BLOOD SUBSTITUTES Most blood substitutes are of limited functionality in that they address relatively few of the physiological and biomechanical functions of whole blood. Recognizing that it is exceedingly difficult to mimic such a complex system, particularly in cases of hypovolemic shock, most investigators have taken the approach of breaking down the functions of blood into their most simple forms and addressing what are deemed to be the most important. A typical approach is to match the biomechanical properties of blood to that of a fluid that is blood compatible. That is, inert within the cardiovascular system to the greatest extent possible while providing matching viscosity and oncotic pressure. Such fluids are referred to as colloids (crystalloids are also used in fluid resuscitation, but will not be discussed in this chapter). In 1963, the National Research Council published these six functional criteria for synthetic blood replacements: 1. 2. 3. 4. 5. 6.

Should not interfere with hemostasis or blood coagulation. No tendency to cause agglomeration or lysis of RBCs or damage to WBCs. Should be metabolized and cause no delayed interference with the function of any organ. Not interfere with the mechanisms involved in the resistance to infection. Not interfere with hemopoesis or formation of plasma proteins. Not interfere with renal function or cardiac output, or cause metabolic acidosis.

To date, almost half a century later, no synthetic biomaterial has met all of these criteria. Even in the most ideal circumstance, the biomechanical functionality of blood substitutes typically comes at some physiological cost, making the development of an ideal synthetic blood substitute a vexing problem. Any discussion of biomaterial-based blood substitutes must clearly delineate two distinct types, non-oxygen and oxygen carrying. The first type of material is represented by polymer solutions that are typically directed at replacement of the serum component of blood. The polymers can be of either synthetic or natural origin. This approach seeks to replace the fluid component of plasma and relies solely on the solubility of oxygen in liquid to provide passive transport and delivery, although oxygen delivery is not the primary goal of these fluids. By the late 1960s, investigators began to envision the possibility of a second, more sophisticated approach whereby the oxygen carrying functionality of RBCs could be mimicked in the form of oxygen carriers such as stroma-free hemoglobin. By the 1970s, the use of fluorocarbon emulsions had been investigated for their extremely high oxygen solubility limits, although oxygen carrying in these emulsions is still considered passive.

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Many biocompatible polymeric materials have been investigated as potential plasma expanders and/or blood substitutes. During more than half a century of research in this field, natural polysaccharides (e.g. pectin), chemically modified polysaccharides (e.g. hydroxyethylated amylopectin and hydroxyethyl starch), polysaccharides produced by bacteria (e.g. alpha-1,6-D-glucosan or dextran), natural and chemically modified proteins (e.g. gelatin and gelatin derivatives), and synthetic polymers (e.g. perflurocarbons and polyvinlypyrolidone) have been investigated. Most materials have not progressed in their development beyond initial animal trials. However, a selected few have been under investigation for decades, and over the years have in some cases advanced to human clinical trials. Albumin, dextran, and hydroxyethyl starch have received most of the attention of researchers starting in the 1950s and have been under almost constant investigation since that time. Reviews on the efficacy of these compounds are mixed, with initial experience having been quite good. Many early reviews extol the virtues of these plasma expanders (Mishler et al., 1977; Davidson et al., 1980; Brecher et al., 1997), while others maintain that the deficiencies are too serious to warrant their use except in specific circumstances (Nearman and Herman, 1991; Roberts and Bratton, 1998; Szeto and Chow, 2005). To illustrate the complexity of fluid resuscitation, consider the requirement to maintain hemostasis. When a blood substitute is first introduced into the patient, its primary function is to replace fluid volume. To accomplish this, the blood substitute must maintain normal colloid oncotic pressure. If oncotic pressure cannot be maintained, fluid leaves that vasculature and escapes into the tissues, causing severe edema when large volumes of fluid have been administered. The types of polymers used as blood substitutes are heterogeneous mixtures of varying molecular weight, with the smallest of these molecules making a relatively higher contribution to oncotic pressure. However, the smallest molecules are also the first to be cleared from the bloodstream (i.e. they have the shortest half life). Moreover, when these smaller molecules are cleared, the overall contribution of the blood substitute to oncotic pressure decreases, resulting in increased edema. While initial resuscitation may be successful, it has been reported that some patients develop fatal pulmonary edema and delayed organ failure (Mendelson, 1975). Another consequence of the in vivo fractionation toward higher molecular weight species is that RBC aggregation is increased. This can lead to increase thrombus formation and decreased tissue perfusion. Oxygen carrying plasma expanders are the closest contemporary technology that has come to mimicking whole blood. Basically, two types of compounds have been the subject of intense investigation, perfluorocarbon liquids and hemoglobin. Perfluorocarbon liquids garnered much attention when first brought to light in dramatic experiments in which animals were ventilated in such liquids with reproducible survivability (Clark and Gollan, 1966; Clark et al., 1970), and in which the entire blood volume was replaced (Geyer, 1975). Since these early experiments, enthusiasm for the use of perfluorocarbon liquids as blood replacements has waned, and has instead focused on liquid ventilation applications where direct contact with the blood is avoided (Modell et al., 1976; Mottaghy et al., 1976; Schweiler and Robertson, 1976; Greenspan et al., 2000). The use of hemoglobin solutions has followed the most contiguous path from initial discovery to human clinical trials. In the early stages of this work, it was discovered that hemoglobin had the ability to transport oxygen in a similar manner to that of RBCs. Unfortunately, the half life of naked hemoglobin is relatively short, and it was concluded that clearance needed to be suppressed in order for efficacy to be extended. The primary mode of clearance is auto-oxidation followed by diafiltration through the kidneys, a problem which can be solved by increasing the molecular weight of the hemoglobin and maintaining its stability. This has been achieved by chemical modification in the form of crosslinking and conjugation to other polymeric molecules such as starch and dextran. Crosslinking hemoglobin molecules with diaspirin, glutaraldehyde, and other reagents has been described, and the results generally show an increase in stability and persistence in the circulation (Olsen et al., 1992). Unfortunately, the oxygen carrying capacity of compounds employing conjugation is negatively affected (Tam et al., 1976). In addition, some of these compounds have been shown to bind nitrous

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oxide (NO), a powerful vasodilator, leading to vasoconstriction and exacerbating delayed organ failure secondary to ischemia. Other reported complications include kidney and liver dysfunction, interference with blood typing, as well as the previously mentioned edema effect and associated coagulopathies. A generation of crosslinked hemoglobins has emerged from this early research with at least one product tested in pivotal preclinical trials (Gould et al., 1990), one in clinical trials (Bjorkholm et al., 2005), and another approved for use in humans (Lok, 2001). In the first study, primates subjected to total exchange transfusion maintained normal MAP, heart rate, cardiac output, and oxygen consumption. In a subsequent phase I clinical trial, no difference was noted between trauma patients receiving the pyridoxylated, glutaraldehyde crosslinked hemoglobin and allogeneic blood, thereby establishing baseline safety in a human clinical application (Gould et al., 1998). An additional in-hospital clinical trial has been conducted with similar results with regard to the safety of the product (Gonzalez et al., 1997). Since the phase I trial, developers of the product have initiated a phase III clinical trial. The study design calls for the pre-hospital use of either normal saline or crosslinked hemoglobin. Once in-hospital, the patients enrolled in the hemoglobin group will continue to receive the product unless opted out of the study. Despite progress in the aforementioned safety trials, investigation of the efficacy of similar hemoglobin derivatives has failed to provide compelling evidence for their clinical implementation. In one pivotal clinical trial of severe hemorrhagic shock, 48-h mortality, 28-day morbidity, and 28-day mortality rates were actually higher in patients receiving diaspirin crosslinked hemoglobin than in those receiving saline (Sloan et al., 1999). The entire field of blood substitutes, with particular emphasis on hemoglobin derivatives and clinical applications, has recently been reviewed by Winslow (2006). It would seem that the application of colloids to resuscitation the hemorrhagic patient is more art than science. Many recent reviews admit that the ideal blood substitute does not exist. One of the basic challenges, maintaining persistence in the vasculature and acting as a plasma surrogate without interfering with the physiological function of whole blood is a delicate balancing act. Almost every natural biomolecule known has been the focus of some experimentation in this field. It has not been until recently, however, that keratins have been assessed (Widra, 1986). In one experiment disclosed in this patent, an anesthetized female beagle was drained of 25% of its blood volume in 5 min. This volume was replaced with an equal volume of Normosol®-R pH 7.4 solution containing 2.5 wt/vol% of keratin. Physical and biochemical characteristics were monitored at various time intervals over a 24 h post-infusion period in the test subject. Blood pressure, serum total protein, red and white cell counts, cell morphology, immune cell counts, and blood chemistry were taken. The animal suffered no serious complications as a result of the keratin infusion and recovered fully. Although these data suggest the utility of using keratin solutions as plasma expanders, no characterization of the keratin used in the study has been reported. Keratins, particularly keratoses, like PEG and alginates have important new beneficial properties derived from their large hydrodynamic radius due to increased hydrophilicity. This characteristic leads to high viscosities at low concentrations and colloid osmotic pressure (COP), allowing one to titrate COP using conventional colloids (Cabrales et al., 2005). Keratins also display remarkable compatibility with the circulatory system, not instigating RBC aggregation at high molecular weights and concentrations. Most keratin solutions, however, are formulated at very low concentrations and do not provide increased COP by themselves. It may be that the perfect colloid does not exist, and that some combination of colloids may offer the best technology.

CARDIOVASCULAR BIOMECHANICS (FUNCTIONAL CAPILLARY DENSITY MODEL) The functional consequences of changing the flow properties of blood from normal conditions are not readily predictable due to the complexity and nature of the vascular wall and network, and the effects of shear stress at

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the endothelial surface in regulating vascular tone. In arterial blood vessels (diameter  100 μm), blood viscosity is proportional to hematocrit (Hct) squared, and in the smaller vessels it is linearly proportional to Hct. In the systemic circulation, Hct is approximately constant down to 100 μm diameter vessels. It falls monotonically down to the capillaries where it is approximately half of the systemic value. The reverse occurs in the venous circulation where it is higher than arterial because of fluid filtration in the microcirculation. In acute conditions, the decrease in Hct is not deemed dangerous until the transfusion trigger (blood hemoglobin content beyond which a blood transfusion is indicated) is reached. However, this exposes the vasculature to low blood viscosity when conventional plasma expanders are used to maintain blood volume. There appears to be no well-defined benefit to lowering blood viscosity, excepting when it is pathologically high, and lowering blood viscosity through hemodilution is considered to have no adverse effects. Richardson and Guyton (1959) determined that changes in blood viscosity are accompanied by compensatory changes in cardiac output, which compensate for changes in intrinsic oxygen carrying capacity of blood due to changes in Hct. This was confirmed systemically (Messmer, 1975), and in the microcirculation (Mirhashemi et al., 1988; Tsai et al., 1991). Empirically, the transfusion trigger is set at 7 g Hb/dl (Hct  22%). Microvascular Hcts are lower than systemic due to the presence of a plasma layer that proportionally occupies a greater relative volume of the vessel lumen, thus blood viscosity is also lower. The transition from macro- to microcirculation in terms of vessel dimensions, Hct, and hemodynamics is gradual. Blood rheological properties also change gradually and blood viscosity in the circulation depends on location. The reduction of Hct with a crystalloid or colloidal plasma expander tends to equalize the rheological properties of blood and viscosity throughout the circulation. When a plasma expander is used to remedy hemorrhage, systemic Hct decreases, significantly reducing blood viscosity in large vessels due to the squared dependence of viscosity on Hct. Viscosity of blood in small vessels is much less affected since Hct is lower than in large vessels. Conversely, small vessel blood viscosity is greatly influenced by the viscosity of the plasma expander. If plasma expander viscosity is low, blood viscosity drops significantly in the small vessels as well as in the large vessels, although for somewhat different reasons. In conventional theory, this reduction in viscosity increases blood flow and may improve the overall rate of oxygen delivery. However, the literature supports the concept that high viscosity plasma is either beneficial or has no adverse effect in conditions of extreme hemodilution. Waschke et al. found that cerebral perfusion is not changed when blood is replaced with fluids of the same intrinsic oxygen carrying capacity over a range of viscosities varying from 1.4 to 7.7 cp (Waschke et al., 1994). Krieter et al. (1995) varied the viscosity of plasma by adding dextran 500 kDa and found that medians in tissue pO2 in skeletal muscle were maximal at a plasma viscosity of 3 cp, while for liver the maximum occurred at 2 cp. In general they found that up to a three-fold increase in blood plasma viscosity had no effect on tissue oxygenation and organ perfusion when blood was hemodiluted. de Witt et al. (1997) found elevation of plasma viscosity causes sustained NO-mediated dilation in the hamster muscle microcirculation. Hct reductions should improve blood perfusion through the increase of blood fluidity. However at an Hct near to and beyond the transfusion trigger the heart cannot further increase flow and as viscosity falls, so does blood pressure. The fall of pressure is deleterious for tissue perfusion because it decreases functional capillary density (FCD) in the normal circulation and in hypotension following hemorrhage (Lindbom and Arfors, 1985). FCD is a critical microvascular parameter indicative of survival during acute blood loss. In hamsters subjected to 4 h 40 mmHg hemorrhagic shock, the fall of FCD accurately predicts outcome and separates survivors from non-survivors when this parameter decreases below 40% of control (Kerger et al., 1996). High viscosity plasma restores MAP in hypotension without vasoconstriction. Moreover, the shift of pressure and pressure gradients from the systemic to the peripheral circulation increases blood flow, which in combination with

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increased plasma viscosity maintains shear stress in the microcirculation. This is needed for shear stressdependent NO and prostaglandin release from the endothelium, and to maintain FCD (Frangos et al., 1985). Conversely, reduced blood viscosity decreases shear stress and the release of vasodilators, causing vasoconstriction and offsetting any benefit of reducing the rheological component of vascular resistance. Since resistance depends on the fourth power of vascular radius and the first power of blood viscosity, the effect of reducing blood viscosity with a low viscosity plasma expander is that it reduces oxygen delivery to the tissues once blood viscosity falls below a threshold value. This threshold has been determined in experimental models as about 2.5 cp (Kerger et al., 1996). Tissue perfusion with reduced blood viscosity may be deleterious at the cellular/endothelial level. There is evidence that genes are activated following changes in the mechanical environment of cells. It is also been established that the endothelium uniquely responds to changes in its mechanical and oxygen environment according to programmed genetic schemes. Among these responses is the mechanism for apoptosis (programmed cell death), which is activated through a genetically controlled suicide process that eliminates cells no longer needed or excessively damaged. In this context, hemodilution with low viscosity plasma expanders may cause cellular and tissue damage due to hypoxia and/or to the reduced vessel wall shear stress. Hypoxia/ischemia may also contribute to endothelial impairment due to inflammatory reactions. Activation of endothelium, platelets, and neutrophils, leading to additional damage through the liberation of cytokines, can induce endothelial apoptosis (Robaye et al., 1991). Studies in the hamster model show that extreme hemodilution (where Hct is 20% of control) with dextran 70 kDa causes hypotension and a drop in FCD to near pathological values (Tsai et al., 1998; Tsai, 2001). This is prevented by increasing plasma viscosity so that the diluted blood has a systemic viscosity of about 2.8 cp, which was achieved by infusing dextran 500 kDa. Thus, high viscosity plasma can be an alternative to the use of blood for maintaining MAP and an adequate level of FCD (Tsai and Intaglietta, 2001). This effect cannot be obtained by causing extreme hemodilution with a low viscosity plasma expander such as 6% dextran 70 Da, which lowers blood viscosity. When blood viscosity is lowered to 2 cp or less, FCD is no longer maintained and vascular diameter decreases. The use of plasma expanders for volume replacement beyond the transfusion trigger may reduce the need for blood transfusions. However, their use increases the risk of microvascular impairment due to lowered blood oxygen delivery and lowered shear stress. The presence of adequate levels of blood, as well as plasma viscosity are critical because they maintain the mechanical conditions in the microcirculation that insure physiological and normal microvascular function as expressed by the operation of normal FCD. Normal FCD is as critical, or perhaps more critical than oxygen delivery because non-functional capillaries prevent the extraction of slowly diffusible byproducts of metabolism from the tissue. Accumulation of these byproducts is toxic and can lead to focal, localized, and irreversible tissue losses that can ultimately lead to the failure of that organ. The more metabolically demanding the tissue, the more susceptible the organ to necrosis and failure during periods of acute ischemia (e.g. brain and heart). The restoration of adequate FCD via increased plasma viscosity is the only mechanism that has the potential of restoring tissue perfusion in all organs, including the heart and brain, thus allowing very small numbers of RBCs to deliver oxygen. In this scenario, oxygen delivery is seldom the limiting factor, while the impairment of microvascular function limits survival. Consequently, an improved plasma expander must possess specific viscogenic properties in order for resuscitation and tissue salvage to be successful.

CELLULAR-BASED BLOOD SUBSTITUTES Not all approaches to replacing the functionality of RBCs are biomaterials based. It has been suggested that the shortage of donated blood and challenges of long-term blood storage can be addressed by the process of

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generating RBC from stem cells. Stem cells have the potential to produce essentially unlimited quantities of tissue-matched RBC if their differentiation can be efficiently and reliably controlled. Several studies have focused on this effort for both RBC production as well as reconstitution of bone marrow. It was shown in 1991 that embryonic stem cells cultured in fetal calf serum demonstrated a capacity for erythropoiesis (Wiles and Keller, 1991). Moreover, these investigators showed that the efficiency of differentiation could be improved by adding erythropoietin to the culture medium. In a more recent study, hematopoietic stem cells isolated from human umbilical cord blood were expanded and differentiated into erythroid cells by sequential application of specific combinations of growth factors in serum-free media (Neildez-Nguyen et al., 2002). While the authors were not able to produce large quantities of mature RBC in vitro, differentiation of hematopoietic precursors into enucleated cells in vivo was demonstrated. Additional studies demonstrate the potential of human embryonic stem cells to differentiate into the hematopoietic lineage (Keller et al., 1993). The authors report that embryonic stem cells differentiate in vitro in a way that recapitulates days 6.5–7.5 of mouse hematopoietic development. Further, those embryonic stem cells differentiated as embryoid bodies develop erythroid precursor cells by day 4 of culture, and that by day 6 greater than 85% of embryoid bodies contain these cells. Using a different cell culture approach, another group of investigators was able to produce similar results (Nakano et al., 1994). The authors report an efficient system for the differentiation of embryonic stem cells into blood cells of erythroid, myeloid, and B cell lineages by coculture with a stromal cell feeder layer. These studies demonstrate the potential of various stem cells to be used in the production of erythroid precursors and perhaps even mature RBC. However, more research is needed to increase the efficiency of differentiation before preclinical testing can be undertaken. If cost effective methods of large scale production can be realized, stem cells may provide a viable source of RBCs for future clinical application.

CLINICAL TRIALS – ETHICAL CONSIDERATIONS As mentioned previously, the advent of clinical trials using crosslinked hemoglobins has raised several serious ethical concerns that have garnered much attention. At issue are two forms of glutaraldehyde crosslinked hemoglobin-based products, one human derived and the other bovine derived. The human derived product is currently being used in a phase III clinical trial. It was recently reported that the company developing the product failed to publicly disclose that in the phase II trial, there was an increased incidence of heart attack in aneurism surgery patients receiving the human hemoglobin product and that the trial was halted before completion (Burton, 2006). The company’s website indicates that the target patient enrollment has been met in the current phase III trial and results are due to be reported in the fall of 2006 (http://www.northfieldlabs.com). The bovine derived product was actually approved for in-hospital use first, in South Africa (Lok, 2001). This product has since been investigated for use in open heart surgery for the purpose of reducing the need for whole blood transfusions (Levy et al., 2002). Leading up to an FDA Blood Products Advisory Committee meeting to discuss the application for a phase III clinical trial on the bovine derived blood substitute, several bioethicists weighed in on the controversy surrounding these products (Dalton, 2006; Kipnis et al., 2006; Guterman, 2006). At issue is the dilemma of informed consent in trauma trials. In these types of clinical studies, the patient is often unconscious or otherwise incapacitated and informed consent is not feasible. FDA regulations allow for such trials to be conducted under a waiver of informed consent provided that no superior treatment exists, and that the sponsor actively pursues “community consultation.” The community consultation requirement necessitates that the sponsoring organization make reasonable efforts to inform the community in which clinical trials are to take place of the details of the trial and other pertinent information, such as results of previous investigations (Schmidt et al., 2006; Richardson et al., 2006).

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Members of the community can opt out of the trial, typically by wearing a wristband or necklace that indicates their desire to avoid enrollment. In the aforementioned clinical trials involving crosslinked hemoglobin products, the nature and completeness of the community consultation that preceded the clinical trials, as well as the design of the studies, was criticized. With regard to the latter, both trials required the use of the hemoglobin product in the field for obvious reasons. However, once the patients reach the hospital the study investigators sought to continue the use of the blood substitute in the absence of refusal from the patient or patient’s advocate, despite the availability of allogeneic blood. Critics argued that this violated the rules of emergency consent waiver and that these patients must discontinue the use of the blood substitute upon admission in lieu of more established therapies. However, such an approach would confound the data and preclude testing of the study hypothesis. The FDA, sponsoring companies, and investigators must come to a consensus on how such trials are conducted so that scientifically relevant data can be acquired and used to advance the development of blood substitutes.

REFERENCES Bjorkholm, M., Fagrell, B., Przybelski, R., Winslow, N., Young, M. and Winslow, R.M. (2005). A phase I single blind clinical trial of a new oxygen transport agent (MP4), human hemoglobin modified with maleimide-activated polyethylene glycol. Haematologica 90: 505–515. Brecher, M.E., Owen, H.G. and Bandarenko, N. (1997). Alternatives to albumin: starch replacement for plasma exchange. J. Clin. Apher. 12: 146–153. Burton, T.M. (2006). FDA to weigh using fake blood in trauma trial. Wall St. J. (East Ed) July 6: B1, B2. Cabrales, P., Tsai, A.G. and Intaglietta, M. (2005). Alginate plasma expander maintains perfusion and plasma viscosity during extreme hemodilution. Am. J. Physiol. Heart Circ. Physiol. 288: H1708–H1716. Clark Jr., L.C. and Gollan, F. (1966). Survival of mammals breathing organic liquids equilibrated with oxygen at atmospheric pressure. Science 152: 1755–1756. Clark Jr., L.C., Kaplan, S., Becattini, F. and Benzing III, G. (1970). Perfusion of whole animals with perfluorinated liquid emulsions using the Clark bubble-defoam heart–lung machine. Fed. Proc. 29: 1764–1770. Dalton, R. (2006). Trauma trials leave ethicists uneasy. Nature 440: 390–391. Data provided by America’s Blood Centers; http://www.americasblood.org. Davidson, I., Gelin, L.E. and Haglind, E. (1980). Plasma volume, intravascular protein content, hemodynamic and oxygen transport changes in dogs: comparison of relative effectiveness of various plasma expanders. Crit. Care Med. 8: 73–80. de Witt, C., Schafer, C., von Bismark, P., Bolz, S.S. and Pohl, U. (1997). Elevation of plasma viscosity induces sustained NO-mediated dilation in the hamster cremaster microcirculation in vivo. Pflugers Arch. 434: 354–361. Frangos, J.A., Eskin, S.G., McIntire, L.V. and Ives, C.L. (1985). Flow effects on prostacyclin production in cultured human endothelial cells. Science 227: 1477–1479. Geyer, R.P. (1975). “Bloodless” rats through the use of artificial blood substitutes. Fed. Proc. 34: 1499–1505. Gonzalez, P., Hackney, A.C., Jones, S., Strayhorn, D., Hoffman, E.B., Hughes, G., Jacobs, E.E. and Orringer, E.P. (1997). A phase I/II study of polymerized bovine hemoglobin in adult patients with sickle cell disease not in crisis at the time of study. J. Investig. Med. 45: 258–264. Gould, S.A., Sehgal, L.R., Rosen, A.L., Sehgal, H.L. and Moss, G.S. (1990). The efficacy of polymerized pyridoxylated hemoglobin solution as an O2 carrier. Ann. Surg. 211: 394–398. Gould, S.A., Moore, E.E., Hoyt, D.B., Burch, J.M., Haenel, J.B., Garcia, J., DeWoskin, R. and Moss, G.S. (1998). The first randomized trial of human polymerized hemoglobin as a blood substitute in acute trauma and emergent surgery. J. Am. Coll. Surg. 187: 113–122. Greenspan, J.S., Wolfson, M.R. and Shaffer, T.H. (2000). Liquid ventilation. Semin. Perinatol. 24: 396–405. Guterman, L. (2006) Artificial-blood study has critics seeing red. Chron. High Educ. 52: A17. http://www.northfieldlabs.com. Keller, G., Kennedy, M., Papayannopoulou, T. and Wiles, M.V. (1993). Hematopoietic commitment during embryonic stem cell differentiation in culture. Mol. Cell Biol. 13: 473–486.

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Kerger, H., Saltzman, D.J., Menger, M.D., Messmer, K. and Intaglietta, M. (1996). Systemic and subcutaneous microvascular pO2 dissociation during 4 h hemorrhagic shock in conscious hamsters. Am. J. Physiol. 270: H827–H836. Kipnis, K., King, N.M. and Nelson, R.M. (2006). An open letter to institutional review boards considering Northfield Laboratories’ PolyHeme trial. Am. J. Bioeth. 6: 18–21. Krieter, H., Bruckner, U.B., Kafaliakis, F. and Messmer, K. (1995). Does colloid induced plasma hyperviscosity in haemodilution jeopardize perfusion and oxygenation of vital organs? Acta Anaesth. Scad. 39: 326–344. Levy, J.H., Goodnough, L.T., Greilich, P.E., Parr, G.V., Stewart, R.W., Gratz, I., Wahr, J., Williams, J., Comunale, M.E., Doblar, D., Silvay, G., Cohen, M., Jahr, J.S. and Vlahakes, G.J. (2002). Polymerized bovine hemoglobin solution as a replacement for allogeneic red blood cell transfusion after cardiac surgery: results of a randomized, double-blind trial. J. Thorac. Cardiovasc. Surg. 124: 35–42. Lindbom, L. and Arfors, K.E. (1985). Mechanism and site of control of variation in the number of perfused capillaries in skeletal muscle. Int. J. Microcirc. Clin. Exp. 4: 121–127. Lok, C. (2001). Blood product from cattle wins approval for use in humans. Nature 410: 855. Mendelson, J.A. (1975). The selection of plasma volume expanders for resuscitation following trauma: a review. Mil. Med. 140: 258–262. Messmer, K. (1975). Hemodilution. Surg. Clin. N. Am. 55: 659–678. Mirhashemi, S., Breit, G.A., Chavez, R.H. and Intaglietta, M. (1988). Effects of hemodilution on skin microcirculation. Am. J. Physiol. 254: H411–H416. Mishler, J.M., Borherg, H., Emerson, P.M. and Gross, R. (1977). Hydroxyethyl starch: an agent for hypovolemic shock treatment. J. Surg. Res. 23: 239–245. Modell, J.H., Calderwood, H.W., Ruiz, B.C., Tham, M.K. and Hood, C.I. (1976). Liquid ventilation of primates. Chest 69: 79–81. Mottaghy, K., Mendler, N., Schmid-Schonbein, H., Schrock, R. and Sebening, F. (1976). A new type of fluorocarbon liquid oxygenator. Eur. Surg. Res. 8: 196–203. Nakano, T., Kodama, H. and Honjo, T. (1994). Generation of lymphohematopoietic cells from embryonic stem cells in culture. Science 265: 1098–1101. Nearman, H.S. and Herman, M.L. (1991). Toxic effects of colloids in the intensive care unit. Crit. Care Clin. 7: 713–723. Neildez-Nguyen, T.M., Wajcman, H., Marden, M.C., Bensidhoum, M., Moncollin, V., Giarratana, M.C., Kobari, L., Thierry, D. and Douay, L. (2002). Human erythroid cells produced ex vivo at large scale differentiate into red blood cells in vivo. Nat. Biotechnol. 20: 467–472. Olsen, K.W., Zhang, Q.Y., Huang, H., Sabaliauskas, G.K. and Yang, T. (1992). Stabilities and properties of multilinked hemoglobins. Biomater. Artif. Cells Immobilization Biotechnol. 20: 283–285. Richardson, L.D., Rhodes, R., Ragin, D.F. and Wilets, I. (2006). The role of community consultation in the ethical conduct of research without consent. Am. J. Bioeth. 6: 33–35. Richardson, T.Q. and Guyton, A.C. (1959). Effects of polycythemia and anemia on cardiac output and other circulatory factors. Am. J. Physiol. 197: 1167–1170. Robaye, B., Mosselams, R., Fiers, W., Dumont, J.E. and Galand, P. (1991). Tumor necrosis factor induces apoptosis (programmed cell death) in normal cells in vitro. Am. J. Pathol. 38: 447–453. Roberts, J.S. and Bratton, S.L. (1998). Colloid volume expanders. Problems, pitfalls and possibilities. Drugs 55: 621–630. Schmidt, T.A., Delorio, N.M. and McClure, K.B. (2006). The meaning of community consultation. Am. J. Bioeth. 6: 30–32. Schwieler, G.H. and Robertson, B. (1976). Liquid ventilation in immature newborn rabbits. Biol. Neonate 29: 343–353. Sloan, E.P., Koenigsberg, M., Gens, D., Cipolle, M., Runge, J., Mallory, M.N. and Rodman Jr., G. (1999). Diaspirin crosslinked hemoglobin (DCLHb) in the treatment of severe traumatic hemorrhagic shock: a randomized controlled efficacy trial. JAMA 282: 1857–1864. Szeto, C.C. and Chow, K.M. (2005). Nephrotoxicity related to new therapeutic compounds. Ren. Fail. 27: 329–333. Tam, S.C., Blumenstein, J. and Wong, J.T. (1976). Soluble dextran–hemoglobin complex as a potential blood substitute. Proc. Natl Acad. Sci. USA 73: 2128–2131. Tsai, A.G. (2001). Influence of cell-free hemoglobin on local tissue perfusion and oxygenation after acute anemia after isovolemic hemodilution. Transfusion 41: 1290–1298.

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Tsai, A.G. and Intaglietta, M. (2001). High viscosity plasma expanders: volume restitution fluids for lowering the transfusion trigger. Biorheology 38: 229–237. Tsai, A.G., Arfors, K.E. and Intaglietta, M. (1991). Spatial distribution of red blood cells in individual skeletal muscle capillaries during extreme hemodilution. Int. J. Microcirc. Clin. Exp. 10: 317–334. Tsai, A.G., Friesenecker, B., Mazzoni, M.C., Kerger, H., Buerk, D.G., Johnson, P.C. and Intaglietta, M. (1998). Microvascular and tissue oxygen gradients in the rat mesentery. Proc. Natl Acad. Sci. USA 95: 6590–6595. Waschke, K.F., Krieter, H., Hagen, G., Albrecht, D.M., Van Ackern, K. and Kuschinsky, W. (1994). Lack of dependence of cerebral flow on blood viscosity after blood exchange with a Newtonian O2 carrier. J. Cereb. Blood Flow Metab. 14: 871–976. Widra, A. (1986). US Patent 4,570,629. Wiles, M.V. and Keller, G. (1991). Multiple hematopoietic lineages develop from embryonic stem (ES) cells in culture. Development 111: 259–267. Winslow, R.M. (ed.) (2006). Blood Substitutes. San Diego, CA: Academic Press.

45 Articular Cartilage Francois Ng kee Kwong and Myron Spector

INTRODUCTION Types of Articular Cartilage Defects That Present in the Clinic Cartilage defects are a common source of pain and/or loss of function in patients presenting to the orthopedic clinic. While, any joint can be affected, the joint most commonly affected is by far the knee. A chondral lesion was found in 63% of a large series of over 31,000 arthroscopic procedures performed in patients with a symptomatic knee (Curl et al., 1997). Articular cartilage damage is often associated with meniscal and anterior cruciate ligament injuries (Shelbourne et al., 2003). These defects can be divided according to their etiology or morphology. Focal injuries typically occur as a result of a sporting injury and hence tend to affect the younger population. Focal defects can be further subdivided into chondral or osteochondral lesions, depending on the depth of the defect. Chondral lesions, also known as partial thickness lesions, lie entirely within the cartilage and do not penetrate into the sub-chondral bone. In the adult, defects of this nature do not regenerate because of the lack of cells which could participate in the repair process. Osteochondral defects penetrate through the vascularized sub-chondral bone and some spontaneous repair occurs as mesenchymal chondroprogenitor cells invade the lesion and form cartilage. However, full-thickness defect repair is only transient and the novel tissue formed does not have the functional properties of native hyaline cartilage (Shapiro et al., 1993). On the other hand, degenerative chondral changes typically occur in the older population as a result of arthritic changes. They often involve a large area of the affected joint, but start off as a focal lesion initially. Rationale for Cell Therapy Articular cartilage has a limited capacity for self-regeneration after injury. This was recognized as early as in 1743 by Hunter who stated that cartilage “once destroyed is not repaired.” This is because none of the normal inflammatory and reparative processes of the body are available to repair the tissue. This itself is a result of its isolation from the systemic regulation, lack of blood vessels, and nerve supply (Mankin, 1982). Furthermore, chondrocytes which are surrounded by an extracellular matrix cannot freely migrate to the site of injury from an intact healthy site, unlike most tissues (Buckwalter and Mankin, 1998), and there is no provisional fibrin clot filling the defect into which cells can migrate. Full-thickness defects induce mesenchymal chondroprogenitor cells to differentiate into repair tissue, but this is predominantly fibrous in nature and degenerates with time. The two major problems that need to be addressed in repair of articular cartilage are the filling of the defect void with a tissue that has the same mechanical properties as articular cartilage and the promotion of successful integration between the repair tissue and the native articular cartilage and calcified cartilage. Even

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a small defect caused by mechanical damage will fail to heal and degenerate over time progressing to osteoarthritis (OA). Conventional surgical techniques of cartilage repair are partially successful in alleviating symptoms, but fail to regenerate tissue anywhere similar in nature to native articular cartilage. There was no promising solution to this problem until Brittberg et al. introduced a cell-based therapy in which culture-expanded chondrocytes were transplanted into defects, raising the expectations of a breakthrough in repairing damaged articular cartilage (Brittberg et al., 1994). Current Cell Therapies Available in the Clinic The possible cell-based tissue repair techniques can be broadly classified into three major categories: (1) targeting local connective tissue progenitors where new tissue is desired, (2) transplanting culture-expanded or modified connective tissue progenitors, and (3) transplanting fully formed tissue generated in vitro or in vivo. In current clinical practice, the first two techniques are already in use while the last one is being actively investigated in animal models and pre-clinical trials. These techniques are generally aimed at delivering chondrogenic cells to the cartilage defect, either in the form of tissues containing precursor cells (e.g. the periosteum or perichondrium) or in the form of autologous chondrocytes isolated from a biopsy of healthy cartilage and expanded in number in vitro. Periosteal Transplantation Rubak initially described this technique in a rabbit model of cartilage defect (Rubak, 1982). He used a periosteal flap to cover the defects. The defects were repaired and filled after 4 weeks with a hyaline-like cartilage whereas the empty control defect showed fibrocartilage-like repair tissue. The first clinical study was published by Niedermann et al. who reported successful results in all of their four initially treated patients (Niedermann et al., 1985). Perichondrial Transplantation Autologous perichondrium has also been employed for cartilage repair (Homminga et al., 1989, 1990, 1991). Perichondrium, taken from the cartilaginous covering of the rib, is placed into the chondral defect of the affected joint. The first clinical study of this approach was performed by Homminga et al. (1990). A major shortcoming of perichondrial grafting is the limited availability of large grafts. Graft size is limited to the rib size, so that several rib perichondrial grafts have to be harvested to fill a large defect. Additionally, endochondral ossification and delamination of the cartilage from the sub-chondral bone plate are potentially significant limitations to the long-term efficacy of this repair. Autologous Chondrocyte Implantation Since first published in 1994 (Brittberg et al., 1994), techniques of cell isolation, expansion in culture, and implantation have remained essentially the same. Cartilage (150–300 mg) is harvested arthroscopically from a minimally load-bearing area of the upper aspect or the medial condyle of the affected knee. The biopsy is then transported to a laboratory facility using a transport media. Chondrocytes are isolated using standard techniques. After a certain period of cell expansion (11–21 days (Peterson et al., 2000), depending upon the growth kinetics) a certain number of cells (e.g. minimally 12 million for Genzyme’s Carticel procedure) are provided in a serum-free and gentamycin-free transport medium.

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Using a medial or lateral parapatellar incision, the defect is debrided to the level of normal-appearing surrounding cartilage. The integrity of the tidemark needs to be maintained in order to avoid infiltration of undifferentiated mesenchymal stem cells (MSCs) which could contribute to the formation of fibrocartilagenous repair tissue (Brittberg et al., 1999). A periosteal flap is harvested from the anterior aspect of the proximal tibia or distal femur, formed to the shape of the lesion, and sutured to the rim of the defect. The chondrocyte suspension is subsequently injected under the periosteal flap and the border of periosteal cover sealed using fibrin glue. Post-operative rehabilitation protocols generally involve continuous passive motion and limited weight bearing for an extended time. Cooperation of the patient in this respect is essential for a favorable outcome, hence difficult to control. This contributes to difficulty in evaluating outcome data. In a variation of this technique, porcine type I/III collagen membrane has been used in place of the periosteal membrane (Bartlett et al., 2005). Its outer surface is smooth, giving a low-friction surface. Its inner surface is rough because of large gaps between collagen fibers into which chondrocytes can be seeded.

CELL THERAPIES An optimal cell source should have the following characteristics: no immunorejection, no tumorigenicity, immediate availability, availability in pertinent quantities, controlled cellular proliferation rate, predictable, and consistent chondrogenic potential as well as controlled integration into the surrounding tissues. Autologous versus Allogeneic An autologous source of stem cells is most desirable as cells are collected from each patient, thereby eliminating complications associated with immune rejection of allogeneic tissue. Even with an autologous system, challenges exist in assuring a safe and reproducible product. Genzyme established a quality assurance program based on US FDA Good Manufacturing Practice regulations, which was reviewed recently (Mayhew et al., 1998). Process variables have to be controlled rigorously and sterility testing and endotoxin testing maintained. Moreover, assessments of cell viability and growth kinetics are a crucial part of non-conformance reporting. According to Genzyme data, 1.64% of the cartilage biopsies received were contaminated (Mayhew et al., 1998). Contamination was recorded only for 0.03% during processing and in 0.16% at release. Endotoxin content ranged between less than 0.15 and 0.5 EU/ml (allowable limit 82.5 EU/ml) and cell viability was 90.9 4.06% at release. Measurement of growth kinetics revealed 0.311 doublings per day. Out of 1377 cartilage biopsies, 86 non-conformances were identified related to biopsy quality, only 12 were related to cell processing. Limitations of the autologous approach in obtaining stem cells and the desire to obtain “marketable products” which could benefit as many patients as possible have provided incentives for the development of generic cell lines, which can be taken off the shelf as, and when, needed for patient treatment. These universal cells would have the following advantages: (i) availability through the development of large cell banks; (ii) consistency and efficacy because only cells with desirable characteristics and controlled critical parameters are selected and amplified; and (iii) sterility and assurance of compatibility through extensive safety testing. Until recently, it was difficult to envision utilization of allogeneic generic cells in orthopedics as it was believed that their transplantation would require immunosuppressive drugs to reduce associated risks of rejection. However, cultured MSCs exhibit a poorly immunogenic phenotype (Tse et al., 2003). In vivo, a single intravenous administration of MSCs led to a modest, but significant, prolongation of skin graft survival (Bartholomew et al., 2002). These data have greatly enhanced the therapeutic appeal of MSCs because they raised the possibility of creating universal cell lines. Indeed, allogeneic adult stem cells are already being investigated in patients with meniscal injuries, in a phase 1 FDA approved clinical trial (http://www.osiristx.com/).

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Intra-operative versus Culture Expanded Intra-operative cell-based therapies have the advantage of being less time consuming and less costly than ex vivo therapies. Ex vivo therapies also have the disadvantage of involving an additional harvesting step. The advantage of an ex vivo technique is that the surgeon can select specific cells (i.e. bone marrow cells or stem cells) and the cellular delivery vehicle for specific clinical problems. It is also safer than an in vivo strategy when working with viruses for gene therapy because no viral particles or DNA complexes are injected directly into the body. In addition, ex vivo strategies have a high efficiency of cell transduction. Articular Chondrocytes Methods for Intra-operative Cell Therapy Osteochondral transplantation has been used clinically for more than 25 years. Large osteochondral allografts have been employed for orthopedic tumor surgery and to a lesser extent for repairing degenerative defects. However, for smaller defects these procedures introduced significant morbidity. More recently osteochondral autografting has been introduced into the clinic as an alternative treatment for small and medium sized defects. Promising reports by Matsusue and Bobic have fueled interest in this method (Matsusue et al., 1993; Bobic, 1999). With this technique an osteochondral plug is harvested from a lower weight-bearing area of the knee joint and transferred to the prepared defect, implanted using a press-fit technique. Culture-Expanded Cells The rationale for using articular chondrocytes for a cell-based therapy is that they already possess the desired phenotype. Chondrocytes comprise the single cellular component of adult hyaline cartilage and are considered to be terminally differentiated, thus being highly specialized. Their main function is to maintain the cartilage matrix, synthesizing-types II, IX, and XI collagen; the large aggregating proteoglycan, aggrecan; the smaller proteoglycans, biglycan and decorin; and specific and non-specific matrix proteins that are expressed at defined stages during growth and development. Freshly isolated articular chondrocytes continue to exhibit their specific phenotype in culture for at least several days to weeks. This makes them a suitable cell type for a cell-based treatment of chondral defects. While the steps involved in the isolation and expansion of articular chondrocytes for autologous chondrocyte implantation (ACI) are quite similar among various commercial and academic laboratories, there may be important differences. One such difference is the use of the patient’s own serum for culturing the cells, as described originally by Brittberg et al. (1994). One commercial enterprise, Genzyme Biosurgery (Cambridge, Massachusetts, USA), uses approved and validated fetal bovine serum (FBS), instead of the patient’s serum, in the culture media. Another potentially important difference is that Genzyme needs to freeze and store the isolated cells in order to allow for verification of adequate insurance coverage prior to the implantation procedure. A recent study has indicated that this freeze-thaw cycle may adversely affect the outcome of the procedure (Perka et al., 2000). Cryopreserved chondrocytes seeded into polymer scaffolds yielded an 85% repair of an osteochondral defect in rabbits, whereas 100% of the defects treated with noncryopreserved cells were filled. One of the disadvantages of employing articular chondrocytes is that they do not readily proliferate in vitro. Cells from a younger population have been found to undergo 0.3 doublings per day, using a standardized and validated approach for culturing cells for later implantation (Mayhew et al., 1998). Even lower proliferation rates are obtained in older patients and arthritic cartilage (Peterson et al., 2000). Another report demonstrates the rapid replicative senescence of articular chondrocytes (Martin and Buckwalter, 2001). Once chondrocytes are deprived of their three-dimensional environment, their phenotype switches to a more

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fibroblastic cell form, expressing types I and III collagen, instead of cartilage-specific type II collagen (Goldring et al., 1986, 1988; Saadeh et al., 1999). Decades ago (in 1922) it was reported that the removal of articular cartilage from the joints of rabbits led to the formation of bone around the margins of the joint (Fisher, 1922). Fisher suggested that the results were due to reactive changes brought about by the removal of cartilage from the joint and not due to a response to necrotic cartilage in the joint. This early study was followed up by another in which osteochondral defects were created in the patella surface of the femurs of rabbit knee joints (Key, 1931). In some instances “the operation was followed by a severe chronic, progressive arthritis which involved not only the femur, but also the tibia and patella.” The author noted that “the most interesting changes were the hyperplastic phenomena which occurred in the lower end of the femur. These changes were not continuous with or even adjacent to the defect, but occurred in the non-traumatized portions of the lower end of the femur, and in many instances both the patella and the upper end of the tibia were also involved.” These changes were present to some degree in every joint. The author noted that the experiments prove “that many of the changes which occur in the hypertrophic arthritis can be produced experimentally in the joints of rabbits by simply creating a defect in the cartilage and that these changes are not dependent upon the presence of dead cartilage within the joint.” Our own studies demonstrated in a canine model that the harvesting of articular cartilage predisposes the other cartilage in the same joint to changes associated with early OA (Lee et al., 2000). While the lesion itself in a knee joint may serve to induce such osteoarthritic changes in the joint, the additional surgical procedure of harvesting cartilage may exacerbate the condition. There is, then, a compelling need for an alternative cell source for a cell-based cartilage repair procedure. MSCs MSCs isolated from the bone marrow and other sources can provide an alternative and abundant supply of cells for cartilage repair procedures. Adult marrow stromal cells are being investigated for the treatment of defects in connective tissues using cell and gene therapy and tissue engineering approaches – see for reviews (Caplan, 1991; Prockop, 1997). Differentiation of such cells can be obtained in vitro by changing the culture conditions after their expansion or in vivo as a consequence of the new “physiological” microenvironment in the transplant area. Whole Marrow Implants Safety of Whole Marrow Injected/Implanted in Human Subjects Whole autologous and allogeneic bone marrow has been injected and implanted into human subjects for decades to treat myriad medical problems with no adverse events associated with the MSC sub-population present. Of note, for example, is the procedure in which up to 1 liter of whole bone marrow is routinely infused into the bone marrow transplant patient. This infusion contains a small but significant proportion of MSCs and does not seem to have any significant side-effects. In an example of one such study in which the MSC population of whole marrow was to provide the principal therapeutic effect (Horwitz et al., 1999), non-manipulated bone marrow from HLA-identical or singleantigen-mismatched siblings was intravenously infused into three children with severe deforming osteogenesis imperfecta after they had received ablative conditioning therapy. The nucleated cell doses ranged from 5.7 to 7.5  108 cells/kg. All three showed engraftment with hemopoietic donor cells. Improvements in clinical outcome were associated with increases in growth velocity and reduced frequencies of bone fracture. The authors concluded that “allogeneic bone marrow transplantation can lead to engraftment of functional

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mesenchymal progenitor cells, indicating the feasibility of this strategy in the treatment of osteogenesis imperfecta and perhaps other MSCs disorders as well.” There were no reports of adverse response to the marrow infusion. For decades an array of “marrow stimulation” techniques, including abrasion arthroplasty, drilling, and micro-fracture, have been used to treat cartilage defects. Each of these methods introduces marrow-derived MSCs into the joint. While these procedures have not yielded lasting symptomatic relief they demonstrate that the presence of endogenous bone marrow-derived MSCs in the joint does not lead to adverse clinical sequelae. Since the early days of bone grafting autogenous marrow has been known to be of value in improving the osteogenic response (Salama et al., 1973). Whole autogenous marrow has been implanted in various sites in the body with no untoward clinical findings. In more recent years bone marrow and bone marrow fractions including the stromal cell population have been injected percutaneously to treat non-unions in human subjects (Connolly et al., 1998). There have been no adverse events reported. An apparatus has become commercially available (Select, DePuy Acromed Inc,) for the intra-operative concentration of MSCs/osteoprogenitor cells from whole marrow.

Efficacy of Whole Marrow for Cartilage Repair in Pre-clinical Animal Studies

The rationale for the benefits to be derived from MSCs also draws from investigations demonstrating the contribution of whole marrow to cartilage repair. In one such study (Solchaga et al., 2002), autologous bone marrow incorporated into a fibronectin-coated hyaluronan-based sponge was implanted into 3-mm diameter osteochondral defects in a rabbit model. Control groups were implanted with the scaffold alone. Except for the 3-week specimens, the histological appearance of the defects was similar in both groups. “Four weeks after surgery, the defects were filled with bone with a top layer of cartilage well integrated with the adjacent cartilage. At each harvest time, the overall histological scores of the specimens did not reveal statistical differences between the treatment groups. However, as revealed by the results of the 3-week sacrifices, bone marrow loading appeared to accelerate the first stages of the repair process.” Coagulated bone marrow aspirates have been used together with gene therapy techniques in a rabbit model of cartilage defect (Pascher et al., 2004). Mixture of an adenoviral suspension with the fluid phase of freshly aspirated bone marrow resulted in uniform dispersion of the vector throughout and levels of transgenic expression in direct proportion to the density of nucleated cells in the ensuing clot. Furthermore, cultures of MSCs previously transduced ex vivo with recombinant adenovirus were readily incorporated into the coagulate when mixed with fresh aspirate. These vector-seeded and cell-seeded bone marrow clots were found to maintain their structural integrity following extensive culture and maintained transgenic expression in this manner for several weeks. These genetically modified bone marrow clots were able to generate similarly high levels of transgenic expression in osteochondral defects with better containment of the vector within the defect. In a rodent pre-clinical model, Gurevitch and colleagues demonstrated that implantation of a composite comprising demineralized bone matrix and a bone marrow cell suspension in a damaged area of a joint resulted in the generation of a new osteochondral complex comprising articular cartilage and sub-chondral bone (Gurevitch et al., 2003). In the same study, the authors implanted the same composite material into an ablated bone marrow cavity and a calvarial defect (Gurevitch et al., 2003). The resulting tissue formed was respectively trabecular bone and stromal microenvironment supporting hematopoiesis and flat bone, respectively. They concluded that the new tissue formation followed differentiation pathways controlled by site-specific physiological conditions, thus developing tissues that precisely met local demands.

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Methods for Culture Expansion Sources: Bone Marrow, Fat, Muscle Use of Bone Marrow as the Source of MSCs: Other Sources of MSCs Human MSCs which have been reported to be present in bone marrow, adipose tissues, dermis, muscles, and peripheral blood (Young et al., 2001) and other connective tissue (Zohar et al., 1997) have the potential to differentiate along different lineages including those forming bone, cartilage, fat, muscle, and nerve. Several studies have compared the characteristics of MSCs from these different sources. One such study (Lee et al., 2004) compared phenotypes and gene expression profile of the human adipose tissue-derived stromal cells (ATSCs) in the undifferentiated states with bone marrow-derived MSCs. Both cell types expressed CD29, CD44, CD90, CD105 and were absent for HLA-DR and c-kit expression. The study confirmed that the marrow-derived MSCs were inducible to differentiate into osteoblasts, adipocytes, and chondrogenic lineages. While the results showed that ATSCs were superior to marrow-derived MSCs with respect to maintenance of proliferating ability, “the proliferating ability and differentiation potential of ATSC were variable according to the culture condition.” That the phenotypes and the gene expression profiles of ATSCs and marrow-derived MSCs were found to be similar may not provide enough of a compelling argument for the use of ATSCs, particularly because of the fact that there are many more safety and efficacy studies of marrow-derived MSCs compared to ATSCs. Culture Procedures

MSCs represent a minor fraction of the total nucleated cell population in the marrow. They can be plated and enriched using standard cell culture techniques. Frequently, the whole marrow sample is subjected to fractionation on a density gradient solution such as Ficoll, after which the cells are plated at densities ranging from 1  104 cells/cm2 to 0.4  106 cells/cm2 (Pittenger et al., 1999; Lodie et al., 2002; McBride et al., 2003). Cells are generally cultured in basal medium such as Dulbecco’s modified Eagle’s medium (low glucose) in the presence of 10% FBS (Pittenger et al., 1999). MSCs in culture have a fibroblastic morphology and adhere to the tissue culture substrate. Primary cultures are usually maintained for 12–16 days, during which time the non-adherent hematopoietic cell fraction is depleted. Optimal expansion of MSCs from marrow requires the pre-selection of FBS. As MSCs are expanded in large-scale culture for human applications it will be important to identify defined growth media, without or with reduced FBS, to ensure more reproducible culture techniques and enhanced safety. Safety of MSCs in Animal Models The use of culture-expanded MSCs in animal models has recently been reviewed (Barry, 2003, #14598). Several studies have focused on the use of monolayer-expanded bone marrowderived MSCs as a renewable and readily accessible source for the treatment of infarcted cardiac tissue. Studies that have injected MSCs in mouse models of myocardial infarcts have not reported adverse effects (Orlic et al., 2001). In other work (Murphy et al., 2003) autologous culture-expanded MSCs were injected into the knee joints of goats in which OA was induced by complete excision of the medial meniscus and resection of the anterior cruciate ligament. Six weeks after induction of OA, a single dose of 10 million MSCs, suspended in a dilute solution of sodium hyaluronan, was delivered to the injured knee by direct intra-articular injection. Control animals received sodium hyaluronan alone. “In cell-treated joints, there was evidence of marked regeneration of the medial meniscus, and implanted cells were detected in the newly formed tissue. Degeneration of the articular cartilage, osteophytic remodeling, and sub-chondral sclerosis were reduced in cell-treated joints compared with joints treated with vehicle alone without cells.” “Animals tolerated the cell injection well, and there was no evidence of local inflammation, immobilization, or unloading of the joint resulting from the cell treatment.”

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Efficacy of MSCs for Cartilage Repair in Pre-clinical Animal Studies That MSCs may yield results comparable to autologous chondrocytes was supported by in vitro studies (Kavalkovich et al., 2002) that have demonstrated that MSC cultures undergoing chondrogenesis synthesize glycosaminoglycan (GAG) at levels significantly higher than explant cultures or primary chondrocyte cultures. Numerous studies in vivo (Caplan, 1991) have supported the supposition that these bone marrow-derived MSCs offer advantages over committed cells (i.e. differentiated cells such as articular chondrocytes) for cellseeded implants developed to facilitate tissue regeneration (e.g. articular cartilage) (Wakitani et al., 1994; Ponticiello et al., 2000). This strategic approach holds that regeneration can be facilitated by the recapitulation of certain phases of embryonic development, and that these stem cells will allow for such, whereas fully differentiated cells will not. Presumably the endogenous regulators in the implant site will serve to induce the implanted undifferentiated stem cells to differentiate along the desired pathway. In one study (Im et al., 2001) using mature rabbits, bone marrow-derived MSCs expanded in culture in monolayer were implanted into a full-thickness osteochondral defect artificially made on the patellar groove of the same rabbit. The semiquantitative histological scores were significantly higher in the experimental group than in the non-cell-treated control group (p  0.05). “In the experimental group immunohistochemical staining on newly formed cartilage was more intense for type II collagen in the matrix and reverse transcriptase-polymerase chain reaction (RT-PCR) from regenerated cartilage detected mRNA for type II collagen in mature chondrocytes. These findings suggest that repair of cartilage defects can be enhanced by the implantation of cultured MSCs.” In another animal study (Wakitani et al., 1994), autologous culture-expanded MSCs incorporated into type I collagen gels were transplanted into 3  6 mm full-thickness (3 mm in depth) defects in the weightbearing surfaces of the medial femoral condyles of rabbit knees. In the contralateral knee, the defect was filled with collagen gels without cells or the defect was left empty. The defect composed 40–50% of the weight-bearing surface of the condyle, “among the largest ever reported in repair studies in rabbits.” “Two weeks after the transplantation of the mesenchymal cells, the whole area of the original defect was occupied by cartilage. … Twelve weeks after the transplantation, the repair cartilage in the defect became a little thinner than the adjacent normal cartilage, which became a little thinner 24 weeks after the transplantation (the longest observation period in the study).” The authors concluded that large, full-thickness defects of the weight-bearing region of the articular cartilage could be repaired with hyaline-like cartilage after implantation of autologous mesenchymal cells. There were no untoward responses reported. In an animal study Zhou et al. implanted autologous culture-expanded MSCs into osteochondral defects in pigs (Zhou et al., 2004). The amount and make-up of the reparative tissue compared favorably to their prior ACI results using the same animal model. No untoward tissue reactions to the implantation of the MSCs were reported. Zhou et al. employed MSCs that were grown in a chondroinductive environment prior to implantation and their defects extended into sub-chondral bone. The fate (survival) of allogeneic marrow-derived and culture-expanded MSCs implanted in osteochondral defects was determined using transgenic rats (Oshima et al., 2005). An autologous transplantation model was simulated using transgenic rats – whose transgenes produce no foreign proteins – as donors, and wildtype rats as recipients. MSC masses were transplanted into osteochondral defects created in the medial femoral condyle of wild-type rats; the cell aggregates were fixed with fibrin glue. “Twenty-four weeks after transplantation, the defects were repaired with hyaline-like cartilage, which was thicker than normal, and with sub-chondral bone. Using the in situ hybridization technique, cells derived from the transplanted ones were detected within both the cartilaginous and the sub-chondral bone layers. … The findings indicate that the transplanted mesenchymal cells contributed to the repair of the osteochondral defects.”

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In another study (Yanai et al., 2005) bone marrow-derived culture-expanded MSCs were implanted into large full-thickness articular cartilage defects in rabbits that underwent joint distraction. The final cell density was adjusted to 5.0  106 cells/ml in a type I collagen gel. The histological scores were significantly higher in the groups with MSC–collagen gel implants. The authors concluded that the repair of large defects of cartilage can be enhanced by joint distraction, collagen gel, and MSCs. Characterization of Phenotype

Identification and Therapeutic Use of the Adherent Cell Population from Bone Marrow Numerous studies have investigated characteristics of the stromal cell population of marrow that includes the MSC (Barry and Murphy, 2004). Many of these studies have characterized the MSC on the basis of selected surface proteins (Barry et al., 1999, 2001; Reyes et al., 2001; Young et al., 2001; Gronthos et al., 2003). Related studies have attempted to isolate more purified sub-populations of MSCs using cell sorting for selected surface markers, including: positive CD105(+)/negative (CD45(–)GlyA(–) (Reyes et al., 2001); endoglin (Majumdar et al., 2003 and Stro-1) (Gronthos and Simmons, 1995). Related studies have focused on the effects of supplementation of the medium with selected growth factors on the characteristics of isolated sub-populations of MSCs (Gronthos and Simmons, 1995). In one recent study (Lodie et al., 2002), the properties of selected MSC sub-populations were compared: positive or negative selection with antibody to CD105 or CD45/GlyA. The results indicated that “in the initial stages of culture, each cell population proliferated slowly, reaching confluence in 10–14 days. Adherent cells proliferated at similar rates whether cultured in serum-free medium supplemented with basic fibroblast growth factor (Solchaga et al., 2005), medium containing 2% FBS supplemented with epidermal growth factor and plateletderived growth factor, or medium containing 10% FBS alone. Cell surface marker analysis revealed that more than 95% of the cells were positive for CD105/endoglin, a putative MSCs marker, and negative for CD34, CD31, and CD133, markers of hematopoietic, endothelial, and neural stem cells, respectively, regardless of cell isolation and propagation method. CD44 expression was variable, apparently dependent on serum concentration.” Of importance was the fact that this study found that there was similarity in the function of the various cell populations with each “expressing the cell type-specific markers beta-tubulin, type II collagen, and desmin, and demonstrating endothelial tube formation when cultured under conditions favoring neural, cartilage, muscle, and endothelial cell differentiation, respectively. On the basis of these data, adult human bone marrow-derived stem cells cultured in adherent monolayer are virtually indistinguishable, both physically and functionally, regardless of the method of isolation or proliferative expansion.” For the purpose of a cartilage repair therapeutic agent there are no data that indicate that any specific sub-population of MSCs would be safer and more efficacious than the entire adherent cell population. Other Cell Types: Synovial Cells Another tissue in which MSCs have been demonstrated is the synovial tissue (De Bari et al., 2001, 2003). De Bari and colleagues have demonstrated that stem cells isolated from periosteum can be expanded in vitro to over at least 15 passages without loss of their phenotypic traits and that the chondrogenic potential of these progenitor cells was independent of donor age.

CELL–SCAFFOLD IMPLANTS Owing to the now well established phenomenon of dedifferentiation of chondrocytes in monolayer culture, there has been increasing interest in three-dimensional systems of culture and delivery of cells to the

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chondral defect. These systems can provide an environment for growth more similar to native tissue and hence contribute to the phenotypic stability of the chondrocytes. A scaffold also provides an increased surface area for cell attachment. Choosing the right scaffold in cartilage repair requires consideration of a number of factors. Bell described the ideal scaffold for tissue engineering as one that provides a transitional framework whereby the cells populating it create a replacement tissue as the scaffolding material disappears (Bell, 1995). Ideally this scaffold should be degraded at the same rate that the cells produce their own framework. The following requirements are necessary for cartilage tissue engineering. The scaffold should: 1. support cartilage-specific matrix production (collagen type II and aggrecan). Our previous studies showed

that there is a considerable difference in performance among scaffolds, even if only changing the collagen type, pore size, or method of cross-linking. 2. provide enough mechanical support for early mobilization of the treated joint. 3. allow for cell migration of cells to achieve bonding to the adjacent host tissue. In a comparison of several matrix materials (polylactic acid, collagen gel, porous collagen), Grande et al. showed a marked variability of the chondrocyte response (Grande et al., 1997). Bioabsorbable polymers such as polyglycolic acid (PGA) enhanced proteoglycan synthesis, whereas collagen matrices stimulated synthesis of collagen. Not only is there a lack of clinical data on matrix applications for cartilage repair, there are only a few preclinical studies in larger animals. Most of the in vivo work has been done in rabbits and has shown comparatively favorable results (Grande et al., 1989; Kawamura et al., 1998; Ponticiello et al., 2000). However, few studies have systematically compared different methods in a larger animal model. Breinan et al. compared the effects of three different treatments on the healing of articular cartilage defects in a canine model previously developed for ACI (Breinan et al., 2000). In the articular surface of the trochlear grooves of 12 adult mongrel dogs, two 4-mm diameter defects were made to the depth of the tidemark. Four dogs were assigned to each treatment group: (i) micro-fracture treatment, (ii) micro-fracture with a type II collagen scaffold placed in the defect, and (iii) a type II collagen scaffold seeded with cultured autologous chondrocytes. After 15 weeks, the defects were studied histologically. Data quantified on histological cross sections included area or linear percentages of specific tissue types filling the defect, integration of reparative tissue with the calcified and the adjacent cartilage, and integrity of the sub-chondral plate. Total defect filling averaged 56–86%, with the greatest amount found in the dogs in the micro-fracture group implanted with a type II collagen matrix. The profiles of tissue types for the dogs in each treatment group were similar: the tissue filling the defect was predominantly fibrocartilage, with the balance being fibrous tissue. There were no significant differences in the percentages of the various tissue types among the three groups. Taking the results of these dog experiments together and comparing the different repair methods 15 weeks post-operatively, there was a significant correlation between the degree to which the calcified cartilage layer and sub-chondral bone were disrupted and the amount of tissue filling the defect. Moreover, when it formed, hyaline cartilage most frequently occurred superficial to intact calcified cartilage. Ochi et al. investigated the clinical, arthroscopic, and biomechanical outcome of transplanting autologous chondrocytes, cultured in atelocollagen gel, for the treatment of full-thickness defects of cartilage in 28 human knees over a minimum period of 25 months (Ochi et al., 2002). Symptomatically, all patients improved over the follow-up period. There were few side-effects, except for hypertrophy of the graft in three knees and partial detachment of the periosteum in three. Biomechanical test revealed that the transplants had acquired hardness similar to that of the surrounding cartilage.

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In a recent human study (Wakitani et al., 2002), autologous culture-expanded MSCs were implanted in patients in a cartilage repair procedure. The study population was 24 knees of 24 OA patients (average age 63 years; range 49–70 years) undergoing high tibial osteotomy. Ten milliliters of heparinized bone marrow blood was aspirated from both sides of the iliac crest. After approximately 10 days in culture when the attached cells became subconfluent, they were detached and subcultured for and additional 20 days 1.3  107 cells were embedded in type I acid soluble collagen from porcine tendon, put onto a collagen sheet, and gelated. This gelcell composite, which was then cultured overnight, was implanted into 12 knees. The other 12 subjects served as cell-free controls.“In the cell-transplanted group, as early as 6.3 weeks after transplantation the defects were covered with white to pink soft tissue, in which metachromasia was partially observed. Forty-two weeks after transplantation, the defects were covered with white soft tissue, in which metachromasia was observed in almost all areas of the sampled tissue and hyaline cartilage-like tissue was partially observed. Although the clinical improvement was not significantly different, the arthroscopic and histological grading score was better in the cell-transplanted group than in the cell-free control group.” There were no adverse responses reported in the study. This study demonstrated the safety and feasibility of autologous culture-expanded bone marrow-derived MSC transplantation for the repair of articular cartilage defects in humans.

SCAFFOLD-FREE CONSTRUCTS A number of animal studies, using chondrocytes without any scaffold as a method of cell-based therapy, preceded the introduction of ACI. There have been fewer studies where stem cells, without any scaffold, have been used as a cell-based therapy for cartilage repair. One such study involved 16 mature white rabbits from which MSCs were aspirated from the bone marrow (Im et al., 2001). These stem cells were then cultured in monolayer and implanted on to a full-thickness osteochondral defect artificially made on the patellar groove of the same rabbit. Another group of 13 rabbits served as a control group and the animals were sacrificed after 14 weeks. The semiquantitative histological scores were significantly higher in the experimental group than in the control group. In the experimental group immunohistochemical staining of newly formed cartilage was more intense for type II collagen in the matrix and RT-PCR from regenerated cartilage detected mRNA for type II collagen in mature chondrocytes. These findings suggest that repair of cartilage defects can be enhanced by the implantation of cultured MSCs.

CURRENT CLINICAL OUTCOMES By far, the most commonly used cell therapy for cartilage repair is ACI, first reported by Brittberg et al. (1994). This was a case series of 23 patients treated in Sweden for symptomatic cartilage defects. Thirteen patients had femoral condylar defects, ranging in size from 1.6 to 6.5 cm2, due to trauma or osteochondritis dissecans. Seven patients had patellar defects. Ten patients had previously been treated with shaving and debridement of unstable cartilage. Cartilage was harvested arthroscopically from a minimally load-bearing area of the upper aspect or the medial condyle of the affected knee. Chondrocytes were isolated and culture expanded in a cell culture laboratory. In a second procedure, following a medial or lateral parapatellar incision, the defect was debrided and a periosteal flap was harvested and sutured to the rim of the defect. Finally, the chondrocyte suspension was injected under the periosteal flap. Follow-up of the patients was over 16–66 months, with a mean of 39 months. Initially, the transplants eliminated knee locking and reduced pain and swelling in all patients. After 3 months, a repeat arthroscopy showed that the transplants were level with the surrounding tissue and spongy when probed, with visible

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borders. A repeat arthroscopic examination showed that in many instances the transplants had the same macroscopic appearance as they had earlier but were firmer when probed and similar in appearance to the surrounding cartilage. Two years after transplantation, 14 of the 16 patients with femoral condylar transplants had good-to-excellent results. Two patients required a second operation, because of severe central wear in the transplants, with locking and pain. A mean of 36 months after transplantation, the results were excellent or good in two of the seven patients with patellar transplants, fair in three and poor in two; two patients required a second operation because of severe chondromalacia. Biopsies showed that 11 of the 15 femoral transplants and 1 of the 7 patellar transplants had the appearance of “hyaline-like” cartilage. These results and the fact that a commercial service for culturing autologous chondrocytes was established led to a dramatic increase in the use of this cell-based therapy for cartilage repair. Recently, there have been a number of randomized trials comparing ACI with the conventional methods of cartilage repair. Knutsen et al. randomized 80 patients with a single symptomatic cartilage defect on the femoral condyle to either ACI or micro-fracture (Knutsen et al., 2004). Two years post-operatively, arthroscopy with biopsy for histological evaluation was carried out. Both methods had acceptable short-term clinical results. There was no significant difference in macroscopic or histological results between the two treatment groups and no association between the histological findings and the clinical outcome at the 2-year time-point. Bentley et al. reported on a prospective, randomized comparison of ACI versus mosaicplasty for osteochondral defects in the knee (Bentley et al., 2003). One hundred patients with a symptomatic lesion of the articular cartilage in the knee were randomized to undergo either ACI or mosaicplasty. The mean followup period was 19 months and involved a clinical examination. The results demonstrated a significant superiority of ACI over mosaicplasty. The 1 year arthroscopic assessment demonstrated excellent or good repairs in 82% of ACIs and only 34% of mosaicplasties. Browne et al. recently reported on a multicenter cohort study to assess the clinical outcomes of patients treated with ACI for lesions of the distal femur (Browne et al., 2005). A modified Cincinnati knee rating system was used to measure outcomes at baseline and at 5 years. Overall, patients reported a statistically significant improvement in their overall score. Additional analysis of the data showed that 62 patients improved, 6 reported no change, and 19 worsened. In recent years, biological (including tissue engineering) therapies for the treatment of cartilage defects have progressed significantly and are becoming important modalities of treatment in orthopedic surgery. However, for all these therapies long-term outcome is unknown, and there is a lack of controlled studies comparing the different treatment options.

SUMMARY Regenerating cartilage tissue in vivo is likely to remain challenging over the next few years. However, cellbased therapies have already shown a great promise in being better at regenerating the damaged tissue than conventional surgical techniques. These techniques can be further improved upon by investigating the role of scaffolds in trials in repairing cartilage defects. Alternative cell sources, such as stem cells derived from bone marrow, may also provide an improvement in the quality of tissue regenerated.

ACKNOWLEDGMENT This work was supported by the US Department of Veterans Affairs.

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REFERENCES Barry, F.P. (2003). Biology and clinical applications of mesenchymal stem cells. Birth Defects Res. C. Embryo Today 69(3): 250–256. Barry, F.P. and Murphy, J.M. (2004). Mesenchymal stem cells: clinical applications and biological characterization. Int. J. Biochem. Cell Biol. 36(4): 568–584. Barry, F.P., Boynton, R.E., Haynesworth, S., Murphy, J.M. and Zaia, J. (1999). The monoclonal antibody SH-2, raised against human mesenchymal stem cells, recognizes an epitope on endoglin (CD105). Biochem. Biophys. Res. Comm.. 265(1): 134–139. Barry, F., Boynton, R., Murphy, M., Haynesworth, S. and Zaia, J. (2001). The SH-3 and SH-4 antibodies recognize distinct epitopes on CD73 from human mesenchymal stem cells. Biochem. Biophys. Res. Comm.. 289(2): 519–524. Bartholomew, A., Sturgeon, C., Siatskas, M., Ferrer, K., McIntosh, K., Patil, S., Hardy, W., Devine, S., Ucker, D., Deans, R., et al. (2002). Mesenchymal stem cells suppress lymphocyte proliferation in vitro and prolong skin graft survival in vivo. Exp. Hematol. 30(1): 42–48. Bartlett, W., Skinner, J.A., Gooding, C.R., Carrington, R.W., Flanagan, A.M., Briggs, T.W. and Bentley, G. (2005). Autologous chondrocyte implantation versus matrix-induced autologous chondrocyte implantation for osteochondral defects of the knee: a prospective, randomised study. J Bone Joint Surg. Br. 87(5): 640–645. Bell, E. (1995). Strategy for the selection of scaffolds for tissue engineering. Tissue Eng. 1: 163–179. Bentley, G., Biant, L.C., Carrington, R.W., Akmal, M., Goldberg, A., Williams, A.M., Skinner, J.A. and Pringle, J. (2003). A prospective, randomised comparison of autologous chondrocyte implantation versus mosaicplasty for osteochondral defects in the knee. J. Bone Joint Surg. Br. 85(2): 223–230. Bobic, V. (1999). Autologous osteochondral grafts in the management of articular cartilage lesions. Orthopade 28(1): 19–25. Breinan, H.A., Martin, S.D., Hsu, H.P. and Spector, M. (2000). Healing of canine articular cartilage defects treated with microfracture, a type-II collagen matrix, or cultured autologous chondrocytes. J. Orthop. Res. 18(5): 781–789. Brittberg, M. (1999). Autologous chondrocyte transplantation. Clin. Orthop. Relat. Res. 367(Suppl 367): S147–S155. Brittberg, M., Lindahl, A., Nilsson, A., Ohlsson, C., Isaksson, O. and Peterson, L. (1994). Treatment of deep cartilage defects in the knee with autologous chondrocyte transplantation. New Engl. J. Med. 331(14): 889–895. Browne, J.E., Anderson, A.F., Arciero, R., Mandelbaum, B., Moseley Jr., J.B., Micheli, L.J., Fu, F. and Erggelet, C. (2005). Clinical outcome of autologous chondrocyte implantation at 5 years in US subjects. Clin. Orthop. Relat. Res. 436: 237–245. Buckwalter, J.A. and Mankin, H.J. (1998). Articular cartilage: tissue design and chondrocyte–matrix interactions. Instr. Course Lect. 47: 477–486. Caplan, A.I. (1991). Mesenchymal stem cells. J. Orthop. Res. 9(5): 641–650. Connolly, J.F. (1998). Clinical use of marrow osteoprogenitor cells to stimulate osteogenesis. Clin. Orthop. Relat. Res. 355(Suppl 355): S257–S266. Curl, W.W., Krome, J., Gordon, E.S., Rushing, J., Smith, B.P. and Poehling, G.G. (1997). Cartilage injuries: a review of 31,516 knee arthroscopies. Arthroscopy 13(4): 456–460. De Bari, C., Dell’Accio, F. and Luyten, F.P. (2001). Human periosteum-derived cells maintain phenotypic stability and chondrogenic potential throughout expansion regardless of donor age. Arthritis Rheum. 44(1): 85–95. De Bari, C., Dell’Accio, F., Vandenabeele, F., Vermeesch, J.R., Raymackers, J.M. and Luyten, F.P. (2003). Skeletal muscle repair by adult human mesenchymal stem cells from synovial membrane. J. Cell Biol. 160(6): 909–918. Fisher, A. (1922). A contribution to the pathology and etiology of osteo-arthritis: with observations upon the principles underlying its surgical treatment. Br. J. Surg. 10: 52. Goldring, M.B., Sandell, L.J., Stephenson, M.L. and Krane, S.M. (1986). Immune interferon suppresses levels of procollagen mRNA and type II collagen synthesis in cultured human articular and costal chondrocytes. J. Biol. Chem. 261(19): 9049–9055. Goldring, M.B., Birkhead, J., Sandell, L.J., Kimura, T. and Krane, S.M. (1988). Interleukin 1 suppresses expression of cartilage-specific types II and IX collagens and increases types I and III collagens in human chondrocytes. J. Clin. Invest. 82(6): 2026–2037. Grande, D.A., Pitman, M.I., Peterson, L., Menche, D. and Klein, M. (1989). The repair of experimentally produced defects in rabbit articular cartilage by autologous chondrocyte transplantation. J. Orthop. Res. 7(2): 208–218.

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Grande, D.A., Halberstadt, C., Naughton, G., Schwartz, R. and Manji, R. (1997). Evaluation of matrix scaffolds for tissue engineering of articular cartilage grafts. J. Biomed. Mater. Res. 34(2): 211–220. Gronthos, S. and Simmons, P.J. (1995). The growth factor requirements of STRO-1-positive human bone marrow stromal precursors under serum-deprived conditions in vitro. Blood 85(4): 929–940. Gronthos, S., Zannettino, A.C., Hay, S.J., Shi, S., Graves, S.E., Kortesidis, A. and Simmons, P.J. (2003). Molecular and cellular characterisation of highly purified stromal stem cells derived from human bone marrow. J. Cell Sci. 116(Pt 9): 1827–1835. Gurevitch, O., Kurkalli, B.G., Prigozhina, T., Kasir, J., Gaft, A. and Slavin, S. (2003). Reconstruction of cartilage, bone, and hematopoietic microenvironment with demineralized bone matrix and bone marrow cells. Stem Cells 21(5): 588–597. Homminga, G.N., van der Linden, T.J., Terwindt-Rouwenhorst, E.A. and Drukker, J. (1989). Repair of articular defects by perichondrial grafts. Experiments in the rabbit. Acta Orthop. Scand. 60(3): 326–329. Homminga, G.N., Bulstra, S.K., Bouwmeester, P.S. and van der Linden, A.J. (1990). Perichondral grafting for cartilage lesions of the knee. J. Bone Joint Surg. Br. 72(6): 1003–1007. Homminga, G.N., Bulstra, S.K., Kuijer, R. and van der Linden, A.J. (1991). Repair of sheep articular cartilage defects with a rabbit costal perichondrial graft. Acta Orthop. Scand. 62(5): 415–418. Horwitz, E.M., Prockop, D.J., Fitzpatrick, L.A., Koo, W.W., Gordon, P.L., Neel, M., Sussman, M., Orchard, P., Marx, J.C., Pyeritz, R.E., et al. (1999). Transplantability and therapeutic effects of bone marrow-derived mesenchymal cells in children with osteogenesis imperfecta. Nat. Med. 5(3): 309–313. Im, G.I., Kim, D.Y., Shin, J.H., Hyun, C.W. and Cho, W.H. (2001). Repair of cartilage defect in the rabbit with cultured mesenchymal stem cells from bone marrow. J. Bone Joint Surg. Br. 83(2): 289–294. Kavalkovich, K.W., Boynton, R.E., Murphy, J.M. and Barry, F. (2002). Chondrogenic differentiation of human mesenchymal stem cells within an alginate layer culture system. In Vitro Cell Dev. Biol. Anim. 38(8): 457–466. Kawamura, S., Wakitani, S., Kimura, T., Maeda, A., Caplan, A.I., Shino, K. and Ochi, T. (1998). Articular cartilage repair. Rabbit experiments with a collagen gel–biomatrix and chondrocytes cultured in it. Acta Orthop. Scand. 69(1): 56–62. Key, J. (1931). Experimental arthritis: the changes in joints produced by creating defects in the articular cartilage. J. Bone Joint Surg. 23: 725–739. Knutsen, G., Engebretsen, L., Ludvigsen, T.C., Drogset, J.O., Grontvedt, T., Solheim, E., Strand, T., Roberts, S., Isaksen, V. and Johansen, O. (2004). Autologous chondrocyte implantation compared with microfracture in the knee. A randomized trial. J. Bone Joint Surg. Am. 86-A(3): 455–464. Lee, C.R., Grodzinsky, A.J., Hsu, H.P., Martin, S.D. and Spector, M. (2000). Effects of harvest and selected cartilage repair procedures on the physical and biochemical properties of articular cartilage in the canine knee. J. Orthop. Res. 18(5): 790–799. Lee, R.H., Kim, B., Choi, I., Kim, H., Choi, H.S., Suh, K., Bae, Y.C. and Jung, J.S. (2004). Characterization and expression analysis of mesenchymal stem cells from human bone marrow and adipose tissue. Cell Physiol. Biochem. 14(4–6): 311–324. Lodie, T.A., Blickarz, C.E., Devarakonda, T.J., He, C., Dash, A.B., Clarke, J., Gleneck, K., Shihabuddin, L. and Tubo, R. (2002). Systematic analysis of reportedly distinct populations of multipotent bone marrow-derived stem cells reveals a lack of distinction. Tissue Eng. 8(5): 739–751. Majumdar, M.K., Keane-Moore, M., Buyaner, D., Hardy, W.B., Moorman, M.A., McIntosh, K.R. and Mosca, J.D. (2003). Characterization and functionality of cell surface molecules on human mesenchymal stem cells. J. Biomed. Sci. 10(2): 228–241. Mankin, H.J. (1982). The response of articular cartilage to mechanical injury. J. Bone Joint Surg. Am. 64(3): 460–466. Martin, J.A. and Buckwalter, J.A. (2001). Roles of articular cartilage aging and chondrocyte senescence in the pathogenesis of osteoarthritis. Iowa Orthop. J. 21: 1–7. Matsusue, Y., Yamamuro, T. and Hama, H. (1993). Arthroscopic multiple osteochondral transplantation to the chondral defect in the knee associated with anterior cruciate ligament disruption. Arthroscopy 9(3): 318–321. Mayhew, T.A., Williams, G.R., Senica, M.A., Kuniholm, G. and Du Moulin, G.C. (1998). Validation of a quality assurance program for autologous cultured chondrocyte implantation. Tissue Eng. 4(3): 325–334. McBride, C., Gaupp, D. and Phinney, D.G. (2003). Quantifying levels of transplanted murine and human mesenchymal stem cells in vivo by real-time PCR. Cytotherapy 5(1): 7–18.

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Murphy, J.M., Fink, D.J., Hunziker, E.B. and Barry, F.P. (2003). Stem cell therapy in a caprine model of osteoarthritis. Arthritis. Rheum. 48(12): 3464–3474. Niedermann, B., Boe, S., Lauritzen, J. and Rubak, J.M. (1985). Glued periosteal grafts in the knee. Acta Orthop. Scand. 56(6): 457–460. Ochi, M., Uchio, Y., Kawasaki, K., Wakitani, S. and Iwasa, J. (2002). Transplantation of cartilage-like tissue made by tissue engineering in the treatment of cartilage defects of the knee. J. Bone Joint Surg. Br. 84(4): 571–578. Orlic, D., Kajstura, J., Chimenti, S., Jakoniuk, I., Anderson, S.M., Li, B., Pickel, J., McKay, R., Nadal-Ginard, B., Bodine, D.M., et al. (2001). Bone marrow cells regenerate infarcted myocardium. Nature 410(6829): 701–705. Oshima, Y., Watanabe, N., Matsuda, K., Takai, S., Kawata, M. and Kubo, T. (2005). Behavior of transplanted bone marrow-derived GFP mesenchymal cells in osteochondral defect as a simulation of autologous transplantation. J. Histochem. Cytochem. 53(2): 207–216. Pascher, A., Palmer, G.D., Steinert, A., Oligino, T., Gouze, E., Gouze, J.N., Betz, O., Spector, M., Robbins, P.D., Evans, C.H., et al. (2004). Gene delivery to cartilage defects using coagulated bone marrow aspirate. Gene Ther. 11(2): 133–141. Perka, C., Sittinger, M., Schultz, O., Spitzer, R.S., Schlenzka, D. and Burmester, G.R. (2000). Tissue engineered cartilage repair using cryopreserved and noncryopreserved chondrocytes. Clin. Orthop. Relat. Res. 378: 245–254. Peterson, L., Minas, T., Brittberg, M., Nilsson, A., Sjogren-Jansson, E. and Lindahl, A. (2000). Two- to 9-year outcome after autologous chondrocyte transplantation of the knee. Clin. Orthop. Relat. Res. 374: 212–234. Pittenger, M.F., Mackay, A.M., Beck, S.C., Jaiswal, R.K., Douglas, R., Mosca, J.D., Moorman, M.A., Simonetti, D.W., Craig, S. and Marshak, D.R. (1999). Multilineage potential of adult human mesenchymal stem cells. Science 284(5411): 143–147. Ponticiello, M.S., Schinagl, R.M., Kadiyala, S. and Barry, F.P. (2000). Gelatin-based resorbable sponge as a carrier matrix for human mesenchymal stem cells in cartilage regeneration therapy. J. Biomed. Mater. Res. 52(2): 246–255. Prockop, D.J. (1997). Marrow stromal cells as stem cells for nonhematopoietic tissues. Science 276(5309): 71–74. Reyes, M. and Verfaillie, C.M. (2001). Characterization of multipotent adult progenitor cells, a subpopulation of mesenchymal stem cells. Ann. NY Acad. Sci. 938: 231–233; discussion 233–235. Rubak, J.M. (1982). Reconstruction of articular cartilage defects with free periosteal grafts. An experimental study. Acta Orthop. Scand. 53(2): 175–180. Saadeh, P.B., Brent, B., Mehrara, B.J., Steinbrech, D.S., Ting, V., Gittes, G.K. and Longaker, M.T. (1999). Human cartilage engineering: chondrocyte extraction, proliferation, and characterization for construct development. Ann. Plast. Surg. 42(5): 509–513. Salama, R., Burwell, R.D. and Dickson, I.R. (1973). Recombined grafts of bone and marrow. The beneficial effect upon osteogenesis of impregnating xenograft (heterograft) bone with autologous red marrow. J Bone Joint Surg. Br. 55(2): 402–417. Shapiro, F., Koide, S. and Glimcher, M.J. (1993). Cell origin and differentiation in the repair of full-thickness defects of articular cartilage. J. Bone Joint Surg. Am. 75(4): 532–553. Shelbourne, K.D., Jari, S. and Gray, T. (2003). Outcome of untreated traumatic articular cartilage defects of the knee: a natural history study. J. Bone Joint Surg. Am. 85-A (Suppl 2): 8–16. Solchaga, L.A., Gao, J., Dennis, J.E., Awadallah, A., Lundberg, M., Caplan, A.I. and Goldberg, V.M. (2002). Treatment of osteochondral defects with autologous bone marrow in a hyaluronan-based delivery vehicle. Tissue Eng. 8(2): 333–347. Solchaga, L.A., Penick, K., Porter, J.D., Goldberg, V.M., Caplan, A.I. and Welter, J.F. (2005). FGF-2 enhances the mitotic and chondrogenic potentials of human adult bone marrow-derived mesenchymal stem cells. J. Cell Physiol. 203(2): 398–409. Tse, W.T., Pendleton, J.D., Beyer, W.M., Egalka, M.C. and Guinan, E.C. (2003). Suppression of allogeneic T-cell proliferation by human marrow stromal cells: implications in transplantation. Transplantation 75(3): 389–397. Wakitani, S., Goto, T., Pineda, S.J., Young, R.G., Mansour, J.M., Caplan, A.I. and Goldberg, V.M. (1994). Mesenchymal cell-based repair of large, full-thickness defects of articular cartilage. J. Bone Joint Surg. Am. 76(4): 579–592. Wakitani, S., Imoto, K., Yamamoto, T., Saito, M., Murata, N. and Yoneda, M. (2002). Human autologous culture expanded bone marrow mesenchymal cell transplantation for repair of cartilage defects in osteoarthritic knees. Osteoarthritis Cartilage 10(3): 199–206. Yanai, T., Ishii, T., Chang, F. and Ochiai, N. (2005). Repair of large full-thickness articular cartilage defects in the rabbit: the effects of joint distraction and autologous bone-marrow-derived mesenchymal cell transplantation. J. Bone Joint Surg. Br. 87(5): 721–729.

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46 Implantation of Myogenic Cells in Skeletal Muscles Daniel Skuk and Jacques P. Tremblay

INTRODUCTION The intramuscular implantation of myogenic cells is an approach to develop a therapeutic tool for myopathies, mainly those of genetic recessive etiology. The point of departure of this approach can be traced to 1978, when Partridge, Grounds, and Sloper proposed that “in subjects suffering from inherited recessive myopathies, muscle function might be restored if normal myoblasts could be made to fuse with defective muscle fibres” (Partridge et al., 1978). Among these myopathies, Duchenne muscular dystrophy (DMD) is the main target of this potential therapeutic tool. This is due to the combination of DMD’s relative frequency (a prevalence of 50 cases per million in the male population) and severity: progressive generalized skeletal muscle degeneration during the childhood and adolescence, leading to paralysis and death. The history of myogenic-cell transplantation in the field of myology is an example of the importance of an appropriate pre-clinical basis to design clinical applications. It was only after few animal experiments in the 1980s (some of which were probably not appropriately conducted) that several groups undertook clinical trials in the early 1990s, most of which on DMD patients (for a review see Skuk, 2004). Lacking appropriate preclinical support to plan the strategies of cell implantation and control of acute rejection, these clinical trials reported scarce and very modest results at the molecular level. The error was to expect too much from the grafted cells: these trials were conducted on the hope that few myogenic cells injected in few sites of a skeletal muscle would be able to diffuse throughout the muscle and spontaneously fuse with most myofiber regions. The subsequent research demonstrated that this hope was unrealistic. An important lesson from this experience is that researchers need to know the actual behavior of the cells, in appropriate animal models, in order to use them for clinical applications. This chapter wishes to introduce the actual characteristics of myogenic-cell transplantation for clinical purposes. Priority will be given to observations obtained in humans and non-human primates. Observations in rodents will be considered when they complemented or supported the observations in humans and monkeys. This is because there is a motley literature concerning myogenic-cell and “stem-cell” implantation in rodents that, either has no clinical relevance, or did not prove so far to be reproducible in appropriate large animal models. The chapter will be organized through three main challenges of cell transplantation: (1) to obtain appropriate cells for implantation, (2) to properly deliver them to the target tissues, and (3) to insure their survival into the recipient. CELLS TO GRAFT In cell transplantation, the implantable elements could be either differentiated cells or precursor cells with the ability to differentiate into the formers. In the skeletal muscle, the differentiated cells of the parenchyma

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(i.e. the myofibers) are useless for transplantation: they are very long syncytia that cannot be properly manipulated nor implanted. However, this possibility is offered by the mononucleated stem cell that is specific to the skeletal muscle (i.e. the satellite cell). One of the main functions of satellite cells in the mature skeletal muscle is to repair the partial or total damage of the myofibers. They are normally quiescent until an injury produces focal or total necrosis of a nearby myofiber. This necrosis triggers a process of regeneration, which involves the removal of the myofiber debris by phagocytic cells and the activation of the satellite cells. Activated satellite cells enter mitosis and give rise to mononucleated muscle precursor cells generally referred as myoblasts (Betz et al., 1966), the adjective “adult” being sometimes added (Yablonka-Reuveni and Nameroff, 1990) to differentiate them from the embryonic or fetal myoblasts that give rise to skeletal muscles during histogenesis. Adult myoblasts proliferate and fuse among themselves to form syncytial myotubes that would give rise to myofibers. Satellite cells can be isolated from skeletal muscle biopsies by standard cell culture techniques, and can be expanded as myoblasts in vitro, maintaining their capacity to fuse into myotubes that will differentiate into myofibers (Konigsberg, 1960). It was this easiness of satellite cells to be isolated from muscle biopsies, and to be proliferated in vitro in order to obtain large quantities of adult myoblasts able to fuse in myofibers, which opened the possibility to use them for strategies of myogenic-cell transplantation. Useful Properties of the Implanted Cells One of the pioneer studies of myogenic-cell transplantation was published by Lipton and Schultz (1979). They reported the two main properties of exogenous myogenic cells implanted into skeletal muscles of mice, that is, (1)they fuse with the myofibers of the recipient and (2)they can form new small myofibers. The first property allows the phenomenon of “gene complementation,” (i.e. myofibers) in which exogenous myogenic cells fused will express at the same time proteins coded by the exogenous and the host nuclei (Watt et al., 1982). Through this property, the implanted myogenic cells can act as vehicles of therapeutic genes (e.g. by introducing a normal genome in the genetically abnormal myofiber of a muscular dystrophy patient). The second property opens the door to the possibility of regenerating skeletal muscle parenchyma when it is lost. A third property, more recently described, is the possibility of giving rise to new satellite cells. Gene Complementation Myotubes or myofibers containing nuclei with different genomic backgrounds are referred as “mosaic” or “hybrid” (Kikuchi et al., 1980). This is the case of a myofiber in which exogenous myogenic cells have fused, since it will contain a mixture of nuclei of recipient’s and donor’s origin, and will express proteins from both origins (Watt et al., 1982). By the way of gene complementation, a protein whose deficiency in the recipient’s genome causes a myopathy can be expressed by the normal donor’s genome. The first experimental demonstration of this principle was reported by Partridge et al. (1989). Following transplantation of normal mouse myoblasts in mdx mice (which lack the protein called dystrophin, as DMD patients) they observed later that several myofibers expressed dystrophin in its normal subsarcolemmal position. The same observation was repeated by other researchers (Kinoshita et al., 1994; Vilquin et al., 1995) and is presently a routine in myogenic-cell transplantation research. Other proteins, which were restored by normal myoblast transplantation in mouse models of muscular dystrophies, were merosin in dy/dy mice (a model of congenital muscle dystrophy with merosin deficiency) (Vilquin et al., 1996) and dysferlin in SJL mice (a model of limb-girdle muscle dystrophy with dysferlin deficiency) (Leriche-Guerin et al., 2002). In humans, occasional observations of improved dystrophin expression following normal myoblast implantation in DMD patients were reported during the clinical trials conducted in the 1990s (Huard et al., 1992; Karpati et al., 1993; Tremblay et al., 1993; Mendell et al., 1995), although these results were not conclusive and most of the patients gave negative results at that time. A recent clinical trial, designed with the basis of pre-clinical experiments in non-human primates,

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Figure 46.1 Allotransplantation of normal adult myoblasts in DMD patients, as done in our recent clinical trial. Several parallel intramuscular cell injections were done in the tibialis anterior using a 100 l Hamilton syringe and 25–27-gauge needles (a). The cells were delivered homogeneously during the needle withdrawal, and the density of cell injections was controlled by placing on the skin a sterile transparent dressing with a grid. A cross-section of a biopsy done 1 month later in a cell-grafted site in one of these patients is shown (b). Fluorescent immunodetection of dystrophin was done, but the negative of the original image is shown for clarity. Most of the myofiber profiles are dystrophin-negative and can be seen due to the presence of some background, while others show an immunolabeling in their periphery, corresponding to the normal location of dystrophin. The distribution of these dystrophin-positive myofibers, created by the cell graft, follows roughly the original injection trajectories, aligned from the top to the bottom of the image.

showed that donor-derived dystrophin can be observed systematically in the muscles of DMD patients implanted with normal myoblasts (Figure 46.1) (Skuk et al., 2004; Skuk et al., in press). An important factor that conditions the technique of myogenic-cell implantation is that the intracellular proteins coded by a single nucleus into a myofiber do not diffuse throughout the syncytium. On the contrary, they remain localized in a region close to its nucleus of origin, this region being known as a “nuclear domain” (Pavlath et al., 1989). This restriction is produced by the limited diffusion of the mRNA, which was reported to be of only 100 μm from the nucleus (Ralston and Hall, 1992), and of the proteins (Hall and Ralston, 1989). Thus, proteins from donor origin will be expressed only in the segments of myofibers where fusion of donor’s myogenic cells was produced. The more-or-less wide expression of the exogenous protein will depend on its capacity to diffuse or to remain anchored to stationary cellular components (Hall and Ralston, 1989). As an example, single injections of β-galactosidase-labeled normal myoblasts into mdx mice produced dystrophin expression throughout segments of roughly 500 μm in the myofibers in contrast with 1500 μm for β-galactosidase (Kinoshita et al., 1998). The wider expression of β-galactosidase may be attributed to the solubility of this enzyme, thus being able to diffuse more than dystrophin, which remains attached to stationary cellular components. Formation of New Myofibers In DMD and other myopathies, increasing muscle weakness is produced by a progressive and irreversible loss of myofibers. An ideal treatment for patients in advanced stages of these diseases may include not only molecular correction but also restoration of muscle mass. The potential of myoblast transplantation to restore functional skeletal muscle mass in mice was reported following acute severe muscle damage, principally when damage was irreversible. Myoblast implantation significantly restored the muscle mass and force, lost after skeletal muscle destruction (Alameddine et al., 1994; Wernig et al., 1995, 2000; Irintchev et al., 1997). However,

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these models were not strictly similar to the conditions of a progressive muscular dystrophy, and the manner in which myogenic-cell transplantation could form new functional tissue in muscles that degenerated to fibrosis and/or fat substitution (as is the case in DMD) remains insufficiently studied. Some studies in mice suggest that it could be possible to create myotubes within the adipose tissue (Satoh et al., 1992). Neo-muscles can be formed ectopically after subcutaneous myoblast implantation in mice, in spite of the absence of a previous endomysial support (Irintchev et al., 1998). In mdx mice, formation of new dystrophin-positive myofibers through the fusion of the implanted myoblasts among themselves was observed following irradiation of the recipient muscle (Kinoshita et al., 1996b). Progressing from those observations to a clinically functional procedure remains a challenge, among other factors because these results were obtained in mice, which have intrinsically a greater muscle regeneration capacity than primates (Borisov, 1999). However, a recent clinical observation encouraging this research was the presence of neo-formed dystrophin-positive small myofibers in DMD patients transplanted with normal myoblasts (Skuk et al., in press). Formation of Donor-Derived Satellite Cells Mouse studies showed that some of the myoblasts injected into skeletal muscles remain as mononuclear myogenic cells, able to participate later in muscle regeneration (Yao and Kurachi, 1993; Gross and Morgan, 1999) and that they give rise specifically to new satellite cells (Heslop et al., 2001). Some observations suggest that this phenomenon could be produced also in humans. Donor-derived mononuclear cells were detected in the muscles of DMD patients, which received myoblast transplantations (Gussoni et al., 1997; Skuk et al., in press), and some of these nuclei were observed in locations susceptible to correspond to satellite cells (Skuk et al., in press). Similar observations were made following human myoblast transplantation in immunodeficient mice (Brimah et al., 2004). This means that the potential therapeutic effect of myogenic-cell implantation is not limited to the early fusion of the implanted cells, and should also ensure a permanent source of normal satellite cells able to participate later in muscle hypertrophy and regeneration. Undesirable Properties to be Avoided The risk of implanting cells capable of developing a neoplasm should be considered in cultured cell transplantation strategies. Rhabdomyosarcomas were observed following myogenic-cell transplantation in mice, but only when permanent cell lines were used for implantation (Wernig et al., 1991, 1995). Although cell lines are often tumorigenic, the propagation of primary cultures into cell lines is difficult (Freshney, 1987), thus the risk of tumorigenicity following primary cultured cells can be considered very low. Indeed, neoplasia was never observed in muscles of monkeys transplanted with primary-cultured myoblasts (Kinoshita et al., 1995, 1996a; Skuk et al., 1999b, 2000, 2002). However, to take precautions in clinical trials of myoblast transplantation, we tried to exclude a tumorigenic potential in the cells to be implanted by using two tests (Skuk et al., 2004). One of them was an in vitro assay comparing the growth of the donor’s cells in soft-agar medium with that of a rhabdomyosarcoma cell line (Tremblay et al., 1991). In this assay, normal cells survive without proliferation, while rhabdomyosarcoma cells proliferate in clusters. An in vivo test was also used, which consisted in grafting the donor’s cells in muscles of immunodeficient mice (Huard et al., 1994; Skuk et al., 1999a). This confirmed that the donor’s cells were able to fuse in vivo, and eliminated a tumorigenic potential.

CELL IMPLANTATION Once appropriate cells for transplantation are isolated and possibly proliferated, the following challenge is to deliver them appropriately to the target tissue. Intramuscular injection remains so far the only method of myogenic-cell delivery that uniformly and reproductively gives rise to hybrid myofibers in human and non-human primates.

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expression of donorderived proteins

recipient's myoblast macrophage

Figure 46.2 A representation of the mechanism allowing the incorporation of the grafted myogenic cells into the recipient’s myofibers in monkeys is shown. In monkeys, the donor’s cells are labeled by introducing a gene coding for β-galactosidase, and the result of the cell transplantation 1 month later (i) is observed through the expression of β-galactosidase into the recipient myofibers (dark staining within the myofibers). A single cell injection, traversing a skeletal muscle fascicle and delivering the cells homogeneously during the needle withdrawal is represented (a). The process of grafted cell uptake into the myofibers is represented in a couple of myofibers isolated from this fascicle (b–g). These two myofibers (b) are physically damaged by the needle and suffer a segmental necrosis (c). This region of necrosis is invaded by circulating monocytes (d), which became macrophages with two main functions: to phagocyte de debris of the necrosed segment and to release factors that help the regenerative process in the myofiber. This regeneration is done by the activation of the recipient satellite cells, which proliferate as myoblasts, migrating to the center of the region being “cleaned” by the macrophages in order to fuse together (e). The grafted myogenic cells placed in the proximity will be recruited in this process (e). The nuclei of the grafted cells recruited into this fusing process will be integrated in the new myotubes that fill the gap lead by the necrosed segment (f), and will later allow the expression of donor-derived proteins throughout a restricted length of the myofiber ( j). This process leads basically to restricted regions of donor-protein expression in the fascicle (h), which will be expressed as “tracks” of β-galactosidase expression in the non-human primate muscles (i).

Density of Cell Injections The main factor conditioning the strategy of intramuscular cell injections is that the injected cells fuse mainly with the myofibers damaged by the injection trajectories. This is clearly observed in non-human primates, where each single myoblast injection leads a narrow track of hybrid myofibers (Figures 46.2 and 46.3) (Skuk et al.,

Implantation of Myogenic Cells in Skeletal Muscles 787

(a)

(b)

(c)

Figure 46.3 Transplantation of adult β-galactosidase-labeled myoblasts in non-human primates. The cells are delivered by parallel close intramuscular injections using Hamilton syringes and 27-gauge needles (a). In the figure, the syringe is attached to a repetitive dispenser to accelerate the procedure without loss of precision. As in humans, the density of cell injections is controlled by placing on the skin a sterile transparent dressing with a grid. One month later, the fusion of the grafted cells with the recipient’s myofibers is analyzed by histochemical detection of β-galactosidase on histological cross-sections at the cell-grafted sites (b and c). The distribution of the β-galactosidase-positive myofibers reminds the pattern of the original cell-injection trajectories (indicated by the arrows). The density of β-galactosidase-positive myofibers is higher in “c” than in “b,” because the density of cell injections was higher: 25/cm2 in “b” against 100/cm2 in “c.”

2000, 2002). An almost similar pattern, although less clear, was observed following injections of normal myoblasts in DMD patients (Figure 46.1) (Skuk et al., 2004; Skuk et al., in press). Since the expression of donorderived proteins is limited to nuclear domains, cell injections must be very close to each other and must reach the whole muscle to obtain an important and homogeneous expression of donor-derived proteins throughout a skeletal muscle. Experiments in monkeys showed that the volume of muscle expressing a donor’s protein is proportional to the density of cell injections (Skuk et al., 2002) (Figure 46.3). The percentage of myofiber profiles expressing β-galactosidase 1 month after the intramuscular injection of β-galactosidase-labeled myoblasts was of 6–15% when the density of cell injections was of 25/cm2, and 23–67% when it was of 100/cm2 (Skuk et al., 2002). The highest percentages (up to 26%) of dystrophin-positive myofibers observed following normalmyoblast allotransplantation in DMD patients were obtained also when 100 cell injections/cm2 were done (Skuk et al., in press). Such a protocol of cell implantation was denominated “high-density injections” (Skuk, 2004) in order to establish a difference from those used in the former unsuccessful clinical trials, which performed few and distant injections throughout large skeletal muscles. Risks of the Procedure A protocol of high-density injections involves risks that need to be determined in order to be avoided. These risks could be local and systemic and, according to the experience in non-human primates, should be limited to the first days post-implantation. Locally, a monkey’s biceps brachium becomes distended the day post-transplantation but reaches its pretransplantation diameter after 5 days (Skuk et al., 2000). This results in a risk of developing a compartment syndrome in muscles enclosed in a rigid osteofascial space. The biceps brachium of monkeys tolerates well this treatment, but muscles such as the tibialis anterior may probably need to be injected in different sessions. Systemically, an extensive muscle damage (rhabdomyolysis) releases intracellular metabolites such as myoglobin and potassium. This implies risks of acute cardiac arrhythmia in the case of severe hyperkalemia and acute renal failure if myoglobinuria is produced. Both phenomena were not observed following high-density

788 THERAPEUTIC APPLICATIONS: CELL THERAPY

cell injections in the biceps brachii of monkeys (Skuk et al., 2000). This problem, thus, may be controlled by maintaining the muscle damage for a single session of cell transplantation under the limits potentially dangerous. As an example, using high-density injections of myoblasts throughout two biceps brachii of a monkey produced an increase of 2000 U/l in serum creatine kinase levels (Skuk et al., 2000), while the risk of developing an acute renal failure is considered to be produced at creatine kinase levels of 16,000 U/l (Ward, 1988). Trying to Improve the Efficiency of Cell Injections Lower densities of cell injections are desirable but, to reach this objective, the volume of muscle expressing the therapeutic protein (e.g. dystrophin) following a single cell injection must be increased. This could be obtained by: (1) developing methods allowing the implanted cells to fuse with myofibers other than those reached by the injection and/or (2) increasing the nuclear domain of the therapeutic protein. The last possibility was rarely investigated: only one study in mdx mice reported a three-fold increase in the nuclear domain of dystrophin after transplantation of myoblasts overexpressing dystrophin fifty-folds (Kinoshita et al., 1998). Two factors explain why the implanted cells fuse mainly with the myofibers reached by the injection: they lack the capacity to move through the tissue and/or there is absence of myofiber damage out of the injection sites; this damage triggering a regeneration process permitting the fusion of the transplanted cells with the damaged fibers. Some mouse studies aimed to promote the diffusion of the grafted myoblasts throughout the tissue, generally by inducing the secretion of enzymes degrading the extracellular matrix (Ito et al., 1998; Caron et al., 1999; El Fahime et al., 2002). Other experiments have been done to increase the number of regenerating myofibers in order to favor the uptake of the grafted cells. For example, local injection of myotoxic substances, such as phospholipases derived from snake venoms (Kinoshita et al., 1994; Vilquin et al., 1995) and local anesthetics (Cantini et al., 1994; Pin and Merrifield, 1997), were used efficiently in mice. Inhibiting the capacity of the recipient’s satellite cells to proliferate could favor the participation of the grafted myoblasts to the regeneration of the myofibers. This was obtained in mice by submitting the recipient muscle to high doses of ionizing radiation prior to cell transplantation (Morgan et al., 1990; Alameddine et al., 1994; Kinoshita et al., 1994; Vilquin et al., 1995; Wernig et al., 2000). Cryoinjury of the recipient muscle necroses myofibers and satellite cells, and was also used in mice as a pre-treatment to favor the implanted cells (Wernig et al., 1995; Irintchev et al., 1997; Brimah et al., 2004). With the exception of local anesthetics, it seems difficult to expect that the other procedures would be accepted for human use, and even in this case increasing muscle damage would reduce the volume of muscle to be treated in a single session, considering the risks of threatening rhabdomyolysis. So far, only the co-injection of myoblasts and of myotoxic phospholipases improved the success of myoblast transplantation in non-human primates. However, this improvement was only observed when cells and the myotoxin were highly concentrated in a small volume of muscle (Skuk et al., 1999b, 2000).

CELL SURVIVAL IN THE RECIPIENT Once a good delivery of the cells is obtained, their survival in the recipient must be ensured. The post-transplantation survival of myogenic cells should be analyzed at two periods: early and long term. Early Survival The prevailing evidences in the field of myogenic-cell transplantation are that most myogenic cells die quite rapidly (during the first 2–3 days) after their intramuscular implantation, independently of the specific immune response. This phenomenon does not prevent the success of myoblast transplantation, because not all cells die (Beauchamp et al., 1999; Skuk et al., 2003) and the proliferation of the surviving cells compensates totally (Skuk et al., 2003) or partly (Beauchamp et al., 1999) the cell death. The process of death and proliferation of the

Implantation of Myogenic Cells in Skeletal Muscles 789

grafted cells during the first days post-transplantation is not well understood, and the studies approaching the subject show contradictions, probably caused by methodological differences (Skuk et al., 2003). Some observations in mice implicated the acute inflammatory reaction in killing the implanted cells (Guerette et al., 1997), however others challenged this hypothesis (Sammels et al., 2004). It was also postulated that the survival of the whole population of grafted myoblasts could be due to a special small subpopulation of cells that specifically avoid the early cell death and proliferate a great deal (Beauchamp et al., 1999; Cousins et al., 2004), but the existence of this subpopulation was not demonstrated, and the hypothesis does not identify the mechanisms responsible for the early cell death. Long-Term Survival The principal challenge of the long-term survival of myogenic-cell grafts in primates is acute rejection, obvious in inadequately immunosuppressed allotransplantations (Kinoshita et al., 1996a; Skuk et al., 2000). Acute rejection in the context of myogenic-cell transplantation was extensively studied in mice, since the first description of lymphocyte infiltration and disappearance of the grafted myoblasts soon after allogeneic transplantation (Jones, 1979). Subsequent studies identified CD8 and CD4 lymphocytes in these infiltrates (Guerette et al., 1995a; Irintchev et al., 1995; Wernig and Irintchev, 1995) and expression of IL-2 receptors, Th-1 cytokine, and granzyme B (Guerette et al., 1995b, 1996). Non-immunosuppressed or insufficiently immunosuppressed monkeys also exhibited CD4 and CD8 infiltration following myoblast allotransplantation, with lymphocyte invasion of myofibers expressing donor proteins (Kinoshita et al., 1996a; Skuk et al., 1999b, 2002). Ensuring Cell Survival in the Recipient The phenomenon of the early death among grafted cells is presently misunderstood and so far cannot be significantly prevented. However, since this mechanism does not devastate the population of implanted cells (which seems well restored by the proliferation of the surviving cells) the only potential benefit of inhibiting it would be, in theory, a reduction in the number of cells to be injected. Acute rejection, on the other hand, precludes the success of myogenic-cell transplantation in allogeneic conditions. Acute rejection is controlled in humans by pharmacological immunosuppression, but a careful selection of the immunosuppressive drug is required for myogenic-cell transplantation, because some of them kill and/or inhibit the differentiation of the grafted cells (for a review, see Skuk and Tremblay, 2003; Skuk, 2004). The best results of myoblast allotransplantation in mice were reported using tacrolimus (Kinoshita et al., 1994), and for this reason this drug became the immunosuppressant of choice for myoblast allotransplantation in monkeys (Kinoshita et al., 1995, 1996a; Skuk et al., 1999b, 2000, 2002) and humans (Skuk et al., 2004; Skuk et al., in press). Since pharmacological immunosuppression has severe secondary effects, one of the main objectives in clinical transplantation is to develop long-term specific unresponsiveness to grafts with preservation of immune reactions against other foreign antigens (immune tolerance). In the context of myogenic-cell transplantation, immune tolerance developed in some mouse strains after a transient immunosuppression (Pavlath et al., 1994), while in others acute rejection was delayed for months (Pavlath et al., 1994; Wernig et al., 1995). The relevance of these observations is relative, because immune tolerance is more easily obtained in mice than in monkeys or humans. In fact, withdrawal of immunosuppression in monkeys caused acute rejection of the myoblast graft (Skuk et al., 2000). Specific protocols to develop immune tolerance in the context of myoblast transplantation are under investigation (Camirand et al., 2002, 2004). Otherwise, another approach proposed to avoid immunosuppression is the autotransplantation of myoblasts genetically corrected ex vivo (e.g. by introducing the dystrophin gene in myoblasts of DMD patients). Mouse experiments support a future development of this approach (Floyd et al., 1998; Moisset et al., 1998).

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CONCLUSIONS To reach a therapeutic objective, cell transplantation needs basically three conditions: the use of an appropriate cell for implantation, a good method to deliver it to the target tissue, and a method to prevent the specific immune response thus insuring their long-term survival in the recipient in allogenic conditions. Non-human primates were used to define these three conditions in a model appropriate for human extrapolation. Concerning the first item, myoblasts derived from the satellite cells constitute so far the only myogenic cells that are easily isolated and proliferated in vitro, and successfully implanted in the skeletal muscles. Concerning the second item, intramuscular implantation through high-density injections is the only method that proved so far to give rise to high percentages of hybrid myofibers in the skeletal muscles. Finally, an appropriate immunosuppression (tacrolimus-based) is the only method successfully tested in monkeys to ensure the survival of myoblast allografts. These three parameters have permitted to consistently restore the normal expression of dystrophin in many myofibers of patients suffering of an inherited myopathy (DMD). Some important challenges remain, and the potential treatment requires further improvements. A main challenge is to reduce the density of cell injections needed for an efficient distribution of the grafted cells throughout a skeletal muscle. Another challenge overpasses the specific field of myogenic-cell transplantation (it is one of the main problems to solve in the global field of transplantation) and is to reduce as most as possible the toxicity of the methods needed to control acute rejection. The identification of the factors that condition the early survival (death and proliferation) of the grafted cells still challenges the researchers in the field. Finally, the capacity to restore the functional parenchyma in skeletal muscles that degenerated to fibrosis and/or fat substitution (the only possibility to give hope of recovering muscle force in advanced myopathic patients) remains unsolved and barely studied.

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Tremblay, J.P., Roy, B. and Goulet, M. (1991). Human myoblast transplantation: a simple assay for tumorigenicity. Neuromuscul. Disord. 1: 341–343. Tremblay, J.P., Malouin, F., Roy, R., Huard, J., Bouchard, J.P., Satoh, A. and Richards, C.L. (1993). Results of a triple blind clinical study of myoblast transplantations without immunosuppressive treatment in young boys with Duchenne muscular dystrophy. Cell Transplant. 2: 99–112. Vilquin, J.T., Asselin, I., Guerette, B., Kinoshita, I., Roy, R. and Tremblay, J.P. (1995). Successful myoblast allotransplantation in mdx mice using rapamycin. Transplantation 59: 422–426. Vilquin, J.T., Kinoshita, I., Roy, B., Goulet, M., Engvall, E., Tome, F., Fardeau, M. and Tremblay, J.P. (1996). Partial laminin alpha2 chain restoration in alpha2 chain-deficient dy/dy mouse by primary muscle cell culture transplantation. J. Cell Biol. 133: 185–197. Ward, M.M. (1988). Factors predictive of acute renal failure in rhabdomyolysis. Arch. Intern. Med. 148: 1553–1557. Watt, D.J., Lambert, K., Morgan, J.E., Partridge, T.A. and Sloper, J.C. (1982). Incorporation of donor muscle precursor cells into an area of muscle regeneration in the host mouse. J. Neurol. Sci. 57: 319–331. Wernig, A. and Irintchev, A. (1995). “Bystander” damage of host muscle caused by implantation of MHC-compatible myogenic cells. J. Neurol. Sci. 130: 190–196. Wernig, A., Irintchev, A., Hartling, A., Stephan, G., Zimmermann, K. and Starzinski-Powitz, A. (1991). Formation of new muscle fibres and tumours after injection of cultured myogenic cells. J. Neurocytol. 20: 982–997. Wernig, A., Irintchev, A. and Lange, G. (1995). Functional effects of myoblast implantation into histoincompatible mice with or without immunosuppression. J. Physiol. (Lond.) 484: 493–504. Wernig, A., Zweyer, M. and Irintchev, A. (2000). Function of skeletal muscle tissue formed after myoblast transplantation into irradiated mouse muscles. J. Physiol. (Lond.) 522: 333–345. Yablonka-Reuveni, Z. and Nameroff, M. (1990). Temporal differences in desmin expression between myoblasts from embryonic and adult chicken skeletal muscle. Differentiation 45: 21–28. Yao, S.N. and Kurachi, K. (1993). Implanted myoblasts not only fuse with myofibers but also survive as muscle precursor cells. J. Cell Sci. 105: 957–963.

47 Islet Cell Transplantation Juliet A. Emamaullee and A.M. James Shapiro

INTRODUCTION Background Diabetes is a disease that results from impaired glucose metabolism. Approximately 90% of diabetes is caused by a defect in insulin production and/or utilization (Type 2 diabetes mellitus; “T2DM”), while the more severe form, Type 1 diabetes mellitus (“T1DM”), is caused by a complete loss of the insulin-producing β-cells within the islets of Langerhans of the pancreas. Diabetes currently affects more than 200 million patients worldwide and is projected to afflict at least 5% of the global adult population by the year 2025 (King et al., 1998). As the incidence of diabetes increases, the cost of treating these patients has skyrocketed, consuming between 7% and 13% of health-care expenditure in developed countries (WHO, 2002). Since the discovery of insulin in 1921, diabetes has become a treatable condition, and the life expectancy of patients with diabetes has been greatly improved. However, even with diligent blood glucose monitoring and insulin administration, the metabolic abnormalities associated with diabetes can lead to many chronic secondary complications, including nephropathy, retinopathy, peripheral neuropathy, coronary ischemia, stroke, amputation, erectile dysfunction, and gastroparesis (National Diabetes Data Group (US) et al., 1995). In the US, patients with diabetes represent 8% of those who are legally blind, 30% of all patients on dialysis due to end-stage renal disease, and 20% of all patients receiving kidney transplants (National Diabetes Data Group (US) et al., 1995). The Diabetes Control and Complications Trial (DCCT) was conducted to determine if intensive blood glucose regulation by frequent insulin injection or pump could prevent these long-term complications in patients with diabetes (DCCT Research Group, 1990, 1993; Keen, 1994). Results from the DCCT and subsequent Epidemiology of Diabetes Interventions and Complications (EDIC) study have clearly demonstrated that this approach improved but did not normalize glycosylated hemoglobin levels (HbA1C) and significantly protected against cardiovascular disease, nephropathy, neuropathy, and retinopathy (DCCT Research Group, 1990; Keen, 1994; Nathan et al., 2003, 2005). However, the consequence of improved glycemic control was a threefold increased risk of serious hypoglycemic reactions leading to recurrent seizures and coma (Keen, 1994; DCCT Research Group, 1995). Recent improvements in the size and sensitivity of insulin pumps have increased their utility, but the creation of implantable devices has been more challenging. Also, while insulin pump therapy can improve HbA1C levels compared to multiple daily injections of insulin, pumps may malfunction and thus still necessitate frequent blood glucose monitoring by the user (Owen, 2006). While advances in the formulation, half-life, and administration of insulin have markedly improved the quality of life- and long-term survival of patients with diabetes, it has long been recognized that the restoration of an adequate islet mass would provide the maximum benefit to diabetic patients, leading to a true physiological correction of the diabetic state. In the early 1960s, great advances were made in the field of renal transplantation due to improved immunosuppressive therapies (azathioprine and corticosteroids), which prompted the first attempts in whole pancreas transplantation (Merrill et al., 1963; Murray et al., 1963). First

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introduced by Kelly and Lillehei in 1966, early attempts were associated with high mortality rates and poor graft survival, with 3% graft function at 1 year post-transplant (Kelly et al., 1967). The risk profile and longterm outcomes in whole pancreas transplantation have been greatly improved by recent improvements in surgical technique, including portal venous and enteric endocrine drainage, and steroid-free maintenance immunosuppression (Newell et al., 1996; Kendall et al., 1997). To date more than 25,000 pancreas transplants have been performed worldwide for end-stage renal disease (simultaneous kidney pancreas or pancreas after kidney transplantation) or less frequently for severe hypoglycemic unawareness (pancreas transplant alone). Data collected in the International Pancreas Transplant Registry (IPTR) have shown that only 50% of patients who have undergone pancreas-alone transplantation remain insulin independent at 5 years, despite recent improvements in surgical technique and immunosuppression (Larsen, 2004; Gruessner and Sutherland, 2005). Also, 30% of the approximately 6,000 cadaveric pancreata donated each year are transplanted due to strict donor criteria and requirements for short cold ischemic time (Larsen, 2004; 2005b). The surgical risks and requirement for lifelong immunosuppression have reserved pancreas-alone transplantation only for those diabetic patients with the most severe and life-threatening disease, despite strong evidence that the procedure can prolong life, reverse established nephropathy, and improve quality of life. Since the major surgical complications in whole pancreas transplantation are related to the exocrine function of the pancreas, which is not necessary to restore euglycemia in diabetic patients, it has long been recognized that β-cell replacement could be achieved with implantation of isolated pancreatic islets. Since this approach involves transplantation of a cellular graft that would be implanted using minimally invasive techniques, it would avoid the risks associated with major surgery, resulting in a more widely available treatment for patients with diabetes. History of Islet Transplantation The concept of islet transplantation actually preceded the discovery of insulin in 1921 by nearly 30 years (Figure 47.1). In 1893, physicians in Bristol attempted to treat a young boy suffering from diabetic ketoacidosis by transplanting fragments of a freshly slaughtered sheep’s pancreas (Williams, 1894). While the graft ultimately failed in the absence of immunosuppression, the patient’s health did temporarily improve, which suggested that cells within the pancreas could restore euglycemia. After the discovery of insulin, it was thought that exogenous insulin replacement would be an effective treatment for patients with T1DM, and therefore islet transplantation was not actively pursued. However, as insulin therapy transformed T1DM from an acute health crisis to a chronic disease, it became apparent that insulin injections could not prevent the onset of debilitating and life-threatening secondary complications. As the first series of whole pancreas transplants in the late 1960s were associated with poor morbidity and mortality, isolated islet transplantation gained a renewed interest (Sutherland et al., 2001). The first successful islet isolations and subsequent transplantation into chemically induced diabetic rodents were pioneered by Dr. Paul Lacy at Washington University in St. Louis, which immediately sparked interest in the implementation of clinical trials (Lacy and Kostianovsky, 1967; Ballinger and Lacy, 1972; Kemp et al., 1973; Reckard et al., 1973). While euglycemia was routinely obtained in animal models of islet transplantation, clinical islet transplantation struggled to find success for most of the 1970s and 1980s. During this time, unpurified islets were infused into the portal vein, leading to many serious complications including portal vein thrombosis, portal hypertension, and disseminated intravascular coagulation (Walsh et al., 1982). While working in Lacy’s group, Dr. Camillo Ricordi developed the “automated method” for high-yield islet isolation in 1989 (Ricordi et al., 1989). This represented a major turning point in the field and led to the report that Lacy’s group had achieved short-lived insulin independence in a patient with T1DM who had received an islet graft following a previous kidney transplant (Scharp et al., 1990). The following year, the group lead by Ricordi at the University of Pittsburgh reported the first series of clinical islet allografts that demonstrated improved insulin-independence rates of 50% at 1 year, in

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Pittsburgh, USA, the first successful series of clinical islet allografts in patients with surgical (non-autoimmune) diabetes showing 50% one year insulin independence.

Bristol, UK, Williams and Harsant attempted first islet xenotransplant with sheep pancreas fragments.

Minneapolis, USA, Two cases of living donor islet allotransplantation attempted unsuccessfully.

1893

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Washington University, Paul E. Lacy was the first to reverse chemically induced diabetes using islet transplantation in a rodent model.

1989

Houston, USA and GRAGIL Consortium, the first successful shipment of islets between centers.

Kyoto, Japan, first successful living donor islet transplant performed.

Edmonton, Canada, 100% insulin independence in the first 7 consecutive patients treated with the Edmonton Protocol.

1990

1996

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Giessen and Geneva (GRAGIL Consortium) both reported a 50% rate of Cpeptide secretion and 20% insulin independence rate at one year with improved peritransplant management and immunosuppresion.

St. Lousis, USA, first short-lived insulin independence achieved in human islet-alone transplantation.

2001

600 patients treated with islet transplants since 2000.

2002

2004

2005

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NIH Immune Tolerance Network completes the first multicenter trial in islet transplantation.

Miami, USA, successfully replicate Edmonton Protocol using islets kept in culture before transplantation for up to three days, eliminating the limitation of immediate islet transplantation.

Figure 47.1 Timeline of notable advances in the history of islet transplantation. subjects who underwent cluster islet–liver transplants for abdominal malignancies in the setting of surgicalinduced (non-autoimmune) diabetes (Ricordi et al., 1989; Tzakis et al., 1990). Although this represented a major advance in the field of islet transplantation, these results could not be reproduced in patients with T1DM, the key patient population in need of β-cell replacement (Ricordi et al., 1992). In the late 1990s, the European GRAGIL consortium reported the first modestly successful insulin-independence rates of 20% at 1 year in patients with T1DM, which could be attributed to improved peritransplant management and immunosuppressive drug regimens (Benhamou et al., 2001). Since the results from Pittsburgh and the GRAGIL consortium were obtained in patients who had previously received a kidney transplant, there was no additional risk in terms of immunosuppression to the patients after receiving an islet graft (Ricordi et al., 1992; Benhamou et al., 2001). An international registry held in Giessen, Germany, has maintained a comprehensive record of previous clinical attempts at islet transplantation globally, and of the total world experience of over 450 attempts at clinical islet transplantation prior to 2000, 8% of subjects achieved insulin independence (Brendel, 2001). After three decades of research, the 1 year insulin-independence rates in clinical islet transplantation were still too low to justify the risks associated with portal infusion and lifelong immunosuppression in the majority of patients with T1DM (Secchi et al., 1991; Gross et al., 1998; Hering, 1999; Benhamou et al., 2001; Brendel, 2001). The Edmonton Protocol Shapiro and colleagues at the University of Alberta developed a new protocol in 1999 that was designed for patients with “brittle diabetes” who experienced extreme difficulty in managing their blood glucose levels (“glucose lability”) and/or severe hypoglycemic unawareness (Shapiro et al., 2000). The so-called “Edmonton Protocol” was unique compared to previous attempts in clinical islet transplantation in its high-targeted islet

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mass, with a mean of approximately 13,000 islet equivalents (IE)/kg recipient body weight, often derived from two (or occasionally more) fresh islet preparations, and in its immunosuppression strategy, with emphasized avoidance of corticosteroids and use of potent immunosuppression with combined sirolimus, tacrolimus, and anti-CD25 antibody to protect against rejection and recurrent autoimmunity (Shapiro et al., 2000). This approach lead to dramatic improvements in islet allograft survival, with all of the first seven patients achieved sustained independence from insulin (Shapiro et al., 2000). More than 85 consecutive patients have received islet transplants at the University of Alberta since 1999, and the 1 year insulin-independence rate remains steady at approximately 80% after completed transplants (13,000 IE/kg). The results obtained at the University of Alberta have been replicated at other centers as part of an international multicenter trial through the Immune Tolerance Network, but each center’s success has varied greatly depending on its previous experience and skill in islet isolation and immunosuppressive management (Shapiro et al., 2003). The Miami group has demonstrated that islets can be cultured for up to 3 days pre-transplant or shipped and transplanted at a remote facility (Houston) with similar success as freshly isolated islets when transplanted using Edmonton-like immunosuppression (Goss et al., 2002, 2004). The GRAGIL Network (a Swiss-French consortium) has also demonstrated the benefits of centralized islet processing facilities which can service a broader network of centers throughout Europe (Benhamou et al., 2001; Kempf et al., 2005). Based upon the success of the Edmonton group, islet transplantation has been funded in Alberta, Canada, as accepted clinical standard of care since 2001. Progress in this area has been slower in the United States, but large registration trials are currently moving forward to secure a Biological License and therefore reimbursement, which will make a significant difference to the availability of islets for transplantation in that country. The recent success of clinical islet transplantation has encouraged many centers around the world to implement a program, and since 2000 more than 550 patients have been transplanted using recent variants of the Edmonton Protocol in almost 50 centers worldwide (International Islet Transplant Registry, 2005a). Despite this success, the current requirement for lifelong immunosuppression in islet-alone transplantation has restricted its availability to patients with T1DM and severe hypoglycemia or glycemic lability. The benefit of islet transplantation in patients with T2DM has not been determined, since many of these patients are overweight and/or insulin resistant and thus would require a large islet mass to meet their metabolic demands. Most patients require two or occasionally three islet implant procedures in order to achieve insulin independence, although insulin independence following single donor infusion has been reported in a cohort of patients at the University of Minnesota (Hering et al., 2004, 2005). While C-peptide secretion (0.5 ng/ml) has been maintained in 88% of islet graft recipients beyond 3 years in Edmonton, emerging data on the longterm insulin-independence rates have shown that only 50% of recipients remain off insulin at 3 years, with 10% off insulin at 5 years post-transplant (Ryan et al., 2005). Although the exact cause of the discrepancy between insulin independence and maintenance of C-peptide status is not fully understood, it is likely that there are multiple events which hinder graft function and survival over time. While rejection (acute or chronic) and recurrent autoimmunity may be responsible for graft loss, it is probable that other, nonimmune-mediated damage occurs, such as chronic toxicity from sirolimus/tacrolimus and failure of islet regeneration or transdifferentiation due to the anti-proliferative effects of sirolimus. Perhaps the most important component of decaying graft function over time is the concept of islet “burn-out” from constant metabolic stimulation, since only a marginal mass of islets actually engraft in most subjects. In clinical islet transplantation thus far, the risks of malignancy, post-transplant lymphoma and life-threatening sepsis have been minimal, but fears of these complications limit a broader application in patients with less severe forms of diabetes including children. Moreover, a number of immunosuppression-related side effects have been encountered, including dyslipidemia, mouth ulceration, peripheral edema, fatigue, ovarian cysts, and menstrual irregularities in female subjects, which can be dose or drug limiting in some patients (Ryan et al., 2002).

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Thus, while dramatic improvements in outcomes following islet transplantation have been observed, extensive refinements in clinical protocols are needed both to improve safety and to enhance success with single donor islet infusions.

CLINICAL ISLET TRANSPLANTATION Patient Assessment and Selection Clinical islet transplantation is associated with a number of risks, including procedural complications such as bleeding or portal vein thrombosis, or those associated with lifelong immunosuppression (i.e. infection or malignancy). For these reasons, patients selected as islet recipients must have severe, life-threatening diabetic complications that justify the risks of transplantation. Two T1DM patient populations have been identified as suitable candidates for islet transplantation: those individuals that experience frequent, severe and recurrent hypoglycemic unawareness, or those patients with highly unstable blood glucose control despite an optimized insulin regimen (glycemic lability). When patients are evaluated for islet transplantation, their metabolic status and diabetes-related secondary complications should be carefully characterized so that those patients who would receive the greatest benefit despite the requirement for lifelong immunosuppression are selected. First and foremost, islet transplantation is reserved for patients with C-peptide negative (0.3 ng/ml) T1DM. Recipients with elevated body mass index (BMI) (30 kg/m2) or those 90 kg are generally excluded, as their metabolic demand may not be met by the transplanted islet mass. As mentioned previously, the current indications for islet-alone transplantation include severe hypoglycemic unawareness and/or glycemic lability. To assess these symptoms, Ryan et al. developed an objective scoring system to measure the severity of both hypoglycemia (the HYPO score), and the lability index (LI), which is based upon the changes in blood glucose over time (Ryan et al., 2004b). Current selection criteria for islet-alone transplantation include a HYPO score 1047 (90th percentile), LI 433 mmol/L2/h/week (90th percentile), or a composite with the HYPO score 423 (75th percentile), and LI 329 (75th percentile) (Ryan et al., 2005). Since patients with poor diabetes compliance or an inadequate baseline insulin regimen are likely to benefit from improved design of their insulin dosing regimens, patients selected for transplant should have a plasma HbA1C 10%. In an effort to reduce the risk of serious procedural and immunosuppressive drugrelated complications, the patient’s cardiac and renal function should be carefully assessed. Selected recipients should have adequate cardiac function including blood pressure 160/100 mmHg, no evidence of myocardial infarction in the 6 months prior to assessment, no angiographic evidence of non-correctable coronary artery disease, and left ventricular ejection fraction (LVEF) 30% as measured by echocardiogram. To eliminate patients who are better candidates for simultaneous kidney–pancreas transplantation or those who may experience adverse renal function as a result of tacrolimus or sirolimus therapy, selected recipients should have no evidence of macroscopic proteinuria (300 mg/24 h) and a calculated glomerular filtration rate (GFR) 80 (70 in females) ml/min/1.73 m2. Proliferative retinopathy should be stabilized prior to transplantation, as acute correction of glycemic control may lead to accelerated retinopathy. Finally, to reduce the risk of antibody-mediated graft rejection, potential recipients should be screened for panel reactive antibody assays (PRA) and determined to be 20%. Islet Transplantation Procedure Although several locations have been tested as potential implantation sites for islet grafts, the high level of graft function and ease of delivery associated with infusion into the portal circulation of the liver have led to this being the transplantation site of choice in clinical protocols (Kemp et al., 1973). There are two accepted approaches for

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(a) Islet transplantation – 2006 Islet isolation

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(b) Islet transplantation – future Islet isolation

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Figure 47.2 The islet transplant procedure-present and future. Islet transplantation, in its current form (a), has provided insulin independence in most diabetic patients at one year post-transplant, but this procedure is currently limited by the availability of suitable cadaveric donors and the requirement for lifelong immunosuppression. In the future (b), islet transplantation could be made available to a broader range of diabetic patients through the usage of alternative tissue sources, such as living donors, xenogeneic donors, or stemcell derived β-cells. Also, as novel immunomodulatory therapies are identified, tolerance induction strategies can be developed that will prolong graft function and allow for the reduction or complete withdrawal of immunosuppressive drug therapy.

implanting purified islets into the liver by way of the portal vein. While surgical laparotomy and cannulation of the portal vein was most often used in the early islet transplant programs, current protocols routinely employ the percutaneous transhepatic approach to implant donor islets in cadaveric islet transplantation (Figure 47.2a) (Ryan et al., 2005). Compared to surgical laparotomy, this procedure is minimally invasive and thus can be performed using local anesthesia, combined with opiate analgesia and hypnotics given as pre-medication. Access to the portal vein is achieved by percutaneous transhepatic approach using a combination of ultrasound and fluoroscopy to guide the radiologist. A branch of the right portal vein is cannulated, and a catheter is positioned proximal to the confluence of the portal vein, which is confirmed with a portal venogram (Owen et al., 2003). The risk of portal vein thrombosis is reduced by inclusion of unfractionated heparin (70 units/kg) in the islet preparation. Islets are then infused, aseptically, into the main portal vein under gravity, with regular monitoring of portal venous pressure (by an indirect pressure transducer) before, during, and after the infusion. An ultrasound examination should be performed at 1 day and 1 week post-transplant to rule out intraperitoneal hemorrhage and to confirm that the portal vein is patent and has normal flow. If a patient must be anti-coagulated prior to transplantation or if a hemangioma is present on the right side of the liver that may be at risk for puncture and bleeding if the percutaneous approach were to be used,

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surgical laparotomy and cannulation of a mesenteric venous tributary of the portal system should be considered. In this situation, complete surgical control is in place to prevent uncontrolled bleeding. Another advantage includes the potential for use of a dual lumen catheter for cannulation of a mesenteric vein (i.e. dual lumen 9Fr Broviac line), which allows for continuous monitoring of portal pressure during islet infusion. Still, this surgical approach should only be considered when the percutaneous transhepatic approach cannot be utilized, as it does present several major disadvantages, including the requirement for a surgical incision, formation of adhesions, and the risk of wound infection and wound herniation, which may be exacerbated when the drug sirolimus is used post-transplant, as this drug interferes with wound healing.

RISKS TO THE RECIPIENT Surgical Complications There are two potentially serious procedural complications in islet transplantation: bleeding from the catheter tract created by the percutaneous transhepatic approach, and portal vein thrombosis, particularly when large volumes of tissue are infused. Adverse bleeding events were noted early in the development of the Edmonton program, but these have been completely avoided in the past 40 consecutive procedures with the routine use of effective methods to seal and ablated the transhepatic portal catheter tract on egress when the catheter is withdrawn. The combination of coils and tissue fibrin glue (Tisseel®) was used previously, but more recently has been replaced by Avitene® paste (1 g Avitene powder mixed with 3 ml of radiological contrast media and 3 ml of saline – approximately 0.5–1.0 ml of this paste is injected into the liver tract) (Villiger et al., 2005). The use of purified islet allograft preparations has not resulted in main portal vein thrombosis in the Edmonton program, but thrombosis of a right or left branch, or peripheral segmental vein has been encountered in approximately 5% of patients. Other rarely observed procedural side effects have included fine needle gallbladder puncture, arteriovenous fistulae (which may require selective embolisation) or steatosis in the hepatic parenchyma, which generally does not present any clinical complications or require intervention (Bhargava et al., 2004). Immunosuppressive Therapy and Complications Islet transplantation for T1DM represents a unique challenge in immunosuppression, as both alloimmunity and islet-specific autoimmunity must be effectively controlled to preserve graft function. An additional important consideration is that many of the immunosuppressive agents used in solid organ transplantation since the 1960s, particularly corticosteroids, are known to be toxic to islets. In the current version of the Edmonton Protocol, the induction agent daclizumab (anti-CD25 (IL-2R) antibody) is administered intravenously immediately prior to transplantation and again at 2 weeks post-transplant (1 mg/kg). Maintenance immunosuppression is achieved using sirolimus with a low dose of tacrolimus, as sirolimus appears to be associated with less nephrotoxicity and diabetogenicity than calcineurin inhibitors (i.e. cyclosporine and tacrolimus). A loading dose of sirolimus (0.2 mg/kg) is given prior to transplant, followed by 0.15 mg/kg, which is then adjusted subsequently to achieve trough levels between 10–12 ng/ml for the first 3 months and 7–10 ng/ml thereafter. Tacrolimus is adjusted to maintain trough levels between 3 and 6 ng/ml. This regimen, described initially at the University of Alberta, has been successfully replicated at other centers as part of a multicenter ITN trial (Shapiro et al., 2003, 2005b). In addition to the Edmonton Protocol immunosuppression described above, alternative regimens have been reported. The Minnesota Group, led by Dr. Bernhard Hering, has utilized anti-thymocyte globulin and etanercept (anti-tumor necrosis factor-α (TNFα) antibody) induction with a combination of sirolimus and mycophenolate mofetil low-dose tacrolimus for maintenance, or hOKT3γ1(Ala–Ala) (humanized antiCD3 antibody) and sirolimus induction with sirolimus and reduced-dose tacrolimus for maintenance (Hering et al., 2004, 2005). In some instances, alternative immunosuppressive agents have been used because of drug intolerance or other side effects. Islet patients often possess mild preexisting renal impairment as a

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result of longstanding diabetes, and this renal dysfunction may be exacerbated with calcineurin inhibitor therapy, even at the low doses involved in the Edmonton Protocol. The drug sirolimus may also have nephrotoxic side effects, which may be compounded when used in combination with a calcineurin inhibitor drug (Kaplan et al., 2004; Senior et al., 2005). For these reasons, renal status must be monitored diligently in all patients following islet transplantation. In addition to its recognized nephrotoxicity, tacrolimus is associated with gastrointestinal side effects which may lead to episodic diarrhea. Neurotoxicity may be seen with tacrolimus but is often avoided in low-dose regimens (Gruessner et al., 1996). Sirolimus is associated with neutropenia and mouth ulceration, but these side effects can be reduced with lower target trough levels and tablet formulations. In the context of islet transplantation, sirolimus has been linked to a number of side effects including dyslipidemia, small bowel ulceration, peripheral edema, and the development of ovarian cysts or menstrual cycle irregularities in female recipients (Molinari et al., 2005; Ryan et al., 2005). While chronically immunosuppressed patients are at risk for developing all types of malignancy, squamous epithelial cancers most commonly occur and are most readily treatable. The lifetime risk of lymphoma is estimated to be 1–2% in transplant recipients, but this risk is likely to be reduced in islet recipients, as these patients are generally not treated with glucocorticoids or OKT3.

FUTURE CHALLENGES Overcoming Tissue Shortage In its current form, islet transplantation is reserved for patients with the most severe forms of diabetes, which in reality constitute a small fraction of all patients with T1DM. Even with the relatively small patient population selected for islet transplantation, the waitlist time for patients in Edmonton, which has access to organs from a large geographic region, ranges from 6 months to 2 years depending on blood group. As islet transplantation becomes more suitable for a broader range of diabetic patients and as the incidence of diabetes increases, there will be an even more severe shortage of islet tissue for transplantation. Presently, clinical islet programs rely on the scarce supply of pancreas organs derived exclusively from heart-beating, brain-dead cadavers. Compared to organs procured for whole pancreas transplantation, which must fall within very strict donor criteria, organs obtained for islet transplantation tend to be more “marginal” and come from older, less stable donors. Furthermore, the pancreas is particularly susceptible to toxicity from the circulating products of severe brain injury, hemodynamic instability, and inotropic support in a brain-dead organ donor. The quality of the pancreas is further degraded by cold ischemic injury during transportation, which inevitably results in islet damage and loss. Contreras et al. demonstrated a marked reduction in islet recovery and in islet viability in experimental islet transplantation using tissue derived following brain death compared to healthy rodent donors, highlighting this issue, and recently his group has confirmed these findings using human islets (Contreras et al., 2003). Similarly, Lakey et al. demonstrated a strong relationship between islet recovery and donor stability (Lakey et al., 1996). Once the pancreas is in the isolation laboratory, the extensive processing and purification steps during processing result in further islet destruction and loss, often resulting in at best 60% recovery of the estimated 107 IE/pancreas (Tsujimura et al., 2004). As a result, nearly all islet recipients require islets derived from two cadaveric donors. Thus, a rapidly growing area of islet transplant research involves the development of improved cadaveric or alternative islet tissue sources for transplantation. Living Donor Islet Transplantation One approach to alleviating islet tissue demand would be to make use of living donors for islet transplantation. Living donor programs in kidney, liver, and lung transplantation have moved forward successfully at most

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leading transplant centers worldwide, in an attempt to meet the growing demand for donor organs and to improve clinical outcomes. Given the rapid, global acceptance of cadaveric islet transplantation over the past 5 years, it is likely that living donor islet transplantation will soon be offered to patients listed in cadaveric islet transplant programs. Despite remarkable progress in clinical islet transplantation since 1999, islet supply and functional viability remain to be significant challenges when islets are derived from cadaveric organ donors, even at the most experienced centers (Contreras et al., 2003). In the living donor setting, the distal half pancreas could be procured under “ideal” circumstances, without exposure of the pancreas to hemodynamic instability or inotropic drugs, and the pancreas would be processed immediately without prolonged cold ischemia. Thus, the potency of islets derived from a living donor source is assumed to be far superior to cadaveric tissue. Living donor islet transplantation represents a unique opportunity to overcome donor organ shortage and procure the islet tissue under perfect conditions, with closer human leukocyte antigen (HLA) matching between donor and recipient. Furthermore, the living donor islet transplant setting will provide a unique opportunity to develop protocols for pre-transplant recipient conditioning for donor-specific tolerance induction. While cadaveric islet transplantation has been an active area of clinical research involving more than 1,000 patients in the past 30 years, only three cases of living-donor islet allo-transplantation have been reported (Sutherland et al., 1980; Matsumoto et al., 2005). The first two clinical attempts at living donor islet allo-transplantation were carried out in 1978 by Sutherland and colleagues at the University of Minnesota (Sutherland et al., 1980). While neither recipient achieved sustained islet function, these pioneering efforts were truly remarkable given the early stage of clinical islet transplant development at the time. The immunosuppression available was primitive by current standards (azathioprine and high-dose steroids), and the islets were isolated using suboptimal conditions, prior to the development of the Ricordi chamber and the sophisticated purification schemes currently used in clinical islet transplantation. The dramatic improvement in clinical outcomes obtained in cadaveric islet transplantation since 2000 has renewed interest in the development of living donor islet transplantation. The first living donor islet transplantation case attempted since the introduction of the Edmonton Protocol was carried out at the University of Kyoto in early 2005, as a collaboration between the Japanese and Edmonton programs (Matsumoto et al., 2005). The recipient, a 27-year-old female, developed C-peptide negative, unstable diabetes following chronic pancreatitis as a child. Her 56year-old mother was approved to be the donor, and islets were purified from the distal pancreas (47% as measured pre-operatively by computed tomography (CT) volumetry) obtained during an open laparotomy. There were no surgical complications in either donor or recipient. The unpurified islet mass (408,114 IE (8,200 IE/kg) in a volume of 9.5 ml after tissue digestion) was transplanted into the portal vein using the percutaneous approach under full systemic heparinisation. Edmonton Protocol-style immunosuppression was started pre-transplant using sirolimus and low-dose tacrolimus (started 7 days pre-transplant), anti-IL2R antibody (given 4 days pre-transplant and on the day of transplant) and anti-TNFα blockade induction (infliximab; given 1 day pre-transplant). Insulin therapy in the recipient was discontinued at 22 days posttransplant, and this patient continues to be insulin independent with excellent glycemic control and a normal HbA1C more than 1 year post-transplant. The donor has presented no evidence of glucose intolerance and has maintained normal HbA1C values since the procedure. While no definitive conclusions can be drawn from this single successful case of living donor islet allotransplantation, results from living donor islet auto-transplantation suggest that the insulin independence may be achieved routinely with significantly less IE/kg recipient body weight than has been required for cadaveric allografts thus far. It is widely accepted that over 70% of patients will remain insulin free following islet auto-transplantation if an islet mass exceeding 300,000 IE ( 2,500 IE/kg) is transplanted, compared to the 13,000 IE/kg that is often required to achieve insulin independence with cadaveric islet preparations (Gruessner et al., 2004). It must be noted, however, that robust long-term follow-up of patients receiving islet

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autografts has not been reported to date. Despite the potential risks for a living donor in terms of surgically induced diabetes and surgical complications, the demand for islet tissue and relative ease of implementation of living donor protocols into established islet transplant programs is likely to move this approach forward rapidly. Xenotransplantation Living donor islet transplantation may circumvent the wait for suitable donor tissue in some diabetic patients, but the risks to the donor and the possibility of insufficient islet yield to obtain insulin dependence remain to be significant concerns. Identification of a renewable xenogeneic source of islets would avoid the requirement for human islet donors altogether and could provide enough tissue to transplant diabetic patients as often as required. Pigs are particularly attractive as a xenogeneic islet donor since they are widely available, produce insulin that is functional in humans, and could be selected for certain donor characteristics. Of all types of experimental xenotransplantation, islet transplantation is probably the closest to clinical application. Over the past decade, a number of small clinical trials in islet transplantation using porcine islets have been reported, but few have resulted in reduced insulin requirements and no patients have achieved prolonged insulin independence (Groth et al., 1994; Elliott et al., 2000; Valdes-Gonzalez et al., 2005). Despite these set-backs, islet xenotransplantation using porcine tissue has remained an active area of research, and progress has been made over the past several years in experimental islet xenotransplantation using pre-clinical non-human primate models (Cardona et al., 2006; Hering et al., 2006; Rood et al., 2006). The generation of α1,3-galactosyltransferase-deficient pigs has provided a source of islet tissue lacking the major xenoantigens causing hyperacute rejection in pig-to-human xenotransplantation (Phelps et al., 2003). Still, it remains to be determined whether the transmission of endogenous retroviruses or other zoonotic infections from pig to human can be completely avoided in xenotransplantation, even with the establishment of highly monitored “clean” pig colonies (Fishman and Patience, 2004). While significant advances have been made in the area of islet xenotransplantation, it is unclear whether enough data has been generated to justify the move toward large scale clinical trials. However, there are verbal reports that clinical trials are ongoing in centers in China and Russia (Rood and Cooper, 2006). Stem-Cell Transplantation Unlike solid organ transplantation, which requires a complex vascularized tissue structure to restore function in a recipient, islet transplantation could be achieved through the development of a renewable source of stemcell derived β-cells. Substantial research efforts have been made in identifying suitable islet precursor cells that could be differentiated into an unlimited source of insulin-producing β-cells, but difficulties in producing physiologically regulated insulin secretion and control of proliferation have delayed progress in this area (reviewed in Bonner-Weir and Weir, 2005; Otonkoski et al., 2005). Some exciting data has been reported using genetically modified human fetal hepatocytes, but data in large animal models is lacking (Zalzman et al., 2003, 2005). The challenge of reproducing the highly differentiated neuroendocrine β-cell phenotype is significant, and more investigation in this area is required before stem-cell derived islets will see clinical application. Even as progress is made in this area, political and ethical issues may prevent the timely application of this technology in human subjects. Improving Engraftment Post-transplant In clinical islet transplantation, islets derived from multiple donors are often required to achieve insulin independence, which suggests that a significant portion of the transplanted islets must fail to engraft and become functional. It has been estimated that up to 70% of the transplanted β-cell mass may be destroyed in the early post-transplant period (Davalli et al., 1995; Biarnes et al., 2002; Ryan et al., 2005). Since this profound loss has

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been observed in both immunodeficient and syngeneic islet transplantation models, islet survival is likely regulated by non-immune-mediated stimuli. Following isolation, the islet microvasculature is completely disrupted, and upon implantation into the portal circulation, hypoxia persists while the islets revascularize, which can take up to 2 weeks (Dionne et al., 1993; Carlsson et al., 2001, 2002; Giuliani et al., 2005). During this engraftment period, the islets are continuously exposed to immunosuppressive drugs including tacrolimus and sirolimus, which are known to adversely impact β-cell survival and function (Hyder et al., 2005). These negative effects are likely compounded by the proximity of the transplanted islets and high concentrations of these drugs in the hepatoportal circulation, further degrading β-cell mass over time (Desai et al., 2003; Shapiro et al., 2005a). Another process which may influence islet engraftment and survival in the early post-transplant period has been termed the “instant blood-mediated inflammatory reaction” (IBMIR). Islets have been shown to naturally express tissue factor, a protein which acts as a receptor and cofactor for Factor VII, an important mediator of the coagulation cascade (Moberg et al., 2002). Isolated human islets release tissue factor along with glucagon and insulin, which ultimately leads to platelet activation and binding at the surface of the islets. This causes the formation of a fibrin capsule around the islet and disruption of the islet morphology (Bennet et al., 1999; Moberg et al., 2002; Ozmen et al., 2002). Most of this process has been characterized using an in vitro tubing loop model, so the true impact of this process in the clinical setting has yet to be fully characterized. However, examination of serum in patients undergoing islet transplantation has shown that a statistically significant increase in the serum concentration of thrombin/anti-thrombin complexes is present almost immediately following portal infusion, with peak levels occurring at 15 min, even when there was no clinical evidence of portal hypertension or intraportal thrombosis (Moberg et al., 2002). Given that platelet activation is one of the primary contributing factors in the generation of an inflammatory response, IBMIR is probably one of the important early processes in islet transplantation that elicits an immune response (Rabinovitch and Suarez-Pinzon, 1998; Moberg et al., 2002). Many studies targeted at enhancing islet survival during the early post-transplant period have been published, and a variety of different strategies have been tested. Some groups have aimed to enhance revascularization with vascular endothelial growth factor (VEGF), but these studies have not yet demonstrated that this approach significantly improves islet graft survival (Narang et al., 2004). Anti-coagulation strategies using injection of activated protein C or inhibition of thrombin have been studied as a means to inhibit IBMIR, but these interventions have shown only a modest benefit in a limited series of in vivo studies in animal models (Ozmen et al., 2002; Contreras et al., 2004). Clinical studies designed to prevent IBMIR are currently under investigation and should provide more insight into this area. Since the processes described above involve both extracellular (i.e. IBMIR) and intracellular (i.e. hypoxia) stimuli leading to β-cell death, another approach to preserve β-cell mass in the early post-transplant period has been to directly inhibit the apoptotic triggers which ultimately lead to loss of islet mass post-transplant. A variety of strategies have been explored in the experimental setting, and while promising data has been generated in vitro, demonstration of in vivo benefit to islet graft survival has been more elusive (Dupraz et al., 1999, 2000; Cottet et al., 2001, 2002; Cattan et al., 2003; Klein et al., 2004). Many studies have described inhibition of a variety of apoptosis-associated proteins, including cFLIP (cellular FLICE-inhibitory protein; prevents caspase-8 activation), A20 (inhibits NF-κB activation), Bcl-2, and Bcl-XL (mitochondria-associated anti-apoptotic proteins) (Dupraz et al., 1999, 2000; Grey et al., 1999, 2003; Cottet et al., 2001, 2002; Klein et al., 2004). A20 has shown promise, as its overexpression reduced the islet mass required in syngeneic islet transplantation in mice (Grey et al., 1999, 2003). Recently investigations using XIAP (X-linked inhibitor of apoptosis protein), which inhibits the downstream effector caspases that function in the final common pathway of apoptosis, have demonstrated promise in both human and rodent models of engraftment and in promoting murine islet allograft

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survival (Emamaullee et al., 2005a, b; Plesner et al., 2005). However, this area of research is currently limited by its requirement for genetic manipulation of islet tissue pre-transplant, which has proven to be quite variable and difficult to achieve in human islets. Also, these genetic alterations are most often regulated with viral vectors, which represent a highly controversial reagent for clinical use, especially in immunosuppressed transplant recipients. As pharmacological compounds which can reproduce or stimulate the expression of anti-apoptotic mediators are identified, this area of research will likely have a positive impact in clinical islet transplantation. Improved Immunomodulation: Toward Donor-Specific Tolerance One unique component of islet transplantation in patients with T1DM is the possibility of recurrent autoimmunity, which may elevate the demand for immunosuppression. Indeed, it has been well established using a rodent model of T1DM, the non-obese diabetic (NOD) mouse, that control of recurrent autoimmune reactivity to β-cells is one of the most difficult obstacles to overcome in islet transplantation (reviewed in Rossini et al., 2001; Pearson et al., 2003). Although it has been quite challenging to study recurrent autoimmunity in clinical patients, some evidence exists to suggest that levels of autoantibodies to glutamic acid decarboxylase (GAD) and IA-2 increase following islet transplantation, although the direct impact of this phenomenon on graft survival is not yet clear (Jaeger et al., 2000; Bosi et al., 2001). If recurrent autoimmunity does alter immunosuppressive drug functional thresholds, this presents yet another problem in the context of islet transplantation, as many of the drugs are directly β-cell toxic. In fact, up to 15% of non-diabetic patients who receive solid organ grafts can develop post-transplant diabetes as a result of calcineurin inhibitor therapy (i.e. tacrolimus) or steroids (i.e. prednisone) (Jindal et al., 1997; Djamali et al., 2003). Most patients that are candidates for islet transplantation have had disregulated diabetes for many years, and as such their renal status may be somewhat impaired (Shapiro et al., 2000). This leads to an increased susceptibility to the deleterious renal side effects of these immunosuppressive drugs, and thus limits the extent to which the dose can be increased to preserve graft function (Ryan et al., 2004a). It is therefore likely that immunosuppressive drugs either contribute directly to β-cell loss over time via toxicity, or indirectly by incomplete protection against recurrent autoimmunity and/or alloreactivity. Direct control of recurrent autoimmunity may enhance long-term graft function in islet transplantation. Attempts have been made to control autoimmunity at the time of diabetes onset, using various immunosuppressive agents such as azathioprine, prednisone, cyclosporin A, or anti-thymocytic globulin, but no significant benefit was observed (Elliott et al., 1981; Eisenbarth et al., 1985; Silverstein et al., 1988; Bougneres et al., 1990). Recent clinical studies using a modified anti-CD3 (hOKT3γ1(Ala–Ala) in patients with new onset T1DM have demonstrated that this treatment significantly improved C-peptide responses in these patients, which persisted for up to 2 years following treatment (Herold et al., 2005). Incorporation of this induction agent into clinical islet transplant protocols has suggested that it may enhance insulin-independence rates following single donor infusion, which may be related to its ability to curtail β-cell autoimmunity in these patients (Hering et al., 2004). Continued development of therapies targeted at regulation of autoimmunity will allow further refinement of immunosuppression protocols for islet transplantation in the future. In all types of transplantation, the ultimate goal is to develop therapeutic protocols that involve a brief period of treatment only during the initial post-transplant period, followed by the complete withdrawal of all immunosuppressive drugs. This phenomenon has been termed “operational tolerance,” since it may involve a passive ignorance of the graft or a more active T-cell tolerance to the graft antigens. In experimental transplantation, the difference in these two types of response is quite important and can be measured using retransplantation of donor type or third party tissue, with tolerance resulting in acceptance of the donor-type graft and rejection of the third party graft. In the clinical setting, however, the distinction may not be so critical, as both ignorance and tolerance would allow for reduction or withdrawal of immunosuppressive

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therapies. The most widely studied pathway to tolerance induction involves the inhibition of T-cell costimulation following T-cell receptor ligation. During an immune response, a T-cell must receive “signal 2” through interactions between its surface molecule CD28 and CD80 or CD86 on the antigen presenting cell to become fully activated. In order to disrupt this interaction, the extracellular portion of CTLA-4, which has a higher affinity for CD80/CD86 than CD28, has been artificially fused with human Fcγ to produce the soluble molecule CTLA4-Ig, designed for therapeutic purposes. CTLA4-Ig has been recognized for its potent immunoregulatory activity in murine models of T1DM, where treatment of young NOD mice dramatically reduced the incidence of T1DM (Lenschow et al., 1995). Our laboratory and others have demonstrated that CTLA4-Ig treatment in allogeneic islet transplantation can prolong graft survival but does not induce tolerance (Kirk et al., 1997; Levisetti et al., 1997; Benhamou, 2002; Casey et al., 2002). A new high-affinity version of CTLA4-Ig called belatacept or LEA29Y has been developed for clinical use and has shown considerable promise in promoting allograft survival in non-human primates and in clinical renal transplantation (Adams et al., 2002, 2005; Vincenti et al., 2005). These studies have generated considerable excitement for this approach and have prompted initiation of clinical trials using belatacept in clinical islet transplantation. A second costimulatory pathway that has been examined in transplantation involves the interaction between CD40 on antigen presenting cells and CD40L (CD154) on T-cells, leading to T-cell activation. This interaction also promotes B-cell differentiation and the activation of antigen presenting cells including macrophages and dendritic cells. Blockade of this pathway using anti-CD154 therapies demonstrated considerable promise in promoting tolerance induction in primate models early on, but further testing of the potent anti-CD154 blocking antibody (Hu5C8) has been halted due to unexpected thromboembolic complications in clinical trials (Kenyon et al., 1999; Kirk et al., 1999, 1997; Kawai et al., 2000). Recent development of therapeutic antibodies targeting the CD40 molecule appears to avoid this negative side effect and should prove to be important in future clinical tolerance induction protocols in islet transplantation (Adams et al., 2005).

SUMMARY AND CONCLUSIONS β-cell replacement through islet transplantation presents the best opportunity to treat T1DM and prevent the long-term serious complications associated with this disease. The concept of islet transplantation is not new, but investigators struggled to find success in achieving insulin independence until the introduction of the Edmonton Protocol in 2000. This has provided hope for many patients with diabetes, but islet transplantation, in its current form, is reserved only for those patients with the most severe disease. While up to 80% of recipients may attain and maintain insulin independence at 1 year post-transplant, insulin independence has not been sustainable over time, with the most recent Edmonton data suggesting that nearly 90% of recipients will have resumed insulin therapy at 5 years post-transplant, albeit with a much lower insulin requirement than before receiving an islet graft. Also, most patients continue to exhibit partial islet function with C-peptide secretion in sufficient amounts to avoid both glycemic lability and hypoglycemic unawareness, which greatly improves the quality of life for many patients. However, the current requirement for islets derived from two or more cadaveric donors severely limits the current availability of this procedure. There are multiple opportunities for intervention throughout the entire process, from pancreas procurement, shipment, and islet processing, through to strategies for enhanced islet survival after implantation. Priority areas for clinical trials currently include expansion of living donor protocols, interventions to impede the IBMIR process, and the use of non-diabetogenic and more “islet-friendly” immunosuppressive and tolerance induction strategies to effectively control both auto- and alloimmunity. Strategies targeted at preserving β-cell mass throughout the process will have a substantial and immediate impact on islet transplantation by reducing the amount of islet tissue necessary to reverse diabetes. Once

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some of these obstacles are overcome, islet transplantation will become available to a broader population of patients with T1DM, especially those early in the progression of their disease who will benefit most as the development of serious chronic secondary complications could be avoided.

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Herold, K.C., et al. (2005). A single course of anti-CD3 monoclonal antibody hOKT3{gamma}1(Ala–Ala) results in improvement in C-peptide responses and clinical parameters for at least 2 years after onset of type 1 diabetes. Diabetes 54: 1763–1769. Hyder, A., et al. (2005). Effect of the immunosuppressive regime of Edmonton protocol on the long-term in vitro insulin secretion from islets of two different species and age categories. Toxicol. In Vitro 19: 541–546. International Islet Transplant Registry (2005a). Jaeger, C., et al. (2000). Islet autoantibodies as potential markers for disease recurrence in clinical islet transplantation. Exp. Clin. Endocrinol. Diabetes 108: 328–333. Jindal, R.M., et al. (1997). Post-transplant diabetes mellitus: the role of immunosuppression. Drug Saf. 16: 242–257. Kaplan, B., et al. (2004). Effect of sirolimus withdrawal in patients with deteriorating renal function. Am. J. Transplant. 4: 1709–1712. Kawai, T., et al. (2000). Thromboembolic complications after treatment with monoclonal antibody against CD40 ligand. Nat. Med. 6: 114. Keen, H., (1994). The Diabetes Control and Complications Trial (DCCT). Health Trends 26: 41–43. Kelly, W.D., et al. (1967). Allotransplantation of the pancreas and duodenum along with the kidney in diabetic nephropathy. Surgery 61: 827–837. Kemp, C.B., et al. (1973). Effect of transplantation site on the results of pancreatic islet isografts in diabetic rats. Diabetologia 9: 486–491. Kempf, M.C., et al. (2005). Logistics and transplant coordination activity in the GRAGIL Swiss-French multicenter network of islet transplantation. Transplantation 79: 1200–1205. Kendall, D. M., et al. (1997). Pancreas transplantation restores epinephrine response and symptom recognition during hypoglycemia in patients with long-standing type I diabetes and autonomic neuropathy. Diabetes 46: 249–257. Kenyon, N.S., et al. (1999). Long-term survival and function of intrahepatic islet allografts in Rhesus monkeys treated with humanized anti-CD154. Proc. Natl Acad. Sci. USA 96: 8132–8137. King, H., et al. (1998). Global burden of diabetes, 1995–2025: prevalence, numerical estimates, and projections. Diabetes Care 21: 1414–1431. Kirk, A.D., et al. (1997). CTLA4-Ig and anti-CD40 ligand prevent renal allograft rejection in primates. Proc. Natl. Acad. Sci. USA 94: 8789–8794. Kirk, A.D., et al. (1999). Treatment with humanized monoclonal antibody against CD154 prevents acute renal allograft rejection in nonhuman primates. Nat. Med. 5: 686–693. Klein, D., et al. (2004). Delivery of Bcl-XL or its BH4 domain by protein transduction inhibits apoptosis in human islets. Biochem. Biophys. Res. Commun. 323: 473–478. Lacy, P.E. and Kostianovsky, M. 1967. Method for the isolation of intact islets of Langerhans from the rat pancreas. Diabetes 16: 35–39. Lakey, J.R., et al. (1996). Variables in organ donors that affect the recovery of human islets of Langerhans. Transplantation 61: 1047–1053. Larsen, J.L. (2004). Pancreas transplantation: indications and consequences. Endocr. Rev. 25: 919–946. Lenschow, D.J., et al. (1995). Differential effects of anti-B7-1 and anti-B7-2 monoclonal antibody treatment on the development of diabetes in the nonobese diabetic mouse. J. Exp. Med. 181: 1145–1155. Levisetti, M.G., et al. (1997). Immunosuppressive effects of human CTLA4Ig in a non-human primate model of allogeneic pancreatic islet transplantation. J. Immunol. 159: 5187–5191. Matsumoto, S., et al. (2005). Insulin independence after living-donor distal pancreatectomy and islet allotransplantation. Lancet 365: 1642–1644. Merrill, J.P., et al. (1963). Successful transplantation of kidney from a human cadaver. JAMA 185: 347–353. Moberg, L., et al. (2002). Production of tissue factor by pancreatic islet cells as a trigger of detrimental thrombotic reactions in clinical islet transplantation. Lancet 360: 2039–2045. Molinari, M., et al. (2005). Sirolimus-induced ulceration of the small bowel in islet transplant recipients: report of two cases. Am. J. Transplant. 5: 2799–2804. Murray, J.E., et al. (1963). Prolonged survival of human-kidney homografts by immunosuppressive drug therapy. N. Engl. J. Med. 268: 1315–1323.

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Narang, A.S., et al. (2004). Vascular endothelial growth factor gene delivery for revascularization in transplanted human islets. Pharm. Res. 21: 15–25. Nathan, D.M., et al. (2003). Intensive diabetes therapy and carotid intima-media thickness in type 1 diabetes mellitus. N. Engl. J. Med. 348: 2294–2303. Nathan, D.M., et al. (2005). Intensive diabetes treatment and cardiovascular disease in patients with type 1 diabetes. N. Engl. J. Med. 353: 2643–2653. National Diabetes Data Group (US), et al., 1995. Diabetes in America. National Institutes of Health, National Institute of Diabetes and Digestive and Kidney Diseases, Bethesda, MD. Newell, K.A., et al. (1996). Comparison of pancreas transplantation with portal venous and enteric exocrine drainage to the standard technique utilizing bladder drainage of exocrine secretions. Transplantation 62: 1353–1356. Otonkoski, T., et al. (2005). Stem cells in the treatment of diabetes. Ann. Med. 37: 513–520. Owen, R.J., et al. (2003). Percutaneous transhepatic pancreatic islet cell transplantation in type 1 diabetes mellitus: radiologic aspects. Radiology 229: 165–170. Owen, S. (2006). Pediatric pumps: barriers and breakthroughs. Diabetes Educ. 32: 29S–38S. Ozmen, L., et al. (2002). Inhibition of thrombin abrogates the instant blood-mediated inflammatory reaction triggered by isolated human islets: possible application of the thrombin inhibitor melagatran in clinical islet transplantation. Diabetes 51: 1779–1784. Pearson, T., et al. (2003). Islet cell autoimmunity and transplantation tolerance: two distinct mechanisms? Ann. NY Acad. Sci. 1005: 148–156. Phelps, C.J., et al. (2003). Production of alpha 1,3-galactosyltransferase-deficient pigs. Science 299: 411–414. Plesner, A., et al. (2005). The X-linked inhibitor of apoptosis protein enhances survival of murine islet allografts. Diabetes 54: 2533–2540. Rabinovitch, A. and Suarez-Pinzon, W.L. (1998). Cytokines and their roles in pancreatic islet beta-cell destruction and insulin-dependent diabetes mellitus. Biochem. Pharmacol. 55: 1139–1149. Reckard, C.R., et al. (1973). Physiological and immunological consequences of transplanting isolated pancreatic islets. Surgery 74: 91–99. Ricordi, C., et al. (1989). Automated islet isolation from human pancreas. Diabetes 38(Suppl 1): 140–142. Ricordi, C., et al. (1992). Human islet isolation and allotransplantation in 22 consecutive cases. Transplantation 53: 407–414. Rood, P.P. and Cooper, D.K. (2006). Islet xenotransplantation: are we really ready for clinical trials? Am. J. Transplant. 6: 1269–1274. Rood, P.P., et al. (2006). Pig-to-nonhuman primate islet xenotransplantation: a review of current problems. Cell Transplant. 15: 89–104. Rossini, A.A., et al. (2001). Islet cell transplantation tolerance. Transplantation 72: S43–S46. Ryan, E.A., et al. (2002). Successful islet transplantation: continued insulin reserve provides long-term glycemic control. Diabetes 51: 2148–2157. Ryan, E.A., et al. (2004a). Risks and side effects of islet transplantation. Curr. Diabetes Rep. 4: 304–309. Ryan, E.A., et al. (2004b). Assessment of the severity of hypoglycemia and glycemic lability in type 1 diabetic subjects undergoing islet transplantation. Diabetes 53: 955–962. Ryan, E.A., et al. (2005). Five-year follow-up after clinical islet transplantation. Diabetes 54: 2060–2069. Scharp, D.W., et al. (1990). Insulin independence after islet transplantation into type I diabetic patient. Diabetes 39: 515–518. Secchi, A., et al. (1991). Effect of pancreas transplantation on life expectancy, kidney function and quality of life in uraemic type 1 (insulin-dependent) diabetic patients. Diabetologia 34(Suppl 1): S141–S144. Senior, P.A., et al. (2005). Proteinuria developing after clinical islet transplantation resolves with sirolimus withdrawal and increased tacrolimus dosing. Am. J. Transplant. 5: 2318–2323. Shapiro, A.M., et al. (2000). Islet transplantation in seven patients with type 1 diabetes mellitus using a glucocorticoidfree immunosuppressive regimen. N. Engl. J. Med. 343: 230–238. Shapiro, A.M., et al. (2003). Edmonton’s islet success has indeed been replicated elsewhere. Lancet 362: 1242. Shapiro, A.M., et al. (2005a). The portal immunosuppressive storm: relevance to islet transplantation? Ther. Drug Monit. 27: 35–37.

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Shapiro, A.M., et al. (2005b). Strategic opportunities in clinical islet transplantation. Transplantation 79: 1304–1307. Silverstein, J., et al. (1988). Immunosuppression with azathioprine and prednisone in recent-onset insulin-dependent diabetes mellitus. N. Engl. J. Med. 319: 599–604. Statistical Data Reported by the US Scientific Registry of Transplant Recipients and the Organ Procurement and Transplantation Network. (2005). Vol. 2005, 2005b Sutherland, D.E., et al. (1980). Transplantation of dispersed pancreatic islet tissue in humans: autografts and allografts. Diabetes 29(Suppl 1): 31–44. Sutherland, D.E., et al. (2001). Lessons learned from more than 1,000 pancreas transplants at a single institution. Ann. Surg. 233: 463–501. Tsujimura, T., et al. (2004). Influence of pancreas preservation on human islet isolation outcomes: impact of the two-layer method. Transplantation 78: 96–100. Tzakis, A.G., et al. (1990). Pancreatic islet transplantation after upper abdominal exenteration and liver replacement. Lancet 336: 402–405. Valdes-Gonzalez, R.A., et al. (2005). Xenotransplantation of porcine neonatal islets of Langerhans and sertoli cells: a 4year study. Eur. J. Endocrinol. 153: 419–427. Villiger, P., et al. (2005). Prevention of bleeding after islet transplantation: lessons learned from a multivariate analysis of 132 cases at a single institution. Am. J. Transplant. 5: 2992–2998. Vincenti, F., et al. (2005). Costimulation blockade with belatacept in renal transplantation. N. Engl. J. Med. 353: 770–781. Walsh, T.J., et al. (1982). Portal hypertension, hepatic infarction, and liver failure complicating pancreatic islet autotransplantation. Surgery 91: 485–487. Williams, P. (1894). Notes on diabetes treated with extract and by grafts of sheep’s pancreas. BMJ 2: 1303–1304. World Health Organization (2002). Diabetes: The Cost of Diabetes, Vol. Fact Sheet No. 236. http://www.who.int/ mediacentre/factsheets/fs236/en/print.html. Zalzman, M., et al. (2003). Reversal of hyperglycemia in mice by using human expandable insulin-producing cells differentiated from fetal liver progenitor cells. Proc. Natl Acad. Sci. USA 100: 7253–7258. Zalzman, M., et al. (2005). Differentiation of human liver-derived, insulin-producing cells toward the beta-cell phenotype. Diabetes 54: 2568–2575.

48 Cell-Based Repair for Cardiovascular Regeneration and Neovascularization: What, Why, How, and Where Are We Going in the Next 5–10 Years? Doris A. Taylor and Andrey G. Zenovich

INTRODUCTION Cardiovascular disease (CVD) has become a major health problem throughout the world, exceeding infection and cancer as the leading cause of death in the Western world (Thom et al., 2006). In the United States, CVD has been the No. 1 killer since 1900, except for 1918, when it momentarily transferred its reign to influenza (Thom et al., 2006). Even though we have experienced a steady decline in mortality in CVD in general, and in acute myocardial infarction (AMI) in particular, since 1980 (Thom et al., 2006), CVD can be viewed as an “impending public health catastrophe” for several reasons. First, although mortality has declined owing to recent major advances in pharmacological therapy of atherosclerosis, in percutaneous and surgical revascularizations, and in therapies aimed at reducing the degree of hypercholesterolemia, hypertension, diabetes and post-AMI left ventricular (LV) remodeling (Pearson et al., 2002; Smith et al., 2006), CVD still accounts for 1 in every 2.7 deaths in the United States, which cumulatively translates into approximately 2.5 million deaths per year (Thom et al., 2006). Secondly, changes in incidence have not paralleled the reduction in mortality, largely because of the increased prevalence of the risk factors for CVD, such as hypertension, obesity and type 2 diabetes (Haffner, 2002; Pearson et al., 2002; Appel et al., 2006; Wyatt et al., 2006). Recent data show that the incidence of CVD in the 30–50 year old age group is actually on the rise, particularly owing to growing prevalence of CVD risk factors (Juonala et al., 2006; Yan et al., 2006). Thirdly, as a result of our getting better at preventing and treating AMI, improved survival post-AMI has been achieved, increasing the prevalence of heart failure (HF) such that nearly 40% of patients manifest HF by year 7 following their first AMI (Miller and Missov, 2001). The root causes of this unsatisfactory dynamic are not only pathophysiological. Pharmacological agents, such as angiotensin converting enzyme (ACE) inhibitors (Jong et al., 2003; Bertrand, 2004), angiotensin receptor blockers (ARB) (Cohn, 2002; Doggrell, 2005; Hernandez et al., 2005), and beta blockers (Thattassery and Gheorghiade, 2004; Jost et al., 2005; Torp-Pedersen

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et al., 2005) have demonstrated their abilities to decrease LV remodeling and reduce the number of HF-related hospitalizations. But the magnitude of their efficacy is very modest, when evaluated from the public health perspective (Levy et al., 2002). In addition, these therapies are underutilized not only geographically but also across diverse racial and economic groups, therefore the efficacy observed in clinical trials is not necessarily similar when these drugs are applied clinically (Lenzen et al., 2005). Furthermore, as survival of HF patients is also increased by technological advances such as implantable cardioverter defibrillators (ICDs) (Moss et al., 2002), which now allow termination of lethal arrhythmias outside the hospital, an increased number of patients survive with HF fueling a rise of health care expenditures. Lastly, the number of people over 65 years of age in the United States will double in the next 25 years as a result of the aging of the “baby-boomer” population. It is estimated that nearly 15% of this population will develop HF due to aging, CVD, and type 2 diabetes (Thom et al., 2006). It is clear that the urgency of this growing public health problem must be solved via a new, more advanced level of understanding of pathophysiology of atherosclerosis and engineering of therapies to be applied throughout the continuum of CVD; that is, to intervene after acute injury to prevent LV remodeling, to treat chronically failing myocardium to stop a progressive loss of cardiac function and worsening of symptomatic status, and also to halt CVD process altogether by restoring vascular health and thereby preventing injury. All these strategies have led to research efforts directed at cell-based repair to achieve functional regeneration of the vasculature at large and of the myocardium. In this chapter, we provide a brief overview of the state of cell-based therapies for vascular repair and provide a perspective on the development of this field in the near future.

THE STATE OF CELL THERAPY Application of stem cells to achieve vascular repair and regeneration remains novel. Developing any new therapeutic product and translating it from bench experiments to bedside efficacy is a process of multiple interrelated steps. The first step is the idea embodied in basic science. Next, that idea must be tested in clinically relevant animal models of disease. If the data indicate a potential therapeutic benefit, further testing occurs in several consecutive clinical studies. However, if unexpected issues arise or new pieces of the puzzle emerge (e.g. delivery-related issues in the case of cell therapy), the process should move back to the bench and, only when resolved, move again to bedside. The ultimate goal of this iterative process is a “clinical product” with broad applicability to patients, easy administration, and an extremely low frequency of adverse events in follow-up. Cell therapy at present is in the iterative stage between bedside and bench. The first set of ideas has successfully moved into clinical studies. Early clinical safety and efficacy data are emerging. New insights into the mechanisms of the effects seen (or lack thereof) are being garnered to produce the next generation of preclinical innovations to optimize the “clinical product.” The concept of a product itself has been undergoing changes as new cells and therapeutic strategies are being evaluated. Ten years ago the concept of tissue or organ (in this case, cardiac) repair with exogenous progenitor cells was unheard of. Only a few studies were published suggesting that injected cells could actually incorporate into the damaged heart (Marelli et al., 1992; Chiu et al., 1995). There were no studies that demonstrated functional improvement of the myocardium. The first preclinical study was published in 1998 (now 9 years ago) showing functional improvement of injured myocardium after transplantation of skeletal myoblasts (SKMBs) (Taylor et al., 1998). Since then, the field has been growing exponentially. Within the following 2 years, a clinical trial in which SKMBs were delivered as an adjunct to coronary artery bypass grafting (CABG) in HF patients began in Europe (Menasche et al., 2001). Now, multiple clinical trials using 5–6 different muscle-, bone-marrow-, and blood-derived cell types have been reported (Tables 48.2–48.3).

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This explosion of trials lies in the simplicity, straightforward nature, and timing of cell therapy. First, cell therapy offers an opportunity for repair of injury, rather than simply an augmentation of the remaining uninjured heart, often to its detriment. In other words, cell therapy provides hope of a permanent solution to a previously unsolvable problem. Next, the treatment makes sense to clinicians and patients. When cells die after AMI leading to HF (Abbate et al., 2006), it seems logical to prevent them from dying, as is a goal with current bone marrow mononuclear cells (BMNCs) trials early post-AMI, or replace them with new, undamaged cells, a current goal of virtually all HF trials. Because cell therapy primarily involves autologous cells, patients endorse it – since ethical concerns (such as with embryonic stem cell research or live allogeneic donors) are not an issue – and often seek out participation in clinical trials. In addition, as is often the case with anything cutting edge in medicine, it is exciting and prestigious to participate both for investigators and patients. In addition, cell therapy came just as the hope of angiogenic therapies (vascular endothelial growth factor, fibroblast growth factor, etc.) failed to deliver the new blood vessel growth they so enticingly promised (Henry et al., 2003; Simons, 2005). When therapeutic angiogenesis did not become the mainstream one-shot-fix-all treatment, there was an urgent need for a new frontier: for something that would fulfill the promise and become the “clinical product.” Cell-based therapies became “what’s next” in cardiovascular medicine. As it did so, the anticipated 5.3 billion dollar market potential in CVD has led scientists, companies, investors, clinicians, and patients into a plethora of first-in-man studies. As a result, within 10 years, we have moved from conception of a field to completed Phase I clinical studies, completed, and ongoing Phase II studies and additional “next generation” bench and pre-clinical studies. But has cell therapy arrived? The good news is the field is moving rapidly, and the possibility of cell-based therapies joining the clinical armamentarium to repair myocardium has begun to be supported by several Phase II patient studies. The bad news is the field is sometimes moving too fast, without a critical evaluation of the science behind the data or with trials that are outdated even by the time they begin to enroll patients. So it is not at all surprising that we have conflicting clinical outcomes: both negative and positive results with the same cells, in the same patients, and in similar pathophysiological contexts. We do not yet understand all the relationships among cells, engraftment, delivery, mechanism of action, and outcomes. We have yet to ask all the right questions. The field is just emerging, and with it our knowledge changes fast. Within the next few years, we should witness the completion of both surgical and percutaneous catheter-based Phase II studies with bone marrow, blood, muscle and possibly even fat-derived progenitor cells. We will likely see the initiation of at least one clinical trial utilizing resident cardiac progenitor cells. It is very likely that as this process further unfolds, the applicability and the maximum benefits of the cells types will segregate with a specific disease state; and as we gain more molecular/mechanistic insights, we will be enabled to choose the right cell for the right stage of the CVD continuum. The likely major impact will be provided by Phase II (and even Phase III) data regarding the use of BMNCs in the acute and subacute setting of AMI and the use of SKMBs or mesenchymal cells (MSCs) in HF patients. So has cell therapy arrived? The short answer is that it is arriving. In summary, this is a dynamic time in a new and exciting field: treatment strategies are being developed and modified almost every week; new cell types are being reported, novel delivery strategies are emerging, and slowly we are dissecting the mechanistic components of cellular cardiomyoplasty.

CELL-BASED REPAIR AS A MEANS OF REGENERATION IN CVD The goal of regeneration as a therapeutic process is to repopulate diseased area of the tissue with exogenous (by direct application) or endogenous (by stimulation of production and homing) cells that restore function

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of the organ and halt the disease process that resulted in tissue damage. In AMI and HF, cardiac regeneration – restoration of myocardial function and cessation of LV remodelling – is the ultimate goal. Functional vasculature has to develop and electrical conductance has to occur in underperfused areas and/or fibrous scar to provide optimal contractile force. However, full regeneration is currently not yet achievable, although it may be possible in the future through tissue engineering methods, such as application of cell-based patches to critically injured areas (Liu et al., 2004). Futurism apart, what can be claimed as a realistic goal of cell therapy in the next 5 years is restoring at least some degree of mechanical function and perfusion to the injured, dilated, and decompensated myocardium – even in the absence of full regeneration. Endogenous tissue repair is inherent in the human body. We realize now that virtually every organ in the body, including heart, is capable of ongoing maintenance throughout much of our lifespan, and only with ongoing disease, aging, or when overwhelmed by a catastrophic event does this process fail. Yet, repair in the heart and vasculature appears to fail more often than in other organs. As we understand more about the elements of this endogenous cardiac process – what initiates it, what controls it, what allows it to work, and what leads to failure – we will refine our targets for repair. Manipulating those targets first pre-clinically and then clinically will become increasingly important as the field advances. Meanwhile, what can we learn from noncardiac repair that could apply to heart? Successful endogenous cell-based repair of many non-cardiac tissues routinely occurs after injury. For example, in bone, endogenous repair occurs if three simple premises are fulfilled: reduction of inflammation, fixation of new cells, and perfusion of the tissue (Ott and Taylor, 2006; Taylor et al., 2006). Although bone injuries are not usually ischemic in origin, the cascade of wound healing is somewhat similar to the process in myocardium: inflammation leads to clearing of necrotic (bone) tissue and formation of a fibrous scar. However, unlike in the heart where the scar stays (Kwong et al., 2006), in bone the scar is replaced by regenerating osteoclasts and osteoblasts (Maddi et al., 2006; Wutzl et al., 2006), and in a relatively short period the bone is completely healed (Karladani et al., 2001). So why does endogenous cardiac repair fail? In bone repair, mechanical stress has to be minimized. Patients’ fractured extremities are routinely immobilized to allow healing. When an injured bone is insufficiently immobilized, an unstable scar forms (and this process frequently takes longer than formation of a stable scar), which could be compared to the process of LV remodeling and progression to HF. Next, perfusion has to be maintained. When perfusion is compromised, tissue necrosis leads to sequestration, chronic inflammation, and pain – a process not unlike chronic post-AMI angina (Hansen-Algenstaedt et al., 2006). Furthermore, in the absence of adequate perfusion, a mature bone does not arise (HansenAlgenstaedt et al., 2005, 2006), similar to the need for functional angiogenesis to support new muscle formation in the myocardium (Simons, 2005). Finally, cells have to home to the site of the lesion. Bone regeneration is ultimately performed by endogenous progenitor cells (osteoblasts and osteoclasts) that migrate into the lesion (Wright et al., 2005). In the injured myocardium, this process breaks down. Rather than intensive colonization with the desired cardiac progenitor cells, the lesion is colonized primarily by fibroblasts, resulting in collagen deposition and scar formation, instead of nascent myocardium (Lu et al., 2004). This comparison with bone makes endogenous cell-based repair in the heart seem difficult, as the need for immobilization cannot be achieved, we do not understand all the components of homing of cells to the lesion, and adequate perfusion is a product of endothelial health, function, and regulation, which is often missing in CVD.

VASCULAR INTEGRITY IS A BALANCING ACT IN WHICH INFLAMMATION IS KEY: INJURY VERSUS REPAIR What we do understand, however, is that atherosclerosis represents the failure of attempts of endogenous vascular repair in response to repeated injuries to the vessel wall (Goldschmidt-Clermont et al., 2005). After Ross’

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seminal work introducing the “response to injury” hypothesis whereby the primum movens in atherogenesis was endothelial denudation (Ross, 1993), it became apparent that endothelial cells directly regulate vascular function, transport of solutes, and antithrombotic properties of the blood–tissue interface (Hansson, 2005; Feletou and Vanhoutte, 2006; Schmieder, 2006). More recently, the revised “response to injury” hypothesis focuses on endothelial dysfunction rather than denudation as a trigger for the inflammatory responses and progression of atherosclerosis (Sun et al., 2000; Hansson, 2005). Once regarded simply as a static barrier between tissue and blood, vascular endothelium is now known to play a key regulatory and integrating role in the initiation and progression of atherosclerosis (Feletou and Vanhoutte, 2006; Schmieder, 2006). We now understand that vasodilatory, antiplatelet, and antithrombotic processes are primarily regulated at the endothelial level, and the loss of normal endothelial function may be the most important driver of the balance in favor of inflammation and thrombosis (Feletou and Vanhoutte, 2006; Schmieder, 2006), which, in turn, contributes to the transition between stable and unstable angina (Kostner et al., 2006). Indeed, atherosclerotic plaques have been shown to be more thrombogenic when they express higher levels of tissue factor, a key endothelial regulator in the extrinsic clotting cascade (Marmur et al., 1993; Libby and Theroux, 2005). Therefore, in CVD, the repair of endothelium with exogenously applied cells must be one of the main steps to achieve restoration of perfusion and slowing and possibly halting both LV remodeling and atherosclerotic lesion progression. We believe that ongoing endogenous repair is a process that reflects the balance between protective and detrimental factors (Goldschmidt-Clermont et al., 2005), and that by providing reparative cells it is possible to tip this balance to allow for repair of the injured tissues (Figure 48.1). Vascular repair is most likely a stepwise process. First, the tissue undergoes injury (e.g. ischemic insult), as shown in Figure 48.1, accompanied by inflammatory response (release of cytokines and chemokines). The inflammatory milieu and injury-specific mediators recruit “detrimental” BMNCs (e.g. CD45, CD11, CD3) and exacerbate injury (Yu et al., 2002; Abbate et al., 2004). At the same time, mobilization of “reparative” BMNCs (e.g. AC133, CD34, CD31, KDR) occurs in an endogenous attempt to promote repair (Fadini et al., 2006; Haider, 2006). If sufficient amounts of reparative cells are recruited (or supplied) the balance tips toward repair and inflammations is halted. Three factors play a crucial role: availability of the “reparative” BMNCs, the capability of these cells to home to the site of injury, and their functional capacity to initiate and propagate repair. Obviously, any of these factors could be the weakest link, and that alone would make repair inefficient allowing injury to prevail. The balance between the injury and repair may also reside in the quantitative, functional, and mechanistic/paracrine relationships of both “protective” and “detrimental” BMNCs at the site of injury (Figure 48.2). Therefore, from the balance point of view, we believe that these relationships are different in mild, moderate, and severe disease (Figure 48.2). Recent studies have shown that the availability of the bone marrow progenitors is reduced in number in severe CVD (Werner et al., 2005; Kunz et al., 2006). Similar findings as well as a profound reduction of functional capacity of these cells have been observed in patients with HF (Valgimigli et al., 2004). The decline in number and function is most likely not the exclusive feature of the advanced stages of atherosclerotic process because similar results have been reported much earlier in the disease process, when endothelial damage is only provoked by hypertension or type 2 diabetes (Hill et al., 2003; Fadini et al., 2005; van Zonneveld, 2006). Inverse associations with aging itself have also been shown (Thorin-Trescases et al., 2005; Shaffer et al., 2006). Our preliminary data in ApoE / mice as well as the recently published WISE study (Bairey Merz et al., 2006) indicate that males and females display different speeds of progression of disease that correlate with progenitor cell profiles. If these data hold, gender differences in repair will need to be extensively studied as the results may have major implications not only for the field of cell therapy but for the entire cardiovascular medicine. Those insights could not only provide guidance for developing next generation clinical studies, but could greatly advance our understanding of pathophysiological changes in men and women.

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Signal

Mobilization of progenitor cells

Injury

Protective cells

Detrimental cells mation

Inflam

Repair

Disease

Endothelial integrity 200 μm

200 μm

Figure 48.1 Schematic representation of the main factors influencing the balance between positive and negative factors (cells, cytokines, chemokines) influencing progression of disease and endogenous repair.

Disease progression

(a) Males

Females

Failure of repair (b)

Reparative PCs Pro-inflammatory PCs

Figure 48.2 Schematic representation of relationships between CVD progression (a) and pro-inflammatory (solid line) and reparative progenitor cells (dashed line) number/function (b). As ongoing injury occurs (panel b, left side), the number of pro-inflammatory cells increases (reflecting inflammation), and the reparative progenitor cells compensates in attempt to repair. However, with failure of repair (shown by arrow), reparatory cells falls, and pro-inflammatory cells increase leading to disease progression (panel b, right side). Dotted lines indicate the proposed dynamic of both reparatory and pro-inflammatory cells when repair is successful. Data from animal experiments in our laboratory and data from the Women’s Ischemia Syndrome Evaluation (WISE) study (Bairey Merz et al., 2006) were used to reflect the time course of CVD in (a). Relationships of pro-inflammatory and reparative cells were portrayed based on recent publications and the preliminary data of our laboratory (b). Abbreviations. PC: progenitor cells. See text for a detailed discussion.

818 THERAPEUTIC APPLICATIONS: CELL THERAPY

Overall, there are several important aspects in cardiac repair. First, the failure of repair occurs earlier in the disease process than previously thought (i.e. before a clinical event) suggesting that owing to age, atherosclerosis, and risk factors, the availability of the progenitor cells and their function decline (Figure 48.2) (Hill et al., 2003). Secondly, repair is most likely evident in the mild–moderate/moderate phase of the disease, where it ultimately fails and injury progresses (Goldschmidt-Clermont et al., 2005). Unfortunately (for detection of abnormalities) many patients in this stage are asymptomatic (Blumenthal et al., 2006). Finally, mobilization of these cells does occur following the acute injury of AMI (Lev et al., 2005; Li et al., 2005; Massa et al., 2005), however, the response is not sufficient for two main reasons: (a) the availability of the cells and their function are greatly reduced (Thum et al., 2006) and (b) a recent study showed that revascularization abolish endogenous mobilization (Muller-Ehmsen et al., 2005). We also now understand that the reduction of inflammation is one of the key steps in successful vascular repair. Recent research has demonstrated that inflammation plays a key role in CVD (Shin et al., 2002; Hansson, 2005; Libby and Theroux, 2005). Indeed, endothelial-dependent relaxation is impaired early in atherogenesis (Feletou and Vanhoutte, 2006). In addition, local administration of proinflammatory cytokines impairs endothelium-dependent dilatation in humans (Bhagat and Vallance, 1997; Shin et al., 2002). Furthermore, clinical measures of coronary endothelial dysfunction are associated with myocardial ischemia in the absence of flow-limiting lesions (Suwaidi et al., 2000) and predict cardiovascular events, including stroke (Targonski et al., 2003). In the early atherosclerotic lesion, immune cells predominate, and their effector molecules significantly contribute to the progression of the plaque (Paoletti et al., 2004; Goldschmidt-Clermont et al., 2005; Libby and Theroux, 2005). Inflammatory response has been shown to be a large part of acute coronary syndromes (ACS)/ AMI (Ikeda, 2003; Paoletti et al., 2004; Libby and Theroux, 2005). The cascade of events leading to clinical atherosclerosis begins with lipid deposition and oxidation, and initiation of an arterial lesion site within the endothelial lining (Libby and Theroux, 2005). These acute processes trigger platelet aggregation and monocyte/macrophage infiltration (Ferns and Avades, 2000; Ikeda, 2003). Smooth muscle cell (SMC) apoptosis is observed within the first few days following the initial assault (Sata et al., 2000). During this time, there is upregulation of pro-inflammatory that recruit various cells including monocytes, macrophages, neutrophils, leukocytes, and smooth muscle precursors to the injury site (Simon et al., 2000; Fujiyama et al., 2003; Ikeda, 2003; Schober et al., 2003; Weber et al., 2004). Studies have demonstrated that by blocking these cytokine signals, a significant reduction in lesion formation is achieved, suggesting that inflammation and/or the cells these signals recruit are a cause of plaque progression (Rogers et al., 1998; Simon et al., 2000; Schober et al., 2003; Weber et al., 2004). A reduction of these signals is associated with re-endothelialization of the artery and the “end to the injury” process (Ferns and Avades, 2000; Schwartz et al., 2004). Extracellular lipids and foam cells form the center of the atherosclerotic plaque, which is surrounded by SMCs and collagen matrix (Paoletti et al., 2004; Libby and Theroux, 2005). However, it is the T-cells, macrophages, and mast cells present in the shoulder region of the plaque ensure atheroma growth (Frostegard et al., 1999). Many of the immune cells show the signs of activation and release pro-inflammatory cytokines (Shin et al., 2002). Recent analysis of soluble inflammatory mediators has been driven by the expectation that they may be used as indicators of presence and/or progression of CVD (Baldus et al., 2003; Labarrere and Zaloga, 2004; Paoletti et al., 2004). ACS occurs when activation of the atherosclerotic plaque leads to accelerated thrombogenesis and blocks the blood flow to the myocardium downstream from the lesion (Libby and Theroux, 2005). Plaque activation is a complex inflammatory process, in which metalloproteinases along with a number of cytokines directly attack collagen and other components of the tissue matrix, which surrounds the lipid core of the plaque (van der Wal et al., 1994). One of the indicators of inflammation is C-reactive protein, which has been found to correlate with (and reflect) the presence of vascular inflammation and endothelial activation (Labarrere and Zaloga, 2004). CRP is elevated in patients at a high risk and/or experiencing ACS (Tsimikas et al., 2006). Increased levels of CRP indicate upregulation of many

Cell Therapies for Repair and Regeneration 819

inflammatory cytokines (IL-6, IL1β, TNFα) and adhesion molecules (ICAM-1, VCAM-1, E-selectin), facilitating the inflammatory process, and promoting cell recruitment and attachment, ultimately advancing CVD (Shin et al., 2002; Paoletti et al., 2004; Jain and Ridker, 2005; Tsimikas et al., 2006). Other inflammatory markers such as MPO, SDF-1, and MCP-1 have also been associated with CVD (Shin et al., 2002; Baldus et al., 2003; Schober et al., 2003). However, reduction of inflammation in clinical settings is problematic. There is no single drug therapy which addresses all components of inflammation or results in the “end to the injury” response. Statins have been shown to be beneficial in reduction of hs-CRP, MCP-1 and some inflammatory cytokines (Jain and Ridker, 2005; Chello et al., 2006). However, whether the effect of statins on inflammatory biomarkers is a function of reduction of low-density lipoprotein (LDL) cholesterol or represents a separate pleiotropic effect (including possible mobilization of progenitor cells), represents a matter of considerable debate (Davidson, 2005). Other therapeutics, such as aspirin and ACE inhibitors have also been shown to elicit a beneficial effect on some components of inflammatory process, but their administration does not abolish the entire ravaging cytokine and soluble marker cascade (Peeters et al., 1998; Lauten et al., 2003; Tsikouris et al., 2004). Non-steroidal agents, cyclooxygenase-2 inhibitors, and glucocorticosteroids are potent anti-inflammatory drugs, but their administration alters the course of ACS and combat restenosis yielded more adverse effects than clinical benefits (Lee et al., 1999; Mukherjee et al., 2001; Niederberger et al., 2001; Tamai et al., 2002; Niederberger et al., 2004; Khan and Mehta, 2005; Krotz et al., 2005; Nicolae et al., 2005; Williams et al., 2005). We have recently shown that repeated intravenous injection of BMNCs from young ApoE/ mice prevented further progression of atherosclerotic plaque lesions in old ApoE null mice (Rauscher et al., 2003). Injected cells differentiated into endothelial cells and engrafted in atherosclerotic lesions of recipient animals. Comparison of bone marrow progenitor cell (BMC) profiles showed a specific depletion of intermediate vascular progenitor cells (CD31/CD45), without parallel changes in more primitive stem cells (sca-1, c-kit, or CD34) or mature vascular cells (VEGFR-2), most likely accounted for the age-related loss of bone-marrowderived vascular repair capacity. In addition, the treatment with BMNCs reduced IL-6, and the effect was maintained up until day 15 post-injection, and then the levels would increase again (Rauscher et al., 2003). Significant regression of plaque that occurred in concert with the reduction of IL-6 demonstrated that exogenous BMNCs might mediate regression of atherosclerotic plaque via anti-inflammatory action/positive interference with immune responses that are associated with the atherosclerotic process. However, when BMNCs were given in sex-matched and mismatched fashion only males that received female marrow showed the highest degree of plaque regression (Nelson, unpublished data). Further dissection of the cytokines involved in BMNC-mediated atheroprotection is in its final stages in our laboratory. Preliminary data suggests that plaque regression may be mediated by a distinct group of cytokines, rather than by IL-6 alone (Zenovich, unpublished data). Because atherosclerosis is an inflammatory disease (Paoletti et al., 2004; Hansson, 2005; Libby and Theroux, 2005), the antiinflammatory properties of BMNC administration and resulting plaque regression offers hope that the treatment with exogenous progenitor cells truly interferes into the pathogenesis of the disease. In addition, recent data on the regression of atherosclerosis with rosuvastatin have demonstrated that this drug is capable of reduction of plaque volume when measured by intravascular ultrasound technique (Nissen et al., 2006). Hopefully one day in the future, a cocktail of a statin drug and bone marrow progenitors will be a therapeutic option in interventional cardiology. While we only have a conceptual understanding of repair, and many questions are yet unanswered, mediation of this process by progenitor cells makes sense. After all, if cells that originate from the bone marrow did not have the ability to fight inflammation and/or pathogens and participate in healing, the organism would not survive for too long. When the number and/or function of these cells decreases because of a systemic disease (such as type 2 diabetes) (Fadini et al., 2005), aging or toxic substances (such as smoking) (Raupach et al.,

820 THERAPEUTIC APPLICATIONS: CELL THERAPY

2006), healthy and functioning endothelium fails and adverse consequences manifest clinically (as atherosclerosis, thrombosis and its micro- and macrovascular sequelae) (Feletou and Vanhoutte, 2006). A recent data analysis modeled the potential health effects of bone-marrow-derived progenitor cell therapy using the follow-up data (1950–1996) of the Framingham Study. To model CVD mortality, progenitor cell therapy was applied at age 30, with the effect assumed to be a 10-year delay in atherosclerosis progression. This study suggests that progenitor cell therapy might increase life expectancy in the population by as much as the complete elimination of cancer (in females, an additional 3.67 versus 3.37 years; in males, an additional 5.94 versus 2.86 years, respectively) (Kravchenko et al., 2005). These exciting findings fuel enthusiasm for cell therapy to halt and ultimately prevent CVD. In summary, in contrast to other tissues in the human body that successfully accomplish the process of repair, cardiac repair becomes inefficient prior to AMI or HF. This failure to endogenously repair/regenerate the vasculature stems from the reduced availability of the circulating bone marrow progenitors and existent inflammatory milieu. Exogenous BMNCs show attractive qualities of plaque regression and reduction of inflammation. Interventions to tip the balance of injury and repair to promote repair may need to be applied earlier in the disease process to ensure efficacy. Further research is needed to optimize that target, as well as to define specific inflammatory mechanisms that play a major role in vascular repair.

REPARATIVE POTENTIAL OF CELL THERAPY Autologous SKMBs The idea of using skeletal muscle to repair the heart evolved well before cell therapy emerged as a possible treatment option. In 1987, after being preconditioned by chronic pacing, the latissimus dorsi muscle was surgically wrapped around the failing heart to provide contractile support to the left ventricle – a procedure named “dynamic cardiomyoplasty” (Chachques et al., 1987). Thereafter, cellular cardiomyoplasty was introduced, when cells derived from the C2C12 SKMB transformed cell line were successfully transplanted into normal mouse hearts (Koh et al., 1993). SKMBs, derived from muscle “satellite cells,” were first described in 1961 as cells that regenerate damaged skeletal muscle (Mauro, 1961). SKMBs expand and form neofibers after muscle injury. It is not surprising that these cells were the first candidates for cardiac repair. In 1994, Magovern et al. reported the first successful transplantation of SKMBs into an injured heart (Zibaitis et al., 1994). The critical finding that transplanted SKMBs survived and formed striated muscle grafts within the damaged cardiac tissue was followed by several independent experimental studies investigating the engraftment potential of these cells. When we showed in 1998, for the first time, that the successful engraftment of SKMBs into injured myocardium improved LV function and attenuated remodeling (Taylor et al., 1998), it was deemed novel by many and unrealistic by some. The mechanisms of how these successfully engrafted SKMBs improved function was unclear. It appeared that the muscle cells could improve contractility of the scarred heart without strict transdifferentiating into cardiomyocytes. SKMBs appeared to yield two populations of cells in injured heart: myogenin-positive transplanted skeletal muscle-like cells in the center of the scar and a second population of myogenin-negative more primitive cardiac muscle-like cells “recruited” around the scar periphery (Atkins et al., 1999a). The transplanted SKMBs appeared to adapt to the surrounding myocardium by forming slow twitch myofibers that were electrically isolated from host cardiomyocytes and yet to improve LV performance (Murry et al., 1996; Atkins et al., 1999b). These results, stating an improved cardiac function without full integration of transplanted cells into the host myocardium and with recruitment of endogenous cells, raised questions about potential mechanisms that

Cell Therapies for Repair and Regeneration 821

provide functional benefit. Numerous mechanisms have been suggested – from changes in LV wall stress and/or geometry to active mechanically induced contraction of the injected cells – but the exact mechanism(s) underlying the beneficial effect are still the matter of considerable debate (Ott and Taylor, 2006). It is likely that the improvement of function is a result of both a direct effect of the transplanted SKMBs on LV performance, and an indirect “paracrine” effect on endogenous cell recruitment and on LV remodeling (van den Bos and Taylor, 2003). As the mechanisms underlying the positive effects of SKMB transplantation are still not fully understood, there is some discordance of thought as to whether autologous myoblasts improve contraction or just prevent further deterioration of the injured heart – despite much preclinical data showing a direct positive effect of the cells above that seen with other or sham treatments (van den Bos et al., 2004). However, from a clinical point of view, whether the cells do one or the other, may not be the right question to ask, as even attenuation of remodeling is a therapeutic avenue much needed in the HF armamentarium, since the current drug therapies are only capable of a very moderate effect on LV remodeling, only approximately 30%–40% (Reiffel, 2005). Despite these and other remaining questions, there are approximately 15 years of preclinical data available have shown that autologous SKMBs transplantation can augment both diastolic and systolic myocardial performance in a number of animal models after both acute and chronic injury. These preclinical data opened the field of CVD to this new therapeutic approach. The advantages of autologous SKMBs for treating patients with CVD/HF extend beyond the evidence of benefit in preclinical studies. By using self-derived SKMBs, it is possible to overcome the major limitations associated with allogeneic cell-based treatments: a critical shortage of donor tissue and the clinical complexities of immunosuppression. By using adult-derived cells, it is possible to avoid the ethical dilemma associated with embryonic stem cells. By using primary cells rather than immortalized or totipotent stem cells, the likelihood of tumor formation after SKMB transplantation is decreased (Tremblay et al., 1991). By using muscle-forming, myogenic cells, the regeneration of contractile muscle in an infarcted cardiac region is more likely. By using relatively ischemia-resistant SKMBs rather than cardiocytes, a higher level of engraftment and survival is likely to occur in infarcted regions – where transplanted cardiocytes perish (Reffelmann et al., 2003). Based on these advantages and the suggestions that regeneration of functional muscle in infarct is possible after autologous SKMBs in pre-clinical models (Taylor et al., 1998; Hutcheson et al., 2000; Fuchs, S. et al., 2001; Ohno et al., 2003; Thompson et al., 2003; Agbulut et al., 2004; Hiasa et al., 2004; Horackova et al., 2004; Ott et al., 2004; Ott et al., 2006) (Table 48.1), clinical studies in both Europe and the United States took off (Field and Reinlib, 2000). Although the initial clinical data (Herreros et al., 2003; Menasche et al., 2003; Smits et al., 2003; Chachques et al., 2004b; Ince et al., 2004; Siminiak et al., 2004; Dib et al., 2005; Siminiak et al., 2005; Gavira et al., 2006;) (Table 48.2) appear encouraging, myoblast transplantation is not without potential limitations. The first limitation is associated with any autologous cell that has to be expanded in the laboratory (i.e. myoblasts, mesenchymal stem cells, endothelial progenitor cells (EPCs)) that the use of autologous cells necessitates sufficient time between injury and injection to allow cell expansion in vitro. In normal healthy donors and in many HF patients, this ranges from several days to several weeks depending on the cell type, which does not seem problematic if the treatment can wait. In the case of AMI, where an early treatment may be beneficial, either alternative cells can be employed, or if myoblasts are truly superior, allogeneic cells may ultimately offer a solution after the acuity has subsided. Of note, the 2–3-week period after AMI is one of the flourishing inflammation (Paoletti et al., 2004; Libby and Theroux, 2005), which in preclinical studies has been associated with increased loss of cells after transplantation (Suzuki et al., 2004). Thus, a delayed treatment with autologous SKMBs could still be beneficial. A second potential limitation to any cardiac cell therapy is the (in)ability of the transplanted cells to electrically integrate with native tissue (Abraham et al., 2005). Myoblasts are the most well studied cell type, yet it is not clear if, and/or how, do they integrate into surrounding myocardium. Nor is it understood what impact various degrees of integration may have on either LV function or continuous normal sinus rhythm. Menasche

822 THERAPEUTIC APPLICATIONS: CELL THERAPY

Table 48.1 Selected Clinically-Relevant Pre-Clinical Models Used for Cell Therapy Cell Type(s)

Species

Model

Outcomes

Investigator, Country

SKMBs

Rabbit

Cryoinjury

LV systolic function and diastolic relaxation improved in 7/12 rabbits, correlated with engraftment of cells.

Taylor et al. (1998) (US)

SKMBs

Pig (targeted cell placement with Da Vinci robotic system)

LAD or LCX embolization

Cells successfully transplanted into apical, anterior and lateral target segments; LVEF, wall thickening, regional wall motion, LV end-diastolic volume improved.

Ott et al. (2006) (US)

SKMBs vs. FBs

Rabbit

Cryoinjury

Diastolic performance improved with FBs and SKMBs. FBs reduced but SKMBs increased systolic function.

(Hutcheson et al., 2000) (US)

SKMBs vs. BMNCs

Rabbit

Cryoinjury

Similar improvement of LV systolic function and equal degree of engraftment with SKMBs and BMNCs. Subset of BMNCs differentiated towards a myogenic phenotype.

Thompson et al. (2003) (US)

SKMBs vs. CFS vs.CMs

Guinea pig

LAD Ligation

CMs and CFs concentrated in infarction territory; CMs formed gap junctions with native cells, CFs did not; SKMBs proliferated and partially differentiated into cardiac phenotype by 2–3 weeks post-procedure. Gap junctions present.

Horackova et al. (2004) (Canada)

Human SKMBs vs. Human AC133

Rat

LAD Ligation

SKMBs increased LV EF by 15 5%, AC133 by 7 3% (controls reduced by 8 4%). Engraftment larger with SKMBs.

Agbulut et al. (2004) (France)

SKMBs vs. BMNCs

Rat

LAD Ligation

Combination of cell types prevented LV remodeling at 8 weeks post-procedure.

Ott et al. (2004) (Austria)

SKMBs vs. VSMCs

Hamster

Cardiomyopathy (Δ-sarcoglycandeficiency)

Attenuation of LV remodeling was greater with SKMBs vs. VSMCs.

Ohno et al. (2003) (Canada)

BMNCs

Mouse

LAD Ligation

Reduction of myocardial infarction size.

Hiasa et al. (2004) (Japan)

BMNCs

Pig

Ameroid Occluder, Placement on LCX

BMNCs secreted angiogenic growth factors, augmented perfusion and LV function.

Fuchs et al. (2001) (US)

Abbreviations. AC133: immature endothelial progenitor cells; CF: cardiac fibroblast; CM: cardiomyocyte; EF: ejection fraction; FB: fibroblast; LAD: left anterior descending artery; LCX: left coronary circumflex artery; LV: left ventricle; VSMC: vascular smooth muscle cell.

and colleagues did not observe any electrical integration of myoblasts pre-clinically (Scorsin et al., 2000). In support of that, Suzuki et al. (2001) reported that in the absence of connexin-43 overexpression, SKMBs did not couple very well with surrounding myocardium. Yet the cells appear to synchronously contract in with surrounding tissue and contribute to overall cardiac performance. Early clinical data (Table 48.2) suggest that

Cell Therapies for Repair and Regeneration 823

Table 48.2 Cell Therapy Trials with SKMBs Year

Investigator, Country

Patients No.

Diagnosis

Average Delivery Route SKMB Dose (106)

Outcomes

2003

Menasche et al. (2003) (France)

10

Post-AMI HF

871

Transepicardial without CABG

NYHA class improved to 1.6 0.1 from 2.7 0.2; LVEF increased to 32 1% from 24 1%.

2003

Herreros et al. (2003) (Spain)

11

Ischemia  prior AMI

221

Transepicardial with CABG

LVEF increased to 53.5 5% from 35.5 2.3%, regional wall motion (by E) and viability (glucose update by PET) improved.

2003

Smits et al. (2003) (Netherlands)

5

Post-AMI (anterior) HF

196

Transendocardial guided by electromechanical mapping of LV

LVEF increased to 41 9% from 36 11%, regional contractility (by MRI) significantly increased.

2004

Chachques et al. (2004b) (France)

20

Post-AMI

300

Transepicardial without CABG

LVEF normalized to 52.0 4.7% (baseline  28 3%), wall motion score improved, glucose update (by PET) increased.

2004

Siminiak et al. (2004) (Poland)

10

>3-month post-AMI

0.4-50 (range)

Transepicardial with CABG

Mean LVEF improved to 42% from 35.2%

2004

Ince et al. (2004) (Germany)

6 (6 controls)

Ischemic HF

210

Transendocardial

LVEF rose to 32.2 10.2% from 24.3 6.7% (in controls decreased to 21.0 4.0%), walking distance and NYHA class significantly improved.

2005

Dib et al. (2005) (US)

30

Post-AMI HF 2.2-300

Transepicardial with CABG or LVAD

LVEF increased to 36% at year 2 post-procedure vs. 35% at year 1 vs. 28% at baseline.

2005

(Siminiak et al., 2005) (Poland)

9

Post-AMI HF 17-106 (range)

Transcoronary

LVEF improved 3–8% in 6 of 9 patients; NYHA class improved in all patients.

2006

Gavira et al. (2006) (Spain)

12 (14 historical controls)

Post-AMI

Transepicardial with CABG

LVEF rose to 55.1 8.2% from 35.5 2.3% at baseline (controls 38. 6 11.0%); wall motion score (by E) improved; myocardial viability (by PET) increased.

221

Studies with 5 patients enrolled. Abbreviations. E: echocardiography; ICM: ischemic cardiomyopathy; LVED: left ventricular end-diastolic diameter; MRI: magnetic resonance imaging; NYHA class: New York Heart Association functional class; PET: positron emission tomography. Note. (Range) denotes the minimum and the maximum amount of cells given in a trial. Dose-escalation or variation was used in studies where range is provided. For methodological and other details, please refer to original publications.

824 THERAPEUTIC APPLICATIONS: CELL THERAPY

SKMBs may be associated with a transient electrical instability in the first weeks after transplantation. How does this occur in the absence of cell coupling? One possibility is suggested by the preliminary modeling data from our group, which shows that the absence of connectivity among SKMBs within the scar provides a tortuous path (through the scar) that can form a nidus for re-entry (Tranquillo and Taylor, unpublished data). Likewise, our data suggest that electrical connection of SKMBs, which have a 10-fold shorter action potential duration than cardiocytes, to the surrounding normal myocardium could similarly provide for re-entry. Which of these hypothesis will ultimately be supported by the majority of clinical data may not be a clear-cut answer, as the location of transplantation (central or peripheral), the homogeneity of the scar itself, the functional properties of the border zone, and the number of cells engrafted (versus administered) are likely to be the major determinants of the outcome. Even in the animals, there is data showing increased incidence of arrhythmia in animals who receive SKMBs into the border zone versus center of the infarct (Atkins et al., 1999c), and there is preliminary data showing the exact opposite (McCue, unpublished data). Finally, and most importantly, the choice of optimal patients may also play a role in electrical outcomes post-SKMBs. The target HF patient population (MADIT-like population) is known to be highly susceptible to arrhythmia (Moss et al., 2002), and it remains to be seen whether the adverse events in clinical trials reflect a cell-associated event (e.g. cell integration, differentiation or even death), related to the electrical status of these seriously ill patients, the HF myocardial environment, or maybe even the location of injections. What appears clear is that the prevalence of arrhythmia may be significantly lowered by administration of low-dose amiodarone (Siminiak et al., 2005), the best clinical agent to normalize inhomogeneous action potential duration. Adverse effects of amiodarone in HF patients are well known to clinicals, so pairing the SKMB therapy with an anti-arrhythmic agent is unlikely to result in a global learning curve. Moreover, a congener of amiodarone – dronedarone – is now in clinical trials, and so far shows a comparable efficacy with a much friendlier safety profile, likely to be associated with removal of two iodine atoms from the molecule (Sablayrolles and Le Grand, 2006). In addition, alterations in the cell culturing process (e.g. the use of human serum for SKMB culture may also be beneficial (Chachques et al., 2004a). Administration of these cells in patients with an already implanted ICD has been tested as another strategy. However, the equivocal outcomes observed in the Myogenesis Heart Efficiency and Regeneration Trial so far (although the trial is still ongoing, last follow-up scheduled for middle of 2007) may be associated with the transplantation of SKMBs very late in the disease process, beyond the point where it may be biologically possible to provide engraftment and derive benefit from the functionality of the transplanted cells. In summary, even though SKMBs transplantation is the most well-defined technique for myocardial repair/regeneration, important questions remain about its long-term safety and efficacy – the questions that will need to be answered for any cell therapy. If cell transplantation becomes clinically a long-term solution to myocardial injury, cells must be able to provide a sustained and functioning revascularization, and mediate a positive contractile effect for years in heart without eliciting negative sequelae. The clinical outcomes data from 5-year patient follow-up will provide some answers in this regard. BMNCs Bone marrow and peripheral blood contain a number of cell populations that have recently been shown capable of differentiation into cells other than blood. They include the hematopoietic stem cells (HSCs), mesenchymal stem cells (MSCs), EPCs, and subsets of each of those, including CD34 progenitors, multipotent adult progenitor cells and CD14- cells (Saulnier et al., 2005; Verfaillie, 2005). Similar cell populations have been isolated from umbilical cord blood (Zhai et al., 2004). These cells have the potential to become endothelial cells and be the drivers of vascular repair and at the same time spare the researchers, clinicians and patients of the ethical and immunological hurdles of embryonic stem cells.

Cell Therapies for Repair and Regeneration 825

Hematopoietic stem cells: Historically, HSCs have been thought of as those that differentiate only down the erythrocyte and leukocyte lineages (Till et al., 1978). These cells are identified as CD34 and/or AC133 for human cells. In mice, these cells were shown to be negative for mature hematopoietic cell lineage markers (lin–) and sometimes positive for stem cell antigen-1 (Sca-1) and c-kit (also known as CD117). Over the past few years, it has been shown that HSCs can, under appropriate conditions, differentiate into various cell types, including cardiomyocytes (Yeh et al., 2003). Although HSCs can become cardiomyocytes under strict in vitro conditions, there have not been any reports showing differentiation into cardiomyocytes when transplanted into an infarcted myocardium (Murry et al., 2004). Perhaps because of this lack of differentiation in infarct, recent studies suggest that HSCs may not have the potential of some other cell types to improve LV function following transplantation into infarcted myocardium (Deten et al., 2005). Mesenchymal stem cells: MSCs are rare multipotent progenitor cells, also known as bone marrow stromal cells. In the past, MSCs were shown to differentiate into a number of cell types including, fat, bone, cartilage, and skeletal muscle precursors both in vitro and in an infarcted rat myocardium (Jiang et al., 2002). There is also some evidence that after the injection into the myocardium, these MSCs differentiate into cardiomyocytelike cells (Kawada et al., 2004). However, current studies suggest that this can only happen when MSCs are in contact with native cardiomyocytes and does not happen in the infarct core (Strauer, B.E. et al., 2002). Therefore, the optimal time of therapy using MSCs may be early in the course of the injury, when surviving cardiomyocytes are still present in the infarcted territory. Despite their inability to form cardiomyocytes to a significant degree, transplanted MSCs engraft at high numbers in an infarcted heart (Schuster et al., 2004), and lead to an increased neovascularization and improved regional contractility and the overall LV diastolic function (Schuster et al., 2004). In fact, a recent study from our group suggests that MSCs and SKMBs improve function after ischemia-induced cardiac injury to a similar degree (Thompson et al., 2003). Furthermore, several other studies suggest that MSCs can home to sites of injury following injection into the coronary or even peripheral vasculature (Strauer et al., 2002; Bittira et al., 2003). However, it has also been reported that intracoronary administration of MSCs can cause microinfarctions and promote damage of otherwise healthy myocardium (Vulliet et al., 2004), which has led to some caution with regards to the design and patient selection in the ongoing clinical trials. More recently, MSCs have been suggested to be immunoprivileged cells capable of allogeneic administration in vivo with very few negative consequences (Jiang et al., 2005), although a certain degree of skepticism about this fact remains in the scientific community. If proven to be true in ongoing trials, this quality could become the most tantalizing aspect in terms of applicability toward cardiac repair. In turn, lack of negative immunological effects and the presence of benefits of functional restoration of the myocardium can lead to a fast development of a commercial cell therapy product for use in many patients. Clinical studies with intravenous administration of allogeneic MSCs in AMI and HF are ongoing. Endothelial progenitor cells: EPCs are bone-marrow-derived cells that are mobilized into peripheral blood and believed to participate in neoangiogenesis (Kalka et al., 2000). Recent research has shown that the number of EPCs in vascular circulation is increased in patients following AMI (Shintani et al., 2001). EPCs are presumed to be mobilized by the ischemic damage in the heart (and other tissues) and migrate to the damaged areas to induce formation of neovasculature. In support of this, it was recently shown that when EPCs were injected either into the rats tail vein or LV cavity after an ischemic myocardial injury, a greater than 2-fold increase in the accumulation of infused EPCs was observed when compared to animals undergoing sham surgery (Aicher et al., 2003). LV dimensions, fractional shortening, and regional wall motion improved in rats that received EPCs and were not observed in the control animals injected with culture media (Kawamoto et al., 2001). Although the mechanism of these benefits has not been clearly elucidated, it is likely that improvements seen in this study were at least in part depended on improved myocardial perfusion and

826 THERAPEUTIC APPLICATIONS: CELL THERAPY

decreased inflammation. To date, EPCs have not been shown to induce or play a role in neomyogenesis within the injured myocardium, but several paracrine properties have recently been attributed to these cells (Kinnaird et al., 2004). Human EPCs are typically thought to primarily express CD133 (AC133), CD34, and VEGF-R2. The quantity of the EPCs circulating in humans decreases with age, and mirrors a rapid increase in CVD-related deaths (Hill et al., 2003; Werner et al., 2005). It has been suggested that this correlation is due to the EPCs contribution to maintaining vascular integrity (Hill et al., 2003). Recent data have shown that the number of circulating EPCs and their ability to migrate is decreased in patients at a high risk for clinical CVD, including AMI (Hill et al., 2003). The reduction in the number and/or functional capacity of EPCs may be a critical factor in the development of major cardiovascular events (Werner et al., 2005). Our group recently published data showing that a reduction in CD31CD45– vascular progenitor cells, thought to be related to EPCs, is associated with aging and disease state in the mouse ApoE–/– model of atherosclerosis (Rauscher et al., 2003). We showed that delivery of functionally viable cells could prevent the progression of atherosclerosis. Completed studies using EPCs and other cell types to treat CVD in humans are shown in Table 48.3. Umbilical cord blood cells: A relatively new source for progenitor cells is umbilical cord blood, which contains most of the bone-marrow-derived cell types. Cord blood cells are easily obtained, albeit not in large volumes, have the potential to develop into multiple lineages, do not pose a myriad of the ethical questions and are less immunogenic than their bone marrow counterparts. As a result, a larger proportion of the population could receive cells from appropriately matched donors. Further, if cord blood cells are isolated and stored at birth, these cells could provide an autologous source of stem cells to treat myocardial damage later in life. Current studies in animal models show that unfractionated cord blood cells injected directly into the infarcted myocardium improve LV ejection fraction, anteroseptal wall thickening, and dP/dt (max), while decreasing infarct size (Henning et al., 2004). In addition, intravenous injection of cord blood cells in mice following ligation-induced injury resulted in an approximately 20% higher capillary density in the border zones of the infarction – a finding not observed in untreated animals (Ma et al., 2005). Recent data have suggested that human cord bloodderived CD34 cells may be capable both of preventing injury progression in nude rats and of partially reversing systolic and diastolic dysfunction in the failing heart, if administered shortly after AMI (Leor et al., 2006). No evidence yet suggests that cord blood cells injected into the infarcted myocardium are able to produce mature cardiomyocytes. Overall, however, it appears that cord blood cells may appear to be an interesting cell of choice to be used in further studies of treatment of myocardial injury. Cardiac-Derived Stem Cells Within the past several years, cardiac-derived stem cells (CSCs) have been identified and are now considered a potential option for cardiac repair. Although the evidence for cardiac repair with these cells is limited, their potential to mature into cardiomyocytes makes them a promising candidate (Laugwitz et al., 2005). These cells have primarily been isolated from neonatal heart (Laugwitz et al., 2005) and, to a very limited extent, from adult myocardium (Anversa and Nadal-Ginard, 2002; Oh et al., 2003). The results of the pre-clinical use are intriguing and suggest that the future of cardiac repair may involve endogenous mobilization or recruitment of these cells – if they can be found in reasonable numbers in the adult myocardium, or can demonstrate adequate transdifferentiation when transplanted into an infarction milieu. CSCs can be isolated from neonatal rat hearts using LIM-homeodomain transcription factor islet-1 (Laugwitz et al., 2005). It is possible to expand these cells in vitro when coupled with a cardiac mesenchymal layer. Further, when these cells are co-cultured with neonatal cardiomyocytes, they are able to electrically integrate with myocardial cells in vitro by forming gap junctions (Laugwitz et al., 2005). CSCs isolated from adult hearts, including those from acutely infarcted, failing, and even the hearts destined to be replaced by transplantation, have been

Table 48.3 Cell Therapy Trials with BMNCs, MSCs, EPCs, and CPCs Investigator (country)

Patient’s number (type of cell)

Diagnosis

Average SKMB dose (106)

Delivery route

LVEF and other outcomes (maximum follow-up time, month/year)

2001

Hamano et al. (2001) (Japan)

5 (BMNCs)

Advanced CAD

300–2,200 (range)

Transepi during CABG

Perfusion improved in three out of five patients (by S) (1 year)

2002

Strauer et al. (2002) (Germany)

10 (BMNCs) (10 controls)

AMI, 5–9 days post

28

IC (after standard Tx)

Infarcted region decreased to 12 7% from 30 13% (by V); LV contractility and EDV improved, perfusion increased (by DE, RV, RC) (3 months)

2002

TOPCARE-AMI: Assmus et al. (2002); Britten et al. (2003) (Germany)

10 (CPCs)  9 (BMNCs) ( 11 controls)

AMI (reperfused)

CPCs: 13; BMNCs: 238

IC, 4-day post-AMI

LVEF improved to 60% from 51.6%; ESV reduced; wall motion in the infarction zone improved (all by V and DE); similar LV functional data by MRI; myocardial viability increased (by PET); CPCs and BMNCs behaved similarly. Migratory capacity predicted LV remodeling in multivariate analysis (4 months)

2003

Fuchs, S. et al. (2003) (US)

10 (BMNCs)

Refractory Angina

78

Transendo EMMguided

Perfusion improved (by SPECT); CCS angina score decreased to 2.0 0.9 from 3.1 0.3 (3 months)

2003

Tse et al. (2003) (China)

8 (BMNCs)

Advanced CAD

40 ml BMNC (0.6–8.9% CD34)

Transendo EMMguided

Regional wall motion, thickening improved, hypoperfused areas lessened (by MRI); angina reduced to 16.4 from 26.5 episodes per week (3 months)

2003

Stamm et al. (2003) (Germany)

6 (AC 133)

AMI

1.02–1.57 (purified)

Transepi with CABG

LVEF, EDV, EDD improved (by E); perfusion (area at-risk) improved in five out of six patients (3 months)

2003

Perin et al. (2003) (study conducted in Brazil)

14 (7 controls)

Severe Ischemic HF

25.5

Transendo EMM guided

Mean LVEF increased to 35.5% from 30% (by E); perfusion improved (by SPECT); NYHA class decreased to 1.1 0.4 from 2.2 0.9; CCS angina reduced to 1.3 0.6 from 2.6 0.8 class. (2 and 4 months) (Continued )

Cell Therapies for Repair and Regeneration 827

Year

Year

Investigator (country)

Patient’s number (type of cell)

Diagnosis

Average SKMB dose (106)

Delivery route

LVEF and other outcomes (maximum follow-up time, month/year)

2004

TOPCARE-AMI: Schachinger et al. (2004); Schachinger et al. (2006a) (Germany)

30 (CPCs)  29 (BMNCs)

AMI

CPCs: 13; BMNCs: 238

IC, 4.9 days after AMI

At 1 year, one patient in each cell group died due to cardiogenic shock, no other MACE or malignant arrhythmias; LV functional data similar to prior report for additional patients. MRI at 1 year showed maintenance of LVEF and reducion of infarct size, no reactive LV hypertrophy. Coronary flow normalized in infarct-related arteries (1 year)

2004

Chen et al. (2004) (China)

34 (MSCs) (35 controls)

AMI

8,000–10,000 (range)

IC, 8 4 h after AMI

LVEF and regional wall motion increased (by V); per fusion defects decreased (by PET); real-time electromechanical mapping of LV showed improvements in mechanical capabilities, electrical properties and functional indices (6 months)

2004

BOOST: Wollert et al. (2004); Meyer et al. (2006); Schaefer et al. (2006) (Germany)

30 (BMNCs) (30 controls)

STEMI

2,460

IC 4.8 1.3-days post-PCI

LVEF increased by 6.7% mostly due to improved regional wall motion in the peripheral area (by MRI) LVEDV and infarct size decreased. Diastolic function improved (by E). LV functional benefits did not persist at 1 year (1 year)

2004

Fernandez-Aviles et al. (2004) (Spain)

20 (BMNCs)

Extensive AMI (reperfused)

78

IC, 13.5 5.5-day post-AMI

LVEF improved by mean of 6%, contractile reserve increased; ESV decreased in (by MRI, DE) (6 months)

2004

Perin et al. (2004) (study conducted in Brazil)

11 (BMNCs) (9 controls)

End-stage ICM

15 injections, 0.2 ml/each (50 ml aspirated)

Transendo EMM guided

LVEF increased at 2 m, did not change at 6 and 12 m, perfusion improved (by SPECT), NYHA class decreased to a mean of 1.4 from 2.2 and CCS fell to 1.2 from 2.6 class. Exercise capacity improved (by treadmill) (6 and 12 months)

2004

Kuethe et al. (2004) (Germany)

5 (BMNCs)

AMI (reperfused, stented)

39

IC, 6.3 0.4 days post-PCI

LVEF and regional wall motion did not change (by E). Coronary flow (by IC Doppler) and contractility indices (by DE) remained similar at follow-up (3 and 12 months)

828 THERAPEUTIC APPLICATIONS: CELL THERAPY

Table 48.3 (Continued)

Silva et al. (2004) (US)

5 (BMNCs)

Pretransplantation HF

15 injections, 0.2 ml/each (50 ml aspirated)

Transendo EMM guided

Exercise capacity (by treadmill oxygen consumption) improved in four out of five patients, disqualifying them from listing for transplantation (2 and 6 months)

2005

Bartunek et al. (2005) (Belgium)

19 (AC133) (16 controls)

AMI

12.6

IC, 11.6 1.4 day post-AMI

LVEF increased to 52.1% from 45% similar to controls (by E), perfusion improved (by SPECT); seven patients in cell therapy group developed restenosis (versus four in control group), two patients in cell therapy group had de novo lesions (4 months)

2005

Dohmann et al. (2005) (Brasil)

14 (BMNCs) (7 controls)

Severe CAD  HF

25.5

Transendo EMM guided

Perfusion increased (by S), NYHA class, functional capacity, and function improved. (2 and 6 months)

2005

IACT: Strauer et al. (2005) (Germany)

18 (BMNCs)

Post-AMI (5m–8.5y)

15–22 (range) each infusion, 4–6 total

IC

LVEF increased by 15% (by V), infarct size fell by 30% (by SPECT), myocardial viability of infarcted zone increased by 15% (by PET) (3 months)

2005

Blatt et al. (2005) (Israel)

6 (BMNCs)

ICM

50 ml of aspirated BMNCs

IC, after induction of ischemia by balloon inflation for 3 min

LVEF improved from mean of 25% to 28% (by E); wall motion (by DE) increased but only in segments with baseline hibernation. NYHA class fell to mean of 2.3 from 3.5; one patient developed post-procedure hypotension and troponin increase (4 months)

2006

ASTAMI: Lunde et al. (2006) (Norway)

47 (BMNCs) (50 controls)

AMI treated with PCI

54–130 (range)

IC

No differences in LVEF (by MRI), perfusion (by SPECT), trend toward infarct size reduction (by MRI) (6 months)

2006

REPAIR-AMI: Schachinger et al. (2006b) (Germany)

101 (BMNCs) (103 controls)

STEMI (reperfused)

236

IC, 3–6 days after AMI

LVEF increased by a mean of 5.5% (by V), patients with LVEF 49% benefited most. At 1 year, BMNC-treated patients exhibited reduction in a combined primary end-point (death, AMI recurrence, revascularization) (4 and 12 months) (Continued )

Cell Therapies for Repair and Regeneration 829

2004

Year

Investigator (country)

Patient’s number (type of cell)

Diagnosis

Average SKMB dose ( 106)

Delivery route

LVEF and other outcomes (maximum follow-up time, month/year)

2006

TOPCARE-CHD: Assmus et al. (2006) (Germany)

34 (CPCs)  35 (BMNCs) (23 controls)

Prior AMI (3 m)

CPCs: 22; BMNCs: 205

IC, with cross-over to the other cell type

LVEF increased significantly in patients that crossed over to BMNCs, absolute increase  2.9% (by MRI). No changes in LVEF with CPCs. NYHA class improved in BMNC group – reduction of 2.0 0.7 from 2.2 0.6, no improvement with CPCs (3 months).

2006

Fuchs et al. (2006) (Israel)

27 (BMNCs)

Refractory Angina  Ischemia

28

TransendoEMM guided

CCS angina class improved to 2.0 0.9 from 3.2 0.5; exercise duration increased to 489 142 s from 418 136 s; ischemia lessened (by SPECT). At 1 year, five patients had revascularization procedures, functional and symptomatic improvements were maintained in other patients. (3 and 12 months)

2006

Tse et al. (2006) (China)

12 (BMNCs)

Advanced CAD

12–16

TransendoEMM guided

No significant changes in LVEF at 3 or 6 months (baseline LVEF 60%). At long-term follow-up, two patients died, one patient received CABG. (3, 6 and 44 10 months)

Studies with 5 patients enrolled.

Abbreviations. CAD: coronary artery disease; CCS: Canadian Cardiovascular Society; DE: dobutamine echocardiography; EDD: end-diastolic (LV) dimension; EDV: end-diastolic (LV) volume; EMM: electromechanical mapping of LV; ; IC: intracoronary; ICM: ischemic cardiomyopathy MACE: major adverse cardiac events; NYHA class: New York Heart Association functional class; PET: positron emission tomography; RC: right heart catheterization; RV: radionuclide ventriculography; S: scintigraphy; SPECT: single photon emission tomography; STEMI: acute ST-elevation myocardial infarction; Transendo: transendocardial; Transepi: transepicardial; Tx: therapy/treatment; V: ventriculography; Note: (range) denotes the minimum and the maximum amount of cells given in a trial. Dose-escalation/variation was used in studies where range is provided. Controls comprise historical and active randomized participants and those patients who received placebo. For methodological and other details, please refer to original publications.

830 THERAPEUTIC APPLICATIONS: CELL THERAPY

Table 48.3 (Continued)

Cell Therapies for Repair and Regeneration 831

identified by their expression of c-kit, MRD-1, and Sca-1 and by their lack of expression of hematopoietic lineage markers (Urbanek et al., 2005). These cells have shown the ability to differentiate down myocyte, smooth muscle, and endothelial cell pathways, but their ability to form mature cells of these types (or cardiomyocytes) is not yet unknown. Endogenous Sca-1 CSCs may differentiate into functional cardiomyocytes (Oh et al., 2003), but such potential within an infarct scar has not been elucidated. To date, methods for harvest, expansion, and in vitro growth of these precursors are very limited. Therefore, because of these factors, it is difficult to judge the clinical potential of these cells. Nonetheless, CSCs biology is interesting enough to make future developments be anticipated with interest and hope. For example, CSCs expanded from endomyocardial biopsies and predifferentiated in vitro could become very strong candidates for cardiac repair.

CLINICAL STUDIES Clinical trials with SKMBs (Table 48.2), BMNCs, MSCs, EPCs and CPCs (Hamano et al., 2001; Assmus et al., 2002; Strauer et al., 2002; Britten et al., 2003; Fuchs et al., 2003; Perin et al., 2003; Stamm et al., 2003; Tse et al., 2003; Chen et al., 2004; Fernandez-Aviles et al., 2004; Kuethe et al., 2004; Perin et al., 2004; Schachinger et al., 2004; Silva et al., 2004; Wollert et al., 2004; Strauer et al., 2005; Bartunek et al., 2005; Blatt et al., 2005; Dohmann et al., 2005; Assmus et al., 2006; Fuchs et al., 2006; Lunde et al., 2006; Meyer et al., 2006; Schachinger et al., 2006a, b; Schaefer et al., 2006; Tse et al., 2006) (Table 48.3) published to date have been summarized. Ongoing trials are listed on the Internet (http://www.clinicaltrials.gov; http://www.thescientist.com). To highlight important points, we chose to comment on several published trials. SKMBs: The first clinical trial using cell therapy to treat CVD was initiated by Menasche (2003) in 2000. In this trial, an average of 871  106 cells (at least 85% were identified as SKMBs by a positive staining for CD56) was injected into non-revascularizable scarred portion of LV as an adjunct to CABG. Over several years following transplantation, significant improvements in LV ejection fraction (EF) and regional wall thickening were observed. Unfortunately, there was no control group. Nonetheless, the data are encouraging. However, 4 out of 10 patients experienced ventricular tachycardia requiring ICD implantation. Fortunately, none of the patients experienced intractable/fatal ventricular tachycardia/fibrillation. The data suggests that concurrent administration of amiodarone can minimize untoward electrical events without compromising efficacy of the cells. The data of Menasche et al. (2003) provided the impetus to begin a new trial. The Myoblast Autologous Graft in Ischemic Cardiomyopathy trial is a Phase II randomized clinical trial to examine the efficacy and safety of CABG  SKMBs versus CABG alone in approximately 300 patients in North America and Europe. The trial was halted in 2006 in part because the design of the trial was no longer considered state of the art (as the number of CABG cases declined), and as a result recruitment was below projected targets. Although no increase in mortality was reported, the published results of this large study are greatly anticipated. In a separate US trial (Dib et al., 2005), SKMBs were injected concurrently with CABG (n  12) or LV assist device as a bridge to transplantation (n  6), myocardial perfusion improved and left ventricular ejection fraction (LVEF) increased. Upon examination of the explanted hearts (for indicated cardiac transplantation), four of the five specimens showed areas of engrafted myoblasts within the infarcted regions. In another clinical trial (Smits et al., 2003), 196/–105  106 SKMBs were injected directly into the infarcted area (via a NOGA-guided catheter system) as sole therapy in HF patients. These patients showed improved regional wall motion and a trend toward increased LV EF over 3–6 months. Taken together, these data suggest that SKMBs can be delivered in HF patients and survive within the infarcted myocardium to achieve improved LV function. Early reports of electrical instability in patients after receipt of autologous SKMBs have led to doubts and overt clinical skepticism about the safety of these cells as a treatment option. However, conflicting data exists, and therefore, several considerations should be made. First, patients who received SKMBs in the earliest clinical

832 THERAPEUTIC APPLICATIONS: CELL THERAPY

studies had advanced HF, where electrical events are an inherent part of the pathophysiology of the disease (and therefore these events are expected). In fact, many of the patients met the MADIT-II criteria (Moss et al., 2002) which were presented after those cell therapy trials began and suggested that all patients who meet those criteria be treated with ICDs. As a result, in more recent clinical studies where myoblasts are being used to treat HF, many investigators have only enrolled patients who had already received ICDs or had ongoing treatment with lowdose anti-arrhythmics. This practice may have significantly reduced the incidence of arhythmogenic events. For example, in the Phase II MAGIC trial, the incidence of electrical instability in patients post-SKMB delivery was approximately 10% (lower than the initial 40% reported by the same group of investigators) (Menasche et al., 2003). Whether this discrepancy occurred because of a better selection of patients in the second study, the coadministration of anti-arrhythmic agents or an improved safety profile of the cells remains to be determined. Furthermore, in clinical studies in the United States, Dib et al. (2005) have not reported an increased incidence of electrical instability after SKMB administration, nor have others in pre-clinical studies (Chachques et al., 1987). Nonetheless, these data suggest that autologous SKMBs for patients with HF have a potential to be a relatively safe and efficacious product, if such holds true in definitive Phase III trials. Bone marrow stem cells: In a trial similar to that performed with SKMBs, patients received up to 1.6  106 AC133 BMNCs into the peri-infarct zone concurrent with CABG (Stamm et al., 2003). However, in contrast to SKMB studies, this study examined patients treated shortly after AMI. A total of six patients were treated, and perfusion in treated areas increased and LV dimensions and EF improved. Further, unlike in the SKMB trials, these improvements occurred without electrical abnormalities. Whether this represents a difference in patient population, cell type or even cell dose remains unresolved. In a more preventive approach, a number of studies have been performed in an attempt to rescue the myocardium and to prevent HF. These studies have primarily focused on percutaneous delivery of bone marrow cells after AMI. In the TOPCARE-AMI studies (Assmus et al., 2002; Britten et al., 2003; Schachinger et al., 2004; Schachinger et al., 2006a), investigators injected 13/–12  106 circulating progenitor cells (CPCs) or 238/–79  106 BMCs into the infarct artery of patients 4.9/–1.5 days (minimum of 4 days) after AMI. At 4 months, LV end-diastolic volume and EF improved in both cell dose groups compared to control patients who underwent standard treatment during the same time but were not randomized into the study. No significant differences between CPC and BM groups were observed. By 1 year, EF remained significantly improved, infarct size was decreased, and no LV remodeling was observed. These data, when combined with the reports by others (Table 48.3) suggest a very favorable response to BMNC therapy following AMI, with improved myocardial performance secondary to improved cardiac perfusion. These encouraging data also provided the impetus for initiation of randomized controlled trials using BMNCs for the treatment of STEMI – REPAIR-AMI, which has brought extremely positive results (Schachinger et al., 2006b). Although the data for the treatment of AMI with bone marrow cells are encouraging, what remains unclear is the response of the myocardium to these cells, when HF pathophysiology predominates. To begin to address this, the TOPCARE-HF study has been initiated. Given the reduced number and migratory capacity of EPCs shown in preclinical studies and the deficits in EPC quantity seen in patients with advanced CVD (Werner et al., 2005), it will be interesting to see if cells from these patients are capable of at improving cardiac function or the HF milieu only allows ischemia-resistant cells, such as SKMBs, to survive. In Germany, in a randomized trial entitled BOne marrOw transfer to enhance ST-elevation infarct regeneration (BOOST trial) (Wollert et al., 2004) compared 30 patients under standard care following AMI (percutaneous coronary intervention (PCI) with stent placement) and 30 patients receiving 24.6108 9.4108 bone marrow cells 4.8 1.8 days after PCI. Six months after therapy, patients receiving cell therapy showed significantly enhanced LVEF when compared to control patients. At 18 months, the speed of LVEF recovery was significant in patients that received cells and PCI. There were no arrhythmic events or increased restenosis in

Cell Therapies for Repair and Regeneration 833

the cell-treated patients. However, in Belgium, a recent clinical controlled trial evaluated the ability of autologous bone marrow cells (mean of 12.6106 AC133 cells) to improve LV function after AMI (Bartunek et al., 2005). In this study, myocardial perfusion was improved, but no improvement in LV function was seen when compared to controls. Most importantly, seven patients treated with cell therapy developed restenosis (versus four in control group), and two had de novo lesions. Recently, three trials with autologous bone marrow cell administration were published. One of them (Lunde et al., 2006) did not demonstrate a profoundly significant effect of BMNCs on LV function (p  0.054) when injected at a median of 6 days (range: 4–7 days) post-AMI. However, on detailed examination of the data, it is clear that the infarct size measured with MRI was reduced compared to the control group at both 2–3 weeks and at 6 months post-therapy, exhibiting a statistically significant trend (p  0.07), if we take a small sample size into consideration. One important aspect in this trial was that even though the groups were carefully matched at randomization, the patients that received BMNCs were prescribed more diuretics (40% in the cell therapy group versus 26% in the control group), which might have negatively impacted the engraftment of the cells. In the second of the recent three published trials, Dimmeler and Zeiher’s group (Schachinger et al., 2006b) achieved a larger sample size and showed a significantly more positive effect in settings and in patient population similar to Lunde et al. In the second trial (Schachinger et al., 2006b), intracoronary infusion of BMNCs was associated with reduction in death, recurrence of myocardial infarction and revascularizations. This trial represents a milestone achievement for the field of cell therapy. There is no doubt that Phase III trial will take place soon. However, the discrepant results may represent that BMNCs need to be aspirated, expanded, prepared, and infused in strict adherence to small technical details. In the third trial (Assmus et al., 2006) (from the same investigators), the type of cells were similar to the TOPCARE-AMI, but the design was crossover, and the delivery target was somewhat different. Both CPCs and BMNCs were infused into the most dyskinetic area of the healed (at least 3 months since index event) infarcted zone in the LV. Infusion of CPCs was much less successful than administration of BMNCs, and crossover to BMNCs was associated with significantly better outcome in terms of LV EF compared with the crossover to CPCs. Taken together, these results show a great deal of evidence toward efficacy and preliminary but very encouraging evidence for BMNC therapy in settings of AMI. Although these data do not strictly address the use of these cells to treat HF, they illustrate what could the future be – early intervention to prevent the progression to endstage HF in addition to optimized pharmacological treatment. Cell therapy and administration of G-CSF: As the field of cell therapy matures, it is important to step back and evaluate the steps that led to progress. Borrowing from the established practices of bone marrow transplantation, several studies in CVD utilized granulocyte colony-stimulating factor (G-CSF) to stimulate production of bone marrow progenitors, then collected peripheral stem cells and infused them intracoronary. Even though G-CSF did not show major effects on LV function, in recent reports from Kang et al. (2004) and Hill et al. (2005), G-CSF administered in patients with chronic angina caused two AMIs and one death. The rates of restenosis following G-CSF administration increased, which can be explained by the augmented circulating cytokine milieu. Stimulating the bone marrow to produce progenitor cells (not necessarily exclusively with G-CSF) may be a part of therapeutic armamentarium in the future. Therefore, before we begin employing cytokine to stimulate bone marrow clinically in CVD patients, we need to better understand the entire cytokine cascade during AMI and during exacerbations of HF and cytokine signaling, as these are the likely therapeutic targets where cell therapy is going to be clinically applied. Such investigations may also shine light on the mechanisms behind the effects observed in clinical trials. These efforts are now underway at our institution. Treating the entire vascular tree with cell therapy: Peripheral arterial disease (PAD) is becoming a specialized focus of cell-based repair. Treatment options in PAD depend on whether the underlying pathology is intermittent claudication or critical limb ischemia. But the current options are limited to exercise, anti-aggregants,

834 THERAPEUTIC APPLICATIONS: CELL THERAPY

thrombolytics, angioplasty, surgical revascularization, and when all fails – limb amputation. Experimental data suggests that the number and/or function of circulating EPCs may reflect progression or stabilization of atherosclerotic lesions and is currently being evaluated as a biomarker for PAD. In earlier stages of the disease, the therapeutic focus clearly lies on repair via decreased claudication and improved vascular endothelium at large. As the disease progresses, restoration of perfusion to minimize tissue damage and achieve symptomatic relief becomes of primary importance, whereas regeneration of functional muscle is secondary. Different cell types and delivery methods are currently evaluated. Preliminary studies applying direct intramuscular injection of BMNCs and MSCs show promising results in increasing microvascular density and tissue perfusion and lead to the move to clinical studies. The Therapeutic Angiogenesis by Cell Transplantation Study investigators (Tateishi-Yuyama et al., 2002) performed a randomized controlled trial in PAD patients and reported a significant increase in transcutaneous oxygen pressure, and pain-free walking time in 22 patients with leg ischemia after intramuscular injection of BMNCs. In a concomitant study, BMNC transplantation was improved endothelial dysfunction by increasing endothelium-dependent vasodilation in patients with limb ischemia (Higashi et al., 2004). Currently, a trial evaluating the therapeutic potential of CD34 cells is underway in patients with intermittent claudication, and results are much anticipated.

WHAT IS REQUIRED FOR SUCCESSFUL REPAIR IN 2007 AND BEYOND? Today, cardiovascular repair seems to be a reachable goal. With the progress made to date, the field appears very promising. Nonetheless, several obstacles remain before we can declare unfettered success. What have we learned from over 10 years of preclinical and 6 years of clinical research in this area? Moving cell therapy from bench to bedside is complex. As new cell types emerge and old ones find new applications, it is important to design a preclinical path that predicts clinical outcome. It will also be important as the field moves forward to compare cells in head-to-head studies. Beginning to dissect the mechanism by which transplanted cells mediate repair is crucial. And finally if we are to ultimately regenerate heart with cell therapy, we must continue to think outside the box and view cells as only one tool in our armamentarium further moving into the 21st century and successfully integrate cell types and delivery routes with new pharmacotherapies. Cells plus genes, small molecules that replace the need for cells, and personalized genomics-based cell therapy are all medicines of the future. All of them may seem too novel and unreachable today, but so did cell therapy in 1998 (just 9 years ago). Considering new options, maximizing our mechanistic understanding of cell effects and standardizing our approaches to cell delivery, conducting clinical trials, and measuring outcomes should provide us the tools to succeed where endogenous repair fails. Below, we have outlined four specific “requirements for success.” Requirement No. 1: Selecting the Appropriate Cell Type for the Appropriate Disease Environment At present, discrepant clinical trials outcomes exist for different types of cells and numbers of cells administered. For example, SKMBs seem to engraft into the myocardium and result in functional regeneration in HF, and BMNCs show a great promise in treating acutely injured myocardium. What is clear from the basic science point of view, is that different environments in the myocardium at the time of injury likely generate different milieus, and therefore the cells that engraft in one environment may not survive in another one. Whether the discrepant clinical results are a result of a rush to clinical trials and applying various cells types in various contexts, or if segregation of the types of cells (at least between SKMBs and BMNCs) for the appropriate types of injury has already happened inadvertently remains to be understood. Because developing a successful therapy, which is based on the biology of the human body and the pathogenesis of disease, requires multiple reiterations between bench and beside, we need to go back to bench research now to compare various cell types

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side by side in various types of ischemic injury in appropriate animal models. This may seem to be a simple process, but in reality all available cells, delivery routes, and injury models are taken together would result in approximately 2,400 comparisons to be done. This clearly is a prohibitive number for a promising therapy. Therefore, we need to come to a consensus on the patients to be treated (i.e. concentrate on the most common types of injury, such as reperfused AMI at up to 4 h from the onset of symptoms, ischemic HF with a mild-tomoderate ischemic process), and conduct comparisons in those models. As the data becomes available, we can then build hypotheses as to what may or may not work in other types of pathology and models. Such experiments will also bring additional insights into our understanding of how and when repair happens. This suggestion may sound contradictory to reality, considering the number of preclinical studies and clinical studies that have been published in the field (i.e. approximately 1,300 MEDLINE hits on keyword search for “heart” and “cell transplantation”). However, only a few head-to-head comparisons of different cell populations have been performed. We clearly lack direct comparisons of different cell types in clearly defined clinically relevant models of disease. In addition, and perhaps most importantly, there is an urgent need for a task force to define the nomenclature of progenitor cells to arrive to a consensus of which cells we are going to call “progenitor cells.” Similar taxonomy efforts have been recently accomplished by Krumholz et al. (2006) for clarification of disease management. A writing group comprising experts in cell biology, taxonomy, cell differentiation, and translational research could very rapidly accomplish this task. Efforts in this direction will advance the field … and may help avoid unfortunate outcomes. Comparing different cell types in various contexts of disease will also help us definite how to improve survival of transplanted cells. Currently, one of the largest hurdles of cell therapy is the limited survival of transplanted cells. Most reports suggest that 70–90% of all transplanted cells die within the first few days of transplantation into infarct scar. Studies have shown that a subset of the transplanted cells survive and multiply, but it is unlikely that this multiplication can make up for the massive early necrosis and apoptosis of cells. Preclinical data suggest a dose response for several cell types, indicating that improving the number of surviving cells is critical to maximize functional outcome. Learning from previous fields is important. A confounding inflammatory response secondary to needle punctures during cell delivery is reminiscent of early percutaneous or transmyocardial myocardial revascularization studies where the “injection” per se promoted inflammation. Although the inflammatory response to needle stick has been reported as mild in most cases, the possibility that needle-based cell delivery is pro-inflammatory should be explored further; similarly if it is we need to define specific cytokines that might be involved. The problem faced is determining whether inflammation is an initiator of the necrosis of transplanted cells or a response secondary to the apoptosis of the transplanted cells. The most likely hypothesis is that the ischemic environment is the driver of these processes. This hypothesis is strengthened by data showing that survival of neonatal cardiomyocytes more than doubled when injected into 2-weekold cardiac granulation tissue or normal myocardium versus myocardial scar tissue in rats (Zhang et al., 2001). Further, preconditioning of cells before transplantation via heat shock or transfecting cells with prosurvival factors (Akt, heat shock proteins, specific growth factors, or certain signaling molecules to provide protection of cells from hypoxia or glucose deprivation) helps increase their survival rate in vivo (Kohin et al., 2001; Zhang et al., 2001). In addition, we have preliminary data indicating that the composition of nutrients in which cells are grown in vitro alters survival in an infarct-like milieu (Davis, unpublished data). Ideally, more work will be focused on this area in the future to better define the relationship between the microenvironment of the infarct scar and outcome of the transplanted cells. In addition to surviving in the ischemic environment at the time of implantation, the ideal cell for myocardial repair will be able to despite ischemia, become a fully functioning cardiomyocyte or an endothelial cell. However, none of the progenitor cells currently used satisfies both of these criteria at the numbers sufficient for

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maximal repair or recovery of function. Therefore, it is important to keep working toward understanding the differentiation of progenitor cells into a cardiomyocyte phenotype at the bench level. The goal, then, might be to design molecular tools to induce a pathway of directed differentiation prior to implantation, so that achievement of a specific phenotype would occur slowly enough to allow neovascularization to become functional (and not collapse) to support nascent myocardium. Lastly, injected cells have significantly different electrical properties than cardiomyocytes. These differences have led so ventricular tachycardia observed in some of the clinical trials – especially where SKMBs were delivered in patients with HF. For cardiovascular cell therapy to reach its potential, it will be critical to electrically integrate transplanted cells into the surviving myocardium. This problem may be approached by genetically altering transplanted cells (to promote electrical coupling), by developing new adjunctive safety measures (such as coadministration of anti-arrhythmics), delivery of cells only in patients who meet the MADIT-II criteria and have ICDs, or preferably, by conditioning the transplanted cells to become true cardiomyocytes that can survive in a regenerating milieu. Only by accomplishing these goals we can better design methods to maximize cell survival and thus to increase the benefits versus the risks of cell therapy and achieve quantum leaps of progress in treating and preventing CVD. Requirement No. 2: Choosing the Best Delivery Route It is clear that choosing the best delivery route is the second prerequisite for success, after choosing the right cell for the right environment. A major experimental obstacle to the clinical efficacy is the poor engraftment seen when cells are administered by intracoronary, intravenous, and intracardiac routes. This is likely due to multiple factors, out of which technical difficulties of injecting exactly into the center or the periphery of the scar or catheter manipulations in the coronary tree cannot be overemphasized. We have recently published data showing that a direct placement of cells with the Da Vinci robotic system results in very accurately directed cell transplantation and does enhance the outcomes of the procedure in terms of improvement of LV function (Ott et al., 2006). We also have preliminary data highlighting the importance of a very careful, targeted needle injection into the center of the scar versus periphery. Not only does LV remodeling differ with location but the arrhythmogenic potential may highly depend on the accurate placement of the transplanted cells (McCue, unpublished data). It is an established dogma that in the real estate business, location, location, location is the most well-known determinant of the success of a transaction. The same may hold for cell therapy. If so, training of the operators gains a pivotal importance. Recently, concerns of myocardial perforation due to operator error halted the GENASIS trial. As we go forward, creating a specialized network of centers for cell therapy, as currently proposed by the NHLBI, could allow for training of interventional cardiologists by experts in delivery techniques. Alternatively, it may also make sense to restrict the number of centers per region that act as referral centers and deliver cell therapy, at least until the techniques come to solid maturity. We have learned that operator volume and experience was a critical determinant of success in CABG and PCI clinical trials and also in routine clinical practice. As the field of cell therapy goes forward, we cannot ignore importance of appropriately trained specialists. However, by the same token, we cannot ignore the need for further studies. The data has suggested that intracoronary delivery, at least in the context of AMI, can provide a comparable level of engraftment of cells to surgical delivery. This again points out the need for head-to-head comparisons of various delivery methods in controlled, designed experiments. We may need to consider again the roles that inflammatory factors play, and how the process of endogenous bone marrow mobilization impacts exogenous administration in AMI or in the beginning of HF, and what the covariates of that process might be. Understanding of the biology of these could allow coupling with an optimal situation-specific delivery system to produce several distinctly different products for the field. Achieving such an understanding and creating such a system will take some time.

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However, the technological progress in the 20th century has been so fast that it will not be surprising if the next 5 years brings major progress in this regard. Requirement No. 3: Rigorous Trial Design and Selection of End-Points Right now clinical trials in cell therapy suffer from several major shortcomings primarily involving design and selection of end-points. Several examples of current limitations can be illustrated. To date, most studies have been accompanied by an additional revascularization procedure, either by PCI or CABG, making any functional improvement due to cellular therapy nearly impossible to distinguish from the current standard of care. The need to establish appropriate controls in a novel area where standard of care is evolving is an active area of debate in the field. We also need to account for the stimulatory effects of drugs, such as statins, PPAR agonists, erythropoietin, estrogen, and possibly others in various disease states. Right now, it is completely unclear which if any of these combinations of drugs alter the number and the function of progenitor cells available for repair. Clearly, if cell therapy is to be adequately evaluated we will need this information for the design of definitive Phase III trials and also going forward with clinical applications. Further, we lack data that evaluate time as an additional factor in treatment, time in disease progression as well as time in dynamics of transplanted cells. Overall, there is a lack of standardization in the current preclinical approach to cell therapy; for example, cell types, doses, pre-clinical models, and end-points all differ. This may also explain the discrepancy between preclinical and clinical results. Attempts to standardize these parameters and to decide on a consensus will move us forward. What we call an “end-point” in cell therapy matters a great deal. So far, clinical trials have been geared toward measuring functional improvement of the LV by assessing global EF. As we know from the HF trials, improvement of regional contractility may not always translate into better HF numbers because of the differences in loading conditions. In addition, recently, we have begun to appreciate observer dependence of such measurements. Even though cardiac MRI offers the best topographic assessment of the heart, the variability is best minimized by conducting clinical trials with centralized core laboratories where the personnel undergoes regular inter- and intra-observer reproducibility assessments. More attention needs to be paid to regional contractility, peri-infarct zone, and scar volume quantification – all best done in an environment of a core laboratory. In addition, we need to evaluate myocardial perfusion. Over the last 10 years, the field of cardiac MRI has matured to offer quantitative assessment of myocardial perfusion (Jerosch-Herold et al., 2004). Several sensitivity and specificity studies showed that assessment of myocardial flow with MRI may offer an edge of superiority over other techniques. Measuring changes in blood flow was proposed to be used as an end-point (Wilke et al., 2001) and it is now becoming apparent that cell therapy will need a measure of blood flow as well. Concurrently, we need to critically evaluate the end-points that are used at the present time and come to an agreement, most likely through an AHA/ACC-sponsored consensus document, similarly to available data standards for AMI, HF, and atrial fibrillation that would outline the standard sets of data to be captured in the cell therapy trials. As that process goes along, some end-points with high subject variability, such as exercise treadmill time, will be critically evaluated and new, biologically relevant and clinically translatable end-points will be introduced. Such process will also enormously aid acceptable of new end-points by Food and Drug Administration (FDA) and will over time accelerate bringing cell therapies to market. Requirement No. 4: Establish the Registry for the Results of Trials and the Biorepository for Blood Samples to Fill the Void of Mechanistic Understanding of Cell Therapy Decades of CVD research have taught us the importance of centralized databases in advancing of our understanding of the disease process. The field of cardiovascular medicine would not have advanced as far as it did in the last 25–30 years if the Framingham Study or the TIMI trials had not been initiated and executed in a

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Figure 48.3 Schematic representation of the proposed model for central registry / biorepository for clinical trials in cell therapy. Abbreviations. FACS: fluorescent-activated cell sorting.

centralized matter. Large databases give the power to control for necessary covariates – a step not possible do accomplish even in a study of a few hundreds of patients. The field of cell therapy has arrived to the point when the next great advancement might be employing a large database of all results of clinical trials to serve as a filter for the hypotheses. With the aid of such a tool, ideas will be segregated before hundreds of thousands of dollars are spent only to find out that a specific factor interfered with the outcome. Creation of such database (Figure 48.3) will pay off a 100-fold over time, as the field of cell therapy is coming out of infancy and maturing into adulthood. We should not underestimate the powers of computer technology and of the Internet available to us to create such a tool. Centralized data collecting efforts in acute HF, such as ADHERE registry (Yancy and Fonarow, 2004), have brought extremely valuable data with regards to the outcomes of clinical management of HF patients. It is time to create a registry for all applications of cell therapy in CVD. Along with the outcomes, we must conduct population studies to define the role of bone marrow progenitors in vascular repair. As we are learning the importance of gender and race in the pathogenesis of CVD, we also need to understand the differences in repair across wide age, gender and race groups, and also understand how the process of repair differs in those groups when different degrees of risk factors are superimposed. Conducting studies of this magnitude will help lessen the chances of not capturing significance when it truly exists – a frequent problem of small samples. As clinical studies go forward and the field matures further, we will need to evaluate other cell types involved (not just EPCs) and measure various evolving markers to supplement the knowledge in the field of vascular repair. Centralized availability of samples will help reduce cost of repeated clinical trials by several orders of magnitude – something we all care about, especially at the times of high national deficit and budget cuts to the NIH. In this regard, a centralized repository of BMCs should be the next national priority (Figure 48.3). Most importantly, measuring various progenitor cell populations,

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their function, carrying out gene profiling experiments, and performing experiments that require specific cell culture conditions needs to be accomplished with strict adherence to standardized protocols to achieve a desired outcome, as we are learning that even smallest technical details matter in cell therapy. Therefore, centralizing sample collection, storage, flow cytometry, and assays makes a great deal of common sense and will help greatly advance the science. Combining the registry for the clinical trials data and the biorepository for the blood and tissue samples seems to be exactly what the field needs to make another decade of major progress and help shape future cell therapy products. The short-term goal of a Repository would be to compare various subsets of circulating bone marrow progenitors in patient populations and to evaluate the impact of age, gender, and race. The long-term goal would be to use the progenitor cell characterization in conjuction with clinical data and examine the dynamics within multiple populations of progenitors in different states of disease when different types of cells are given. Provided that CD34 cells, for example, only represent approximately 1% of all circulating cells, a centralized Repository will provide in many ways a “win–win” situation by minimizing the effort required and maximize the benefits yielded. In summary, we believe that the outlined four requirements represent major issues in the cardiovascular cell therapy field today. As the field develops further and products moves closer to market, resolution of each requirement will increase the likelihood of successful outcome. The ultimate success, however, will be achievement of prevention of atherosclerosis and CVD altogether, reduction of hospitalization and major adverse cardiac events, and in prolonging a healthier life for patients who currently have limited options available to them.

SUMMARY This is an exciting time in the field of treating CVD. Cell transplantation opened a new frontier, providing physicians with techniques and treatment alternatives for a large patient population that extends beyond revascularization and metabolic control to reverse damage that, in many cases, has already been done and may not truly be controllable. The concept of repairing or regenerating ischemic cardiac tissue is a truly fantastic possibility, and while many question its validity, it has an excellent chance to eventually become a clinical reality, if we address every requirement for its success. While some more conservative researchers consider large human trials premature at this point, cell therapies, especially the recent trials, have shown clinical benefit. Due to the small study sizes and an inability, at this point, to standardize therapy, we are limited in our power to determine the best cell type, dose, and administration techniques, and to answer many other relevant questions. But the relevance of this therapy is evident to both researchers and clinicians. Both animal studies and clinical trials thus far have evoked the scientific enthusiasm and promising results to warrant large-scale controlled clinical trials to determine the best and safest application of this technology, and to gain a better understanding of its mechanism(s). To bring this field forward we now have to come together and outline a plan for future studies. The diversity of cell types, application techniques, and disease stages can be a hurdle and an opportunity – only collaborations will allow us to move forward as a field instead of expanding the information that cannot be combined or compared. We have the opportunity to create a new era in the treatment of CVD. Doing so will require continued bench to bedside and back to bench evaluations as we learn from early clinical studies, find a consensus on preclinical models and the design of clinical trials to maximize the potential of a 21st century approach to repairing the injured heart. Finally, even as the field progresses, we have a responsibility to promise patients (and the press) only what we can deliver, that is to tell the truth about cardiac repair. BMNCs, MSCs, SKMBs, or other types of cells hold a great promise to modify pathophysiological process in specific ways. It is crucial to understand for clinicians,

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patients, and the press that specificity precludes panacea. As we go forward, some applications will succeed, and some will fail. Cells may not be found guilty of failures. On the contrary, the disease contexts may come to be the primary determinants of efficacy. We have already experienced a similar process with angiogenic growth factors in CVD, and we now know that those trials should have more carefully targeted the disease process, as the results uniformly showed that sicker patients had larger therapeutic benefits. As investigators, we need to be realistic of the expectations we place on cell therapy, and ultimately we need to underpromise and overdeliver, based on rigorous science … otherwise, the great potential will eventually be destroyed. Cell therapy is, however, a new and very promising alternative that warrants much further exploration, inspiration, and investment of our time and resources.

ACKNOWLEDGMENTS This work has been supported in part by NHLBI/NIH award to Dr. Taylor (R01-HL-063346), Minnesota Partnership for Biotechnology and Medical Genomics award, and by funding from the Center for Cardiovascular Repair, University of Minnesota. Authors sincerely thank Harald C. Ott, MD, for his continuous contributions to the ongoing success of the Center for Cardiovascular Repair.

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Tse, H.F., Thambar, S., Kwong, Y.L., Rowlings, P., Bellamy, G., McCrohon, J., Bastian, B., Chan, J.K., Lo, G., Ho, C. L., and Lau, C.P. (2006). Safety of catheter-based intramyocardial autologous bone marrow cells implantation for therapeutic angiogenesis. Am. J. Cardiol. 98: 60–62. Tsikouris, J.P., Suarez, J.A., Simoni, J.S., Ziska, M. and Meyrrose, G.E. (2004). Exploring the effects of ACE inhibitor tissue penetration on vascular inflammation following acute myocardial infarction. Coron. Artery Dis. 15: 211–217. Tsimikas, S., Willerson, J.T. and Ridker, P.M. (2006). C-reactive protein and other emerging blood biomarkers to optimize risk stratification of vulnerable patients. J. Am. Coll. Cardiol. 47: C19–C31. Urbanek, K., Torella, D., Sheikh, F., De Angelis, A., Nurzynska, D., Silvestri, F., Beltrami, C.A., Bussani, R., Beltrami, A.P., Quaini, F., Bolli, R., Leri, A., Kajstura, J. and Anversa, P. (2005). Myocardial regeneration by activation of multipotent cardiac stem cells in ischemic heart failure. Proc. Natl Acad. Sci. USA. 102: 8692–8697. Valgimigli, M., Rigolin, G.M., Fucili, A., Porta, M.D., Soukhomovskaia, O., Malagutti, P., Bugli, A.M., Bragotti, L.Z., Francolini, G., Mauro, E., Castoldi, G. and Ferrari, R. (2004). CD34 and endothelial progenitor cells in patients with various degrees of congestive heart failure. Circulation 110: 1209–1212. van den Bos, E.J. and Taylor, D.A. (2003). Cardiac transplantation of skeletal myoblasts for heart failure. Minerva Cardioangiol. 51: 227–243. van den Bos, E.J., Davis, B.H. and Taylor, D.A. (2004). Transplantation of skeletal myoblasts for cardiac repair. J. Heart Lung Transplant. 23: 1217–1227. van der Wal, A.C., Becker, A.E., van der Loos, C.M., Tigges, A.J. and Das, P.K. (1994). Fibrous and lipid-rich atherosclerotic plaques are part of interchangeable morphologies related to inflammation: a concept. Coron. Artery Dis. 5: 463–469. van Zonneveld, A.R.T. (2006). Endothelial progenitor cells: biology and therapeutic potential in hypertension. Curr. Opin. Nephrol. Hypertens. 15: 167–172. Verfaillie, C.M. (2005). Multipotent adult progenitor cells: an update. Novartis Found. Symp. 265: 55–61. Vulliet, P.R., Greeley, M., Halloran, S.M., MacDonald, K.A. and Kittleson, M.D. (2004). Intra-coronary arterial injection of mesenchymal stromal cells and microinfarction in dogs. Lancet 363: 783–784. Weber, C., Schober, A. and Zernecke, A. (2004). Chemokines. Key regulators of mononuclear cell recruitment in atherosclerotic vascular disease. Arterioscler. Thromb. Vasc. Biol. 24: 1891–1896. Werner, N., Kosiol, S., Schiegl, T., Ahlers, P., Walenta, K., Link, A., Bohm, M. and Nickenig, G. (2005). Circulating endothelial progenitor cells and cardiovascular outcomes. N. Engl. J. Med. 353: 999–1007. Wilke, N., Zenovich, A., Jerosch-Herold, M. and Henry, T. (2001). Cardiac magnetic resonance imaging for the assessment of myocardial angiogenesis. Curr. Interv. Cardiol. Rep. 3: 205–212. Williams, P.C., Coffey, M.J., Coles, B., Sanchez, S., Morrow, J.D., Cockcroft, J.R., Lewis, M.J. and O’Donnell, V.B. (2005). In vivo aspirin supplementation inhibits nitric oxide consumption by human platelets. Blood 106: 2737–2743. Wollert, K.C., Meyer, G.P., Lotz, J., Ringes-Lichtenberg, S., Lippolt, P., Breidenbach, C., Fichtner, S., Korte, T., Hornig, B., Messinger, D., Arseniev, L., Hertenstein, B., Ganser, A. and Drexler, H. (2004). Intracoronary autologous bone-marrow cell transfer after myocardial infarction: the BOOST randomised controlled clinical trial. Lancet 364: 141–148. Wright, L., Maloney, W., Yu, X., Kindle, L., Collin-Osdoby, P. and Osdoby P. (2005). Stromal cell-derived factor-1 binding to its chemokine receptor CXCR4 on precursor cells promotes the chemotactic recruitment, development and survival of human osteoclasts. Bone. 36: 840–853. Wutzl, A., Brozek, W., Lernbass, I., Rauner, M., Hofbauer, G., Schopper, C., Watzinger, F., Peterlik, M. and Pietschmann, P. (2006). Bone morphogenetic proteins 5 and 6 stimulate osteoclast generation. J. Biomed. Mater. Res. 77: 75–83. Wyatt, S.B., Winters, K.P. and Dubbert, P.M. (2006). Overweight and obesity: prevalence, consequences, and causes of a growing public health problem. Am. J. Med. Sci. 331: 166–174. Yan, L.L., Liu, K., Daviglus, M.L., Colangelo, L.A., Kiefe, C.I., Sidney, S., Matthews, K.A. and Greenland, P. (2006). Education, 15-year risk factor progression, and coronary artery calcium in young adulthood and early middle age: the coronary artery risk development in young adults study. JAMA 295: 1793–1800. Yancy, C. and Fonarow, G. (2004). Quality of care and outcomes in acute decompensated heart failure: The ADHERE Registry. Curr. Heart Fail. Rep. 1: 121–128. Yeh, E.T., Zhang, S., Wu, H.D., Korbling, M., Willerson, J.T. and Estrov, Z. (2003). Transdifferentiation of human peripheral blood CD34-enriched cell population into cardiomyocytes, endothelial cells, and smooth muscle cells in vivo. Circulation 108: 2070–2073.

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49 Retinal Pigment Epithelium Derived from Embryonic Stem Cells Irina Klimanskaya Blindness is one of the most devastating ailments in humans, and in the United States alone, over 9 million people suffer loss of vision from retinal hereditary and degenerative diseases which lead to the loss of photoreceptors. Among these, are ailments such as retinal dystrophies, macular degeneration, and retinitis pigmentosa, which are commonly associated with dysfunctions of the retinal pigment epithelium (RPE) that plays a key role in the support of the photoreceptor which does not have its own blood supply. RPE is a highly specialized tissue located between the choroids and the neural retina, and its functions include absorption of stray light that allows a better resolution of images; ion and metabolite transport between the neurosensory retina and the choroids; storage, metabolism, and delivery of vitamin A to the photoreceptor; phagocytosis of shed photoreceptor fragments and serving as a barrier between the neurosensory retina and the choroids (Besharse and Defoe, 1998; Thompson and Gal, 2003).

DEVELOPMENT OF RPE In vertebrate development RPE shares the same progenitor, neuroectoderm of the optic vesicle, with the neurosensory retina. The cells competent to give rise to RPE or neural retina are morphologically very similar and express Otx2, Pax6, Rx1, and Six3, transcription factors necessary for eye development (Carpenter et al., 2001; Reubinoff et al., 2001; Ying et al., 2003; Ben-Hur et al., 2004; Meyer et al,. 2005; Takagi et al., 2005, reviewed by Ben-Hur, 2006). Signals coming from the tissues surrounding the optic vesicle seem to be instructing different populations of these cells to selectively differentiate into RPE or neural retina. The cells of the distal part of the optic vesicle next to ectoderm which produces fibroblast growth factor (FGF)1 and FGF2 are thought to be giving rise to neural retina, as FGF signaling was shown to support neural retina formation (Pittack et al., 1997; Nguyen and Arnheiter, 2000) and convert RPE into neural retinal cells (Pittack et al., 1997; Vogel-Hopker et al., 2000; Zhao et al., 2001). The dorsal part of the optic vesicle adjacent to mesoderm receives RPE-inductive signals, such as activin A expressed by extraocular mesenchyme (Feijen et al., 1994; Fuhrmann et al., 2000; Chow and Lang, 2001). There are several transcription factors involved in the signaling cascade, essential for RPE specification. In the following model suggested by Martinez-Morales et al. (2004) the RPE induction is triggered by Pax6 and Otx1/Otx2 combined activity, when Pax6 and Wnt signaling induce expression of microftalmia-associated transcription factor (Mitf) which plays an important role in development of melanin-containing cells, melanocytes, and RPE, and can support the further formation of RPE together with Otx proteins, and according to Baumer and co-authors, expression of Mitf is controlled by the redundant activities of Pax6 and Pax2 (Baumer et al.,

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2003). The paired and homeodomain transcription factor Pax6 is known to be a key regulator of eye formation in multiple species (Gehring and Ikeo, 1999; Ashery-Padan and Gruss, 2001; Kumar, 2001) and is expressed throughout the optic cup (Walther and Gruss, 1991; Grindley et al., 1995). Its overexpression is sufficient to induce ectopic eyes formation in fly and frog embryos (Halder et al., 1995; Chow et al., 1999) and no functional eye structures develop in its absence, as shown in experiments in mouse, human, rat, frog, and fly (reviewed by Gehring and Ikeo, 1999). Its activity seems to be sufficient to induce RPE formation (Baumer et al., 2003), and in mature RPE it is downregulated, remaining in the lens, corneal, and conjunctive epithelia, iris, and amacrine and ganglion cells of the neural retina (Walther and Gruss, 1991; Hitchcock et al., 1996).

CHARACTERISTICS OF RPE Mature RPE is characterized by the polygonal morphology of the cells forming a “cobblestone” monolayer, granules of melanin in the cytoplasm, and several specific molecular markers. A water-soluble 36 kD cellular retinaldehyde-binding protein (CRALBP) is found in apical microvilli of RPE and in Muller glia (Bunt-Milam and Saari, 1983). It is a product of RLBP1 gene, and its mutations were shown to be associated with rod-cone dystrophy (Eichers et al., 2002), retinistis pigmentosa (Maw et al., 1997; Morimura et al., 1999), and impaired dark adaptation in a mouse model (Saari et al., 2001). Another important RPE hallmark is RPE65, a 65 kD cytoplasmic protein involved in retinoid metabolism (Hamel et al., 1993; Redmond et al., 1998; Ma et al., 2001). Mutations in RPE65 were shown to be associated with Leber’s congenital amaurosis and retinitis pigmentosa as well as with similar dystrophies in animal models of childhood blindness (Marlhens et al., 1997; Morimura et al., 1999; Acland et al., 2001; Van Hooser et al., 2000; reviewed by Perrault et al., 1999) and in a canine RPE65–/– model. Bestrophin, another specific RPE marker localized in the basolateral plasma membrane, is a 68 kD product of the best vitelliform macular dystrophy gene (VMD2) (Petrukhin et al., 1998; Marmorstein et al., 2000) and is associated with macular dystrophies (Marquardt et al., 1998; Allikmets et al., 1999; Kramer et al., 2000). Mer tyrosine kinase protooncogene (MERTK) is associated with retinal dystrophies and phagocytosis pathways (D’Cruz et al., 2000; Gal et al., 2000; Feng et al., 2002). One more molecular marker of RPE is pigment epithelium-derived factor (PEDF) a 48 kD secreted protein with angiostatic properties (Steele et al., 1993; Jablonski et al., 2000; Karakousis et al., 2001). Bestrophin and RPE65 are subject to translational control, and RPE65 protein was shown to be absent from cultured RPE, while the gene expression product can be detected by reverse transcriptase polymerase chain reaction (RT-PCR) (Hamel et al., 1993; Liu and Redmond, 1998; Bakall et al., 2003). TRANSDIFFERENTIATION OF RPE IN CULTURE A specific feature of the RPE is its apparent plasticity. The common origin of RPE and neurosensory retina may be the reason why under certain conditions both in vivo and in vitro RPE can transdifferentiate into neuronal progenitors (Opas and Dziak, 1994), neurons (Vinores et al., 1995; Chen et al., 2003), and lens epithelium (Eguchi, 1986). Amphibian RPE is able to regenerate the retina via dedifferentiation into neural progenitors followed by their differentiation into retinal neurons (Moshiri et al., 2004). So far there have been no indications of mammalian RPE being capable of regenerating retina, although in vitro it can undergo a similar transdifferentiation process. One of the factors which can stimulate the change of RPE into neurons is basic fibroblast growth factor (bFGF) (Opas and Dziak, 1994), and this process is associated with the expression of transcriptional activators normally required for the eye development, including rx/rax, chx10/vsx-2/a/x, ots-1, otx-2, six3/optx, six6/optx2, Mitf, and Pax6 (Fischer and Reh, 2001). These authors also showed the presence of proliferative pigmented cells in the margin of the post-natal chicken retina expressing proliferating cell nuclear antigen and incorporating bromdeoxyuridine; these cells also express Pax6 and Mitf and can lose pigmentation and turn into neuronal cells in response to simultaneously added FGF and insulin (Fischer and Reh, 2001). In vitro, depending

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on the combination of growth factors and substratum, RPE can be maintained as an epithelium, or rapidly dedifferentiate and become proliferative (Opas and Dziak, 1994; Zhao, 1997); the epithelial phenotype can then be reestablished in long-term quiescent cultures (Grierson et al., 1994).

TRANSPLANTATION OF RPE While there is no current technology allowing to restore degenerated photoreceptors which would establish connections with the optic nerve, transplantation of RPE has the potential to prevent the loss of the remaining photoreceptors before a degenerative disease would result in blindness, and it has been extensively studied for over 30 years in various animal models, including rabbits (Lopez et al., 1987; Brittis et al., 1987 El Dirini et al., 1992; Crafoord et al., 1999), monkey (Gouras et al., 1985; Berglin et al., 1997), dogs (Veske et al., 1999; Verdugo et al., 2001), and Royal College of Surgeons (RCS) rats, the latter being an excellent working model of retinal degeneration which is likely to be associated with the impaired phagocytosis function of RPE and its detachment from Bruch’s membrane (Lopez et al., 1989; Sauve et al., 2002; Wang et al., 2005a; Philips et al., 2003; Girmann et al, 2003, 2005). Several studies have been conducted in human subjects (Benson et al., 1998; Algevre et al., 1999; Binder et al., 2004). The potential sources of RPE for transplantation to treat retinal degenerative diseases and prevent the loss of the photoreceptor are autologous (Binder et al., 2004; van Meurs et al., 2004) and fetal RPE (Weisz et al., 1999; Radtke et al.,2002, 2004). Each source is not free of potential problems: autologous RPE may already have impaired function due to the patient’s own age and disease and fetal cells can vary from batch to batch and need to be characterized for safety before transplantation, which means that they need to be propagated in culture for a certain time and may lose some of their important functions. Ethical issues related to all fetal material should also be considered. Sourcing suitable donor cells is critical to any transplantation, and such cells need to be non-tumorigenic, free of pathogens and of contaminating other cell types, and every batch of donor tissue meant for clinical applications needs to be assessed for all these parameters. Donor tissue should be readily available with minimal batch-to-batch variation, and it is desirable that each batch be tested first in an animal model (Lund et al., 2001a; Wang et al., 2005b). Cultured cell lines could be an attractive cell source if they meet all of the above criteria. Lund and co-authors have shown that RPE cell lines can be very effective in supporting the photoreceptor function in RCS rats (Lund et al., 2001b). Two RPE cell lines established from adult donors, spontaneously derived ARPE-19 and SV40 large T (tumor) antigen-transformed h1RPE7, were used in the study. Both lines showed attenuation of the photoreceptor and visual function loss up to 5 months after transplantation into subretinal space. This demonstrates the potential of these or similar cell lines for treatment of retinal degenerative diseases associated with RPE dysfunction. However, their application for human therapy depends on further evaluation of their performance at extended passages, stability of karyotype, preservation of RPE phenotype, and function in long-term cultures through multiple passages. One potential hurdle for therapeutic applications of RPE, in contrast to the success of transplantation experiments in laboratory animals, is that there are age-related changes in Bruch’s membrane in humans which may lead to poor RPE survival and differentiation, as has been shown in organ culture experiments (Gullapalli et al., 2004, 2005). Bruch’s membrane is a multi-layered structure underlying RPE, which consists of the RPE basement membrane, inner and outer collagenous layers and elastin layer between them, and choricapillaris basement membrane (Zarbin, 2003). The RPE basement membrane appears to be the most favorable matrix for RPE attachment, compared to other layers of Bruch’s membrane (Tezel and Del Priore, 1999; Tezel et al., 1999). Age- and disease-related changes in Bruch’s membrane include increased thickness, exposure of underlying collagenous or even elastic layers, deposition of extracellular matrix (ECM) proteins and lipids which impair the attachment, and survival of cultured RPE because the graft may be interacting with

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less favorable substrate (Zarbin, 2003; Gullapalli et al., 2004). These authors have studied resurfacing of the Bruch’s membrane by fetal RPE, using eyes from old donors and donors with age-related macular degeneration (AMD). The explants were treated to expose different layers of the Bruch’s membrane to imitate possible age and disease-related changes, and cultured fetal RPE cells between passages 2 and 4 were plated on the explants and incubated in organ culture up to 7 days, then examined histologically by light and scanning electron microscopy. It appeared that on aged submacular Bruch’s membrane there were multiple defects in RPE surface coverage and morphology, notable cell death, and the poorest resurfacing was noted in the eye of a donor with AMD. Other experiments performed by the same group compared integrin profiles in cultured versus uncultured fetal and aged RPE and have shown that there is a downregulation of integrin subunits forming adhesion receptors for laminin, collagens, and fibronectin in uncultured aged RPE (Zarbin, 2003). These studies suggest that certain “adjustment” of the cells’ phenotype may be required prior to grafting in order to provide a better “match” between the RPE adhesion receptors and the ECM composition of the recipient’s Bruch’s membrane. Culturing RPE can allow to modulate integrin expression and thus make the cells attach to Bruch’s membrane more efficiently and improve their survival and differentiation. However, multiple passages and prolonged time in culture can amend many RPE features and lead to cell senescence; for instance, loss of α5 integrin was reported in “post-confluent” quiescent RPE cultures (Proulx et al., 2003, 2004); therefore, multiple passages in culture can produce phenotypically or functionally impaired cells. These factors may impede production of fetal or adult RPE cells of high quality and in high enough quantities for characterization, safety assessment, and transplantation. An attractive solution to the shortage of reproducibly safe and functional cells for therapeutic applications could be the use of human embryonic stem cells (hESC), which can serve as a nearly unlimited cell source for regenerative medicine. After the first work reporting derivation of stable hESC lines was published (Thomson et al., 1998), this field has been almost exponentially growing, many new hESC lines having been derived around the world (Andrews et al., 2005; Hoffman and Carpenter, 2005), and various differentiation derivatives of these lines generated and characterized in vitro (Conley et al., 2004; Gerecht-Nir and Itskovitz-Eldor, 2004; Liew et al., 2005; Sathananthan and Trounson, 2005).

GENERATION OF RPE FROM ES CELL ES cells are progeny of the inner cell mass (ICM) of a blastocyst, the same totipotent entity that gives rise to a whole new organism, and can remain pluripotent virtually indefinitely. Pluripotent hESC express a set of molecular markers, such as octamer-binding protein 4 (Oct-4), stage-specific embryonic antigens (SSEA)-3 and SSEA-4, tumor rejection antigens (TRA)-1-60, TRA-1-80, alkaline phosphatase, Nanog (Carpenter et al., 2003; Hoffman and Carpenter, 2005) and when growing in culture on a feeder layer or feeder-free in defined conditions maintain a very specific morphology (Figure 49.1a). They form flat colonies comprised of small, tightly packed cells with a high ratio of nuclei to cytoplasm, with clear boundaries between the cells and usually sharp, refractile colony borders. Similar to the cells of the ICM that differentiate into predetermined lineages, ES cells in culture easily differentiate, and the first detectable signs of such commitment are the loss of the unique ES cell morphology (Figure 49.1c and d) and downregulation of molecular markers of pluripotency. Differentiation in vitro can be spontaneous when formation of various lineages happens rather unpredictably in different cultures, or it can be directed toward a variety of derivatives by the use of various agents that can selectively activate different pathways. To date, the differentiation of human and mouse ES cells into numerous cell types has been reported (reviewed by Smith, 2001). However, such directed differentiation does not usually produce a pure population of a single type of derivatives because its orchestration usually requires a cross-talk of signals from different co-differentiating cell types which are close to each other in culture,

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Figure 49.1 Growth and differentiation of hESC. (a, b) undifferentiated colonies of hESC. (a) phase contrast, (b) Oct-4 staining, (c, d) spontaneous differentiation of hESC, Scale bar: (a, b) 200 μm, (c, d) 500 μm. similar to patterning in vivo. Thus most differentiation models currently published have only partial efficacy, and isolation of the desired differentiation derivative is required for obtaining a pure cell population. There are currently several reports on derivation of RPE-like cells from ES cells of various any species. In 2002 Kawasaki et al. (2002) published a study describing derivation of dopamine neurons and pigmented epithelia from primate ES cells in the presence of stromal cell-derived inducing activity (SDIA) coming from a feeder layer of mouse skull bone marrow PA6 cells. The authors found patches of pigmented epithelial cells in their cultures, which stained positively for Pax6 and showed cortical actin distribution; such cells could be mechanically isolated under the microscope and passaged to confluency on PA6 cells or on collagen. This work was continued, and 2 years later another report was published describing in vitro and in vivo characterization of such pigmented epithelial cells (Haruta et al., 2004). The cells expressed mRNA for RPE65, CRALBP, and MERTK, which are all characteristic markers for RPE cells, and were able to perform phagocytosis of latex beads and support the function of the photoreceptor when transplanted in RCS rats. In another study, SDIA promoted the formation of various structures, pigmented epithelial cells among them, in the cultures of mouse ES cells differentiating in the presence of PA6 feeder cells (Hirano et al., 2003). Our group has generated RPE from hESC and found that hESC can reliably differentiate into RPE without any such aid, presumably as the “default” neural lineage commitment taken another step further (Klimanskaya et al., 2004). Such pigmented cells usually appear within 6–8 weeks after passaging ES cells, independently of the presence of mouse feeder cells in the starting culture. We were able to isolate putative RPE cells from every hESC line we used for these experiments (the total is currently over 20), establish primary cultures of up to 8 passages, and evaluate the molecular profiles of such cultured cells by immunostaining, RT-PCR, real-time PCR, Western blot, and gene expression analysis. In our system hESC are routinely cultured in serum-free medium, containing serum replacement (Invitrogen) and Plasmanate (Bayer) on mouse embryonic fibroblast feeder layers in the presence of 10–20 nm/ml human leukemia inhibitory factor (LIF) and 8–16 ng/ml of human bFGF; for passaging we use

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Figure 49.2 Formation of RPE in spontaneously differentiating hESC cultures. (a) phase contrast, (b) Pax6 staining, (c) tubulin βIII staining, (d) merged, (a–c, e) petri dish with pigmented culters, (f) embryoid body culture with pigmented cells on the surface. The arrows show the position of pigmented cells. Scale bar: (a–d) 100 μm, (e) 3 mm, (f), 200 μm. The parts of the figure are reproduced from Klimanskaya et al. (2004). with the permission from the publisher Mary Ann Liebert, Inc.

trypsin or mechanical dispersion of the colonies (Klimanskaya and McMahon, 2005). After 5–7 days in culture, we usually see signs of differentiation, when the typical ES cell morphology is lost and various differentiated cell types appear. Most colonies usually show signs of neural lineage commitment, including cells which stain positively for tubulin βIII, Pax6 (Figures 49.1b and 49.2), and glial fibrillary acidic protein (GFAP). These observations are in agreement with numerous observations in the literature that ES cells in culture select the neuronal pathway of differentiation most readily, which could be chosen by default (Tropepe et al., 2001; Smukler et al., 2006) or in response to autocrine activity of FGF (Ying et al., 2003; Bouhon et al., 2005) or as a result of elimination of other inductive signals (Ying et al., 2003). After 7–10 days of culture, when the majority of the cells have lost their ES cell morphology and molecular markers of pluripotency, the medium is changed to “differentiation medium” which has no LIF, no FGF, no Plasmanate. It is possible that one of the reasons for massive neural induction is the presence of bFGF in the ES culture medium when the initial commitment is being made. The plates are then cultured until the clusters of pigmented epithelial cells begin to appear, which usually happens in 6–8 weeks. Such clusters keep slowly increasing in size, while new clusters continue to emerge. The same process can be initiated in conventional embryoid body culture (EB), in which case pigmented epithelial cells would appear on the surface of EBs, and then this transition of non-pigmented cells to pigmented epithelium would slowly take over the whole EB. Of note, that in such differentiating systems we found clusters of cells positively stained for neural lineage markers Pax6 and/or tubulin βIII, often in close conjunction with pigmented epithelium; moreover, we frequently saw a pattern that looks like a transition from neural progenitors to RPE within the same cluster: as Pax6 and tubulin βIII staining becomes weaker and disappears toward the center in such clusters (Figure 49.2),

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pigmentation and epithelial morphology become more prominent (Figure 49.2). Such Pax6-positive cells could be similar to the multipotent cells of the optic vesicle, and the loss of Pax6 from more mature cells may be analogous to the same process during RPE maturation in eye development. The cells stained positively for tubulin βIII may be precursors of neural retina cells, and our histological examination of the same specimens revealed clusters that resemble disorganized rods and cones. RT-PCR performed on the same samples showed the presence of these cell types (data not shown). In addition, we usually saw cells of various types, still unidentified, in the same differentiating cultures of hESC, surrounding the clusters of RPE and their presumptive progenitors. It was possible that the cells producing signals promoting RPE specification in clusters of Pax6-positive progenitors, similar to the signaling of ocular mesoderm in patterning ocular tissues, could be found among such differentiated cells next to Pax6-positive clusters. These weeks-old cultures were comprised of several layers of cells with a lot of ECM deposition, which made it difficult to disperse them into a single cell suspension to select the desired cell type using fluorescence-activated cell sorter (FACS) or magnetic beads. Instead, we used an approach when the multi-layer of cells was loosened with trypsin or collagenase and the pigmented cells were picked under the dissecting microscope using a glass capillary. Collected cells were plated on laminin or gelatin in RPE culture medium containing serum replacement and fetal bovine serum (FBS) with optional bFGF, and in 48 h we found clusters of attached, spread cells which were beginning to proliferate (Figure 49.3b). Proliferating cells lost pigment and acquired a fibroblastic phenotype (Figure 49.2c), strongly resembling the transdifferentiated RPE which dedifferentiated as they proliferated and returned to typical RPE morphology after they established a monolayer (Figure 49.3d), which usually took 2–3 weeks (Reh et al., 1987; Vinores et al., 1995; Sakaguchi et al., 1997; Chen et al., 2003). Such RPE transdifferentiation has been shown to result in formation of neuronal, amacrine, and photoreceptor cells (Zhao et al., 1995), glia (Sakaguchi et al., 1997), neural retina (Galy et al., 2002), and neuronal progenitors (Opas and Dziak, 1994). bFGF accelerated transdifferentiation and RPE proliferation, thus allowing the cells to reach confluence and

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Figure 49.3 Morphology of hES-derived RPE in culture. (a) Differentiated cluster of RPE cells surrounded by other cell types, (b) isolated RPE cells after 5 days in culture, (c) transdifferentiated passaged hES–RPE, (d) morphology of hES–RPE monolayer in culture. Scale bar: (a, b, c) 200 μm, (d) 100 μm. The parts of the figure are reproduced from Klimanskaya et al. (2004). with the permission from the publisher Mary Ann Liebert, Inc.

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begin to revert to the RPE phenotype much sooner. ES-derived RPE (ES–RPE) in the transdifferentiated state expressed the neural markers Pax6 and tubulin βIII, strongly resembling immature neural cells, and our comparative gene expression profiling showed their similarity to neural stem cells (Klimanskaya et al., 2004).

CULTURE AND PROPERTIES OF HES-DERIVED RPE For any regenerative medicine application it is important that the cells can be produced in quantities large enough for thorough characterization and safety testing, and because the life span of differentiated cells is limited, these quantities need to be obtained before the cells get close to senescence and lose any of their RPE functions. RPE are relatively “slow” cells: even in the presence of bFGF that accelerates their transdifferentiation and proliferation, it may take up to 2–3 weeks at each passage at a 1:3 ratio before they “mature” and fully re-gain the RPE phenotype. We tested whether it is possible to passage the ES–RPE while they are still transdifferentiated, so the desired cell quantity can be achieved in less time, and after that allow them to differentiate, rather than letting the cells to go through a full cycle of dedifferentiation–proliferation–differentiation at every passage. We found that the cells that were “rushed” – passaged immediately after they established confluency – lost their ability to re-gain the RPE phenotype several passages earlier than the RPE which went through the whole cycle prior to each passaging. Such cell behavior requiring slow propagation may seem to impede scaling up the cells for characterization and therapy, especially because the yield of RPE from any given differentiating dish is quite low: from several clusters of RPE cells usually found in one 35 mm plate of differentiating ES cells (each cluster usually has several hundred cells; some large older ones may have several thousand) two or three confluent wells of a four-well plate can be produced in 3–4 weeks. After that the cells are usually subcultured at 1:3–1:6 ratio at 2–3 week intervals, so as to generate several million of cells at P2–P3, which would barely be enough for thorough assessment and one transplantation application, takes several weeks. However, with ES cells as starting material there is a simple solution: setting up large scale differentiating ES cell cultures, high numbers of RPE can be obtained at the earliest passages. Furthermore, better understanding of the signaling mechanisms guiding the differentiation of ES cells in culture would allow us to devise a step-wise procedure employing an efficient combination of growth factors and ECM to manipulate the differentiation process and increase the yield of RPE cells in less time. Since

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Figure 49.4 Transdifferentiation in hES–RPE cultures: immunolocalization of tubulin βIII and Pax6 in recently passaged dedifferentiated ((a–d) 3 days after passaging) and “mature” ((e–h) 3 weeks after passaging) cultures of RPE-like cells. (a, e) tubulin βIII, (b, f) Pax6, (c, g) corresponding phase contrast microscopy field, (d) merged, (a–c), (h) merged (e–g). Scale bar, 100 μm. The figure is reproduced from Klimanskaya et al. (2004). with the permission from the publisher Mary Ann Liebert, Inc.

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animal co-culture, as in studies with PA6 mouse stromal cells, is undesirable for potential therapeutic applications, some of the approaches might, for instance, focus on producing Pax6-positive retinal progenitors and exposing them to activin A or co-culturing ES cell-derived neuroectoderm with “autologous” (derived from the same ES cell line) mesoderm. After the culture of differentiation derivatives of ES cells is established, the next important step is to characterize the cells at the molecular and functional level. In our studies we are using bestrophin, CRALBP, RPE65, and PEDF as RPE molecular markers. Pax6, although seen by some authors as a molecular marker of RPE (Kawasaki et al., 2002), is normally downregulated in mature RPE, so it could rather indicate the presence of immature cells. Our ES–RPE cells showed a remarkable resemblance to cultured or intact human RPE: CRALBP, PEDF, and bestrophin were detected by Western blot and by immunofluorescence (bestrophin and CRALBP) and enzyme-linked immunosorbent assay (ELISA) (PEDF). Translationally controlled RPE65 was not found at the protein level, although the real-time RT-PCR has detected high levels of RPE65 mRNA. Interestingly, the level of its expression correlated with the differentiation: in more mature cultures its expression was several times higher than in recently passaged cells (Klimanskaya et al., 2004). Functional tests for characterization of potential transplantation candidates could include RPE-specific phagocytosis using an assay with labeled rod fragments (Finnemann et al., 1997; Finnemann, 2003) and vitamin A metabolism assay. Another useful “quality assurance” assay could be cell adhesion assessment: literature and our own data indicate that there may be certain variations in integrin expression, and it is yet to be determined whether this is genetic variation or a result of culture condition and in this case could be modulated by plating cells on various ECMs or by using bioactive substances. For instance, loss of integrin alpha 5 in postconfluent culture was described by Proulx et al. (2003) and this may impair adhesion of the cells used for transplantation. In our ES–RPE cells we detected a certain variation in expression of α5 integrin which could be the outcome of such post-confluent cultures. Recently αVβ5 was identified as a receptor involved in retinal adhesion (Nandrot et al., 2006) in addition to its role in phagocytosis (reviewed by Finnemann, 2003), and it is unclear yet how much variation in expression of various integrins may be tolerated by the RPE cells without significant loss of their adhesion qualities and functionality. However, all the molecular and functional in vitro characterization data need to be assessed with understanding of tissue culture limitations which may be reversible. The same cells that fail one criteria in culture may turn out to be excellent performers in the natural eye environment. Currently there is not enough data in the literature correlating the molecular profile and in vitro assayed functions of cultured and freshly isolated RPE with its ability to support the function of the photoreceptor in animal models. This is applicable to any stem cell derivative: the molecular profiles, morphology, and function could appear quite similar to their in vitro counterparts, but even a seemingly subtle difference may turn out to be crucial for performance of the transplanted derivatives in vivo – or it may be reversible when the shortcomings of the culture system are replaced with natural environment. Gene expression profiling performed on hEC-derived RPE versus fetal human RPE tissue showed their remarkable similarity (Klimanskaya et al., 2004). The data were also compared with previously published data (Rogojina et al., 2003) on human RPE cell ARPE-19 (spontaneously derived and shown to attenuate the loss of visual function in RCS rats, Lund et al., 2001) and D407 (transformed), and it appeared that hES–RPE showed much more resemblance to fetal RPE than any of these in vivo originated RPE cell lines, the latter two lacking expression of many RPE-specific genes. For instance, neither of these two lines expressed bestrophin, lecitin retinol acyltransferase, retinal G-protein coupled receptor, or RPE65. Our preliminary data show that different hES–RPE lines used for transplantation in RCS rats had different efficacy lines in photoreceptor support, and we are currently correlating these differences with the phenotypes of the lines, number of passages after differentiation, molecular markers, and adhesion receptor expression. Some of the possible reasons for

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different transplant performance could be genetic differences between the lines, or differences in culture conditions and number of passages, which in turn would determine the differentiation potential of the cells. We found that at early passages after derivation hES–RPE transdifferentiate and differentiate to RPE phenotype quite easily, producing a homogeneous monolayer of pigmented polygonal cells. With more divisions and higher passage numbers we start seeing more elongated non-pigmented cells, and after the cells undergo 6–8 passages, they lose the ability to revert to RPE phenotype after transdifferentiation. For any regenerative medicine applications it is of utmost importance to be able to predict the therapeutic value of any particular culture of stem cell derivatives, and for RPE, assessment of morphology and expression of molecular markers, even at the gene chip scale, may not be sufficient due to high plasticity of these cells. For instance, fully differentiated cells may have a disadvantage attaching to the Bruch’s membrane and proliferating, while highly transdifferentiated cells may act unpredictably in the subretinal space environment. On the other hand, the in vivo microenvironment may be able to instruct the cells much better than it can be done in culture conditions. If the transplanted cells attach and survive, they still have to prove functionality by restoring the function of the photoreceptor, and this may depend on their intrinsic abilities, which may not be necessarily directly dependent on their molecular profiles. It is important to perform a series of experiments correlating the performance of the cells in animal models with their in vitro assessment, and find parameters that would ensure their grafting, survival, and function. A possible solution would be to use gene expression profiling approach in combination with in vitro functional assessment of differently isolated and cultured derivatives, and/or on progenies of several different ES cell lines and to compare the behavior of the same cells in animal models in order to identify the crucial molecular markers for the survival and function of the candidate cells. Still the paramount test of the cells’ quality is their performance in clinical and preclinical trials. RPE considered for clinical trials need to be proven safe and efficient in appropriate animal models. Several animal studies with other hES-derived cells (neural progenitors) showed that they formed tumors, even though they were Oct-4 negative (Arnhold et al., 2004), and RPE would need to successfully pass safety tests in animals in addition to pathogen clearance. Controlled efficacy studies are also a part of the preclinical trials that can be performed in an adequate animal model. Our own studies have shown that in RCS rat model hES-derived RPE are able to extensively resuce the photoreceptor in this model of inherited retinal degeneration (Lund et al., 2006) reaching 100% improvement over the untreated control. If hES-derived RPE can successfully pass preclinical safety and efficacy tests, it can become one of the first hES-derived cell products for regenerative medicine applications.

ACKNOWLEDGMENTS I thank Sandy Becker for critical reading of the manuscript and helpful comments and apologize to all those whose work was not mentioned due to space limitations.

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50 Cell Therapies for Bone Regeneration Rehan N. Khanzada, Chantal E. Holy, F. Jerry Volenec, and Scott P. Bruder

INTRODUCTION Historical Overview Bone damage, either due to pathology or trauma, is a very common occurrence that requires costly medical and/or surgical intervention, and is associated with significant morbidity (Cancedda et al., 2004). Of all fractures that occur in the United States each year, about 15% require some type of bone grafting to improve the healing process. To date, graft materials include autograft (bone taken from one part of the patient’s body and replaced in another site that requires bone healing), allograft (bone taken from a donor) or synthetic materials. The earliest evidence of an orthotopic autograft dates back to the Bronze Age. A circular disk of bone was removed from a human’s calvarium to relieve intracranial pressure and placed elsewhere as an autograft. Written accounts from Egypt, China, and India dating back many centuries describe similar autograft-based experimentation. One Indian text from 700 CE describes a procedure for nasal reconstruction that is very similar to modern methods. While autograft is currently considered the gold standard for bone regeneration due to its success rate, it requires secondary bone harvesting procedures that can cause high morbidity (Gupta et al., 2001) and is the only available in small supply. The first use of allografts and xenografts (bone from a donor of different species) dates back to over 300 years ago, when Job van Meekeren historically performed the first bone graft procedure using canine xenograft in 1668. The need for bone grafting became critical during World War II, as the US Navy established bone banks to treat fractures sustained in war. However, despite significant progress in allograft preparation and cleansing technologies, allografts still carry the risk of disease transmission. Synthetic grafts of all types have therefore been developed. While these grafts are available in high volumes and do not carry risks of disease transmission, their effectiveness in vivo does not consistently meet that of autograft. Research on synthetic grafts for bone regeneration has thus evolved into state-of-the-art science, especially after the discovery of mesenchymal stem cells (MSCs) capable of forming bone (Friedenstein et al., 1968) and bone morphogenetic proteins (BMPs) (Urist, 1965). The Clinical Need for Therapeutic Solutions to Bone Regeneration One of the reasons for so many graft choices is the vast quantity of bone graft required: an estimated 1.5 million bone graft operations were performed in the United States in 2004 to enhance the healing of spinal fusions, internal fixation of fractures, maxillofacial reconstruction, long-bone repair, and lost bone due to trauma or ablative surgery. However, selecting the right graft for the right patient is one of the key challenges

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Figure 50.1 Schematic representation of iliac crest autograft harvesting. The iliac crest represents the largest source of autologous bone, and requires a secondary surgery that can lead to morbidity.

for orthopedic and spine surgeons. In fact, the choice of a bone graft is based on four main factors: the size of the defect, the location of the defect, the biology of the defect site, and whether structural support is required (Gamradt and Lieberman, 2003). Autograft harvested from the iliac crest is most often used in treating these conditions, as it is histocompatible, and does not transport any diseases. A schematic representation of iliac crest autograft harvesting is shown in Figure 50.1. Hydroxyapatite and collagen within the native bone serve as osteoconductive frameworks, while stromal cells within the bone marrow, and to some extent, along the microcavities of the bone, contain osteogenic cells that lead to reproducible bone formation when placed in a surgical site. In addition, growth factors within the bone and adjacent hematoma provide osteoinductive factors (Sutherland and Bostrom, 2005). There are some drawbacks and potential complications associated with autograft harvested from the iliac crest. Although severe complications from iliac crest bone harvesting are rare, the incidence of donor site pain reported in the literature ranges from 25% to 49%, with 19% to 27% of patients experiencing chronic site pain 2 years postoperatively (Younger and Chapman, 1989; Fernyhough et al., 1992). To better understand the causes of this morbidity, Gupta et al. reviewed literature reports spanning 34 years and including 1,020 patients. The authors found no correlation between the patients’ pain ratings and any of the following parameters: incision site, surgical approach, harvesting technique, or demographics including patient age or gender. In addition to the issue of unpredictable morbidity, limited harvest supply of autograft is sometimes problematic for patients undergoing procedures that require large graft volumes. Autograft also has poor handling characteristics, as it is typically morcelized during harvesting and does not have any structural integrity (Figure 50.2). Cadaveric allograft is sometimes used but there are continued concerns about graft resorption, inadequate revascularization, and possible transmission of blood-borne diseases. Allografts may be demineralized to expose native growth factors, which increase the grafts’ in vivo efficacy (Zhang et al., 1997). Structural allografts, on the other hand, are frozen or freeze-dried, which destroys cells within the allografts, thereby reducing potential complications from immune responses but also destroying the grafts osteogenic activity (Goldberg and Stevenson, 1987).

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Figure 50.2 Photograph of human iliac crest autograft, morcelized, and ready for re-implantation. Autograft from the iliac crest is the “gold standard” for bone regeneration.

To address shortcomings of both autograft and allograft, completely synthetic options are being developed with an eye toward creating synthetics that would mimic autograft, and thus a strong understanding of the biological processes required for bone formation has become critical. Biological Ingredients Bone repair and regeneration is a complex process consisting of a tightly regulated cascade of cellular interactions. As part of the acute inflammatory response, bioactive molecules are released, which promote the influx of MSCs to the fracture site. These MSCs adhere to osteoconductive scaffolding within the fracture site and, in response to local growth factors, proliferate and differentiate into osteoblasts capable of secreting osteoid, which is subsequently mineralized to form new bone (Whang and Lieberman, 2003). Thus, the ideal bone graft for bone repair and regeneration requires three key ingredients: (1) surface areas allowing cell attachment (i.e. osteoconductive scaffold), (2) cells capable of forming bone (i.e. osteogenic cells), and (3) biological stimulants. While synthetic methods can be used to develop large amounts of osteoconductive surfaces, finding enough osteogenic cells to populate grafts may be seen as a limiting factor for success. The search for rich and easily accessible sources of osteogenic cells is therefore spurring significant interest. Delivery of Osteogenic Cells Early work by Burwell (1964) demonstrated that the main repository of cells capable of forming bone within iliac crest bone grafts was the bone marrow. Owen’s studies (1985) using in vitro cell growth confirmed that isolated cells from bone marrow had osteogenic and adipogenic potential (Figure 50.3). The term osteogenic was thus coined to define a cell capable of forming bone or capable of differentiating into a bone-forming cell; osteogenicity further referred to materials containing osteogenic cells. Following the report of bone marrow’s

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Bone cell

Cartilage cell

Tendon cell

Stem Cell

Nerve cell

Muscle cell

Figure 50.3 Schematic representation of mesenchymal stem cell development pathways. As discovered in the 1980s by Owen et al. mesenchymal stem cells can develop along multiple pathways.

osteogenicity, orthopedic surgeons since the 1980s routinely used fresh bone marrow for repairing large bone defects. As bone marrow was analyzed for osteogenic cell content, Muschler and Midura (2002) and others demonstrated that less than 1% cells within the marrow had osteogenic potential. In addition, Muschler et al. (2001) demonstrated that the number of osteogenic cells was variable from one patient to another, especially as a function of age and gender. Therefore, new methodologies that would take advantage of bone marrow’s osteogenicity and alleviate issues of cell count variability were also investigated, as described below. As bone marrow has been defined as an easily accessible source of osteogenic cells, two additional ingredients for bone regeneration are thus required: an optimized carrier and biological stimulants. Carriers and Growth Factors Osteoconductive graft materials refer to scaffolds that provide the appropriate framework for bone growth and osteoblast attachment. These scaffolds provide appropriate three-dimensional shape and structure to restrict cell movement in an implant site. For successful bone healing, these scaffolds need to have direct contact with viable bone and support bony ingrowth and vascularization without excessive inflammatory response. Examples of osteoconductive scaffolds include naturally occurring materials such as mineralized cancellous chips and fibrin clots, and synthetics such as tricalcium phosphates, hydroxyapatites, collagen sponges, and various polymers. The most appropriate scaffold for a given clinical application depends on the pathological condition being treated, its anatomic location, and the biomechanical stresses and loads that apply to that specific site. Osteopromotive graft materials have the ability to provide stimulatory signals at various stages and enhance the bone repair and regeneration process. These materials do not have the capacity to induce new bone growth by themselves and work best with osteoinductive and osteogenic graft materials in orthotopic applications. An example is platelet-rich plasma (PRP), which is prepared by collecting and concentrating platelets from a patient’s whole blood immediately before surgery. These platelets contain a rich source of various growth factors that play an important role in bone repair and regeneration (Kevy and Jacobsen, 2004).

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Osteoinductive graft materials have the capacity to induce bone growth in ectopic sites. These materials function as biological stimulants (i.e. growth factors), which activate MSCs toward chemotaxis, proliferation, and differentiation into osteoblasts that leads to new bone growth (Urist, 1965). BMPs are the key osteoinductive proteins identified to date. These are either available at native levels in allogeneic demineralized bone matrix (DBM) or as recombinant human proteins. Osteoinductivity of DBM varies between donors and as a function of bone processing (Zhang et al., 1997); therefore, ongoing osteoinductivity testing of DBM is required to ensure high potency of these graft materials. The BMPs are members of the TGF-β superfamily of growth factors known to play a critical role in initiating endochondral bone formation. They are low molecular weight proteins that induce a quick biological response at the implantation site. The US Food and Drug Administration has recently approved recombinant human BMP-2 and BMP-7 for specific and limited clinical indications. Since these are potent osteogenic agents, they require an optimized delivery system in order to provide appropriate biological response. Currently, supraphysiologic doses of these recombinant BMPs are required for induction of bone formation (Yoon and Boden, 2002). Current delivery methods for these growth factors are either collagen-based sponges alone or calcium–phosphate granules, used as bulking agents, surrounded with collagen sponges (Barnes et al., 2005). While these combination matrices may be more effective than the collagen alone, these delivery systems still require significant optimization as current recommended clinical doses are excessively high, very costly, and have unknown long-term effects (Gamradt and Lieberman, 2004). Despite such high doses, BMPs may not produce sufficient osteogenic response where there is poor bone quality, scar tissue, large defect size, or inadequate vascularization (Cook et al., 1995). While more research is required to optimize the use of BMPs, the availability of these growth factors represents progress for bone grafting, as grafts containing all three ingredients (matrix, cells, and biological stimulants) could thus be envisioned. Practical and effective approaches for the preparation of such grafts have therefore been investigated, as described below, starting with optimized methods for obtaining osteogenic cells.

CURRENT SOURCES OF OSTEOGENIC CELLS AND CELL ISOLATION TECHNIQUES Source of Osteogenic Cells Osteogenic cells, as mentioned above, can be defined as cells that are, or will differentiate into, osteoblasts capable of forming bone. There are two major sources for osteogenic cells: (1) tissues containing MSCs (e.g. bone marrow) and (2) differentiated bone tissue. MSCs, the main source of osteogenic cells, can be further defined as cells that retain the capacity to differentiate along osteogenic, adipogenic, fibroblastic, and chondrogenic lines (Lennon et al., 1996; Bruder and Caplan, 2000). The most accessible source for MSCs is the bone marrow. More recently, differentiated connective tissues have also been shown to contain osteogenic cells that could form bone in specific culture conditions (Zuk et al., 2002). A secondary source of osteogenic cells that was recently described consisted of mature bone fragments, obtained as debris from reaming procedures (Wenisch et al., 2005). These samples did not contain any bone marrow, and thus the osteogenicity of the bone fragments was described as specifically due to bone-lining cells. This source of osteogenic cells has not been widely investigated, and thus will not be further described in this manuscript. Autologous Bone Marrow As described above, bone marrow was shown in the mid-1960s to contain both hematopoietic stem cells and MSCs (Friedenstein et al., 1968). Recently, a third type of precursor cell was identified within bone marrow: the “side population” (SP); these cells are defined by their ability to regenerate the hematopoietic compartment as

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well as to differentiate into osteoblasts through a mesenchymal intermediate (Olmsted-Davis et al., 2002). These findings suggest a population of cells, precursor to MSCs, and elucidates yet another step in the early development of osteoblasts. Unlike the bony site of an autograft harvest, bone marrow is self-renewing and can be obtained from the iliac crest in a non-invasive procedure with a simple needle aspiration. This usually does not cause morbidity (Connolly, 1995) and is the most inexpensive method for bone repair and regeneration. To increase the effectiveness of bone marrow, Muschler et al. (1997) investigated bone marrow aspiration techniques, and described a methodology to maximize the osteogenic cell content within a given bone marrow aspirate. This technique involved aspirating no more than 2 ml of bone marrow from a given site. The aspiration needle was then moved either further into the bone marrow cavity or at a different location, in which an additional 2 ml could then be aspirated. Aspirating more than 2 ml per site resulted in dilution of the bone marrow with peripheral blood, and thus diluting the cellularity of the final aspirate. While this technique was shown to ensure the highest possible cellularity within the bone marrow aspirate, researchers are also looking at other methods to completely move away from potential dilution issues by developing osteogenic cell banks that would provide the same number and efficacy of cells without requiring aspiration of bone marrow from patients prior to bone grafting. Allogeneic Bone Marrow Allogeneic stem cell sources hold great promises as universal cell banks that may be developed for bone and other tissue repair. It was hypothesized early on that allogeneic MSCs might be applicable for bone repair and regeneration if one could successfully mute immunoreactive groups on the MSCs. However, in vivo preclinical studies seemed to indicate that, surprisingly, allogeneic MSC implantation failed to provoke an immune response. In one instance, analysis of circulating antibody levels against MSCs 9-week postimplantation in a canine cranial site supported the hypothesis that neither autologous nor allogeneic MSCs induced a systemic response by the host. Authors concluded that autologous and allogeneic MSCs had the capacity to regenerate bone within craniofacial defects (De Kok et al., 2003). More recently, undifferentiated human MSCs were shown not to express immunologically relevant cell surface markers. They also seemed to inhibit the proliferation of allogeneic T-cells in vitro. Evidence seemed to indicate that these cells did not elicit an immune response after allogeneic or xenogenic transplantation. Thus, MSC could be described as immunoprivileged or immunomodulating cells (Niemeyer et al., 2004). These findings confirmed that allogeneic stem cells may indeed become a possible therapeutic tool and as such are currently being developed for bone and soft tissue repair. However, due to the potentially arduous regulatory path required for allogeneic stem cells to meet approval by federal agencies, other autologous sources of osteogenic cells are also being investigated. Novel Tissue Sources for Osteogenic Stem Cells: Muscle, Fat and Other Connective Tissues Cell derived from connective tissues such as muscle and fat were shown to “behave” similarly to bone-marrow-derived MSCs. These cells had the ability to differentiate into bone under appropriate biological cues (Betz et al., 2005). Both muscle and fat cells were found to contain MSCs that were readily expanded in culture and underwent osteogenic differentiation. These MSCs were obtained conveniently from muscle biopsy or liposuction, procedures involving less morbidity than traditional bone graft harvesting. Interestingly, these cells were shown to differentiate not only into bone, but also fat, cartilage, and muscle tissues, with growth factors specific to each culture condition (Zuk et al., 2002). Muscle-derived stem cells were also retrovirally transduced to express osteogenic factors BMP-2 and BMP-4, and were capable of differentiating into bone and accelerating repair of skull defects in mice (Huard and Peng, 2004).

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Fat was found to be the most convenient tissue other than bone marrow for harvesting MSCs, as it is easily biopsied, cultured, expanded, and transduced. Moreover, adipose tissue was shown to contain proliferative properties that did not decline with age. In addition, this tissue was described as a richer, and more effective, source of osteoprogenitor cells than bone marrow, when genetically modified to express BMP-2 (Dragoo, 2003). More recently, Cowan et al. used these cells to investigate the in vivo osteogenic capability of adiposederived stromal (ADAS) cells to heal critical-sized mouse calvarial defects. These ADAS cells were harvested from the subcutaneous anterior abdominal wall yielding 800 mg of fat tissue, and were shown to be multipotent, and available in large numbers. In vitro, they were observed to attach and proliferate rapidly. Authors also reported a yield of cells that, by itself, was much higher than that of typical bone marrow; however, it is worthwhile to note that mouse bone-marrow-derived cells are technically difficult to isolate and manipulate, and that this fact is not usually observed with tissues from other species. In vivo, ADAS cells seeded onto apatite coated PLGA scaffolds regenerated bone in critical sized calvarial defects (Cowan et al., 2004). With osteogenic cells available and clearly identifiable in vitro, researchers also investigated the potential to genetically modify these cells to secrete the biological stimulants required for bone formation. That way, two of the three ingredients for bone formation could be provided within a given cell population. This gene therapy approach is briefly described below and in other chapters. Gene Therapy Gene therapy deals with the transfer of genetic material into cells, which in turn, secrete specific proteins in selected sites. In this model, growth factor(s) are synthesized in situ as a result of gene transfer and would be presented to the surrounding tissue in a natural, cell-based manner (Nussenbaum and Krebsbach, 2004). Local MSCs would then undergo osteogenic differentiation, or form another appropriate tissue (Lou, 2004). Gene therapy involves three fundamental elements: a sequence of DNA encoding a protein of interest, a vector that facilitates the entry of genetic material into cells, and target cells into which the gene is inserted. Two different types of therapeutic conditions can be envisioned for gene therapy: (1) conditions that require continuous, sustained delivery of specific proteins and (2) conditions that require a transient bone inducing agent (e.g. trauma cases and spinal fusions). A clinical example of a case requiring continuous delivery of proteins includes osteogenesis imperfecta. In this case, a patient would be implanted with MSCs genetically modified with a retrovirus containing the gene for normal type I collagen. These cells would re-establish themselves in the bone marrow and thus provide mesenchymal progenitors with the appropriate collagen Type I building capabilities (Pereira et al., 1995). Retroviruses are currently the key vectors for continuous gene expression but, as they depend on cell replication for transcription, they can only be used in highly proliferative cells (Tibor, 2003). For one-time bone repair applications (e.g. bone fractures, spinal fusions), gene therapy must be limited to short-term gene expression, and thus, non-viral vectors are currently under investigation. These are typically easier to generate, more stable than viruses, and less immunogenic (Gamradt and Lieberman, 2004). These vectors are however far less effective than viral vectors to transduce cells. Others options to transduce cells also include adenoviruses, whose limitation include the potential to provoke immune responses (Musgrave et al., 2002). In addition to short- and long-term protein expression, two types of gene therapy approaches are currently under investigations: a so-called in vivo approach, as well as an ex vivo approach. In vivo gene therapy involves the direct transfer of genes into patients, while the ex vivo gene therapy involves transducing cells in vitro and then implanting those cells at a specific site. Both in vivo and ex vivo gene therapies for bone formation have been successfully demonstrated in several different animal models, including rat femurs, mice skulls, and anterior spinal fusion in pigs. But while successful in preclinical trials for one-time bone repair

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applications, the use of gene therapy for these particular conditions may be excessive, since biologically simpler and more economical approaches can effectively treat bone conditions like fractures and fusions. Gene therapy, therefore, may be better suited for pathologies that, like osteogenesis imperfecta, require continuous protein expression and for which no satisfactory treatment option exists. Cell Isolation Techniques Use of Autologous Bone Marrow on Optimized Matrices As far back as the late 1960s, bone marrow was used as a research tool to isolate MSCs that would form bone in vitro. Maniatopoulos et al. (1988) first described a methodology to isolate osteoprogenitor cells from rat bone marrow. This technique involved explanting the entire femur of the rats and flushing the bone marrow in an osteogenic cell media. After 7–14 days, alkaline–phosphatase positive colonies would become visible, indicating potential differentiation of osteogenic cells. Cell isolation methodologies were then developed to culture human bone-marrow-derived cells. Jaiswal et al. (1998) first established a reproducible system for the in vitro osteogenic differentiation of human marrow-derived MSCs. Muschler et al. (1997) utilized similar cell isolation techniques to quantify osteogenic precursors in bone marrow aspirates of patients. Both rodent and human cultures were found to be strongly sensitive to media as well as surface conditions. The term “osteoconductive” was found to be particularly important in vitro (as well as in vivo), since cell growth on specific polymers was found to be inhibited, while that on, for example, collagen or poly-L-lysine coated surfaces, it was found to be optimal (Liu et al., 1999; Karp et al., 2003). These studies highlighted the importance of osteoconduction and optimized carriers for cell proliferation and differentiation. Point-of-care Osteogenic Cell Enrichment As discussed previously, age, disease, and other factors can reduce bone marrow cellularity prompting the idea of cellular enrichment methodologies that could be used at point of care. Three major approaches were described in the literature to increase cell numbers within graft materials: (1) enzymatic tissue digestion, (2) bone marrow centrifugation, and (3) selective cell retention. Enzymatic Tissue Digestion

Enzymatic tissue digestion so far has strictly been used as an in vitro method to release osteoblasts from bone tissue. It can be hypothesized that enzymatic tissue digestion could be used to release cells from bone fragments obtained during surgery, for re-implantation in defect sites. In brief, enzymatic tissue digestion as described in the literature involves mincing bone tissues (typically from rodent femurs or calvaria) and washing those minced fragments with series of collagenase/trypsin enzyme solutions. These solutions degrade connective tissues between cells and release osteoblasts (Thomas et al., 2004). This technique was shown to offer high-level cellular yields containing precursor, differentiating, and osteoblastic cells in enzyme-enriched cell preparations (Vinay et al., 1981). Released osteoblasts and committed osteoprogenitor cells exhibited properties of bone that included characteristic morphology, synthesis of bone-related proteins, and calcification after 3–4 weeks in culture (Webster et al., 2000). In another example, rat calvarial cells formed mineralized nodules within 2 weeks in vivo (Irie et al., 1998). Similarly, osteoblasts obtained from canine diaphyseal bones developed bone-like tissue in vitro. (Boyan et al., 1999). While successful in yielding osteoprogenitors, this method was also shown to have significant shortfalls: the collagenase had the potential to harm osteoblasts by removing proteins from their membrane, thereby affecting their ability for attachment. When compared to other cell isolation techniques, enzymatic tissue digestion therefore produced the lowest amount of functional osteoblasts.

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Bone Marrow Centrifugation

The concept of isolating bone marrow cells using centrifugation dates all the way to Friedenstein et al. (1968). Early research used Ficoll gradients to separate the cellular content of bone marrow. In 1989, Connolly et al. established a protocol to recover close to 100% of all nucleated cells within a bone marrow aspirate. Briefly, bone marrow was centrifuged at 400 times gravity for 10 min. The band containing the nucleated cells was removed and counted (Connolly et al., 1989). While cells obtained by centrifugation showed less metabolic activity (i.e. produce smaller amounts of lactic acid, consume smaller amounts of glucose, and contain less intracellular protein) than those cells obtained by enzymatic tissue digestion, the overall osteoblast concentration yield was far greater using the centrifugation technique than with tissue digestion (Thomas et al., 2004). Selective Cell Retention

Following the centrifugation techniques, Muschler et al. (1997) developed a cell enrichment methodology that used the principles of an affinity chromatography column to retain anchorage-dependent connective tissue osteoprogenitors on porous biological matrices. Unlike the centrifugation technique that concentrated all nucleated cells, Muschler’s technology only retained osteoprogenitor cells that would develop along the connective tissue paths. In this line of work, Muschler et al. realized that most of the nutrients available within the graft sites were taken over by cells that did not affect bone regeneration. Reducing the number of nonessential cells and increasing that of osteoprogenitors could ensure nutrients and oxygen availability for boneforming cells. The selective cell retention technology used a process in which fresh bone marrow was passed through a porous, three-dimensional bone matrix under controlled flux conditions. This technique allowed attachment of nearly 90% of osteoprogenitor cells to the matrix surface, with no selective retention of other nucleated or hematopoietic cells. This technique produced a bone graft substitute with an average of 3.6-fold increase of osteoprogenitor cells per unit volume. Culture Expansion As described above, MSCs could be isolated from bone marrow and expanded in cell culture, which could raise prospects of cellular concentrations much greater than the 3.6-fold increase observed with cell-enriched techniques. This culture expansion of MSCs in the laboratory was shown to provide an abundant supply of osteogenic cells for bone repair and regeneration. MSCs derived from bone marrow retained their undifferentiated phenotype through an average of 38 doublings, resulting in over a billion-fold expansion. These cells were then differentiated into osteoblasts by culturing with dexamethasone, ascorbic acid, and β-glycerophosphate (Bruder et al., 1997). The in vitro reports of bone marrow cell isolation, culture, and differentiation into osteoblasts further prompted questions on how fine tuning and optimization of cell-based bone graft could maximize the in vivo efficacy of the grafts. Multiple preclinical and clinical evaluations were thus conducted.

PRECLINICAL AND CLINICAL RESULTS Preclinical Studies Bone grafts with biological stimulants are developed under strict FDA guidelines that often require a substantial number of preclinical animal studies followed by human clinical trials. On the other hand, bone graft extenders that contain only osteoconductive materials have a faster path to clinical availability. In both cases, preclinical and clinical data are critical to convince medical professionals of the efficacy of new products.

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While animal studies can significantly improve understanding of biological mechanisms, they need to be conducted with a critical understanding of species-specific anatomical differences. Autologous Bone Marrow with Optimized Matrices The use of bone marrow has more than 40 years of ongoing preclinical research for bone repair and regeneration. Specifically, more than 100 peer-reviewed papers have analyzed the benefit of using autologous bone marrow for bone repair and regeneration in preclinical orthopedic and spinal fusion studies. However, bone marrow research carried some specific challenges, including: (1) limited bone marrow volumes available in some species (e.g. rats); (2) bone cellularity profiles different in animals vs. humans (e.g. rabbits have poorly cellular bone marrow in their iliac crest bones but highly cellular marrow in their long bone, a pattern opposite to that found in humans); and (3) lack of cell culture techniques to fully characterize bone marrow from different species (e.g. bone marrow from sheep requiring completely different culture conditions than that of other species). Bone marrow research in vivo has mostly provided positive results: Ohgushi et al. (1989) demonstrated that bone marrow cells delivered on a hydroxyapatite carrier could heal critical sized defects in rats. Tiedeman et al. (1991) used bone marrow aspirate and DBM to heal critical sized tibial defects in dogs. Novel composite carriers that combine collagen and hydroxyapatite (Figure 50.4 – HEALOS®) were also shown to effectively

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Figure 50.4 Bone marrow carriers: (a) low-magnification micrograph of a bone marrow carrier; (b) scanning electron micrograph of bone marrow nucleated cells attaching on mineralized collagen fibers. Nucleated cells are also described as anchorage-dependent cells. As such, they will adhere to osteoconductive surfaces, as shown in this scanning electron micrograph of cells on mineralized collagen (Healos®).

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Figure 50.5 Postmortem radiographs of rabbit spine segments implanted with: (a) autograft; (b) bone marrow on a carrier (courtesy: Tay et al., 1998). Animals were sacrificed 12-week postsurgery for further analyses of the fusion sites. Autograft and bone marrow on carrier (Healos®) were equivalent in this study.

regenerate bone when combined with autologous bone marrow (Figure 50.5 – Tay et al., 1998). This last study, however, was the subject of a controversy that highlights the need for a clear understanding of bone marrow cellularity in animal models. Tay et al. used a rabbit posterolateral fusion model to evaluate the carrier with bone marrow and showed that, when combined with heparinized or non-heparinized bone marrow, bone marrow constructs yielded fusion rates comparable to autograft. As with all animal models, several challenges needed to be addressed to generate clinically relevant data; in particular, rabbits were known as suboptimal species for bone marrow research, their iliac crest having poor cellularity. Tay et al. therefore created a secondary surgical site to harvest bone marrow from the rabbits’ long bone. The cellularity of the marrow aspirate obtained by Tay averaged 238 million cells/ml. Unlike humans, this bony site was shown to provide the most cellular marrow. In a recent publication using the same carrier, Kraiwattanapong et al. (2005) went back to the rabbit model to compare the efficacy of bone marrow aspirate to human recombinant BMP-2 (rhBMP-2). Two groups were compared: the first group was implanted with the carrier and autologous bone marrow from the iliac crest of the rabbits, while the second group was implanted with BMP-2 on a collagen–ceramic material. The average cellularity of the marrow aspirate was 30 million cells/ml. The authors used radiographs and manual palpation to report no fusion in the bone marrow group and 100% fusion in the rhBMP-2 animals. In a second publication involving rhBMP-2 and bone marrow, Minamide et al. (2005) evaluated four different graft materials in groups of seven animals each: (1) autologous bone; (2) collagen–ceramic material with rhBMP-2; (3) bone-marrow-derived, culture-expanded cells at a concentration of 1 million cells/ml on hydroxyapatite; and (4) bone-marrow-derived, culture-expanded cells at a concentration of 100 million cells/ml on hydroxyapatite. Manual palpation and radiography indicated 4/7 fusions in the autograft group, 7/7 fusions in the rhBMP-2 group, 0/7 fusions in the 1 million cells/ml group and 5/7 fusions in the 100 million cells/ml group. Minamide et al. (2005) thus concluded that, if expanded, bone marrow cells were capable of forming bone similar to autograft. Interestingly, while conclusions of these 3 studies seem contradictory at first, a closer look in the use of bone marrow and the cellularity of the marrow aspirates indicated that in fact, these papers demonstrated the same message: bone marrow requires osteogenic cells to form bone, and the use of suboptimal marrow, depleted in cells, did not result in bone formation: Both Kraiwattanapong and Minamide used bone marrow from the rabbits iliac crest that had very low cell counts. When Minamide increased the cellularity of the grafts to 100 million cells/ml by culture expansion, fusion rates were comparable to autograft. Using the rabbit’s long bone, Tay obtained fresh bone marrow with 238 million cells/ml and observed a fusion rate similar to autograft without the need to culture-expanded cells.

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Interestingly as well, these publications may seem to indicate that the bone healing performance of the carrier with the highly cellular bone marrow, as seen by Tay et al. was comparable to that seen with the rhBMP-2, as reported by Kraiwattanapong. However, in absence of a true head-to-head comparison between rhBMP-2 and optimized bone marrow grafts, the relative efficacy of those two types of graft cannot be effectively demonstrated. Others however have attempted to understand how bone marrow graft would perform compared to BMPs. The earliest study evaluating bone marrow versus BMP-2 dates back to 1981 (Takagi and Urist, 1982). In this early rat segmental defect study, grafts containing rhBMP alone or bone marrow alone did not perform as well as grafts containing both rhBMP and bone marrow. More recently, Den Boer et al. (2003) investigated the healing potential of ceramic grafts containing either bone marrow or OP-1 (rhBMP-7), in a 3-cm segmental bone defect in sheep tibia. Five treatment groups were included: no implant, autograft, hydroxyapatite alone, hydroxyapatite loaded with rhOP-1, and hydroxyapatite loaded with autologous bone marrow. At 12 weeks, torsional strength and stiffness of the healing tibiae were about two to three times higher for autograft and hydroxyapatite plus rhOP-1 or bone marrow compared to hydroxyapatite alone and empty defects. The mean values of both combination groups were comparable to those of autograft. Healing of bone defects, treated with porous hydroxyapatite was enhanced by the addition of rhOP-1 or autologous bone marrow. The results of these composite biosynthetic grafts were equivalent to those of autograft. Cell-Enriched Grafts Connolly et al. (1989) first suggested concentrating bone-marrow-derived cells on bone grafts for performance enhancement. A 4 cell concentrate of rabbit bone marrow using the centrifugation method significantly improved the bone-forming rate in vivo in a rabbit intraperitoneal chamber model. The selective cell retention technology developed by Muschler was facilitated by a novel, single-use, disposable device that could be used at the point of care (Figure 50.6). Prototypes of this device were tested in many preclinical studies, as described below. Muschler et al. utilized the selective cell retention technology to create bone grafts for posterolateral fusion in dogs. In a preliminary study, the authors tested the hypothesis that the biologic milieu of a bone marrow clot significantly would improve the efficacy of such a graft. An established posterior spinal fusion model was used to compare cell-enriched cancellous bone alone, cancellous bone plus a bone marrow clot, and a cell-enriched cancellous bone plus bone marrow clot. Results from union score, quantitative computed tomography, and mechanical testing all demonstrated that the bone matrix plus enriched bone marrow clot was superior to all other groups. These data also confirm that cell enrichment significantly improved graft performance. In a subsequent study, the actual cell concentrate was compared directly to whole bone marrow. Groups included (1) matrix alone (demineralized cortical bone powder), (2) matrix plus marrow, and (3) matrix with enriched marrow cells. Enriched matrix grafts delivered a mean of 2.3 times more cells and approximately 5.6 times more progenitors than matrix mixed with bone marrow. Again, union scores and fusion volumes both confirmed that selective cell retention improved healing outcomes (Muschler et al., 2005). Using the same selective cell retention methodology, Brodke et al. used a canine critical-sized segmental defect to evaluate the healing efficacy of cell-enriched grafts versus autograft (Brodke et al., pending). Canine demineralized bone matrix (cDBM) and cancellous chips were enriched in osteoprogenitors and placed in 21-mm long osteoperiosteal femoral defects for 16 weeks, at which point the animals were sacrificed and the femurs removed and analyzed (Figure 50.7). The results showed equivalency between both the cell-enriched grafts and autograft. Both resulted in 100% bridging bone across the defect spans. Histology sections also demonstrated bone formation across the defects in all autograft and cell-enriched cases (Figure 50.8).

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Figure 50.6 Photograph of a bone marrow osteogenic cell concentration point-of-care device (Cellect™). This device was developed to allow medical professionals to intra-operatively concentrate nucleated stem cells from the bone marrow on bone grafting materials.

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Figure 50.7 Radiographs of a canine critical size femoral defect: (a) immediately post-operative; (b) 16-week post-operative without any graft; (c) 16-week post-operative treated with autograft; and (d) 16-week post-operative treated with cell-enriched canine demineralized allograft. In this study, autograft performed similarly to cell-enriched allograft (Brodke et al., 2006).

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Figure 50.8 Histological micrographs of canine femoral defects 16-week postsurgery: treated with: (a) autograft and (b) cell-enriched canine allograft. Bone trabeculae can bee seen throughout the defects in both cases. In addition, evidences of remodeling can be observed in the cell-enriched graft can be observed. Arrows indicate where the defect was created (Brodke et al., 2006). Finally, a sheep model was also used to evaluate the healing potential of cell-enriched grafts using autologous bone marrow. In this model, the bone grafting efficacy of tri-calcium phosphate (CaP) grafts in three different configurations was evaluated: (1) CaP alone; (2) CaP saturated with whole bone marrow, and (3) CaP enriched with osteoprogenitor cells. In this model, CaP enriched with osteoprogenitor cells reached 33% fusion while autograft only fused at 25%. CaP alone did not result in any fusions and CaP saturated with whole bone marrow only reached 8% fusion (Gupta et al., 2004). These results indicate once more that osteoprogenitor enrichment resulted in increased osteogenicity. These favorable results seemed to imply that other cell concentration methods, as obtained using, for example, in vitro cell culture, might also provide positive outcomes in vivo. These prospects led to the use of culture-expanded grafts for preclinical in vivo testing. Culture-Expanded Grafts As discussed previously, MSCs can be isolated and expanded in vitro. Preliminary research in bone tissue engineering involved the use of autologous, culture-expanded MSCs. These cells would be first harvested from a patient, culture expanded and implanted back into the same patient. This strategy was described in multiple publications, of which two are described in detail below. Bruder et al. (1998) investigated the ability of MSC loaded implants to repair canine femoral defects. The healing of a 21-mm osteoperiosteal defect was studied using ceramic implants loaded with autologous cultureexpanded MSCs at a density of 7.5  106 cells/ml, and compared those to defects left empty. At 16 weeks, atrophic non-union occurred in all defects left empty. In contrast, radiographic union was established rapidly at the interface between the host bone and the implants in samples that had been loaded with MSCs. A large collar of bone formed around the implants; this collar became integrated and contiguous with a callus that formed in the region of the periosteum of the host bone. The collar of bone remodeled during the study ultimately resulting in a size and shape that was comparable with that of the segment of bone that had been resected. Culture-expanded autologous MSCs were also evaluated by Fialkov et al. (2003) using a polylactideco-glycolide (PLGA) foam in critical size rabbit defects. After 8 weeks in vivo, quantitative and qualitative assessments confirmed bone formation in the critical sized defects filled with cell-enriched grafts, while limited bone formation was observed in the animals implanted with foams alone. While the tissue engineering strategy of re-implanting autologous culture-expanded cells seemed successful in vivo, other more cost- and time-effective venues involving allogeneic stem cells were explored. This

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strategy would alleviate the need for a primary cell biopsy from a patient followed by a cell growth phase, since off-the-shelf cell concentrates would be available at all times. In vivo studies thus evaluated the use of allogeneic and xenogeneic cells for bone formation. Discoveries that MSCs did not express immunologically relevant cell surface markers further heightened enthusiasm for this strategy. In a recent study, Arinzeh et al. investigated the effectiveness of allogeneic MSCs to heal a criticalsized bone defect in the femoral diaphysis in dogs without the use of immunosuppressive therapy. Similar to that used by Bruder et al., a critical-sized segmental bone defect of 21 mm in length was created in the midportion of the femoral diaphysis of 12 adult dogs. Each defect was treated with allogeneic MSCs loaded onto a hollow ceramic cylinder such that a complete mismatch between donor stem cells and recipient dogs was achieved. For defects treated with allogeneic mesenchymal stem cell implants, no adverse host response could be detected at any time point. Histologically, no lymphocytic infiltration occurred and no antibodies against allogeneic cells were seen. In addition, at 16 weeks, new bone had formed throughout the cell-enriched implants. These results demonstrated that allogeneic MSCs loaded on ceramic implants did not generate an immune response and were effective for bone repair (Arinzeh et al., 2003). The positive results obtained in preclinical settings with either plain bone marrow or cell-enriched marrow prompted surgeons to use bone marrow in clinical applications. This has been facilitated by the fact that no regulatory or indeed, risk/complication concerns deterred surgeons from this procedure. As a result, while no long-term, multi-center, prospective, randomized, blinded clinical studies have been completed on the efficacy of bone marrow versus autograft, multiple clinical reports have described the use of bone marrow and bone marrow cell-enriched grafts to improve bone healing. Clinical Studies Autologous Bone Marrow with Optimized Matrices In one of the first published studies, Salama and Weissman (1978) published a preliminary reported on 28 patients undergoing long-bone repair under conditions covering a wide range of indications. Bone marrow was implanted and in all cases provided very satisfactory results. Connolly et al. injected bone marrow directly into bone grafting sites, thereby alleviating the need for open surgery. Comparing healing patterns in 100 patients, the study reported an 80% healing rate following marrow grafting (Connolly, 1998). In a clinical study for collagen–calcium phosphate graft material (Collagraft®), Chapman et al. (1997) described the efficacy of bone marrow with the carrier in the treatment of long-bone fractures. No significant difference between the autograft and the bone marrow carrier groups was observed. In a more challenging clinical application, Garg et al. (1993) reported the use of percutaneous autogenous bone marrow grafting in 20 cases of non-united fracture. After 5 months, 17 cases progressed to healing. Delayed union and non-union cases were also treated with bone marrow by Sim et al. (1993), who described 11 cases that healed within a median time of 10 weeks following injection of bone marrow. Bone marrow with allograft was used in a variety of pediatric cases, including cysts, fibromas, long-bone non-unions, and tibia lengthening procedures (Wientroub et al., 1989). In this study, all cases showed good new bone formation with no adverse reaction. In a subsequent pediatric tibial non-ossifying fibromas case, Tiedeman et al. (1991) also reported successful healing after injection of demineralized bone powder with autologous bone marrow. Grafting bone marrow was also proven effective in medically compromised patients, for example, cancer patients, with delayed union or non-unions. Healey et al. (1990) reported bone marrow injections in eight patients with primary sarcomas. Bone formation was observed in seven patients after marrow injection, while complete healing was observed in five patients. Healey et al. concluded that these encouraging results warranted further clinical studies and that his findings suggested a useful technique for the treatment of delayed unions and non-unions in difficult clinical circumstances.

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In clinical spinal fusion applications, bone marrow combined with either DBM or osteoconductive substrates also resulted in improved bone fusion rates. Results similar to autograft were observed in a retrospective posterior spinal fusion study examining 88 consecutive patients and comparing: (1) autologous iliac crest bone graft, (2) freeze-dried corticocancellous bone without marrow, and (3) DBM plus autologous bone marrow. Success rates of 88% and 89% were observed in the autologous iliac crest bone graft and the DBM plus autologous bone marrow groups, respectively. The highest failure rate (28%) was obtained in the freeze-dried corticocancellous bone without marrow group. The authors concluded that augmentation of demineralized bone with bone marrow resulted fusion rates similar to those of iliac crest bone graft (Price et al., 2003). Similar findings have been reported with the use of synthetic grafting materials. Bone marrow aspirate combined with a mineralized collagen matrix (HEALOS®) produced similar fusion rates to those observed with autograft in a posterior spine fusion study (Kitchel et al., pending). Cell-Enriched Grafts The concept of using cell-enriched grafts was initially based on the assumption that a minimum number of osteoprogenitor cells was required to successfully form bone, and that in some severely compromised patients, this cell number may not be reached using whole bone marrow. This hypothesis was investigated by Hernigou et al. (2005), who evaluated 60 non-union patients implanted with concentrated autologous bone marrow. In this study, concentration was achieved by centrifugation, and a 4.2-fold cell concentration ratio was typically achieved (from 612 134 progenitors/cm3 before concentration to an average of 2,579 1,121 progenitors/cm3 after concentration). Union was obtained in 88% cases (53 patients), and the bone marrow that had been injected into the non-unions of those patients contained an average of 54,962 17,431 progenitors, or more than 1,500 progenitors/cm3. In contrast, the total number of osteoprogenitors (19,324 6,843 or less than 700 progenitors/cm3) injected into the non-union sites of the seven patients who did not heal was significantly lower (p  0.01) than that of patients who healed. Therefore, in this study, a minimum of 1,000 progenitors/cm3 and 30,000 progenitors in total seemed to be required to achieve healing. This study represented the first clinical attempt to quantify the required cell numbers for successful union. Muschler’s selective cell retention technique, based on a 3- to 4-fold increase in osteoprogenitors, was also evaluated in vivo. A pilot clinical trial conducted at the Cleveland Clinic reviewed spinal fusion outcomes of 21 patients that received DBM enriched using Muschler’s technique; 20 out of 21 patients showed radiographic evidence of fusion at 12 months (Lieberman, 2004). This study was followed by a prospective, multi-center, randomized study; 51 patients across 5 centers were included in the study. All underwent one or two-level posterolateral fusions. Grafts were prepared using iliac crest bone marrow aspirate. Selective cell retention of the marrow was prepared on DBM. After 12 months, VAS scores were decreased favorably by an average of 55% for back pain, 58.5% for the right leg pain and 65.7% for the left leg pain. Fusion rates were 84.2% (Wang et al., 2005). While encouraging, there is still a need for additional data to fully demonstrate the potential of cellenriched grafts, and their role in bone graft surgery.

CONCLUSION: FUTURE DEVELOPMENTS AND CHALLENGES The osteogenic potential of bone marrow and its role and efficacy in long-bone repair and spinal fusion procedures has been demonstrated in a large body of preclinical and clinical studies over the past 50 years. Bone marrow, combined with an osteoconductive substrate, was shown to produce fusion rates similar to those reported with the use of iliac crest bone graft. As our understanding of the complex phenomena of bone formation increases, there will be an increasing number of potent grafts and cell therapies available to help in bone-related surgical procedures.

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Short- and medium-term research may include new osteopromotive and osteoinductive growth factors for bone repair and bone fusion, as well as improved delivery systems for existing growth factors. In the long term, focus may be shifted to injectable formulations that can form bone in situ and would altogether alleviate the need for invasive surgeries. Another example of ongoing research includes so-called “cell-painting” technologies, which involve introduction of specific proteins on the outer surface of selected cells to guide and target them to a defect site (Dennis et al., 2004; Caplan, 2005). These powerful technologies, as well as advances in gene therapy, may address severe pathologies for which no other satisfactory cure currently exists.

REFERENCES Arinzeh, T.L., Peter, S.J., Archambault, M.P., van den Bos, C., Gordon, S., Kraus, K., Smith, A. and Kadiyala, S. (2003). Allogeneic mesenchymal stem cells regenerate bone in a critical-sized canine segmental defect. J. Bone Joint Surg. 85A: 1927–1935. Barnes, B., Boden, S.D., Louis-Ugbo, J., Tomak, P.R., Park, J.S., Park, M.S. and Minamide, A. (2005). Lower dose of rhBMP-2 achieves spine fusion when combined with an osteoconductive bulking agent in non-human primates. Spine 30(10): 1127–1133. Betz, O., Vrahas, M., Baltzer, A., Lieberman, J.R., Robbins, P.D. and Evans, C.H. (2005). Gene transfer approaches to enhancing bone healing. In: Lieberman, J.R. and Friedlaender, G.E. (eds.), Bone Regeneration and Repair. Totowa, NJ: Humana Press, pp. 158–162. Boyan, B.D., Caplan, A.I., Heckman, J.D., Lennon, D.P., Ehler, W. and Schwartz, Z. (1999). Osteochondral progenitor cells in acute and chronic canine nonunions. J. Orthop. Res. 17: 246–255. Brodke, D., Pedrozo, H.A., Kapur, T.A., Attawia, M., Kraus, K.H., Holy, C.E., Kadiyala, S. and Bruder, S.P. (2006). Bone grafts prepared with selective cell retention technology heal canine segmental defects as effectively as autograft. J. Orthop. Res. 24(5): 857–866. Bruder, S.P. and Caplan, A.I. (2000). Bone regeneration through cellular engineering. In: Lanza, R.P., Langer, R. and Vacanti, J. (eds.), Principles of Tissue Engineering. San Diego, CA: Academic Press, pp. 683–693. Bruder, S.P., Jaiswal, N. and Haynesworth, S.F. (1997). Growth kinetics, self-renewal, and the osteogenic potential of purified human mesenchymal stem cells during extensive subcultivation and following cryopreservation. J. Cell. Biochem. 64: 278–294. Bruder, S.P., Jaiswal, N., Ricalton, N.S., Mosca, J.D., Kraus, K.H. and Kadiyala, S. (1998). Mesenchymal stem cells in osteobiology and applied bone regeneration. Clinical Orthopaedics & Related Research 355(Suppl): S247–S256. Burwell, R.G. (1964). Studies in the transplantation of bone VII. The fresh composite homograft–autograft of cancellous bone: an analysis of factors leading to osteogenesis in marrow transplants and in marrow-containing bone grafts. J. Bone Joint Surg. Br. 46(1): 110–140. Cancedda, R., Quarto, R., Bianchi, G., Mastrogiacomo, M. and Muraglia, A. (2004). Engineered cells in scaffolds heal bone. In: Sandell, L.J. and Grodzinsky, A.J. (eds.), Tissue Engineering in Musculoskeletal Clinical Practice. Rosemont, IL: AAOS, p. 115. Caplan, A.I. (2005). Mesenchymal stem cells: cell-based reconstructive therapy in orthopedics. Tissue Eng. 11(7–8): 1198–1211. Chapman, M.W., Bucholz, R. and Cornell, C. (1997). Treatment of acute fractures with a collagen–calcium phosphate graft material. J.Bone Joint Surg. 79A(4): 495–502. Connolly, J.F. (1995). Injectable bone marrow preparations to stimulate osteogenic repair. Clin. Orthop. 313: 8–18. Connolly, J.F., Guse, R., Lippiello, L. and Dehne, R. (1989). Development of an osteogenic bone-marrow preparation. J. Bone Joint Surg. Am. 71: 684–691. Cannolly, J.F. (1998). Clinical use of marrow osteoprogenitor cells to stimulate osteogenesis. Clinical Orthopaedics & Related Research. 355(Suppl.): S257–S266. Cook, S.D., Wolfe, M.W., Salkeld, S.L. and Rueger, D.C. (1995). Effect of recombinant human osteogenic protein-1 on healing of segmental defects in non-human primates. J. Bone Joint Surg. Am. 77: 734–750.

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Cowan, C.M., Shi, Y.Y., Aalami, O.O., Chou, Y.F., Carina, M., Thomas, R., Quarto, N., Contag, C.H., Wu, B. and Longaker, M.T. (2004). Adipose-derived adult stromal cells heal critical-size mouse calvarial defects. Nat. Biotechnol. 22(5): 560–567. De Kok, I.J., Peter, S.J., Archambault, M., Van den Bos, C., Kadiyala, S., Aukhil, I. and Cooper, L.F. (2003). Investigation of allogeneic mesenchymal stem cell-based alveolar bone formation: preliminary findings. Clin. Oral Implant. Res. 14(4): 481–489. Den Boer, F.C., Wippermann, B.W., Blokhuis, T.J., Patka, P., Bakker, F.C. and Haarman, H.J. (2003). Healing of segmental bone defects with granular porous hydroxyapatite augmented with recombinant human osteogenic protein-1 or autologous bone marrow. J. Orthop. Res. 21: 521–528. Dennis, J.E., Cohen, N., Caplan, A.I. and Goldberg, V.M. (2004). Targeted delivery of progenitor cells for cartilage repair. J. Orthop. Res. 22: 735. Dragoo, J.L. (2003). Bone induction by BMP-2 transduced stem cells derived from human fat. J. Orthop. Res. 21(4): 622–629. Fernyhough, J.C., Schimandle, J.J., Weigel, M.C., Edwards, C.C. and Levine, A.M. (1992). Chronic donor site pain complicating bone graft harvesting from the posterior iliac crest for spinal fusion. Spine 17(12): 1474–1480. Fialkov, J.A., Holy, C.H., Shoichet, M.S. and Davies, J.E. (2003). In vivo bone engineering in a rabbit femur. J. Cranifac. Surg. 14(3): 324–332. Friedenstein, A.J., Petrakova, K.V., Kurolesova, A.I. and Frolova, G.P. (1968) Heterotopic transplants of bone marrow. Analysis of precursor cells for osteogenic and hematopoietic tissues. Transplantation 6(2): 230–247. Gamradt, S.C. and Lieberman, J.R. (2003). Bone graft for revision hip arthroplasty. Clin. Orthop. Relat. Res. 417: 183–194. Gamradt, S.C. and Lieberman, J.R. (2004). Genetic modification of stem cells to enhance bone repair. Ann. Biomed. Eng. 32: 136–147. Garg, N.K., Gaur, S. and Sharma, S. (1993). Percutaneous autologenous bone marrow grafting in 20 cases of ununited fracture. Acta. Orthop. Scand. 64(6): 671–672 Goldberg, V.M. and Stevenson, S. (1987). Natural history of autografts and allografts. Clin. Orthop. 225: 7–16. Gupta, A.R., Shah, N.R., Patel, T.C. and Grauer, J.N. (2001). Perioperative and long-term complications of iliac crest bone graft harvesting for spinal surgery: a quantitative review of the literature. Int. Med. J. 8(3): 163–166. Gupta, M.C., Theerajunyaporn, T., Schmid, M.B., Holy, C.E., Kadiyala, S. and Bruder, S.P. (2004). Use of mesenchymal stem cells enriched grafts in an ovine posterolateral lumbar spine model. IMAST. Healey, J.H., Zimmerman, P.A., Jessop, A.B., McDonnel, M. and Lane, J.M. (1990). Percutaneous bone marrow grafting of delayed union and non-union in cancer patients. Clin. Orthop. Relat. Res. 256: 280–285. Hernigou, Ph., Poignard, A., Beaujean, F. and Rouard, H. (2005). Percutaneous autologous bone-marrow grafting for nonunions: influence of the number and concentration of progenitor cells. J. Bone Joint Surg. Am. 87: 1430–1437 Huard, J. and Peng, H. (2004). Induction of bone formation by stem cells. In: Sandell, L.J. and Sandell, A.J. (eds.), Tissue Engineering in Musculoskeletal Clinical Practice. Rosemont, IL: AAOS, p. 131. Irie, K., Zalzal, S., Ozawa, H., McKee, M. and Nanci, A. (1998). Morphological and immunocytochemical characterization of primary osteogenic cell cultures derived from fetal rat cranial tissue. Anat. Rec. 252(4): 554–567. Karp, J.M., Shoichet, M.S. and Davies, J.E. (2003). Bone formation on two-dimensional poly(DL-lactide-co-glycolide) (PLGA) films and three-dimensional PLGA tissue engineering scaffolds in vitro. J. Biomed. Mater. Res. A 64(2): 388–396. Kevy, S.V. and Jacobson, M.S. (2004) Comparison of methods for point of care preparation of autologous platelet gel. J. Extra-Corp. Technol. 36(1): 28–35. Kitchel, S.H. (2006). A preliminary comparative study of radiographic results using mineralized collagen and bone marrow aspirate vs. autologous bone in the same patients undergoing posterior lumbar interbody fusion with instrumented posterolateral lumbar fusion. Spine J.: Official Journal of the North American Spine Society 6(4): 405–411. Kraiwattanapong, C., Boden, S.D., Louis-Ugbo, J., Attallah, E., Barnes, B. and Hutton, W.C. (2005). Comparison of Healos/bone marrow to INFUSE(rhBMP-2/ACS) with a collagen-ceramic sponge bulking agent as graft substitutes for lumbar spine fusion. Spine 30(9): 1001–1007. Lennon, D.P., Haynesworth, S.E., Bruder, S.P., Jaiswal, N. and Caplan, A.I. (1996). Human and animal mesenchymal progenitor cells from bone marrow: identification of serum for optimal selection and proliferation. In Vitro Cell. Dev. Biol. Anim. 32: 602–611.

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Lieberman, I. (2004). Local cell delivery strategies. Bone Summit, Cleveland, OH. Liu, L.S., Thompson, A.Y., Heidaran, M.A., Poser, J.W. and Spiro, R.C. (1999). An osteoconductive collagen/hyaluronate matrix for bone regeneration. Biomaterials 20(12): 1097–1108. Lou, J. (2004). Bone engineering with mesenchymal stem cells and gene therapy. In: Sandell, L.J. and Grodzinsky, A.J. (eds.), Tissue Engineering in Musculoskeletal Clinical Practice. Rosemont, IL: AAOS, p. 123. Maniatopoulos, C., Sodek, J. and Melcher, A.H. (1988). Bone formation in vitro by stromal cells obtained from bone marrow of young adult rats. Cell Tissue Res. 254: 317–330. Minamide, A., Yoshida, M., Kawakami, M., Yamasaki, S., Kojima, H., Hashizume, H. and Boden, S.D. (2005). The use of cultured bone marrow cells in Type I collagen gel and porous hydroxyapatite for posterolateral lumbar spine fusion. Spine 30(10): 1134–1138. Muschler, G.F. and Midura, R.J. (2002) Connective tissue progenitors: practical concepts for clinical applications. Clin. Orthop. 395: 66–80. Muschler, G.F., Boehm, C. and Easley, K. (1997) Aspiration to obtain osteoblast progenitor cells from human bone marrow: the influence of aspiration volume. J. Bone Joint Surg. Am. 79(11): 1699–1709. Muschler, G.F., Nitto, H., Boehm, C.A. and Easley, K.A. (2001). Age-and gender-related changes in the cellularity of human bone marrow and the prevalence of osteoblastic progenitors. J. Orthop. Res. 19(1): 117–125. Muschler, G.F., Nitto, H., Matsukura, Y., Boehm, C., Valdevit, A., Kambic, H., Davros, W. Powell, K. and Easley, K. (2005). Selective retention of bone marrow-derived cells to enhance spinal fusion. Clin. Orthop. Rel. Res. 432: 242–251. Musgrave, D.S., Fu, F.H. and Huard, J. (2002). Gene therapy and tissue engineering in orthopedic surgery. J. Am. Acad. Orthop. Surg. 10: 6–15. Niemeyer, P., Seckinger, A., Simank, H.G., Kasten, P., Sudkamp, N. and Krause, U. (2004). Allogenic transplantation of human mesenchymal stem cells for tissue engineering purposes: an in vitro study. Orthopade 33(12): 1346–1353. Nussenbaum, B. and Krebsbach, P.H. (2004). Practical matters in the application of tissue engineered products for skeletal regeneration in the head and neck region. In: Sandell, L.J. and Grodzinsky, J. (eds.), Tissue Engineering in Musculoskeletal Clinical Practice. Rosemont, IL: AAOS, p. 154. Ohgushi, H., Goldberg, A.I. and Caplan, A.I. (1989). Repair of bone defects with marrow cells and porous ceramic. Experiments in rats. Acta. Orthop. Scand. 60: 334–339. Olmsted-Davis, E.A., Gugala, Z., Gannon, F.H., Yotnda, P., McAlhany, R.E., Lindsey, R.W. and Davis A.R. (2002). Use of a chimeric adenovirus vector enhances BMP2 production and bone formation. Human Gene Ther. 13(11): 1337–1347. Owen, M. (1985). Lineage of osteogenic cells and their relationship to the stromal system. In: Peck, W.A. (ed.), Bone and Mineral. Amsterdam: Elsevier, pp. 1–25. Pereira, R.F., Halford, K.W., O’Hara, M.D., Leeper, D.B., Sokolov, B.P., Pollard, M.D., Bagasva, O. and Prockop, D.J. (1995). Cultured adherent cells from marrow can serve as long-lasting precursor cells for bone, cartilage, and lung in irradiated mice. Proc. Natl Acad. Sci. USA 92(11): 4857–4861. Price, C.T., Connolly, J.F., Carantzas, A.C. and Ilyas, I. (2003). Comparison of bone grafts for posterior spinal fusion in adolescent idiopathic scoliosis. Spine 28(8): 793–798. Salama, R. and Weissman, S.L. (1978). The clinical use of combined xenografts of bone and autologous red marrow. J.Bone Joint Surg. 60B(1): 111–115. Sim, R., Liang, T.S. and Tay, B.K. (1993). Autologous marrow injection in the treatment of delayed and non-union in long bones. Singapore Med. J. 34: 412–417. Sutherland, D. and Bostrom, M. (2005). Grafts and bone graft substitutes. In: Lieberman, J.R. and Friedlaender, G.E. (eds.), Bone Regeneration and Repair. Totowa, NJ: Humana Press, pp. 133–136. Takagi, K. and Urist, M.R. (1982). The role of bone marrow in BMP-induced repair of femoral massive diaphyseal defects. Clin. Orthop. Rel. Res. 171: 224–231. Tay, B.K., Le, A.X., Heilman, M., Lotz, J. and Bradford, D.S. (1998). Use of a collagen–hydroxyapatite matrix in spinal fusion. Spine 23(21): 2276–2281. Thomas, C.B., Kellam, J.F. and Burg, K. (2004). Comparative study of bone cell culture methods for tissue engineering applications. J. ASTM Int. 1: 1–17. Tibor, T.G. (2003). Short overview of potential gene therapy approaches in orthopedic spine surgery. Spine 28(3): 207–208.

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Tiedeman, J.J., Connolly, J.F., Strates, B.S. and Lippiello, L. (1991). Treatment of nonunion by percutaneous injection of bone marrow and demineralized bone matrix. An experimental study in dogs. Clin. Orthop. 268: 294–302. Tiedeman, J.J., Huurman, W.W., Connolly, J.F. and Strates, B.S. (1991). Healing of a large nonossifying fibroma after grafting with bone matrix and marrow. Clin. Orthop. 265: 302–305. Urist, M.R. (1965). Bone: formation by autoinduction. Science 150: 893–899. Vinay, P., Gougoux, A. and Lemieux, G. (1981). Isolation of a pure suspension of rat proximal tubules. Am. J. Physiol. 241: F403–F411. Wang, J.C., Youssef, J.A., Lieberman, I.H., Brodke, D.S., Lauryssen, C., Haynesworth, S.E. and Muschler, G.F. (2005). A prospective, multi-center study of selective osteoprogenitor cell retention for enhancement of lumbar spinal fusion. IMAST, Banff, Canada. Webster, T.J., Ergun, C., Doremus, R.H., Siegel, R.W. and Bizios, R. (2000). Enhanced functions of osteoblasts on nanophase ceramics. Biomaterials, 21(17): 1803–1810. Wenisch, S., Trinkaus, K., Hild, A., Hose, D., Herde K., Heiss, C., Kilian, O., Alt, V. and Schettler, R. (2005). Human reaming debris: a source of multipotent stem cells. Bone, 36(1): 74–83. Whang, P.G. and Lieberman, J.R. (2003). Clinical issues in the development of cellular systems for use as bone graft substitutes. In: Laurencin C.T. (ed.), Bone Graft Substitutes. West Conshohocken, PA: ASTM International, pp. 142–155. Wientroub, S., Goodwin, D., Khermosh, O. and Salama, R. (1989). The clinical use of autologous marrow to improve osteogenic potential of bone grafts in pediatric orthopedics. J. Pediatr. Orthop. 9(2):186–190. Yoon, S.T. and Boden, S.D. (2002). Osteoinductive molecules in orthopaedics: basic science and preclinical studies. Clin. Orthop. 395: 33–43. Younger, E.M. and Chapman, M. (1989). Morbidity at bone graft donor sites. J. Orthop. Trauma 3: 192–195. Zhang, M., Powers, R.M. and Wolfinbarger, L. (1997). Effect(s) of the demineralization process on the osteoinductivity of demineralized bone matrix. J. Periodontol. 68: 1085–1092. Zuk, P.A., Zhu, M., Ashjian, P., De Ugarte, D.A., Huang, J.I., Mizuno, H., Alfonso, Z.C., Fraser, J.K., Benhaim, P. and Hedrick, M.H. (2002). Human adipose tissue is a source of multipotent stem cells. Mol. Biol. Cell 13: 4279–4295.

51 Cell-Based Therapies for Musculoskeletal Repair Wan-Ju Li, Kiran Gollapudi, David P. Patterson, George T.-J. Huang, and Rocky S. Tuan INTRODUCTION Common musculoskeletal disorders include osteoarthritis (OA), rheumatoid arthritis (RA), intervertebral disk (IVD) degeneration, anterior cruciate ligament (ACL) injuries, and muscular dystrophy. These diseased conditions can result from trauma, work-related injuries, immunological malfunction, aging, or genetics. Arthritis is highly prevalent among adults. As the population ages, arthritis is expected to affect an estimated 67 million adults in the United States by 2030, according to the report by Centers for Disease Control and Prevention (CDC, 2006). Approximately 13–16 million people are diagnosed with OA per year in the US or Europe, and joints are replaced due to OA at the rate of one every 1–2 min. A total of 300,000 and 500,000 joint replacements are performed per year in Europe and the US, respectively. Although OA is not normally life threatening, it is progressive, disabling, and can greatly impact an individual’s quality of life. Current approaches to musculoskeletal care emphasize prevention, medical treatment, and surgical intervention as a final resort. Progress has been made in the past decades with biological therapeutic approaches to reverse or slow disease progression by specifically targeting molecules involved in the disease process. Despite advances in these approaches, once tissue damage reaches a certain stage, self-repair does not take place. Particularly with cartilage, the repair mechanism is almost non-existent. Currently, there is no available remedy to repair the eroded cartilage in the arthritic joints and the degenerated IVD. While joint replacements provide significant functional restoration and symptomatic improvement, they are compromised by the possibility of prosthetic failure and associated complications (e.g. peri-implant osteolysis). The emerging disciplines of cell-based therapy and tissue engineering have suggested the prospect of regenerative medicine as a promising approach to the treatment of damaged and diseased musculoskeletal tissues. This chapter will outline mesenchymal cell biology in the context of musculoskeletal tissues, including osteogenec, chondrogenec, myogenec, tenogenec, and ligament cell types. Particular emphasis will be made on the definition of mesenchymal stem cells (MSCs), their niches, isolation and in vitro characterization, lineage differentiation and regulation, and immunomodulatory properties. Cell-based applications using these cells to produce specific tissues will be reviewed, covering clinical disorders, gene therapy approaches, in vivo studies, and current applications, as well as potential pitfalls and future improvement. BIOLOGY OF CELLS IN MUSCULOSKELETAL TISSUES Osteoblasts and Osteocytes Bone provides structural support for the body, facilitates movement, protects internal organs, and acts as a mineral reserve. Derived from embryonic mesoderm, bone contains a matrix consisting of hydroxyapatite

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mineral and macromolecular extracellular matrix (ECM) components. Osteoblasts (OB) are the bone forming cells found on all surfaces of bone. They produce osteoid consisting of collagen type I, fibronectin, proteoglycans, and other specialized proteins, which mineralizes by the deposition of hydroxyapatite crystals [Ca10(OH)2(PO4)6]. Osteocytes are former OB that have become encased by the bone matrix and are the most abundant cells in mature bone. Their presence highlights the fact that bone is a living, dynamic tissue. Osteocytes cease to produce osteoid but communicate with other cells (osteocytes and OB) through small pores called canaliculi, and are believed to play a role in mechanotransduction activities of bone with feedback to the remodeling process. Osteoprogenitors, located in the bone marrow and periosteum, can be induced to become OB via growth factors, such as the transforming growth factor-β (TGF- β) superfamily and in particular the bone morphogenetic proteins (BMPs). BMPs regulate chemotaxis, mitosis and differentiation, and are critical in initiating fracture healing. The transcription factors Cbfa1/Runx2 and Osterix are essential for OB differentiation in both intramembranous and endochondral ossification. Runx2 is involved in activating gene expression of collagen type I and other bone proteins, such as osteopontin and osteocalcin. Osterix is believed to be downstream of Runx2 as it is not expressed in Runx2 null mice. Other transcription factors shown to be involved in OB proliferation and differentiation include Msx and Dlx proteins. However, mice containing null alleles for these genes do produce bone, unlike Runx2 and Osterix null mice (Eames et al., 2003). Chondrocytes There are three main types of cartilage: hyaline cartilage, elastic cartilage, and fibrocartilage. The external ear is an example of elastic cartilage, and the meniscus and annulus fibrosus of the IVD, which will be discussed later, are examples of fibrocartilaginous tissues. Articular cartilage is an example of hyaline cartilage, the most common type. The major function of articular cartilage is to provide a smooth surface for reduced friction and to support large loads during movement. Articular cartilage consists of a fluid phase composed of water and electrolytes as well as a solid phase consisting of ECM and chondrocytes. Chondrocytes, making up less than 10% by volume of articular cartilage, are the only cell type in the tissue and are responsible for the maintenance of the ECM. Collagen type II is the predominant collagen in articular cartilage and is responsible for its tensile strength; however, other minor collagens, such as collagen types V, VI, IX, and X, are also present. The most abundant component of articular cartilage is water (60–85% by volume), which is held in place by the highly charged proteoglycans, and allows for its compressive behavior. Thus, collagens and proteoglycans provide cartilage with its mechanical properties, and their content is a key component in assessing the functional quality of engineered articular cartilage. During developmental chondrogenesis, mesenchymal cells are recruited and migrate to areas of chondrogenesis, and subsequent mesenchymal–epithelial cell and cell–cell matrix interactions promote cellular condensation. Mesenchymal cells then differentiate into chondrocytes, heralding the deposition of ECM proteins and the activation of transcription factors such as Sox-9. Sox-9 is required for the expression of cartilagespecific collagen type II in normal skeletal development (Bi et al., 2001). L-Sox-5 and Sox-6, which are other members of the Sox family, cooperate with Sox-9 to turn on the collagen type II gene and are also essential for cartilage formation (Lefebvre and Smits, 2005). Muscle Cells Muscle is a specialized contractile tissue, derived from embryonic mesoderm and allows for force exertion and locomotion. There are three types of muscle: skeletal (voluntary) muscle found in the musculoskeletal system, smooth muscle (involuntary) found within walls of organs and structures, and cardiac muscle, a specialized tissue found only in the heart. Skeletal muscle cell differentiation from embryonic mesoderm is driven by

889

890 THERAPEUTIC APPLICATIONS: CELL THERAPY

transcription factors MyoD, Myf5, myogenin, Mrf4, and MEF2 in a highly coordinated fashion (Sartorelli and Caretti, 2005). Skeletal muscle is composed of bundles of dense muscle fibers that are highly oriented and able to generate longitudinal contraction. These bundles (muscle cells) are multinucleated and derived from myoblasts (muscle cell precursors). Satellite cells are specialized myoblast populations capable of muscle regeneration. They can be obtained and cultured from muscle biopsy. Upon initiation by local growth factors, these normally quiescent and undifferentiated cells become mitotic, differentiate, and eventually self-assemble into muscle fibers themselves. M-Cadherin, Pax7, and neural cell adhesion molecule (NCAM) are known satellite cell markers that can be used to localize and follow satellite cells in vivo. Pax7 null mice lacked satellite cells but retained a unique population of interstitial stem cells in muscle that express the stem cell markers, CD 34, and Sca-I (Tamaki et al., 2002). These stem cells, tracked by expression of green fluorescent protein (GFP), were found to originate in the bone marrow (Dreyfus et al., 2004). Tendon and Ligament Cells Tendons and ligaments are dense fibrous structures that connect muscle to bone and bone to bone, respectively. Both are composed mostly of collagen type I, produced by specialized elongated fibroblasts (known as tenocytes in tendons) that lie between the collagen fibers. These tissues are very hypocellular compared to other connective tissues, presenting problems for injury repair. Both arise from mesodermal compartments distinct from those that give rise to myogeneic cells. Not many cellular markers have been identified, but the transcription factor Scleraxis appears promising as a specific and early marker for tendons and ligaments (Tozer and Duprez, 2005).

EMBRYONIC AND ADULT STEM CELLS Embryonic Stem Cells Embryonic stem (ES) cells are pluripotent cells, derived experimentally from the inner cell mass of the embryonic blastocyst. Human ES (hES) cells are typically obtained from 4- to 5-day-old blastocysts of embryos after in vitro fertilization. These cells are potentially immortal in vitro without loss of differentiation potential, and when reimplanted into a host embryo, they give rise to pluripotent daughter cells that differentiate into all tissue types. The use of hES cells for research and clinical applications is complicated by controversies surrounding the legal and ethical status of human embryos and is currently restricted by regulations on federal funding. Despite these challenges, both mouse and hES cells have been examined for applications in musculoskeletal regeneration, albeit limited. Human ES cell differentiation and organization can be influenced by a supportive three-dimensional (3D) environment such as poly(lactic-co-glycolic acid) (PLGA) polymer scaffolds and directed by growth factors such as retinoic acid, TGF-β, activin-A, or insulin-like growth factor (IGF). These growth factors induce differentiation into 3D structures with characteristics of developing neural tissues, cartilage, or liver, respectively (Levenberg et al., 2003). hES cells have been injected into the joint space of immunocompromised rats to promote cartilage repair (Wakitani et al., 2004). Osteogeneic potential of hES cells in the presence of chemical stimuli in vitro has also been demonstrated (Karp et al., 2006). In a mouse model, chondrogene differentiation was observed using mES cells via embryoid bodies (EBs) modulated by members of the TGF-β family (TGF-β1, BMP-2 and -4) (Kramer et al., 2000). mES cells differentiate into chondrocytes, which progressively develop into hypertrophic and calcifying cells. At a terminal differentiation stage, cells expressing an OB-like phenotype appear either by transdifferentiation from hypertrophic chondrocytes or directly from OB precursor cells. Under the influence of ascorbic acid, β-glycerophosphate, and 1,25-dihydroxy vitamin-D3, mES cells are

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induced to differentiate in vitro into OB that produce a mineralized matrix. In 3D scaffold systems such as hydrogels, both hES and mES cells were shown to have chondrogene capacity (Elisseeff et al., 2006). Mesenchymal Stem Cells While ES cells hold promise as a cell source for tissue regeneration, ethical issues and the possibility of teratoma formation need to be resolved before they can be used in clinical applications. Notably, our body has other stem cell populations that possess the capabilities of self-renewal and multidifferentiation to repair damaged tissues or maintain tissue homeostasis when the repair is needed. With our improved knowledge of stem cells combined with advances in culture techniques, it is possible that we can harness their potential for the treatment of degenerative musculoskeletal diseases. Unlike pluripotent ES cells which have the ability to form tissues from all three germ layers, adult stem cell populations are believed to be limited in their differentiation capacity. In general, stem cells derived from a particular tissue are programmed to differentiate into various progenies that belong to the same developmental germ layer origin. Although recent studies suggest that adult stem cells can differentiate across germ line boundaries, it remains debatable if differentiation plasticity between the cells of different germ layers actually exists or is simply an artifact resulting from contamination of heterogeneous cell populations or cell fusion. More evidence is needed to assess the possibility of differentiation across the three germ layers. However, plasticity within a germ layer is more strongly supported by evidence from several research groups. The ethically acceptable nature of adult stem cells, combined with proven differentiation abilities, as compared to ES cells and tissue-committed cells, makes them an attractive option for use in cell-based therapy. For cell-based musculoskeletal tissue regeneration, adult MSCs are an attractive candidate progenitor cell type since they may be isolated from various adult tissues, and can differentiate into different mesenchymal lineage cells, such as bone, cartilage, fat, muscle, ligament, tendon, and stroma (reviewed by Tuan et al., 2003) (Figure 51.1). Reflecting their origin and cell functions, these cells have also been named or described as

MSC

Proliferation Biomaterial scaffold interaction

Growth factor induction

Differentiation gene regulation

Mechanical stimulation Differentiation

OB

CC

AC

MB

TC

SM

Figure 51.1 Multidifferentiation potential of MSCs and factors regulating their biological activities in vitro. CC, chondrocyte; AC, adipocyte; MB, myoblast; TC, tenocyte; SM, stroma.

892 THERAPEUTIC APPLICATIONS: CELL THERAPY

marrow stromal cells, or mesenchymal stromal cells, which are all abbreviated as MSCs. MSCs were first discovered by a German pathologist in the 1860s and he described the cell morphology as “fibroblast-like.” More than a century later, in 1976, Friedenstein further identified MSCs as colony-forming unit-fibroblasts, which were able to commit to osteogene differentiation. The multidifferentiation potential of MSCs was demonstrated in vivo as early as in 1980, in which MSCs were induced to become bone and cartilage. Caplan (1991) and Pittenger et al. (1999) showed that MSCs underwent osteogene, chondrogene, and adipogene differentiation in response to different biochemical signals. Numerous subsequent studies have produced a significant body of evidence demonstrating the phenotypic and functional characteristics of MSCs. According to the statistic data retrieved from Medline, research publications containing the key word “MSC” have increased 15-fold in the past 5 years, highlighting the increasing interest in MSC studies. While many studies are focused on characterization and differentiation potential of MSCs, few are concentrated on the molecular regulation of MSCs. Future MSC research should focus on the intrinsic and extrinsic mechanisms mediating the molecular switch between undifferentiated and differentiated MSCs. MSC Isolation MSCs can be isolated from a variety of mesenchymal tissues such as bone marrow, fat, trabecular bone, cartilage, muscle, peripheral blood, and umbilical cord blood (reviewed by Tuan et al., 2003). Dependent on the species and tissue types, different isolation protocols and culture methods have been developed. Among these MSC sources, bone marrow is the best studied tissue. The isolation process for bone marrow-derived MSCs includes several steps aimed at reducing contamination by other cell types. Erythrocytes can be removed by density gradient centrifugation using Percoll or Ficoll after bone marrow aspirate is obtained from the iliac crest, tibia, or femur. Fluorescence-activated cell sorting (FACS) and magnetic-activated cell sorting (MACS) are sometimes used to select a more defined cell population based on cell surface markers. With their ability to adhere to tissue culture plastic, MSCs are further discriminated from non-adherent hematopoietic cells after several medium changes. The low frequency of 1 MSC per 10,000–100,000 bone marrow cells in vitro indicates that the MSC is a rare cell population (Pittenger et al., 1999). Previous studies have shown that MSC yield is affected by age and health of a donor. The trend is that MSC yield decreases with donor age. Patients with degenerative diseases, such as osteoporosis and OA, tend to have lower MSC yield. Unfortunately, it is this group of people who would benefit most from MSC-based treatment. Therefore, an alternative could be the use of allogeneic MSCs. Although immune reaction is a concern associated with using allogeneic cells, the finding that MSCs have low immunogeneic potential as well as immunosuppressive properties suggests that this concern may not be significant. Immunoregulation of MSCs will be discussed in detail later. Another MSC source gaining recent attention is adipose tissue, since fat is abundant, easy to access, and considered surgical waste during a cosmetic surgery operation. The procedure of marrow harvesting is painful, and over-harvesting bone marrow may be a risk to the patient’s health. In contrast, liposuction is less painful and considered a relatively safe procedure. The method to isolate MSCs is similar to a typical primary cell isolation, in which MSCs are released from collagenase-digested fat and purified after being plated in plastic culture to remove unattached hematopoietic cells. A study comparing MSC proliferation between bone marrow and fat tissue suggests that fat tissue is a more effective source than bone marrow for MSC yield (Lee et al., 2004). As mentioned above, bone marrow-derived MSC number, proliferation, and differentiation potential may decrease with the donor age. A possible solution is to isolate MSCs from tissue of a younger donor. Fetal tissues, such as umbilical cord blood (Mareschi et al., 2001), cord vein (Romanov et al., 2003), placenta (Fukuchi et al., 2004), and amniotic fluid (In’t Anker et al., 2003), have been processed to isolate MSCs. A comprehensive study by Kern et al. (2006) in which they compared the morphology and functions of MSCs

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isolated from bone marrow, adipose tissue, and umbilical cord blood showed that MSCs isolated from these three tissue sources are morphologically and immunophenotypically similar. However, umbilical cord bloodderived MSCs form the fewest colonies and show the highest proliferation capacity, whereas adipose tissuederived MSCs form the greatest number of colonies and bone marrow-derived MSCs show the lowest proliferation capacity. The findings suggest both adipose tissue and umbilical cord blood can be used as alternatives to bone marrow for MSC isolation. Recently, various tissue-specific adult stem cells have been successfully isolated and characterized from non-mesenchymal tissues. These cells from ectoderm (epidermis) or endoderm origin (pancreas) morphologically and functionally behave similarly to bone marrow-derived MSCs. Although the lack of specific markers makes verification of their identity as MSCs difficult, and their phenotype could be a result of in vitro culture conditions, the possibility of MSCs residing in different tissues throughout the body raises a number of interesting questions. For example, how and where do these cells reside in the different tissues? Is there a common tissue pool, such as bone marrow, that houses the MSCs which are able to integrate to different tissues in response to specific biological needs or activities? In Vitro MSC Behavior After being plated in culture, MSCs adhere to the substratum and start to divide, forming colonies. The typical growth pattern is that MSCs experience a few days of a lag phase before undergoing a log phase of growth and then reach a stationary phase. Generally, these cells can undergo 50 population doublings in 10 weeks without losing their multidifferentiation capability. Colter et al. (2000) reported that the initial cell seeding density affects the expansion capacity and doubling time of mouse MSCs. When plated at low initial plating density (1.5– 3 cells/cm2), MSCs have 2,000-fold expansion in 10 days, whereas when plated at high density (12 cells/cm2), the cell number only increases 60-fold in the same period of time. In addition, the doubling time increases from 12 to 24 h with high cell seeding density. MSCs are a heterogeneous population, demonstrated by the finding that there are two MSC morphologies in a colony; small spindle-shaped and large flat fibroblast-like cells. Interestingly, a third type of extremely small, rapidly self-renewing (RS) cells was recently identified in MSC colonies. During the three phases of growth, RS cells give rise to large flat cells in the log phase. It is believed that the large flat MSCs are more mature cells, replicating slowly, whereas small RS cells proliferate rapidly. MSC Identification Despite a significant number of studies, to date, there is no specific marker available for the identification of MSCs. There have been some cell surface markers available for MSC identification, but none are exclusive to MSCs. Because MSCs can not be positively identified, their biological activities in vivo cannot be verified as with hematopoietic stem cells (HSCs). Our knowledge of MSCs is thus primarily from in vitro experimental results. The in vitro results of MSC identification are likely dependent on culture environment and purity of cell population, which could explain the differences between the findings reported in the literature. Nevertheless, the use of multiple markers, such as cell surface cluster of differentiation (CD) markers, ECM proteins, cell adhesion molecules, integrins, and cytokines as well as genetic or proteomic fingerprinting can help one identify MSCs. CD cell surface molecules are the most commonly used markers to identify MSCs. Both positive and negative CD markers have been used to identify MSCs and exclude endothelial cells, HSCs, and hematopoietic lineage cells. Positive MSC markers include Stro-1, SH2 (CD105), SH3 (CD73), SH4 (CD73), CD 29, CD 44, CD 54, CD 90, CD 105, CD 133, CD 166, and p75LNGFR, whereas negative markers are CD 11, CD 14, CD 19, CD 31, CD 34, CD 45, CD 79, and HLA-DR (Deans and Moseley, 2000). To shorten the list, the International Society for Cellular Therapy (ISCT) has provided minimum criteria for defining MSCs

894 THERAPEUTIC APPLICATIONS: CELL THERAPY

(Dominici et al., 2006). Acceptable MSCs meet the minimum requirements of CD 73, CD 90, and CD 105 positive and CD 14, CD 34, CD 45, and HLA-DR negative expression. MSC Niche The term MSC niche refers to both physical and chemical environment that MSCs reside in, which includes other cell types, ECM molecules directly or indirectly contacting with MSCs, and soluble factors regulating MSC activities. The cellular and non-cellular components interact with each other in this highly complex 3D environment, responsible for the maintenance of MSC stemness properties as well as the regulation of symmetric and asymmetric cell division. The concept of the stem cell niche was first introduced in the 1970s (Schofield, 1978) and has been elucidated by in vitro co-culture experiments. Ball et al. (2004) demonstrated that, by co-culturing MSCs and endothelial cells, MSCs were induced to a phenotype similar to smooth muscle cells, whereas MSCs became myofibroblast-like cells when co-cultured with dermal fibroblasts. Their study suggests that the interaction between MSCs and their neighboring cells may regulate the fate of MSCs and that the type of progeny of MSCs may be determined by the interacting cells in the niche. The bone marrow microenvironment is a principal MSC niche in the body. It is not only a complex 3D structure, but also allows for many interactions between cellular and non-cellular components, such as HSCs, MSCs, stroma, hematopoietic cells, mesenchymal origin cells, ECM components, growth factors, and cytokines. Tellingly, the success rate of HSC engraftment improves when co-transplanting with MSCs in vivo. The ex vivo expansion of HSCs increases dramatically when MSCs are co-cultured with HSCs, suggesting that HSCs and MSCs maintain a close biological interaction in the naïve marrow niche. The signals needed for MSC proliferation and differentiation come from soluble factors as well as both cell–cell and cell–matrix interactions. OB have been known to play an important role in the regulation of hematopoiesis but their role in MSC osteogenesis still remains inconclusively defined. A previously published report shows that OB have synergistic interactions on MSC proliferation and alkaline phosphatase activity but not calcium deposition (Kim et al., 2003). In comparison, our preliminary results show that MSC co-culturing with OB enhances their osteogene differentiation (unpublished observation). MSC–ECM interactions, both physical and chemical, are likely to be critical in the regulation of MSC physiology in the niche. Matsubara et al. (2004) cultured and maintained MSCs on basement membrane-like ECM and observed profound effects on MSC proliferation and differentiation. With the support of basement membrane-like ECM, MSCs better maintain their multidifferentiation potential after many cell divisions, which suggests that the interactions between MSCs and basement membrane-like ECM may recapitulate some of the MSC–ECM interactions in bone marrow. Cell–ECM interactions are also the center of the study in biomaterial-based cell therapy and tissue engineering. Artificial ECM, a biomaterial scaffold used to replace damaged or malfunctioning ECM, is designed to fully function as native ECM and interact with MSCs for successful tissue regeneration. The interactions of MSC and a biomaterial scaffold will be discussed later. MSC Regulation Soluble factors such as growth factors and cytokines play a significant role in physiological regulation of MSCs in the niche. These biochemical signals guide MSCs either to stay as undifferentiated cells or to differentiate into tissue-specific progenitor cells by activating specific signal pathways. A recent important finding is the involvement of the Wnt signaling pathway. Gregory et al. (2003) demonstrate that MSCs enter the cell cycle and inhibit osteogeneic differentiation after Dickkopf-1 (Dkk-1) deactivates the Wnt pathway. Boland et al. (2004) further identify that Wnt 3a working through the canonical pathway promotes MSC proliferation but discourages osteogenesis, whereas Wnt 5a via the non-canonical pathway promotes osteogenesis. Chondrogenesis of MSCs also involves the activation of the Wnt signaling pathway (Tuli et al., 2003). During

Cell-Based Therapies for Musculoskeletal Repair 895

chondrogene differentiation of MSCs, TGF-β1 activates the mitogen activated protein (MAP) kinase pathway which is demonstrated to be involved in cross-talk signaling with the Wnt signaling pathway. The regulation of the Wnt signaling pathway likely induces chondrogenesis by enhancing cell–cell interaction through N-cadherin expression. The most important and valuable characteristic of MSCs is their multipotential differentiation capacity. Previous studies have shown that MSC differentiation can be induced and regulated by soluble signal factors, both protein- and non-protein-based molecules. TGF-βs and BMPs induce MSCs to undergo chondrogenesis in a serum-free medium. Each member of the TGF-β family has a different level of induction efficiency. TGF-β3 has a higher efficiency in inducing chondrogenesis compared to TGF-β1 but they both contribute similarly to chondrogenesis in long-term culture. To induce osteogenesis, β-glycero-2-phosphate, ascorbic acid, dexamethasone, and 1,25-dihydroxy vitamin-D3 are required to enhance alkaline phosphatase activity and matrix mineralization. For adipogenesis, MSCs are treated with isobutyl-1-methylxanthine and insulin, resulting in adipocytes with the presence of lipid droplets in the cytoplasm (Pittenger et al., 1999). In addition to biochemical factors, physical factors, such as mechanical loading as well as matrix geometry and elasticity, have also been found to play a role in MSC biology. Cells generally receive and transduce physical cues from the surrounding environment to the cell nucleus through cytoskeletal changes or signaling pathways. McBeath et al. (2004) demonstrated that cell shape regulates commitment of MSCs by using a micropatterned substrate to control cell shape and size of cultured MSCs. They found that cell shape is a key regulator in MSC differentiation with the shape-dependent control of lineage commitment mediated by ROCK-mediated cytoskeletal tension. MSCs forced to spread and flatten in the large substrate pattern differentiate into OB, whereas those forced to unspread and become round in the small pattern differentiate into adipocytes. Recently, Engler et al. (2006) demonstrated the effect of matrix stiffness on differentiation of MSCs, mediated by non-muscle myosin II. MSCs on a soft matrix with stiffness similar to brain stiffness differentiate into neurons, on a matrix with intermediate stiffness close to muscle stiffness become myoblasts, and on a stiff matrix comparable to bone commit to OB. In addition, they also show that typical soluble factors for lineage commitment do not alter the differentiation lineage previously activated by matrix stiffness. This study suggests matrix stiffness appears to be important in MSC lineage commitment. MSC Immunoregulation Immunoregulation by and of MSCs can be viewed from two perspectives: (1) immunosuppressive effects of allogeneic MSCs, and (2) inflammatory cytokine effect on MSC activity and differentiation. Due to the interest in using allogeneic or xenogeneic MSCs to compensate for the paucity and time constraints associated with expanding autologous MSCs, there has been considerable progress in the understanding of the MSC immunoregulatory effect. While xenogeneic MSCs are rejected by the host after transplantation, allogeneic MSCs are well tolerated by the recipient hosts. Many in vivo studies have confirmed the immunosuppressive effects of MSCs (Chen XI et al., 2006). The potential mechanisms underlying this immunosuppression can be explained by downregulation of T, dendritic, natural killer (NK), and B cells. This immunosuppressive characteristic suggests that MSCs can be potentially used in vivo for enhancing the engraftment of other tissues (e.g. HSCs), or for prophylactic prevention and even possibly as a treatment of graft-versus-host-disease or autoimmune diseases such as RA (Jorgensen et al., 2003). Limited information is available on the effects of pro-inflammatory cytokines on MSCs. Liu and Hwang (2005) demonstrated that human cord blood-derived MSCs secrete cytokines and growth factors. Most importantly, continuous supplementation of IL-1β in the cord blood-derived-MSC culture facilitates adipogenec maturation of cord blood-MSCs. A preliminary study using porcine MSCs showed that interferon-α-2b may act to differentiate MSCs into OB (Abukawa et al., 2006). Our recent study suggests that MSCs are relatively

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resilient to pro-inflammatory cytokines in terms of apoptotic response (Okafor et al., 2006). In the context of autologous implantation for chondrogenesis, a study has shown that inflammatory reactions against scaffold materials and serum components lead to the production of cytokines such as IL-1α that may inhibit cartilage tissue formation (Rotter et al., 2005). These few studies suggest the importance of understanding the effect of tissue inflammation, either caused by diseases or in response to scaffold materials after implantation, on the differentiation and cell behavior of MSCs.

CELL-BASED THERAPIES AND TISSUE ENGINEERING Tissue Engineering Three general strategies have been adopted for the engineering of new tissues: (1) delivery of cells, (2) local or systemic delivery of tissue inducing agents, and (3) delivery of biomaterial scaffolds containing both cells and inductive agents. The use of a biomaterial scaffold-based carrier facilitates the delivery of therapeutic cells or agents to the target site and subsequent growth and regeneration of new tissue. A general strategy is that target cells (differentiated/undifferentiated) expanded in vitro are cultured in 3D, biomaterial scaffolds (natural/synthetic) under conditions that favor the desired phenotype. After the appropriate culture period, the cell-seeded composite exposed to biological and physical stimuli develops into a natural tissue-like cellular construct (Figure 51.2). The biomaterial scaffolds are believed to play a critical role in the tissue engineering process, by providing a 3D structure for cellular functions such as attachment, migration, proliferation, and differentiation. The ultimate success of this process is determined by the biological and functional similarity of the engineered tissue to native tissue. Despite the promising prospects of tissue engineering, regenerating tissues that serve a predominantly biomechanical function, such as bone, articular cartilage, and tendon, presents significant challenges. The

Tissue culture expansion

Autologous, allogenec tissue donor

ES, MSC, Tissue-specific cells

Rotatory wall vessel, spinner flask, perfusion pump bioreactor

Natural or synthetic fiber, gel, foam, sphere

Signaling

Growth factor, cytokine, non-peptide agent, physical stimulus

Figure 51.2 General strategy of cell-based therapy and tissue engineering.

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concept of “functional tissue engineering” has emerged, specifically prescribing the production of a tissue that meets functional and in our case especially, mechanical requirements of the repair tissue. This requires a clear delineation of the structure and functions of living tissues, design of biomaterial scaffolds, and mechanical optimization of regenerative tissues. It is necessary to incorporate each of these steps to regenerate a mechanically sound tissue for safer and more efficacious repairs and replacements for the patient. Gene Therapy Gene therapy involves the gene-based modification of cells for the correction of defective genes or the regulation of gene expression. It has a great potential in cell therapy-based treatments due to the fact that a number of musculoskeletal diseases are caused by genetically malfunctioning cells; thus successful tissue engineering strategies must include the repair of the “defective” cells in any engineered construct. The self-renewal capacity and multipotentiality of MSCs suggest their suitability for cell-based and gene therapy applications in regenerative medicine. For gene therapy, viral transduction has a high efficiency of delivering genes into MSCs. Adeno-associated viral-mediated gene transfer has been tested to be effective in delivering genes into MSCs and to repair bone disorders such as osteogenesis imperfecta (Chamberlain et al., 2004). Lentiviral vectors have also been shown to be effective in delivering genes into MSCs (Lu et al., 2005). Non-viral methods such as transfection using Nucleofection™ has been demonstrated to be promising in delivering functional genes into MSCs (Haleem-Smith et al., 2005). We will now discuss the application of these cell-based strategies towards regenerating musculoskeletal tissues. A brief list of representative studies in the field of musculoskeletal regeneration is provided in Table 51.1. Bone A wide variety of patients with significant bone defects that necessitated amputation in the past now benefit from various orthopedic strategies. Congenital defects of bone, growth plate fractures and defects, fractures resulting in malunion or non-union, the genetic disorder osteogenesis imperfecta (brittle bone disease), and bone loss from tumor resection (primary bone tumors or tumors metastatic to bone) are just a handful of musculoskeletal problems that could be addressed by regenerative medicine. Currently, using orthopedic prosthetics is a severe but highly functional option. Distraction osteogenesis, a surgical procedure for bone reconstruction and lengthening, was developed in the 1950s by Ilizarov and is still used today. Bone autografting is a therapeutic option developed in the 19th century and considered to be the current gold standard, but has limitations, particularly in the size of the defect to be grafted. Autografts contain the patient’s own OB and osteocytes, but require a second surgical site for the bone harvesting, most often the iliac crest of the pelvis. This increases patient morbidity such as post-operative pain and risk of infection. Allograft bone from bone banks or cadavers avoids the pitfalls of autograft bone but does not possess the strength or cellularity of autograft bone. As the number of surgeries requiring bone grafting continues to rise, the development of functional tissue-engineered bone grafts becomes increasingly significant. Four critical factors to successful bone tissue engineering are osteoconduction, osteoproduction, osteoinduction, and mechanical stimulation. Osteoconduction refers to the integration of the scaffold or graft material into the site and its eventual remodeling and replacement. Osteoproduction is the production of bone material by cells, and osteoinduction is the use of growth factors that draw additional osteogene cells to the site. For both in vivo and ex vivo bone tissue, mechanical stimulation appears to be a critical factor in the development of biologically and mechanically optimal bone tissue. Tissue engineering of bone must take into account the tremendous mechanical strength and elasticity of bone . For load-bearing long bones such as the femur, mechanical stability of the construct is crucial, whereas for finer tissues such as fingers or

Tissue type

Model

Cells

Strategy and observation

Reference

Bone

Rat Human Human Human

MSCs MSCs MSCs MSCs

Cells transduced with BMP-2 improved healing of a critical defect Porous ceramic seeded with MSCs repaired large bone defects Direct grafting of cells to non-union defects achieved union Cells with platelet-rich plasma were polymerized and used successfully for alveolar graft osteoplasty

Lieberman et al. (1999) Quarto et al. (2001) Hernigou et al. (2005) Hibi et al. (2006)

Growth plate

Rabbit

MSCs

Direct loading of cells into growth plate defects reduced growth arrest in the tibia

Chen et al. (2003)

Cartilage

Human

Chondrocytes

Brittberg et al. (1994)

Pig

Chondrocytes

Human

MSCs

Autologous cells injected into deep cartilage defects produced good to excellent results in 14 of 16 patients Cells seeded in gelatin microbeads mixed with type I collagen gel repaired full-thickness cartilage defects Collagen gel containing cells implanted into cartilage defects improved arthroscopic and histologic scores

Rat

MSCs

Dog

Mesangioblasts

Tendon

Rabbit

MSCs

Collagen gels seeded with MSCs implanted into patellar tendon defects showed improved biomechanics

Awad et al. (2003)

Meniscus

Rat

MSCs

MSCs embedded in fibrin glue contributed to the healing of meniscal defects

Izuta et al. (2005)

Intervertebral disk

Human

Chondrocytes

Autologous cells delivered to the nucleus pulposus improved pain and disability scores

Meisel et al. (2006)

Muscle

Myogenic-induced cells were directly injected to damaged muscle, contributing to muscle repair Vascular delivery of wild-type mesangioblasts led to significant clinical amelioration of muscular dystrophy

Chiang et al. (2005) Wakitani et al. (2002) Dezawa et al. (2005) Sampaolesi et al. (2006)

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Table 51.1 Representative reports of in vivo cell-based musculoskeletal repair

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craniofacial applications, plasticity takes an increased significance. Tissue-engineered bone used for clinical applications should meet both biological and mechanical requirements. Various scaffold strategies have been used for MSC-based bone tissue engineering. Matching the strength of bone is a leading concern, and many strategies have been employed. Fully or partially demineralized bone matrix (DBM) from processed allograft bone contains collagen, growth factors, and other proteins and has been seeded with MSCs to create promising engineered constructs. DBM shares many structural and functional similarities to autologous bone and, as expected, supported osteogene function of MSCs (Mauney et al., 2004). Coral, composed mostly of calcium carbonate and with a similar structure to bone, has been seeded with periosteum as a therapeutic strategy (Vacanti et al., 2001). Porous ceramics, such as those composed of tricalcium phosphate and hydroxyapatite, have been used in conjunction with MSCs to produce bone replacement tissues successfully in patients who failed traditional therapies (Quarto et al., 2001). Optimization of the scaffold strategy will require understanding the mechanism of its action. DBM is capable of withstanding shear forces and does not impair elasticity in the implant, and partially mimics the autologous environment in bone, although allogeneic antigens and pathogens may not have been fully removed. Ceramics provide good osteoconductivity and good integration into the defect site by bonding to tissues without rejection or inflammatory reactions, but unfortunately lack tensile strength, limiting applications involving torsion, shear stress, or bending. Natural coral has been investigated for decades as a bone graft substitute, and is biocompatible, osteoconductive and biodegradable. Improved ex vivo construct manufacturing requires combining biomaterial strategies with bioreactors that can produce shear and compressive forces to provide a dynamic culture system. Dynamic culture of cell-seeded scaffolds, for example, using spinner flasks, has been shown to result in more even cell distribution and a 121% increase in cell density (Mauney et al., 2004). Direct cell therapy has also been tested for musculoskeletal applications. Percutaneous autologous bone-marrow grafting, the re-introduction of aspirated bone marrow directly to the site of a non-union in the tibia, has been described in human patients with good results (Hernigou et al., 2005). Growth plate (physis) injuries in children can result in shortening or angular deformity with the formation of bony bridges across the growth plate between the epiphysis and metaphysis. Direct implantation of MSCs into growth plate defects resulted in a significant reduction of growth arrest in rabbit tibia (Chen et al., 2003). Gene therapy also holds promise for bone tissue engineering. A number of strategies have been tested for bone repair. Proof of concept was established with improved healing of a critical defect in a rat femur with delivery of rat MSCs transduced with the gene for BMP-2 to the site of the defect (Lieberman et al., 1999). In mice, it was demonstrated that systemically injected MSCs transduced with IGF-1 established themselves in bone marrow. The MSCs demonstrated chemotactic ability by responding to the local fracture environment and locating preferentially to the fracture site, where they also accelerated the healing process (Shen et al., 2002). Gene therapy with an MSC-based vehicle is also being harnessed for the treatment of a genetic disease. Engineered adeno-associated viral vectors were successfully used to disrupt the expression of mutated collagen type I gene in MSCs derived from individuals with osteogenesis imperfecta. Subcutaneous implantation of transduced MSCs into immunodeficient mice produced improved bone matrix (Chamberlain et al., 2004). If host MSCs could be augmented or replaced, future OB could then produce osteoid of higher quality. A gene therapy approach has also been developed with muscle-derived mesenchymal progenitor cells. These cells act as vehicles producing osteoinductive proteins and have been demonstrated to heal critically sized bone defects (Young et al., 2002). Mouse primary myoblasts over-expressing Runx2 via aretroviral system were implanted in conjunction with collagen scaffolds in the hind legs of mice and resulted in trabecular bone growth (Gersbach et al., 2006). Future improvement of bone tissue engineering depends critically on understanding the biological signals necessary for bone induction and optimizing the pharmacokinetics of their delivery. Optimized vascularization

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is essential, as cell labeling experiments show a considerable loss of OB in the first week following transplantation in porous cancellous bone matrices, presumably due to suboptimal initial vascularization (Kneser et al., 2006). Scaffolds that incorporate growth factors such as vascular endothelial growth factor (VEGF) as well as endothelial cells have been shown to increase vascular formation in constructs in vivo, but integration with the host vascular system remains a challenge (Rouwkema et al., 2006). Articular Cartilage The demand for engineered articular cartilage arises from the prevalence of traumatic injuries and degenerative diseases of articular cartilage such as OA. Other than total joint arthroplasty, which is associated with surgical risks as well as a finite life span, current therapeutic modalities for OA patients include pharmacologic intervention, lavage, shaving, laser abrasion, drilling or microfracture of subchondral bone to stimulate healing, autologous periosteal and perichondrial grafting, autogeneic or allogeneic osteochondral transplantation, and autologous chondrocyte implantation. These procedures, although variably effective, often cannot repair cartilage to a disease-free state. With an increasing elderly population and the predicted rise in the incidence of OA, novel cell-based therapies are a promising avenue to meet the therapeutic needs of these patients. Ultimate success of a tissue-engineered cartilage construct requires the presence of cells which can produce the proper cartilaginous ECM. In principle, fully differentiated chondrocytes are the ideal cell candidate, since they are programmed for the cartilage phenotype. Indeed, Brittberg et al. (1994) showed that full thickness chondral defects could be repaired using autologous chondrocytes derived from a minor load-bearing area and injected under a periosteal flap. This procedure, commonly referred to as autologous chondrocyte implantation/transplantation, is marketed as Carticel™ by Genzyme Biosurgery and is the only Food and Drug Administration (FDA) approved cell-based therapy for cartilage repair. This therapy has had promising results; however, the cost effectiveness as well as the superiority of this procedure over other available procedures is debatable. Furthermore, the use of chondrocytes is hampered by their limited availability, the potential donor site morbidity, and the propensity for chondrocytes to dedifferentiate when in monolayer culture. Alternatively, stem cells, including ES cells and MSCs, may be induced to differentiate into the chondrogene lineage. EB-derived ES cells cultured as pellets with TGF-β3 showed increased gene expression of cartilagespecific ECM markers as well as a significant increase in collagen and proteoglycan production after 14 days compared to untreated controls. These changes toward a chondrogenec phenotype were accompanied by a downregulation of hematopoietic and neural markers. MSCs derived from various sources present another stem cell source for cartilage repair. Wakitani et al. (2002) showed that in OA patients undergoing high tibial osteotomy, autologous bone marrow MSCs seeded in a collagen gel had histological and arthroscopic improvement compared to controls; however, there was no statistically significant clinical improvement. A variety of both natural and synthetic scaffolds have been tested to serve as carriers for the aforementioned cells (reviewed by Kuo et al., 2006). The advantage of using a carrier scaffold is demonstrated by a study by Chiang et al. (2005) in which chondrocytes seeded in gelatin microbeads mixed with a collagen type I gel improved repair in a porcine cartilage defect model and maintained a chondrocyte phenotype better than cells delivered alone. Examples of scaffolds that have been studied for cartilage tissue engineering include fibrous scaffolds of biodegradable polymers such as poly-glycolic acid (PGA), poly-lactic acid (PLA), and their copolymer PLGA, as well as fibrin, agarose, collagen, alginate, gelatin, poly-ethylene glycol (PEG), and hyaluronanbased gels. Combinations of these materials have also been used. However, the currently available scaffolds do not fully meet the necessary requirements and can have variable effects on cell behavior and function. Novel techniques are being developed to engineer scaffolds with biomimetic properties to optimize biomaterial framework for cartilage regeneration. For example, bovine chondrocytes cultured in a PEG-based hydrogel cross-linked with a matrix metalloproteinase (MMP) sensitive peptide, which allows for remodeling by the

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seeded cells, had increased expression of collagen type II and aggrecan, compared to hydrogels lacking the MMP sensitive peptide (Park et al., 2004). In addition, electrospun nanofibers, structurally similar to natural ECM, have recently been developed as a novel scaffold for tissue engineering. MSCs seeded into a poly(ε-caprolactone) (PCL) nanofibrous scaffold and treated with TGF-β1 were able to differentiate into a chondrocytic phenotype (Li et al., 2005b). More recently it has been shown that changing the fiber diameter in these scaffolds can influence chondrocyte proliferation and ECM matrix production, demonstrating that even slight dimensional changes in scaffolds of the same material can affect seeded cells (Li et al., 2006). Supplementation of growth medium with specific signaling molecules and growth factors is commonly used to promote cell proliferation, differentiation, and ECM production. These molecules include members of the TGF-β superfamily (TGF-βs, BMPs, and growth differentiation factors (GDFs), IGFs, platelet-derived growth factors (PDGFs), and Wnts. These factors activate intracellular signaling pathways that are presumably similar to those involved in developmental morphogenesis (reviewed by Chen FH et al., 2006). Isoforms of TGF-β as well as BMPs are some of the most potent positive modulators used in tissue engineering. Although TGF-β1 has been widely used as an anabolic factor for both chondrocytes and MSCs, TGF-β2 and TGF-β3 have been shown to be superior inducers of chondrogenesis in MSCs (Barry et al., 2001). Among the BMPs, BMP-2 has been shown to be more effective than other BMPs (-12 and -13, and -4 and -6) for chondrocyte and MSC-based engineered constructs, respectively (Sekiya et al., 2005). Similar to scaffold design, the current trend in growth factor supplementation is to use a combination approach; however, not all combinations are favorable. For example, FGF-2 and IGF-1 showed no synergism, and actually decreased protein synthetic rate by canine articular chondrocytes seeded in a collagen type II–glycosaminoglycan scaffold (Veilleux and Spector, 2005). The challenge of applying soluble factors for cartilage regeneration is that the effects of these factors depend not only on their optimal combination but also on their amount, the timing of administration, and the target cell type. Since cartilage is a tissue whose function and form are related to physical stimuli, recent attempts to optimize the functional properties of engineered cartilage have focused on using bioreactors to incorporate mechanical loading environments in vitro to mimic in vivo environments. Mechanical loading environments tested include dynamic deformation, hydrostatic pressure, fluid flow, and shear stress. Beneficial effects of mechanical loading on cartilage constructs have been widely reported. For example, dynamic deformational loading of cell-seeded agarose disks improved the mechanical properties, as well as the sulfated glycosaminoglycan (sGAG) and hydroxyproline contents, compared to controls (Mauck et al., 2000). Recent evidence also supports a positive role for mechanical loading on MSCs. Human MSCs grown in pellet culture with TGF-β3 showed increased expression of cartilage markers in the presence of cyclic hydrostatic pressure. These positive effects were dependent on both the dose and length of time cells were exposed to loads (Miyanishi et al., 2006). This study underscores the need for further studies to determine the optimal type, timing, and amount of mechanical loading, and to establish the appropriate cells, scaffolds, and soluble factors with which it should be combined in order to regenerate truly functional cartilage. Gene therapy is gaining recognition as a tool for cartilage regeneration. Lapine articular chondrocytes transfected with plasmid vectors expressing IGF-1 were combined with alginate and implanted into a rabbit osteochondral defect model (Madry et al., 2005). IGF-1 expressing implants improved articular cartilage and subchondral bone repair compared to controls. Although gene therapy strategies are promising, long-term studies are needed to evaluate repaired cartilage. It is likely that in the future, optimal cartilage regeneration will result from the combination of genetically modified cells with tissue engineering. Meniscus The menisci are two semilunar fibrocartilaginous structures located between the tibia and femur in the knee joint. The meniscus functions as a shock absorber, joint stabilizer, and joint lubricator. Furthermore, it has an

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outer vascular area which possesses the ability to repair itself and an inner avascular zone where the majority of tears occur. The cells in meniscus have both fibroblast- and chondrocyte-like properties and are termed fibrochondrocytes. Also, the cell phenotype as well as the ECM components vary between the different zones. In addition to having a different cell type, fibrocartilage differs from hyaline cartilage in that the cells predominantly secrete collagen type I and it has a relatively lower sGAG content giving it different mechanical properties. Despite these differences, many attempts at meniscal regeneration have incorporated some of the same scaffolds and growth factors used in articular cartilage repair and have been tested in a variety of animal models. Recent evidence has pointed to the potential for the use of MSCs in meniscal regeneration. GFPlabeled rat MSCs were seeded in fibrin glue and used to treat meniscal defects (Izuta et al., 2005). GFP-positive cells were detected up to 8 weeks post implantation and promoted meniscal repair. Successful cell-based meniscal repair would be of substantial therapeutic value, but further optimization of regenerative conditions and factors is still needed. Another challenge for meniscus tissue engineering is to regenerate a meniscus exhibiting significant anisotropic mechanical properties reflective of a highly oriented underlying ECM. We recently applied electrospinning technology to fabricate a biomaterial scaffold mimicking meniscal ECM fiber alignment, which directs fibrochondrocyte orientation and has controllable, anisotropic properties. These findings suggest the potential application of aligned-nanofiber-based scaffolds for meniscal tissue engineering (Li et al., 2007). Osteochondral Tissue Defects of both articular cartilage and the underlying subchondral bone are often associated with pain and joint instability, risk factors for the development of OA. Using a tissue-engineered osteochondral graft is a promising alternative to the current use of autologous osteochondral grafts which are limited by tissue availability, lack of appropriate geometric configuration, and donor site morbidity. Engineering an osteochondral graft has been challenging due to the technical difficulties involved in generating a single unit consisting of two tissues with different properties, which naturally require different conditions for development and optimal functionality. General approaches to osteochondral tissue engineering have involved choosing a scaffold for the bone layer alone, two separate scaffolds for the bone and cartilage layer, or a single scaffold for both layers and combining these scaffolds with one or two cell sources. These cell sources may have either chondrogene, osteogene, or bipotential differentiation capacity. The disadvantage to using two separate scaffolds or cells is that there may be impaired integration between the osteo and chondral components. Various approaches have been employed to circumvent this problem. For instance, a pellet of trabecular bone-derived mesenchymal progenitor cells previously induced to undergo chondrogenesis was press-coated onto a PLA scaffold (Tuli et al., 2004). Progenitor cells from the same patient which were induced to undergo osteogenesis were then seeded onto the other end of the scaffold, followed by culturing under conditions that supported both osteogenesis and chondrogenesis, resulting in an interface resembling the native osteochondral junction. Ideally, a single cell source would be able to differentiate into both lineages on a single scaffold. Recent evidence that electrospun PCL nanofibrous scaffolds allow MSCs to undergo multilineage differentiation suggests that this scaffold is a promising candidate for tissue engineering osteochondral grafts (Li et al., 2005a). Skeletal Muscle While muscle-based disorders are not as prevalent as OA or bone defects, there is a serious clinical need for muscle tissue engineering. These needs include muscle atrophy and muscular dystrophy, as well as muscle loss from trauma or surgical resection. Several disease states, such as cancer, infectious disease, heart failure and AIDS can produce a body wasting syndrome known as cachexia, in which muscle atrophy is severe. Muscular dystrophy is a group of mostly inherited neuromuscular disorders that cause muscle wasting, and can lead to death in patients with severe mutations in the dystrophin gene. For the treatment

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of traumatic muscle loss, free tissue transfer is an option, but autologous muscle transfer causes not only donor site morbidity, but can produce loss of function at the donor site. Requirements of functional engineering of muscle would be the recapitulation of functional motion and integration with host connective tissues. On the cellular level, parallel alignment of fibers and integration of functional neuromuscular junctions are also important to achieve. As with other musculoskeletal tissues, mechanical stimulation is essential during myogenesis, and influences metabolic activity and gene expression, as well as fiber alignment. Also, myoblasts, similar to many types of cells, become difficult to differentiate as they are culture-expanded. Myoblast cell lines, once immortalized, cannot approximate myogenesis as well as primary myoblasts. Therefore, understanding the biology of precursor cell types, such as satellite cells and other muscle-derived stem cells, is crucial. While satellite cells act as the local regenerative cells in muscle, their limited expansion potential in vitro limits their current usefulness relative to MSCs. The therapeutic use of MSCs in muscle disorders was demonstrated when GFP-labeled human bone marrow-derived MSCs were induced to undergo myogenesis and then transplanted by local injection into muscles of immunosuppressed rats. Histological section 4 weeks later showed mature muscle characteristics in most GFP-positive myofibers. In addition, some of the cells appeared to become Pax7 expressing satellite cells, which could respond to local damage and contribute to muscle repair (Dezawa et al., 2005). For genetic disorders such as muscular dystrophy, there have been many attempts to reintroduce a wild type dystrophin gene into muscle via cell-based ex vivo gene therapy. Mesangioblasts, a type of vessel-associated stem cell, from wild type canine were delivered intra-arterially in a muscular dystrophy canine model with significant recovery of dystrophin gene expression, muscle morphology, and muscle function. Autologous genetically corrected mesangioblasts were not as effective (Sampaolesi et al., 2006). As with all tissue engineering, uncovering the optimal growth factors and stimulation environment to effectively produce muscle tissue is essential. As would be expected for the tissue responsible for motor functions, mechanical stimulation is crucial in the development of in vitro functional muscle tissue. Mechanical forces applied in vitro yield significant differences in morphological and functional appearance of muscle tissues. Mean myofiber diameter and elasticity both improve, as well as the ratio of muscle fiber to ECM (Bach et al., 2004). Electrical stimulation mimicking nerve stimulation during myogenesis and regeneration of injured skeletal muscle also drives differentiation. Chronic electrical stimulation can change gene expression of muscle-specific genes, with increased VEGF expression and blood flow also shown after stimulation. Due to the unique contractile requirement of skeletal muscle, optimal scaffold strategies may differ from those of bone or cartilage production. PGA meshes, collagens, and alginates have been used, as in other tissues. Acellular muscle has also been exploited as a potential scaffold, and 3 weeks after seeding with myoblasts, isometric contractile force testing revealed longitudinal contractile forces upon stimulation (Borschel et al., 2004). Additionally, co-culture of fibroblasts with myoblasts has been explored as a way to produce a primary matrix. Tendon and Ligament Injuries to ligaments and tendons heal by forming tissues of inferior quality, due to hypocellularity of the tissue as well as scar formation that is weaker than normal tissue. Current strategies to replace tendons and ligaments consist of tissue autografts and allografts. Both have the traditional problems associated with graft technology, but additionally exhibit slow or poor functional integration to the surgical site, requiring at minimum several months for recovery. In addition, tendons and ligaments are prime examples of how in the musculoskeletal system, mechanical and structural properties are crucial to the function of the tissue. There are significant challenges in developing tissue engineering strategies, due to the fact that using autologous tenocytes can cause a tendon defect, limiting options for cell-based therapy. MSCs have been used to produce engineered tendon and ligament tissues. Collagen gels seeded with autologous MSCs were implanted into full thickness, full length, central defects in patellar tendons of rabbits.

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This approach significantly improved the biomechanical properties of tendon repair tissues as compared to natural repair controls (Awad et al., 2003). In one approach to overcoming poor functional integration of graft tissues, MSCs were used at the tendon–bone junction in tendon grafting for ACL repair in the rabbit knee. This resulted in a zone of fibrocartilage closely resembling that of the normal ACL (Lim et al., 2004). Dermal fibroblasts have also been used to engineer tendon. Autologous dermal fibroblasts, seeded on PGA fibers, were implanted in a porcine model and shared similar tensile strength with constructs developed with autologous tenocytes (Liu et al., 2006). Gene therapy using MSCs as a vehicle was demonstrated with transduction of BMP-12 into a mouse MSC cell line, leading to tendon and cartilage-like tissue formation after injection in the thigh muscle of nude mice (Lou et al., 1999). Tensile strength and stretch loading are believed to be vital in producing the proper alignment of ligament and tendon tissues in ex vivo engineering. It is known that collagen is arranged along the axis of the loading force, implying its requirement for structurally competent tissues (Liu et al., 2006). Tensional and torsional stimulation of MSC-based constructs, mimicking cues the cells receive in vivo, enhanced ECM fiber deposition. Our recent findings reveal that both mechanical and biological stimuli act to regulate differentiation and matrix synthesis and remodeling during MSC tenogenesis (Kuo and Tuan, submitted for publication). Designing bioreactors that replicate the most essential of these conditions is critical to allow for ex vivo production of ready-to-implant tendons and ligaments with increased load to failure over endogenous tissues. Intervertebral Disk Degenerative disk disease is a leading cause of back pain and disability. The IVD has a complex structure comprised of a proteoglycan-rich nucleus pulposus and an outer annulus fibrosus. Disk degeneration is characterized by proteoglycan loss in the nucleus pulposus with concurrent degradation of the annulus fibrosus. As cell-based regenerative efforts for the IVD are still in its infancy, appropriate cell, scaffold, and growth factor combinations are under investigation and have been tested in various animal models. In Europe, based on results from a canine model and pilot studies in humans, a prospective, randomized trial for comparing autologous disk chondrocyte transplantation (ADCT) plus discectomy to discectomy alone is being performed. Nucleus pulposus cells derived from therapeutic discectomy are cultured and delivered 12 weeks after discectomy. An interim analysis after 2 years shows promising results with sustained pain relief and improvement of disability scores in the ADCT patients (Meisel et al., 2006). Autologous disk cells are, however, in limited supply. MSCs have been recently shown to differentiate into cells similar to IVD cells, and present an alternative cell source for regenerative purposes (Steck et al., 2005). In addition, gene therapy strategies can be used to repair IVD, as demonstrated by a study in which adenovirus expressing Sox-9 was used to infect human IVD cells in vivo, and was also injected directly into rabbit IVD following a stab injury (Paul et al., 2003). In vitro, infected disk cells exhibited increased production of collagen type II, and in vivo, IVDs injected with these cells showed decreased scarring compared to those that were untreated or injected with mock adenovirus. Craniofacial Tissues The temporomandibular joints (TMJ) in the craniofacial system, connecting the mandible to the cranium, play a vital role in coordinating our eating and speech activities. Osteoarthritic TMJ is correlated with not only aging but also dental function. Craniofacial defects can result from trauma, neoplasm, and most commonly from infection. Additionally, teeth penetrating through the mucosal epithelial barrier are exposed to the moist harsh environment of the mouth and subject to infection via dental caries and periodontal

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pathogens. Loss of teeth resulting from severe periodontal diseases or root canal infection is a common cause of losing the jaw bone. Recent clinical success of the dental implant renders the need for augmenting lost mandible before implant even more important, because the success of the dental implant is dependent upon sufficient quality bone. Cell-based therapy to regenerate periodontal defects using autologous MSCs has been recently demonstrated in a dog model (Hasegawa et al., 2006), in which 4 weeks after MSC transplantation, the periodontal defects were almost regenerated with periodontal tissue. Cementoblasts, OB, and osteocytes are found in the regenerated periodontal tissue, suggesting that transplanted MSCs could survive and differentiate into periodontal tissue cells and repair the tissue. Similar accomplishments have also been observed in humans except the assessment is only performed clinically with radiographs and periodontal probing (Yamada et al., 2006). More extensive mandibular repair using cell therapy has been recently demonstrated in both animal models and humans. Mandibular distraction osteogenesis is enhanced by direct delivery of MSCs into the defect in rats (Qi et al., 2006). Autologous MSCs from patient’s iliac crest plus platelet-rich plasma (PRP) was applied to reconstruct an alveolar cleft defect or to augment alveolar bone with success (Hibi et al., 2006). In addition, the mixture of MSCs and PRP facilitates osteointegration in dental implants, and may replace the use of autologous particulate cancellous bone and marrow for the same purpose (Yamada et al., 2004). Ex vivo gene therapy for de novo jaw bone regeneration has been demonstrated in a swine model using adenovirus BMP-2-mediated gene transfer to expanded autologous MSCs. Functional bone capable of sustaining axial compressive loads is formed in the maxillary bone defect filled with the BMP-2-transduced cells cast in collagen gel (Chang et al., 2003). Jaw bone regeneration utilizing recombinant protein therapy is considered more straightforward and simpler. The recombinant human BMP-2/collagen sponge implant converts undifferentiated MSCs into OB and promotes an intense local neovascular response. The type of bone synthesis depends upon the mesenchymal substrate and the local cellular environment. Using this simple technique, bone defects can be resynthesized with good outcomes and a significant reduction in donor site morbidity. A drawback of recombinant protein therapy is that a single dose of exogenous protein is not as robust for osteoinduction when compared to the results from preclinical animal studies (Nussenbaum and Krebsbach, 2006). A proof of concept to regenerate cartilage of TMJ has been demonstrated using rat MSCs to engineer human-shaped mandibular condyle in immunocompromised mice (Alhadlaq and Mao, 2003). Challenges for TMJ tissue engineering include promotion of ECM synthesis and tissue maturation by stem cell-derived chondrogene and osteogene cells that have been encapsulated in biocompatible and bioactive scaffolds. Enhancement of the mechanical properties of a tissue-engineered mandibular condyle for ultimate in situ implantation into the human TMJ is another challenge for tissue engineers. Tooth regeneration with cells from tooth buds, and other dental tissue regeneration including dentin and periodontal ligament using stem cells from the respective tissues, has also been proposed and tested in animal models (Ohazama et al., 2004).

CONCLUSIONS AND PROSPECTS It appears promising that regenerative medicine via cell-based therapy and tissue engineering will become a widespread therapeutic modality. The essential procedures consist of utilizing and manipulating ex vivo expanded multipotent stem cells, and delivering them into hosts under a designed condition or package to grow new musculoskeletal tissues (Figure 51.3). However, there are many challenges before reaching this goal. Specifically, we need to understand (1) the native environment of stem cells, that is, the niche and stem cell markers, to allow stem cell isolation and culture expansion in a more specific and predictable manner; (2) the molecular regulation at various stages of the stem cell differentiation program such that a better control of cell

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MSCs

Nanofibrous scaffold

MSC–nanofiber construct

Cartilage

Bone

Figure 51.3 Production of tissue-engineered cartilage and bone based on MSC-seeded nanofibrous scaffolds. MSCs were cultured in chondrogene and osteogene media for 3 and 6 weeks, respectively. Bar  25 mm. lineage commitment can be achieved; (3) which delivery method or scaffold system is optimal for regeneration of a specific tissue; and (4) the interactions of cells with their carrier or scaffold system and the newly produced ECM, so that optimal tissue regeneration can be accomplished. There is also a need for testing the long-term compatibility of allogeneic MSCs in the host before cell banks can be established to provide an adequate cell source for cell-based applications and tissue engineering. There has been significant experience accumulated in the regeneration of bone, whereas other musculoskeletal tissue regeneration, such as cartilage, ligament, and tendon, is still at its infancy. Another aspect of tissue engineering that requires further development is the design of bioreactors that can replicate conditions to allow ex vivo production of ready-to-implant musculoskeletal tissues, which is critical to achieve functional tissue engineering.

ACKNOWLEDGEMENTS Supported by NIH NIAMS Intramural Research Program (AR Z0141131) and Howard Hughes Medical Institute-NIH Research Scholar Program.

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Li, W.J., Mauck, R.L., Cooper, J.A., Yuan, X. and Tuan, R.S. (2007). Engineering controllable anisotropy in electrospun biodegradable nanofibrous scaffolds for musculoskeletal tissue engineering. J. Biomech. 40: 1686–1693. Lieberman, J.R., Daluiski, A., Stevenson, S., Wu, L., McAllister, P., Lee, Y.P., Kabo, J.M., Finerman, G.A., Berk, A.J. and Witte, O.N. (1999). The effect of regional gene therapy with bone morphogenetic protein-2-producing bone-marrow cells on the repair of segmental femoral defects in rats. J. Bone Joint Surg. Am. 81: 905–917. Lim, J.K., Hui, J., Li, L., Thambyah, A., Goh, J. and Lee, E.H. (2004). Enhancement of tendon graft osteointegration using mesenchymal stem cells in a rabbit model of anterior cruciate ligament reconstruction. Arthroscopy 20: 899–910. Liu, C.-H. and Hwang, S.-M. (2005). Cytokine interactions in mesenchymal stem cells from cord blood. Cytokine 32: 270–279. Liu, W., Chen, B., Deng, D., Xu, F., Cui, L. and Cao, Y. (2006). Repair of tendon defect with dermal fibroblast engineered tendon in a porcine model. Tissue Eng. 12: 775–788. Lou, J., Tu, Y., Ludwig, F.J., Zhang, J. and Manske, P.R. (1999). Effect of bone morphogenetic protein-12 gene transfer on mesenchymal progenitor cells. Clin. Orthop. Relat. Res. 333–339. Lu, F.-Z., Fujino, M., Kitazawa, Y., Uyama, T., Hara, Y., Funeshima, N., Jiang, J.-Y., Umezawa, A. and Li, X.-K. (2005). Characterization and gene transfer in mesenchymal stem cells derived from human umbilical-cord blood. J. Lab. Clin. Med. 146: 271–278. Madry, H., Kaul, G., Cucchiarini, M., Stein, U., Zurakowski, D., Remberger, K., Menger, M.D., Kohn, D. and Trippel, S.B. (2005). Enhanced repair of articular cartilage defects in vivo by transplanted chondrocytes overexpressing insulin-like growth factor I (IGF-I). Gene Ther. 12: 1171–1179. Mareschi, K., Biasin, E., Piacibello, W., Aglietta, M., Madon, E. and Fagioli, F. (2001). Isolation of human mesenchymal stem cells: bone marrow versus umbilical cord blood. Haematologica 86: 1099–1100. Matsubara, T., Tsutsumi, S., Pan, H., Hiraoka, H., Oda, R., Nishimura, M., Kawaguchi, H., Nakamura, K. and Kato, Y. (2004). A new technique to expand human mesenchymal stem cells using basement membrane extracellular matrix. Biochem. Biophys. Res. Commun. 313: 503–508. Mauck, R.L., Soltz, M.A., Wang, C.C., Wong, D.D., Chao, P.H., Valhmu, W.B., Hung, C.T. and Ateshian, G.A. (2000). Functional tissue engineering of articular cartilage through dynamic loading of chondrocyte-seeded agarose gels. J. Biomech. Eng. 122: 252–260. Mauney, J.R., Blumberg, J., Pirun, M., Volloch, V., Vunjak-Novakovic, G. and Kaplan, D.L. (2004). Osteogenic differentiation of human bone marrow stromal cells on partially demineralized bone scaffolds in vitro. Tissue Eng. 10: 81–92. McBeath, R., Pirone, D.M., Nelson, C.M., Bhadriraju, K. and Chen, C.S. (2004). Cell shape, cytoskeletal tension, and RhoA regulate stem cell lineage commitment. Dev. Cell 6: 483–495. Meisel, H.J., Ganey, T., Hutton, W.C., Libera, J., Minkus, Y. and Alasevic, O. (2006). Clinical experience in cell-based therapeutics: intervention and outcome. Eur. Spine J. 15(Suppl 15): 397–405. Miyanishi, K., Trindade, M.C., Lindsey, D.P., Beaupre, G.S., Carter, D.R., Goodman, S.B., Schurman, D.J. and Smith, R.L. (2006). Dose- and time-dependent effects of cyclic hydrostatic pressure on transforming growth factor-beta3-induced chondrogenesis by adult human mesenchymal stem cells in vitro. Tissue Eng. 12: 2253–2262. Nussenbaum, B. and Krebsbach, P.H. (2006). The role of gene therapy for craniofacial and dental tissue engineering. Adv. Drug Deliv. Rev. 58: 577–591. Ohazama, A., Modino, S.A., Miletich, I. and Sharpe, P.T. (2004). Stem-cell-based tissue engineering of murine teeth. J. Dent. Res. 83: 518–522. Okafor, C.C., Haleem-Smith, H., Laqueriere, P., Manner, P.A. and Tuan, R.S. (2006). Particulate endocytosis mediates biological responses of human mesenchymal stem cells to titanium wear debris. J. Orthop. Res. 24: 461–473. Park, Y., Lutolf, M.P., Hubbell, J.A., Hunziker, E.B. and Wong, M. (2004). Bovine primary chondrocyte culture in synthetic matrix metalloproteinase-sensitive poly(ethylene glycol)-based hydrogels as a scaffold for cartilage repair. Tissue Eng. 10: 515–522. Paul, R., Haydon, R.C., Cheng, H., Ishikawa, A., Nenadovich, N., Jiang, W., Zhou, L., Breyer, B., Feng, T., Gupta, P., et al. (2003). Potential use of Sox9 gene therapy for intervertebral degenerative disc disease. Spine 28: 755–763. Pittenger, M.F., Mackay, A.M., Beck, S.C., Jaiswal, R.K., Douglas, R., Mosca, J.D., Moorman, M.A., Simonetti, D.W., Craig, S. and Marshak, D.R. (1999). Multilineage potential of adult human mesenchymal stem cells. Science 284: 143–147.

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Qi, M., Hu, J., Zou, S., Zhou, H. and Han, L. (2006). Mandibular distraction osteogenesis enhanced by bone marrow mesenchymal stem cells in rats. J. Craniomaxillofac. Surg. 34: 283–289. Quarto, R., Mastrogiacomo, M., Cancedda, R., Kutepov, S.M., Mukhachev, V., Lavroukov, A., Kon, E. and Marcacci, M. (2001). Repair of large bone defects with the use of autologous bone marrow stromal cells. N. Engl. J. Med. 344: 385–386. Romanov, Y.A., Svintsitskaya, V.A. and Smirnov, V.N. (2003). Searching for alternative sources of postnatal human mesenchymal stem cells: candidate MSC-like cells from umbilical cord. Stem Cells 21: 105–110. Rotter, N., Ung, F., Roy, A.K., Vacanti, M., Eavey, R.D., Vacanti, C.A. and Bonassar, L.J. (2005). Role for interleukin 1alpha in the inhibition of chondrogenesis in autologous implants using polyglycolic acid–polylactic acid scaffolds. Tissue Eng. 11: 192–200. Rouwkema, J., de Boer, J. and Van Blitterswijk, C.A. (2006). Endothelial cells assemble into a 3-dimensional prevascular network in a bone tissue engineering construct. Tissue Eng. 12: 2685–2693. Sampaolesi, M., Blot, S., D’Antona, G., Granger, N., Tonlorenzi, R., Innocenzi, A., Mognol, P., Thibaud, J.L., Galvez, B.G., Barthelemy, I., et al. (2006). Mesangioblast stem cells ameliorate muscle function in dystrophic dogs. Nature 444: 574–579. Sartorelli, V. and Caretti, G. (2005). Mechanisms underlying the transcriptional regulation of skeletal myogenesis. Curr. Opin. Genet. Dev. 15: 528–535. Schofield, R. (1978). The relationship between the spleen colony-forming cell and the haemopoietic stem cell. Blood Cells 4: 7–25. Sekiya, I., Larson, B.L., Vuoristo, J.T., Reger, R.L. and Prockop, D.J. (2005). Comparison of effect of BMP-2, -4, and -6 on in vitro cartilage formation of human adult stem cells from bone marrow stroma. Cell Tissue Res. 320: 269–276. Shen, F.H., Visger, J.M., Balian, G., Hurwitz, S.R. and Diduch, D.R. (2002). Systemically administered mesenchymal stromal cells transduced with insulin-like growth factor-I localize to a fracture site and potentiate healing. J. Orthop. Trauma 16: 651–659. Steck, E., Bertram, H., Abel, R., Chen, B., Winter, A. and Richter, W. (2005). Induction of intervertebral disc-like cells from adult mesenchymal stem cells. Stem Cells 23: 403–411. Tamaki, T., Akatsuka, A., Ando, K., Nakamura, Y., Matsuzawa, H., Hotta, T., Roy, R.R. and Edgerton, V.R. (2002). Identification of myogenic-endothelial progenitor cells in the interstitial spaces of skeletal muscle. J. Cell Biol. 157: 571–577. Tozer, S. and Duprez, D. (2005). Tendon and ligament: development, repair and disease. Birth Defects Res. C Embryo Today 75: 226–236. Tuan, R.S., Boland, G. and Tuli, R. (2003). Adult mesenchymal stem cells and cell-based tissue engineering. Arthr. Res. Ther. 5: 32–45. Tuli, R., Tuli, S., Nandi, S., Huang, X., Manner, P.A., Hozack, W.J., Danielson, K.G., Hall, D.J. and Tuan, R.S. (2003). Transforming growth factor-beta-mediated chondrogenesis of human mesenchymal progenitor cells involves N-cadherin and mitogen-activated protein kinase and Wnt signaling cross-talk. J. Biol. Chem. 278: 41227–41236. Tuli, R., Nandi, S., Li, W.J., Tuli, S., Huang, X., Manner, P.A., Laquerriere, P., Noth, U., Hall, D.J. and Tuan, R.S. (2004). Human mesenchymal progenitor cell-based tissue engineering of a single-unit osteochondral construct. Tissue Eng. 10: 1169–1179. Vacanti, C.A., Bonassar, L.J., Vacanti, M.P. and Shufflebarger, J. (2001). Replacement of an avulsed phalanx with tissueengineered bone. N. Engl. J. Med. 344: 1511–1514. Veilleux, N. and Spector, M. (2005). Effects of FGF-2 and IGF-1 on adult canine articular chondrocytes in type II collagen–glycosaminoglycan scaffolds in vitro. Osteoarthr. Cartilage 13: 278–286. Wakitani, S., Imoto, K., Yamamoto, T., Saito, M., Murata, N. and Yoneda, M. (2002). Human autologous culture expanded bone marrow mesenchymal cell transplantation for repair of cartilage defects in osteoarthritic knees. Osteoarthr. Cartilage 10: 199–206. Wakitani, S., Aoki, H., Harada, Y., Sonobe, M., Morita, Y., Mu, Y., Tomita, N., Nakamura, Y., Takeda, S., Watanabe, T.K., et al. (2004). Embryonic stem cells form articular cartilage, not teratomas, in osteochondral defects of rat joints. Cell Transplant. 13: 331–336. Yamada, Y., Ueda, M., Naiki, T. and Nagasaka, T. (2004). Tissue-engineered injectable bone regeneration for osteointegrated dental implants. Clin. Oral Implants Res. 15: 589–597.

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52 Hepatocyte Transplantation Stephen C. Strom and Ewa C.S. Ellis INTRODUCTION The concept of regenerative medicine implies that the clinician works with the innate healing and regenerative process of the body to affect an improvement in a patient’s health. Perhaps more than with any other organ, the liver offers the greatest opportunity for regenerative medicine. This is because, unlike most other tissues, the liver has the capacity to regenerate following massive chemical or physical insult and tissue loss (Michalopoulos and DeFrances, 1997). Our very existence may well rely on the ability to regenerate liver mass. The liver is an incredibly complex organ which performs quite diverse biological functions, from glycogen storage and catabolism to maintain blood sugar levels, to the production and secretion of critical plasma proteins including albumin, clotting factors, and protease inhibitors. In addition the liver is the major site in the body for the metabolism and excretion of hormones, metabolic waste products such as ammonia as well as exogenous compounds such as toxins, drugs, and a variety of other compounds to which we are exposed through odiet and environment. These processes are so critical to survival that the loss of any of these functions has serious and often lethal consequences for the individual. Until recently, the only option for treating chronic liver disease or metabolic defects in liver function has been whole organ transplantation. Recently, hepatocyte transplantation has been performed. Although still an experimental therapy, there are some potential advantages for a cell therapy approach to treat liver disease. Some of the advantages and problems with the current treatments for liver disease are listed in Table 52.1. Despite the unquestioned success of this technique orthotopic liver transplantation (OLT) requires major surgery and has a significantly long recovery period. The financial costs associated with OLT and lifelong immunosuppression is considerable. There is a high incidence of complications from the surgical procedure and the Table 52.1 Current treatments for liver disease Orthotopic liver transplantation Major and expensive surgery Extensive recovery period High incidence of complications Expensive maintenance therapy Shortage of donor organs Timing is critical Hepatocyte transplantation Less invasive and less costly procedure Complications, fewer, and less severe Timing of procedure is easier Alternative cell sources Patient retains native liver Graft loss is not necessarily lethal Option remains for whole organ transplant

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concomitant immunosuppression which is required following the organ transplant. Complications can range from simple infections to renal failure, hyperlipidemia, and an increased incidence of skin and other types of cancers following long-term immunosuppression. As with all other organs, the number of liver donors does not nearly equal the number of patients on the waiting list. Patients may wait 2 or more years for a liver transplant, and there is a death rate of greater than 10% per year of patients on the waiting list. Timing is critical for whole organ transplant. An ABO-compatible liver donor must be available when a patient requires the transplant. Some of the limitations associated with whole organ transplants are addressed with hepatocyte transplants (Table 52.1). Hepatocyte transplants do not require major surgical procedures as they are performed by infusion of cells into the blood supply to an organ such as the liver or spleen. Thus, hepatocyte transplants are less invasive and less costly procedures. Because major surgery is not required there are fewer complications associated with the procedure. Since cell infusions are minor procedures, there is essentially no recovery period needed. If patients were healthy prior to the procedures, such as a stable metabolic disease patient, they would likely feel no adverse effects from the procedure other than from the placement of a catheter. Hepatocytes can be banked and cryopreserved, so theoretically, cells could be available anytime for a patient transplant. The timing of a hepatocyte transplant depends on the status of the patient rather than on the availability of a suitable organ. Currently, the source of hepatocytes for hepatocyte transplants is mainly discard organs not suitable for whole organ transplant (Nakazawa et al., 2002). Currently, there are not enough hepatocytes to transplant all recipients who would likely benefit from the procedure. However, some inventive new ideas have been proposed, such as to use segment IV, which can be made available from a split-liver procedure (Mitry et al., 2004) to make more hepatocytes available for transplants. Alternative sources of hepatocytes could also be available in the future. Although many options are discussed, the most prominent sources are xenotransplants from pigs or other species, immortalized hepatocytes and most recently stem cell-derived hepatocytes (Strom and Fisher, 2003). Future developments in these areas may make the number of cells available for hepatocyte transplants virtually unlimited. A significant benefit of hepatocyte transplantation is that the patients retain their native liver. In cases of cell transplants for metabolic disease, the patient’s native liver still performs all of the liver functions with the exception of the function which initiates the disease. Patients with ornithine transcarbamylase deficiency (OTC) have mutation in an enzyme involved in the urea cycle which prevents the metabolism and elimination of ammonia. Although the native liver is not proficient in ammonia metabolism, it is still capable of performing other liver functions including the secretion of clotting factors, albumin, drug metabolism, and all other metabolic and synthetic processes. A cell transplant need only support the ammonia metabolism for the patient, and will not be required to provide complete liver support. Because all liver functions are not dependent on donor cells, loss of the cell graft or failure of the cells to function properly will not necessarily be life threatening, especially for a stable metabolic disease patient. Finally, a whole organ transplant always remains as an option for the cell transplant patient. Even if the cell transplant fails to function or is rejected, nothing done as part of the cell transplant procedure would likely interfere with a subsequent whole organ transplant. Fisher et al. (1998) reported that prior hepatocyte transplantation did not sensitize the cell transplant recipient to either the donor cells or to an eventual liver graft. Thus, despite sometimes transplanting hepatocytes directly into an immunological response organ, the spleen, no immunological reactions are initiated which are deleterious to the cell transplant or an eventual whole organ transplant. There are potential disadvantages of hepatocyte transplants as well. First, there are no reports of long-term complete corrections of metabolic liver disease in patients following cell transplantation alone. Because it is a new field, much additional experimentation will be required to determine the full efficacy of cell therapy of liver disease and the length of time the cell graft will function. Also, like whole organ transplants, it is believed that cell transplant recipients will require the administration of immunosuppressive drugs. It is likely that lower doses of the drugs will be needed to prevent rejection of cell transplants than are required for whole organ transplants.

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Because of this, fewer and less severe side effects from immunosuppressive drugs would be expected, but definitive studies are lacking.

BACKGROUND STUDIES Choice of Sites for Hepatocyte Transplantation Hepatocyte transplants have been conducted for over 20 years. A number of good reviews are available for details of the experiments and the original references which may be omitted in this review (Strom et al., 1999, 2006; Malhi and Gupta, 2001; Ohashi et al., 2001; Fox, 2002; Fox and Roy-Chowdhury, 2004). The large numbers of preclinical studies conducted on hepatocyte transplants firmly establish that the transplants are safe and effective. The most common sites for the transplantation of hepatocytes are the spleen and the liver; however, transplants to the peritoneal cavity, stomach, or omentum have been reported. Long-term survival of the cells is readily measured following transplants into the spleen or liver. The majority of cells transplanted into the peritoneal cavity intellectual property (IP) are rapidly lost. Following IP transplants, only those cells which nidate near blood vessels and can attract sufficient nutrition survive long term. Despite the ease of the procedure, IP transplants of hepatocytes have only limited efficacy. Transplants of hepatocytes to the spleen or the liver have been shown to function for the lifetime of the recipient (Mito et al., 1979; Gupta et al., 1991; Ponder et al., 1991; Holzman et al., 1993). Studies by Mito and coworkers clearly show long-term survival of hepatocytes and that over time the spleen of an animal can be “hepatized” to where 80% of the mass of the organ can replaced with hepatocytes (Mito et al., 1978, 1979; Kusano et al., 1981, 1992; Kusano and Mito, 1982). The concept of establishing ectopic liver function in the spleen is similar in theory to the bioartificial liver (BAL). In BAL, the hepatocytes are seeded into and maintained in some form of an extracorporal device. The patient’s blood or plasma is pumped to the device where it interacts with the hepatocytes across membrane barriers and is then returned to the patient by a second series of pumps. There are reports that BAL can provide short-term synthetic and metabolic support (Gerlach et al., 2003; Demetriou et al., 2004). The ease of transplant of hepatocytes and the abundance of the patient’s own natural basement membrane components coupled with the naturally high blood flow make the spleen a useful site for the establishment of short- or long-term ectopic liver function. It is likely that hepatocyte transplants will be easier, cheaper, more efficient, and will provide the same, or better, level of support as extracorporal devices. For transplants into the liver, the preferred route for administration of cells is via the portal vein. Cells are infused into the blood supply which feeds the liver and the hepatocytes are distributed to the different lobes in proportion to the blood flow they receive from the portal vein. Portal vein injections are difficult in small animals, so an alternative method is used in these studies. Hepatocytes are injected directly into the splenic pulp. The proportion of the cells which remain in the spleen is determined by the extent to which the outflow through splenic veins is impeded. In the studies of Mito et al. (1979), where the spleen was “hepatized” the authors briefly occlude the splenic outflow which helps retain the cells in the spleen. Alternatively, when the spleen is used as a method to affect a portal vein injection, the splenic veins are left open. It was reported that up to 52% of the cells injected into the spleen traverse to the liver via the splenic and portal veins within a few minutes (Gupta et al., 1991; Ponder et al., 1991). Integration of Hepatocytes Following Transplantation Integration of hepatocytes into recipient liver is a complex process which requires the interaction of donor and native hepatocytes to form an integrated tissue. The process may be considered in four steps (Table 52.2) (Gupta et al., 1995, 1999b, 2000; Koenig et al., 2005). Although they are presented as separate, there is considerable overlap of the steps in both time and space. Some of the most spectacular photographs of the entire

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Table 52.2 Integration of donor hepatocytes into native liver following transplantation Filling vascular spaces with donor cells Disruption of the sinusoidal endothelium Donor cell integration in host parenchyma Remodeling of liver via modulation of extracellular matrix

process are provided by Koenig et al. (2005). Following infusion into the portal vein hepatocytes must traverse the endothelium to escape the vascular system. Although the liver has fenestrated endothelium, under normal conditions the pores which are in the range of 150 nm are far too small to provide a simple transit of parenchymal hepatocytes which range in size of 20–50 μm. Infusions of hepatocytes quickly fill the portal veins and embolize secondary and tertiary portal radicals (Gupta et al., 1999a). Portal pressures increase as flow is restricted by hepatocyte plugs in the portal veins. Venograms which were normal prior to cell transplantation become markedly attenuated and show greater filling of vessels proximal to the portal vein including the mesenteric and splenic vein. If the number of hepatocytes transplanted is in the range of 5% of the total number of hepatocytes in the native liver, the portal hypertension is transient and resolves within minutes to hours. A proportion of transplanted cells begin to fill sinusoidal spaces and the space of Disse as the endothelium in the region of the transplanted cells begins to degenerate. It is likely that both physical and humoral (growth factors, cytokines) factors are involved in this process. Microscopic analysis of tissue sections reveal that endothelium is breached in many places and donor hepatocytes leave the portal veins in regions where endothelium is incomplete and broken. Reports suggest that most of the hepatocytes which eventually integrate into recipient liver will have traversed the endothelial barrier by 24 h post transplant. Cells which remain in the portal vessels are eventually removed by macrophages between 16 and 24 h post transplant. Other reports suggest that cells may continue to integrate into parenchyma for 2–3 days following transplantation (Shani-Peretz et al., 2005). Transient hypoxia in the region of the occluded vessels leads to changes in both the endothelium as well as both recipient and donor hepatocytes. Endothelium and donor and native hepatocytes all express vascular endothelial growth factor (VEGF) in the areas of hepatocyte integration (Gupta et al., 1999b; Shani-Peretz et al., 2005) a factor known to be induced by hypoxia. It is interesting that VEGF was previously known as vascular permeability factor (VPF). Expression and secretion of VEGF/VPF a potent angiogenesis factor is thought to contribute to the reformation of new sinusoids and restoration of the endothelial barrier following cell transplantation. Passage through the endothelial barrier allows donor hepatocytes to become integrated into recipient parenchyma. Full integration of donor hepatocytes and restoration of full hepatic function is difficult to ascertain. However, careful studies of the expression of antigens and activities localized to specific membrane fractions clearly demonstrate that donor hepatocytes fully integrate into the hepatic plate of native liver and for hybrid structures between native and donor cells within 3–5 days following transplantation. The antibody to CD26 recognizes the dipeptidylpeptidase IV (DPPIV) antigen which is localized to the basolateral membrane of hepatocytes. Antibodies to connexin 32 can be used to visualize gap junctions between adjacent hepatocytes. Likewise, canicular ATPase activity can be used to identify bile cannicular regions between adjacent hepatocytes. The proper localization of these different antigens and activities requires that the hepatocyte be fully integrated into the hepatic plate and polarized. By 3–7 days post transplant, hybrid structures could be visualized in recipient liver containing both donor (DPPIV) hepatocytes and recipient ATPase activity (Gupta et al., 1995) or donor DPPIV co-localized with connexin 32 (Koenig et al., 2005). Both studies clearly demonstrate proper integration of donor hepatocytes as well as the reestablishment of intracellular communication (connexin 32) between donor and recipient hepatocytes. Hybrid structures between donor and recipient hepatocytes were shown to be functional as shown by the transport and excretion of a fluorescent conjugated bile acid (Gupta et al., 1995). Hepatic transport of indocyanine and sulfobromothalein into the bile following hepatocyte transplantation was also reported by

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Hamaguchi et al. (1994). Hepatocyte transplants were conducted on Eizai-hyperbilirubinemic rats. These animals have a defect in multidrug resistance protein2 (MRP2), which prevents the normal transport of bile acid conjugates and their excretion into bile. This is a relevant animal model of metabolic disease as the condition is similar to Dubin–Johnson syndrome in humans. The correction of this transport defect by hepatocyte transplantation is definitive proof of the complete functional integration of donor hepatocytes into recipient liver. As part of the integration process there is significant remodeling of the hepatic parenchyma. Koenig et al. (2005) has reported the activation and release of matrix metaloprotease-2 (MMP-2) in the immediate area of donor cells. It is not clear if the proteases are produced by the donor or recipients cells or even which cell type is the source of the protease, but the degradation of extracellular matrix components helps to create space for the donor cells. Expression of MMP-2 was detected in and surrounding foci of proliferating donor hepatocytes 2 months following cell transplantation. Increased production and release of MMP-2 were also observed at the growth edge of nodules of fetal rat hepatocytes proliferating in adult liver following transplantation (Oertel et al., 2006). While all of the components of the process are not completely understood, it is clear that hepatocytes can be transplanted into the vascular supply of the liver, breach the endothelial barrier, remodel and integrate into hepatic parenchyma, and establish communication with adjacent cells and the biliary tree all within 3–5 days in a process of remodeling which completely retains normal host hepatic architecture.

CLINICAL HEPATOCYTE TRANSPLANTATION Hepatocyte transplantation has been employed in the clinics in three types of procedures (Table 52.3). Cell transplants have been used to provide short-term liver support to patients who are dying of their disease before a suitable organ could be found. As these patients are already listed for a whole organ transplant, the hepatocyte infusion is used sometimes referred to as a “bridge” to transplant. A second use for hepatocyte transplants grew out of the attempts to bridge people to OLT. It was discovered that some of the patients receiving hepatocyte transplants recovered completely following the hepatocyte transplants and no longer required whole organ transplant. The third general use for hepatocyte transplants is for the correction of metabolic liver disease. Each technique will be discussed separately. Hepatocyte Bridge With the bridge technique, hepatocytes are provided to a patient in acute liver failure or those experiencing acute decompensation following chronic liver disease. The majority of these patients are already listed for OLT, and they are in danger of dying before a suitable organ could be found. Hepatocyte transplants have been conducted on these patients in an effort to keep them alive long enough to receive OLT. The primary goal of the bridge transplant is not to prevent whole organ transplant, but rather to support and sustain the patient until an organ becomes available. Preclinical studies with several different models of acute or chronic liver failure have demonstrated that hepatocyte transplantation can support liver function and improve survival (Sutherland et al., 1977; Sommer et al., 1979; Makowka et al., 1981; Demetriou et al., 1988; Mito et al., 1993; Takeshita et al., 1993; Arkadopoulos et al., 1998b; Kobayashi et al., 2000; Ahmad et al., 2002; Aoki et al., 2005). The results with human hepatocyte transplantation in the clinics also show an increase in the survival of patients following hepatocyte transplantation. There are now several reports and review articles which provide details of the patients and the transplant procedures (Habibullah et al., 1994; Strom et al., 1997a, b, 1999, 2006; Bilir et al., 2000; Ohashi et al., 2001; Soriano, 2002; Fox and Roy-Chowdhury, 2004; Fisher and Strom, 2006). The results indicate that there is a 65% survival rate for patients receiving hepatocyte transplants. Although randomized control studies could not be conducted, the preliminary results with approximately 25 patients indicate a survival advantage to those patients receiving cell transplants. In addition to increase survival, there are consistent reports that clinical parameters such as ammonia levels, intracranial pressures, and cerebral blood flow are improved following hepatocyte transplantation (Strom et al., 1997a, b, 1999; Soriano et al., 1998; Bilir et al., 2000; Fisher, 2004; Fisher and Strom, 2006). These results indicate

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that desperately ill patients who receive hepatocyte transplants are morel likely to survive long enough to receive OLT than the non-transplant controls. Most of the patients who would be candidates for the hepatocyte bridge technique suffer from chronic liver disease and have advanced cirrhosis. Because of the cirrhotic changes in the liver and the accompanying portal hypertension, hepatocytes were not transplanted into the liver (portal vein) in most of the clinical studies. Preclinical studies were conducted where cirrhosis was induced in rats by the administration of phenobarbital and carbon tetrachloride (Gupta et al., 1993). When hepatocytes were subsequently transplanted into animals with increased portal pressures and cirrhosis, there was significantly greater intrapulmonary translocation of donor cells presumably because of portosystemic shunting. These results suggest that serious complications could arise if portal infusion of hepatocytes were conducted on cirrhotic patients with portal hypertension. Indeed, shunting of transplanted hepatocytes to pulmonary vascular beds has been reported in one clinical study (Bilir et al., 2000). To avoid this possible complication, Fisher et al. recommends that hepatocytes be transplanted into the spleen in cirrhotic patients via the splenic artery (Strom et al., 1997b; Fisher and Strom, 2006). Despite the obvious success of the splenic artery route for hepatocyte transplantation, a recent report suggests that transplantation of hepatocytes by direct splenic puncture results in superior engraftment and fewer serious complications, although long-term engraftment was not studied (Nagata et al., 2003b). Although the method for splenic delivery of cells may not be settled, it is clear that in cases where physical and/or anatomic abnormalities are present in the native liver, the preferred route for hepatocyte transplantation is to an ectopic site, the spleen. The promising results reported to date suggest that hepatocyte transplantation is beneficial to patients suffering from severe hepatic insufficiency while awaiting OLT. A logical extension of these results might be for the use of hepatocyte transplants earlier in the process. Rather than wait until the patient is near death and with no immediate prospect for a whole organ transplant, a more preemptive approach might be warranted. Hepatocyte transplants could be performed when patients awaiting OLT become unstable. This would presumably stabilize the patient and avoid or at least delay more serious complications of liver failure. Early intervention might avoid more costly hospitalization and other treatments. Hepatocyte Transplantation in Acute Liver Failure As described above, hepatocyte transplants have been used as a bridge to OLT. Most of the patients who have been referred for bridge transplants suffered from chronic liver disease and had cirrhotic changes in liver architecture. There is a subgroup of patients referred for OLT who experience acute liver failure. In these patients there is massive loss of hepatocytes over a short period of time leading to hepatic insufficiency. Except for the dramatic loss of hepatocytes there is no long-standing pathological change in liver architecture. Since the liver has the capacity for robust regeneration following loss of liver mass (Michalopoulos and DeFrances, 1997), there is considerable interest in trying to correct acute liver failure with hepatocyte transplantation. The hypothesis is similar to the bridge technique, where hepatocyte transplantation is used to provide support at a time of critical and otherwise lethal liver failure. The expectation is that if the patient survives the acute loss of tissue mass, their native liver will regenerate. If the native liver regenerates, there will no longer be a need for OLT. An exogenous source of hepatocytes by transplantation would provide support of liver function to prevent lethal hepatic failure. Both donor and native hepatocytes would be expected to participate in the regeneration response. Once the native liver has been fully restored there might not be a need for donor-derived hepatocytes. If the chimeric liver generated following the transplant is composed predominantly of native hepatocytes, the patient could be safely removed from immunosuppressive therapy. In this manner, the patient receives, what amounts to, a temporary liver cell transplant. If cell therapy is sufficient, the patient will be spared whole organ transplantation and lifelong immunosuppression. Several preclinical studies support the hypothesis that hepatocyte transplantation can provide sufficient liver function to maintain an animal experiencing acute liver failure. Studies have shown that hepatocyte transplants dramatically improve survival of animals with acute liver failure induced by

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Table 52.3 Opportunities for hepatocyte transplantation “Bridge” for patients to whole organ transplantation Cell support for acute liver failure “Cell therapy” for metabolic disease

D-galactosamine (Sutherland et al., 1977; Sommer et al., 1979; Makowka et al., 1981; Baumgartner et al., 1983); 90% hepatectomy (Cuervas-Mons et al., 1984; Demetriou et al., 1988; Mito et al., 1993; Kobayashi et al., 2000), or ischemic liver injury (Takeshita et al., 1993; Arkadopoulos et al., 1998a). There are now reports of reversal of acute liver failure in four patients following hepatocyte transplantation (Fisher et al., 2000; Soriano, 2002; Fisher and Strom, 2006; Ott et al., 2006). The causes of acute liver failure ranged from hepatitis B-induced liver failure to acetaminophen intoxication, to liver toxicity following eating poisonous mushrooms to liver failure of unknown etiology in a pediatric patient. In each case patients presented with classic symptoms of acute liver failure, and most were immediately listed for OLT. The number of cells transplanted varied between different procedures but ranged from approximately 1 to 5 billion total viable cells. In all cases cells were transplanted into the portal vein to get a direct transplant into the liver. In general, patients were given fresh frozen plasma prior to placement of the catheter to prevent bleeding. The results presented by Fisher et al. (2000) are typical of the response to hepatocyte transplantation. There is usually a rapid fall in ammonia levels following the transplant. Circulating levels of clotting factors stabilize following the transplant and then slowly increase over the next 2 weeks. Fisher et al. reports that Factor VII levels were 1% of normal prior to transplant and increased to 25% by 7 days and 64% of normal by week 2 post cell transplant. The recovery of the clotting factors is usually rapid enough that following the cell transplant, no additional fresh frozen plasma is required. Patients are generally discharged within 2–4 weeks and are judged to experience a complete recovery. The cell transplant recipients ranged in age from 3 to 64 years in age, indicating that even older patients have sufficient regenerative capacity to be supported by hepatocyte transplantation. As is observed with donor tissue allografts, hepatocyte allografts produce and secrete human leukocyte antigen-I (sHLA-I) immediately upon implantation. If there is a mismatch between the donor and recipient the donor specific sHLA-I can be detected in the circulation and quantified by enzyme-linked immunosorbent assay (ELISA). Donor specific HLA class I alleles can be identified and quantified by polymerase chain reaction (PCR) analysis of tissue samples taken at biopsy. When it is determined that the preponderance of cells in the patients liver are native, the patients can slowly be removed from immunosuppressive therapy as was described by Fisher et al. (2000). In the cases described to date, the patients recovered completely from liver failure following hepatocyte transplantation without serious adverse consequences and without whole organ transplant and lifelong immunosuppression. Although the numbers of patients are small, the treatment of acute liver failure by hepatocyte transplant has some significant advantages which make further investigation of this novel therapy appropriate (Table 52.3).

Hepatocyte Transplantation for Metabolic Liver Disease A common indication for whole organ transplantation in pediatric patients is metabolic liver disease. In these cases, there is usually a genetic defect in an enzyme or protein which is produced in the liver which inactivates a critical liver function. Although all other liver functions are generally normal, the liver is removed and replaced with a liver which can perform the missing function. Because there is usually only one genetic defect associated with each metabolic liver disease, a gene therapy approach to correct the defect would seem appropriate. Unfortunately, gene therapy has met with considerable problems which have prevented successful use of this experimental technique. Hepatocyte transplantation has been used in attempts to correct the metabolic defects associated with several types of metabolic liver disease (Table 52.4).

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Table 52.4 Clinical transplants for metabolic liver disease Familial hypercholesterolemia (3) Crigler–Najjar(5) Ornithine transcarbamylase deficiency (4) Arginosuccinate lyase deficiency (1) Factor VII deficiency (2) Glycogen storage disease (2) Infantile Refsum disease (1) Progressive familial intrahepatic cholestasis (2) Alpha-1 antitrypsin deficiency (2)

In an approach similar to gene therapy, with hepatocyte transplants one tries to seed the patient’s liver with cells which are proficient in the enzyme or function missing in the native liver. The goal is to repopulate the liver of the transplant recipient with sufficient numbers of hepatocytes to provide the missing liver function by donor cells. Large numbers of hepatocytes cannot be infused into the portal system because of the problems with embolism of the portal veins and portal hypertension. We have used as a general rule to infuse approximately 2  108 cells/kg body weight of the recipient (Fox et al., 1998; Horslen et al., 2003). Infusions of these cell numbers has not resulted in any long-term complications. There is always a transient increase in portal pressures which resolves within hours (Strom et al., 1997a; Fox et al., 1998; Bohnen et al., 2000; Soriano, 2002; Horslen et al., 2003; Sokal et al., 2003; Horslen and Fox, 2004). While quite experimental, this number was arrived at by an extrapolation from preclinical studies with non-human primates. Grossman et al. (1992) reported that the infusion of between 1–2  108 cells/kg into baboons who had previously received a left or right lobectomy was accomplished without serious complications and with only transient increases in portal pressures. Because only a few percent of liver mass can be transplanted at any one time, single hepatocyte transplants cannot be expected to replace a large percentage of liver with donor cells. For this reason, the metabolic diseases which are candidates for cell transplants are those in which the restoration of 10% less of total liver function or activity is likely to correct the disease. The liver has highly redundant functions. Thus, it is recognized that 10% of a normal amount of gene product or enzyme activity would likely correct the symptoms of most metabolic liver diseases. Exceptions exist, like hypercholesterolemia, where more than 50% replacement of liver with donor cells would likely be needed to correct circulating low density lipoprotein levels. However, for most metabolic liver disease and all of those listed in Table 52.4, it is believed that the replacement of the liver with 10% donor hepatocytes would either be completely corrective or at least ameliorate most of the symptoms of the disease. In general, hepatocyte transplants work best when the donor cells have a selective growth advantage. There are a number of animal models of liver disease where the native hepatocytes show an increased death rate as compared to normal liver (Sandgren et al., 1991; Rhim et al., 1994; Overturf et al., 1996; De Vree et al., 2000). In these situations, when cells without the defect are transplanted into the diseased liver, the donor cells have a strong and selective growth advantage as compared to the native hepatocytes. Over time the liver may become nearly completely replaced with donor cells. In certain human diseases there might be sufficient selective pressure to strongly favor the replacement of large parts of the liver with donor cells. Such diseases include tyrosinemia Type 1, Wilson’s disease (Irani et al., 2001), progressive familial intrahepatic cholestasis (PFIC) (De Vree et al., 2000), alpha-1 antitrypsin deficiency (A1AT) (Rudnick and Perlmutter, 2005). In these diseases, integration of only a small proportion of liver mass by hepatocyte transplantation would likely be necessary because the donor cells would be expected to continue to proliferate in the host liver, and over time replace the diseased cells. Although there are clear examples of this in studies of transplants of laboratory animals, there are no studies with human patients showing comparable results. Most metabolic diseases such as Crigler–Najjar (CN), OTC deficiency, and all of those diseases listed in Table 52.4 would not be expected to show such selective growth pressure for donor cells. For diseases such

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as these, multiple transplants over time will be required to populate the liver with 10% donor cells (Rozga et al., 1995). A large number of studies with different animal models have shown the efficacy of hepatocyte transplantation to correct metabolic liver disease (reviewed in Malhi and Gupta, 2001 and Strom et al., 2006). Metabolic defects in bilirubin metabolism (Matas et al., 1976; Groth et al., 1977; Vroemen et al., 1986; Demetriou et al., 1988; Moscioni et al., 1989; Holzman et al., 1993; Hamaguchi et al., 1994), albumin secretion (Mito et al., 1979; Kusano and Mito, 1982; Demetriou et al., 1993; Rozga et al., 1995; Moscioni et al., 1996; Oren et al., 1999), ascorbic acid production (Onodera et al., 1995; Nakazawa et al., 1996), tyrosinemia Type 1 (Overturf et al., 1996), copper excretion (Yoshida et al., 1996; Irani et al., 2001; Allen et al., 2004), PFIC (De Vree et al., 2000) as well as other defects in biliary transport similar to Dubin–Johnson syndrome in humans (Hamaguchi et al., 1994) have been shown to be amenable to correction by hepatocyte transplantation. These encouraging results suggested that similar defects in human patients could be corrected by hepatocyte transplantation. The diseases listed in Table 52.4 have been the focus of human trials of hepatocyte transplants. The numbers in parenthesis are the number of patients who have received transplants. Hepatocyte transplants were previously shown to result in a rapid correction of ammonia levels (Strom et al., 1997b, 1999; Bilir et al., 2000; Soriano, 2002). For this reason, urea cycle defects which result in life-threatening hyperammonemia were the first metabolic disease target for hepatocyte transplants (Strom et al., 1997b; Bohnen et al., 2000). In the initial study, 1 billion viable cells were transplanted into the portal vein of a 5-year-old recipient. Portal pressures increased from 11 cm of water prior to cell transplant to 19 cm immediately following the cell infusion, but recovered rapidly. The patient’s ammonia levels normalized without medical intervention within 48 h of cell infusion and his glutamine levels returned to normal. Although OTC activity was undetectable prior to cell transplant, measurable OTC activity was detected in a biopsy performed at 28 days. In these studies 10% of the cells were labeled with indium111 prior to infusion into the patient to monitor distribution of the cells. Quantitative analysis of the scientigraphic images showed an average distribution ratio of liver:spleen of 9.5:1. Measurements made prior to cell infusion indicated that free indium was released from hepatocytes at a rate of 10% per hour, and free indium is rapidly cleared from circulation by reticuloendothelial systems such as the spleen. Thus, most of the tracer in the spleen following cell infusion were thought to be free indium, not hepatocytes. Pulmonary radiotracer uptake was consistent with background counts, indicating the absence of portosystemic shunting despite the modest increase in portal pressures observed at the time of transplant. This first transplant for metabolic liver disease indicated that hepatocyte transplantation into the portal vein could be conducted safely in patients with no significant liver pathology with only a moderate and reversible increase in portal pressures. From the rapid normalization of ammonia levels following hepatocyte transplant, it was concluded that cell transplantation can partially correct the hyperammonemia associated with the disease. Subsequent studies have verified that partial corrections of ammonia levels are possible by cell transplants alone (Horslen et al., 2003; Dhawan et al., 2004; Stephenne et al., 2005). While complete corrections of OTC deficiency have not been accomplished these studies indicate that cell transplants provide much needed metabolic control of ammonia levels. Even in the absence of complete correction, liver cell transplantation should be considered as a bridge to whole organ transplantation for OTC patients to prevent the neurological problems associated with uncontrolled hyperammonemia (Bohnen et al., 2000; Stephenne et al., 2005). A number of groups have attempted to correct CN syndrome, Type 1 with hepatocyte transplants. The first case was in many ways typical of the results obtained by other groups and will be discussed in greater detail (Fox et al., 1998). This disease is caused by a defect in the enzyme which is responsible for the conjugation and eventual excretion of bilirubin. The absence of the enzyme results in severe hyperbilirubinemia which can lead to central nervous system (CNS) toxicity including kernicterus. Following the transplantation of approximately 7.5 billion cells into the liver of a 10-year-old female, there was a slow and continuous decrease in circulating bilirubin levels over the first 30–40 days, and bilirubin conjugates were readily detected in the bile. Overall, there was

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approximately a 60–65% decrease in bilirubin levels as compared to pretransplant levels. Because the bilirubin conjugates could only be produced by the donor cells, their detection in the bile demonstrates the robust biochemical function of the transplanted cells and established that donor hepatocytes integrated into the hepatic parenchyma and quickly established connections with the recipient’s biliary tree. Several important finding were gained from this transplant. First, large numbers of hepatocytes could be safely transplanted into the portal vein without complication. Although the total numbers of hepatocytes in liver are difficult to assess, a transplant of 7.5 billion cells represents and estimated 3.5–7.5% of the liver mass, which was transplanted without complication over approximately a 15-h period. Second, the apparent engraftment and function of hepatocytes in the clinical trials seems to exceed that found in previous animal studies. The transplantation of 3.5–7.5% of liver mass resulted in the restoration of approximately 5% of a normal amount of bilirubin conjugation capacity in the liver. Third, a long-term correction in bilirubin levels was observed. This patient was followed for more than 1.5 years. Fourth, single transplants of hepatocytes are effective in creating partial corrections of the disease, but given the limitation of transplanting 2  108 cells/kg body weight, one cannot transplant sufficient numbers of hepatocytes to achieve a complete correction of metabolic liver disease with one transplant. It is estimated that complete corrections would require 2–4 transplants if each were as successful and efficient as the first. Finally, this was the first unequivocal demonstration of the long-term success of hepatocyte transplantation. Although patients were bridged to transplant and clinical parameters such as ammonia levels rapidly changed following transplantation, many of the previous patients underwent subsequent OLT the long-term metabolic function of the transplanted cells was difficult to assess. These studies firmly established that hepatocyte transplants were an effective means to correct metabolic liver disease. The results of hepatocyte transplants of other patients with CN largely confirm those seen with the first patient (Dhawan et al., 2004; Ambrosino et al., 2005). Muraca et al. (2002) reported partial correction of glycogen storage disease, Type 1 following hepatocyte transplantation. Improvement was documented by the patient’s ability to maintain blood glucose between meals as well as sustained and higher glucose levels with meals. Sokal et al. (2003) employed hepatocyte transplants to achieve a partial correction of infantile Refsum disease an autosomal recessive inborn error in peroxisome metabolism of very-long chain fatty acid metabolism, bile acid, and pipecolic acid. The authors reported improvement in fatty acids metabolism, a reduction in circulating pipecolic acid and bile salt levels. An overall improvement in the health of the patient was evident by the report of significant increase in muscle strength and weight gain. Dhawan et al. (2004) reported that hepatocyte transplantation partially corrected a severe deficiency in the production and secretion of coagulation Factor VII. Following cell transplant, the Factor VII requirement decreased nearly 80% of that administered prior to HTx. Most recently, Stephenne et al. (2006) reported the complete correction of a 3.5-yearold female patient with neonatal onset arginosuccinate lyase (ASL) deficiency. Like OTC deficiency, ASL patients are at risk of brain damage from hyperammonemia. The patient received three-sequential hepatocyte transplants over a 5-month period. Both freshly isolated and previously cryopreserved hepatocytes were used. At 1 year post transplant the patient displayed 3% of normal ASL activity in hepatic biopsy samples. Engraftment of donor cells could be demonstrated by fluorescence in situ hybridization for Y chromosome. These results confirm that hepatocyte transplantation can achieve sustained engraftment of donor cells and sustained metabolic and clinical control.

HEPATOCYTE TRANSPLANTATION NOVEL USES, CHALLENGES, AND FUTURE DIRECTIONS Hepatocyte Transplants for Non-organ Transplant Candidates Most of the patients who have received a hepatocyte transplant were already listed for a whole organ transplant. The need for liver support is not limited to this group. There are large numbers of patients for whom OLT is not an option. Patients in this group could include alcoholic cirrhotic patients who have not met the required

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abstinence period, acute liver failure patients resulting from suicide attempts and cancer patients. Early case reports suggested that hepatocyte transplants into the spleen could be useful to restore liver function to end-stage cirrhotic patients (Strom et al., 1999). Although both of the patients in the reported study eventually died of concomitant renal failure which was left untreated, the patients were sufficiently improved following the cell transplants that they were able to be discharged from the hospital. Fox and coworkers created an animal model to study the efficacy of hepatocyte transplants to support liver function in cirrhosis in a more controlled setting. Their studies clearly demonstrated that hepatocyte transplants significantly improve liver function and survival of rats experiencing chronic liver failure following repeated injections of carbon tetrachloride (Ahmad et al., 2002; Cai et al., 2002; Nagata et al., 2003a). With millions of patients currently infected with hepatitis viruses there is clearly a need for additional means to support liver function in these patients. Not withstanding the difficulties of such clinical studies in cirrhotic patients, cell transplantation should be thoroughly evaluated as possible support therapy. The single most important factor preventing the use of hepatocyte transplants in additional medical centers is the limited availability of hepatocytes. The normal source of cells for hepatocyte transplants are livers with greater than 50% steatosis, vascular plaques, or other factors which render the tissue unsuitable for whole organ transplantation (Strom et al., 1997a, b; Fox et al., 1998; Bilir et al., 2000; Fisher et al., 2000; Muraca et al., 2002; Nakazawa et al., 2002; Soriano, 2002; Horslen et al., 2003; Mitry et al., 2003; Strom and Fisher, 2003; Ott et al., 2004). Hepatocyte transplants will not be able to progress past the small proof of concept studies in humans until sufficient numbers of hepatocytes become available (Strom and Fisher, 2003). Xenotransplants (Nagata et al., 2003a), immortalized human hepatocytes (Kobayashi et al., 2000; Cai et al., 2002; Wege et al., 2003a, b) and stem cell-derived hepatocytes (Avital et al., 2002; Miki et al., 2002, 2005; Davila et al., 2004; Ruhnke et al., 2005) and fetal hepatocytes have been proposed as alternative sources of cells for clinical transplants. To date, no alternative cell source has been found which meets all of the requirements for safety and efficacy. Because of the increased interest in stem cell-derived hepatocytes and scientific investigations into their production, it is likely that they will be a significant source of cells for future hepatocyte transplants. Better utilization of existing liver tissue could increase the numbers of hepatocytes available immediately. In the United States there are no regulations requiring that donor organs be allocated to transplantation research centers for hepatocyte isolation, and relatively few organs go to centers where hepatocyte transplant is a possibility. Most of the organs not used for whole organ transplant are provided to commercial firms where hepatocytes are isolated for resale or for in-house metabolism and toxicology studies. While most uses of donor liver tissue have merit, simple allocation procedures could be instituted to route the organs to transplant centers for initial review and selection of the most suitable cases for cell isolation. Split-liver procedures have made it possible to use caudate lobe and segment IV for hepatocyte isolation. Depending on the surgical procedure, these portions of liver tissue may remain untransplanted and have been shown to be useful for hepatocyte isolation (Mitry et al., 2004). Although, currently quite hypothetical, in the future most or all livers which are currently transplanted could be split. A portion such as the left lateral segment or the entire left lobe could be made available for cell isolation while the remaining liver tissue is utilized as a tissue graft. Because hepatocyte transplantation is not currently the standard of care, such proposals are not currently feasible. However, if the efficacy of hepatocyte transplants were firmly established, the risk and the extra time needed for the split procedure would be outweighed by the benefit of the cell transplants. Cell transplants rather than OLT could free-up the organs which are now used for acute liver failure and metabolic disease patients.

SUMMARY Hepatocyte transplantation studies conducted in animal models of liver failure and liver-based metabolic disease have proven safe and effective means to provide short- or long-term synthetic and metabolic support of

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liver function. For certain organ transplant candidates such as those with metabolic liver disease, cell transplantation alone could provide relief of the clinical symptoms. Cell transplant studies in patients with acute or chronic liver failure or genetic defects in liver function clearly demonstrate the efficacy of hepatocyte transplantation to treat liver disease. In virtually all cases a clinical improvement in the condition of the patient could be documented. No serious complications of hepatocyte transplant have been reported. Although all of the initial reports concerning hepatocyte transplants are encouraging, it must be realized that there are still no reports of long-term and complete corrections of any metabolic disease in patients. The recent report of a complete correction of a patient with a urea cycle defect is most encouraging; however, the length of time that human hepatocytes will function following transplantation has not been determined. Studies in animal models of liver disease have documented that donor hepatocytes transplanted into the spleen or the liver function for the lifetime of the recipient and participate in normal regenerative events. Although it is likely that human hepatocyte transplantation will result in lifelong and normal function of donor cells, this needs to be clearly demonstrated in a clinical study. Future work will have to be conducted to establish optimal transplant and immunosuppression protocols to minimize complications and maximize engraftment and function. A major problem for clinical hepatocyte transplant is the inability to track donor cells following transplantation. Except for the short-term tracking of hepatocytes pre-labeled with radioactive substances such as indium111 (Bohnen et al., 2000), there are no reports of quantitative and facile methods to detect donor cells. Relatively non-invasive methods will be needed to optimize transplant and immunosuppressive protocols as well as for day-to-day monitoring of the cell graft. None of the problems cited here seem insurmountable. There are now reports of successful hepatocyte transplants from laboratories in many different countries. The cooperative spirit which has developed between the investigators at the different transplant centers should benefit the research field and especially the future recipients of hepatocyte transplants.

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Mitry, R.R., Dhawan, A., Hughes, R.D., Bansal, S., Lehec, S., Terry, C., Heaton, N.D., Karani, J.B., Mieli-Vergani, G. and Rela, M. (2004). One liver, three recipients: segment IV from split-liver procedures as a source of hepatocytes for cell transplantation. Transplantation 77: 1614–1616. Moscioni, A.D., Roy-Chowdhury, J., Barbour, R., Brown, L.L., Roy-Chowdhury, N., Competiello, L.S., Lahiri, P. and Demetriou, A.A. (1989). Human liver cell transplantation. Prolonged function in athymic-Gunn and athymicanalbuminemic hybrid rats. Gastroenterology 96: 1546–1551. Moscioni, A.D., Rozga, J., Chen, S., Naim, A., Scott, H.S. and Demetriou, A.A. (1996). Long-term correction of albumin levels in the Nagase analbuminemic rat: repopulation of the liver by transplanted normal hepatocytes under a regeneration response. Cell Transplant. 5: 499–503. Muraca, M., Gerunda, G., Neri, D., Vilei, M.T., Granato, A., Feltracco, P., Meroni, M., Giron, G. and Burlina, A.B. (2002). Hepatocyte transplantation as a treatment for glycogen storage disease type 1a. Lancet 359: 317–318. Nagata, H., Ito, M., Cai, J., Edge, A.S., Platt, J.L. and Fox, I.J. (2003a). Treatment of cirrhosis and liver failure in rats by hepatocyte xenotransplantation. Gastroenterology 124: 422–431. Nagata, H., Ito, M., Shirota, C., Edge, A., McCowan, T.C. and Fox, I.J. (2003b). Route of hepatocyte delivery affects hepatocyte engraftment in the spleen. Transplantation 76: 732–734. Nakazawa, F., Onodera, K., Kato, K., Sawa, M., Kino, Y., Imai, M., Kasai, S., Mito, M., Matsushita, T. and Funatsu, K. (1996). Multilocational hepatocyte transplantation for treatment of congenital ascorbic acid deficiency rats. Cell Transplant. 5: S23–S25. Nakazawa, F., Cai, H., Miki, T., Dorko, K., Abdelmeguid, A., Walldorf, J., Lehmann, T. and Strom, S. (2002). Human hepatocyte isolation from cadaver donor liver. In Proceedings of Falk Symposium, Hepatocyte Transplantation, Vol. 126, Kouwer Academic Publishers, Lancaster, UK, pp. 147–158. Oertel, M., Menthena, A., Dabeva, M.D. and Shafritz, D.A. (2006). Cell competition leads to a high level of normal liver reconstitution by transplanted fetal liver stem/progenitor cells. Gastroenterology 130: 507–520; quiz 590. Ohashi, K., Park, F. and Kay, M.A. (2001). Hepatocyte transplantation: clinical and experimental application. J. Mol. Med. 79: 617–630. Onodera, K., Kasai, S., Kato, K., Nakazawa, F. and Mito, M. (1995). Long-term effect of intrasplenic hepatocyte transplantation in congenitally ascorbic acid biosynthetic enzyme-deficient rats. Cell Transplant. 4(Suppl 1): S41–S43. Oren, R., Dabeva, M.D., Petkov, P.M., Hurston, E., Laconi, E. and Shafritz, D.A. (1999). Restoration of serum albumin levels in nagase analbuminemic rats by hepatocyte transplantation. Hepatology 29: 75–81. Ott, M.C., Barthold, M., Alexandrova, K., Griesel, C., Shchneider, A., Attaran, M., Arsenieva, M., Penkov, B., Net, M., Peralta, V., Bredehorn, T., Manyalich, M., Kafert-Kasting, S., Manns, M. P., Dimitrova, V., Nachkov, Y. and LArseniev, L. (2004). Isolation of human hepatocytes from donor organs under cgmp conditions and clinical application in patients with liver disease. 7th International Congress of Cell Transplantation Society, Boston 142. Ott, M., Schneider, A., Attaran, M. and Manns, M.P. (2006). Transplantation of hepatocytes in liver failure. Dtsch. Med. Wochenschr. 131: 888–891. Overturf, K., Al-Dhalimy, M., Tanguay, R., Brantly, M., Ou, C.N., Finegold, M. and Grompe, M. (1996). Hepatocytes corrected by gene therapy are selected in vivo in a murine model of hereditary tyrosinaemia type I. Nat. Genet. 12: 266–273. Ponder, K.P., Gupta, S., Leland, F., Darlington, G., Finegold, M., DeMayo, J., Ledley, F.D., Chowdhury, J.R. and Woo, S.L. (1991). Mouse hepatocytes migrate to liver parenchyma and function indefinitely after intrasplenic transplantation. Proc. Natl Acad. Sci. USA 88: 1217–1221. Rhim, J.A., Sandgren, E.P., Degen, J.L., Palmiter, R.D. and Brinster, R.L. (1994). Replacement of diseased mouse liver by hepatic cell transplantation. Science 263: 1149–1152. Rozga, J., Holzman, M., Moscioni, A.D., Fujioka, H., Morsiani, E. and Demetriou, A.A. (1995). Repeated intraportal hepatocyte transplantation in analbuminemic rats. Cell Transplant. 4: 237–243. Rudnick, D.A. and Perlmutter, D.H. (2005). Alpha-1-antitrypsin deficiency: a new paradigm for hepatocellular carcinoma in genetic liver disease. Hepatology 42: 514–521. Ruhnke, M., Nussler, A.K., Ungefroren, H., Hengstler, J.G., Kremer, B., Hoeckh, W., Gottwald, T., Heeckt, P. and Fandrich, F. (2005). Human monocyte-derived neohepatocytes: a promising alternative to primary human hepatocytes for autologous cell therapy. Transplantation 79: 1097–1103.

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Sandgren, E.P., Palmiter, R.D., Heckel, J.L., Daugherty, C.C., Brinster, R.L. and Degen, J.L. (1991). Complete hepatic regeneration after somatic deletion of an albumin-plasminogen activator transgene. Cell 66: 245–256. Shani-Peretz, H., Tsiperson, V., Shoshani, G., Veitzman, E., Neufeld, G. and Baruch, Y. (2005). HVEGF165 increases survival of transplanted hepatocytes within portal radicles: suggested mechanism for early cell engraftment. Cell Transplant. 14: 49–57. Sokal, E.M., Smets, F., Bourgois, A., Van Maldergem, L., Buts, J.P., Reding, R., Bernard Otte, J., Evrard, V., Latinne, D., Vincent, M.F., Moser, A. and Soriano, H.E. (2003). Hepatocyte transplantation in a 4-year-old girl with peroxisomal biogenesis disease: technique, safety, and metabolic follow-up. Transplantation 76: 735–738. Sommer, B.G., Sutherland, D.E., Matas, A.J., Simmons, R.L. and Najarian, J.S. (1979). Hepatocellular transplantation for treatment of D-galactosamine-induced acute liver failure in rats. Transplant. Proc. 11: 578–584. Soriano, H.E. (2002). Liver cell transplantation: human applications in adults and children. In Proceedings of Falk Symposium, Hepatocyte Transplantation, Vol. 126, pp. 99–115. Kouwer Academic Publishers, Lancaster, UK. Soriano, H.E., Kang, D.C., Finegold, M.J., Hicks, M.J., Wang, N.D., Harrison, W. and Darlington, G.J. (1998). Lack of C/EBP alpha gene expression results in increased DNA synthesis and an increased frequency of immortalization of freshly isolated mice [correction of rat] hepatocytes. Hepatology 27: 392–401. Stephenne, X., Najimi, M., Smets, F., Reding, R., de Ville de Goyet, J. and Sokal, E.M. (2005). Cryopreserved liver cell transplantation controls ornithine transcarbamylase deficient patient while awaiting liver transplantation. Am. J. Transplant. 5: 2058–2061. Stephenne, X., Najimi, M., Sibille, C., Nassogne, M.C., Smets, F. and Sokal, E.M. (2006). Sustained engraftment and tissue enzyme activity after liver cell transplantation for argininosuccinate lyase deficiency. Gastroenterology 130: 1317–1323. Strom, S. and Fisher, R. (2003). Hepatocyte transplantation: new possibilities for therapy. Gastroenterology 124: 568–571. Strom, S.C., Fisher, R.A., Rubinstein, W.S., Barranger, J.A., Towbin, R.B., Charron, M., Mieles, L., Pisarov, L.A., Dorko, K., Thompson, M.T. and Reyes, J. (1997a). Transplantation of human hepatocytes. Transplant. Proc. 29: 2103–2106. Strom, S.C., Fisher, R.A., Thompson, M.T., Sanyal, A.J., Cole, P.E., Ham, J.M. and Posner, M.P. (1997b). Hepatocyte transplantation as a bridge to orthotopic liver transplantation in terminal liver failure. Transplantation 63: 559–569. Strom, S.C., Chowdhury, J.R. and Fox, I.J. (1999). Hepatocyte transplantation for the treatment of human disease. Semin. Liver Dis. 19: 39–48. Strom, S., Bruzzone, P., Cai, H., Ellis, E., Lehmann, T., Mitamura, K. and Miki, T. (2006). Hepatocyte Transplantation: Clinical Experience and Potential for Future Use. Cell Transplantation. 15: S105–S110. Sutherland, D.E., Numata, M., Matas, A.J., Simmons, R.L. and Najarian, J.S. (1977). Hepatocellular transplantation in acute liver failure. Surgery 82: 124–132. Takeshita, K., Ishibashi, H., Suzuki, M. and Kodama, M. (1993). Hepatocellular transplantation for metabolic support in experimental acute ischemic liver failure in rats. Cell Transplant. 2: 319–324. Vroemen, J.P., Buurman, W.A., Heirwegh, K.P., van der Linden, C.J. and Kootstra, G. (1986). Hepatocyte transplantation for enzyme deficiency disease in congenic rats. Transplantation 42: 130–135. Wege, H., Chui, M.S., Le, H.T., Strom, S. and Zern, M.A. (2003a). In vitro expansion of human hepatocytes is restricted by telomere-dependent replicative aging. Cell Transplant. 12: 897–906. Wege, H., Le, H.T., Chui, M.S., Liu, L., Wu, J., Giri, R., Malhi, H., Sappal, B.S., Kumaran, V., Gupta, S. and Zern, M.A. (2003b). Telomerase reconstitution immortalizes human fetal hepatocytes without disrupting their differentiation potential. Gastroenterology 124: 432–444. Yoshida, Y., Tokusashi, Y., Lee, G.H. and Ogawa, K. (1996). Intrahepatic transplantation of normal hepatocytes prevents Wilson’s disease in Long-Evans cinnamon rats. Gastroenterology 111: 1654–1660.

53 Bioartificial Livers Randall E. McClelland and Lola M. Reid INTRODUCTION The development of bioartificial livers resulted from the need to extend the lives of patients confronted with liver failure, given that the liver, like the heart, is an organ system that does not come in pairs, as do lungs or kidneys, and is solely responsible for functions (Gebhardt, 1992; Alberts et al., 1995; Anderson et al., 1996) that are critical for survival. Reports of liver treatment can be dated from the 1950s when low protein diets were recommended to improve mental impairment and hepatic encephalopathy (Soulsby, 1999) and the 1960s for novel concepts of liver assist devices (Kimoto, 1959; Allen and Bhatia, 2002). Some of these artificial assist systems have entered into preliminary Food and Drug Administration (FDA) trials due to their abilities to support patients suffering from liver failure or less serious liver malfunctions. These systems include charcoal filters for ammonia detoxification (Malchesky, 1994; Sussman and Kelly, 1996), mechanical dialysis permitting toxin transfers (Malchesky, 1994), and plasmapheresis for removal of diseased circulating substances (Gislason et al., 1994; Malchesky, 1994). Investigations have shown that many liver assist devices are successful but sometimes limited in broad functional capabilities (Sussman and Kelly, 1996). Although improvements of liver assist devices are frequently updated to improve market potential – such as making use of advanced design parameters found in kidney dialysis machines (Colton, 1999; Wright et al., 2002) – the multitude of tasks performed by a healthy in vivo liver cannot be adequately reproduced by these systems. Bioartificial livers are used also as an adjunct to liver transplantation. Organ transplantation can be highly successful with the caveat that the recipient patient must remain on immunosuppressive drugs to eliminate organ rejection. The need for bioartificial livers, even with the existence of successful organ transplantation strategies, is due to the fact that donor livers are not readily available. There are approximately 5,000 donor organs/year and 3–4 times that number of people on waiting lists to get them (NIDDK, 2002; UNOS, 2002). Bioartificial livers are already used in clinical programs to enable patient survival until an organ donor is available. Thus, the use of bioartificial livers remains an important option even though current bioreactor designs remain limited due to the inadequacy of organ donor sources and to remaining engineering challenges required to optimize nutrient transfer to cells. STRATEGIES FOR EX VIVO MAINTENANCE OF CELLS Strategies for maintaining differentiated cells ex vivo have been dominated, for approximately 100 years, by 2-dimensional (2D) culture formats. These formats include monolayer seeded cells submerged in nutrient medium with passive gas exchange controlled by regulated incubator environments. Cells attach rapidly, within hours, and are easily evaluated by microscopy for morphology and by biochemical or immunochemical analyses for functions. As illustrated by the light micrographs in Figure 53.1, both adult rat and cryopreserved adult human hepatocytes, in such monolayers, display similar morphologies when plated on collagen type I. In Figure 53.1a and b, the newly seeded cells are spherical and with distinct cell boundaries and within a few days have attached, flattened, and established linkages to neighboring cells as displayed in Figure 53.1c

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Adult rat cells Fresh liver Seed day, day 0

Adult human cells Cryopreserved liver

(b)

(c)

(d)

Hepatocytes & STO, day 3

Hepatocytes, day 3

(a)

(e)

(f)

Figure 53.1 Comparisons of the morphology of cultures of freshly isolated adult rat hepatocytes (left panels) versus cryopreserved adult human hepatocytes (right panels). As shown on the day of seeding, day 0, both freshly isolated rat hepatocytes and cryopreserved human hepatocytes ((a) and (b)) display segmented and rounded cells with distinct borders. By day 3, hepatocyte only cultures, the rat tissues have merged (c) while the human hepatocytes continue with their migration routes (d). For the day 3 hepatocyte and stromal feeder cocultures, the rat tissues are intertwined with stromal cells – displayed as thinly lined cultures – which are concurrently visualized (e); while human hepatocytes have merged into well-defined segmented tissue structures (f). and d. Under these conditions, the cells can be maintained for 7 day time periods. Longevity of the cultures can be increased by approximately a week and tissue-specific functions maintained more stably by altering the microenvironment such as by using feeder cells (Figure 53.1e and f), by using defined mixtures of regulatory signals, and by minimizing or eliminating serum from the media (Reid, 1990; MacDonald et al., 2002). The 2D cultures are constrained by the numbers of cells (e.g. 1–10 million) that can be feasibly handled and by the muted differentiated functions typical of these systems. Alternate strategies have emerged to handle cells in a 3D format and at high densities (e.g. 109–12 cells) for maximal differentiation (MacDonald et al., 1999, 2002). Use of a 3D format necessitates dynamic perfusion of cells and tissues to achieve mass transfer of nutrients. In vivo this is achieved by angiogenesis and vascularization (indicated in schematic form in Figure 53.2a and b). Over the past approximately 50 years, there has been steady progression towards bioreactors that offer variations in designs to provide 3D culture options (MacDonald et al., 1999). For these 3D systems, adequate and stable nutrient concentrations must be accessible at the cultures surface and interior core locations in order to prevent cell death associated with heterogeneous nutrient gradients. As shown in both images, the 3D outer membranes or shell casings must enclose networks of 2D transport channels to support large tissue masses. In general, this 2D scenario is found throughout the liver as sinusoidal pathways supporting aggregates of cells folded onto each other – forming

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Figure 53.2 A cartoon of the liver lobule showing cords of hepatocytes demarcated with respect to functional zones 1–3 using three distinct colors and surrounded by a membrane surface (a). The blood supply supporting the liver cells comes in at hepatic arteries and portal veins (:O), then flows across the liver’s sinusoids, and empties into the central vein (CV). A bioreactor is displayed using a perfusion bed design contained inside and outer casing (b). Pathways of nutrient transport are shown as input and output sources that mimic sinusoids. The layered brick background represents matrix-embedded hepatocytes. A cross-sectional view shows nutrient source alignments throughout this system. 3D tissue masses. Bioreactors are unable to duplicate geometrically complex designs of vascular beds that occur in vivo (e.g. the liver’s sinusoids). So, numerous cell seeding techniques (MacDonald and Wolfe, 1999) and unique microchannel arrays (McClelland and Coger, 2000, 2003, 2004) are exploited to decrease transport resistances with the goal of achieving representative nutrient concentration gradients.

BIOARTIFICIAL LIVER DESIGNS Bioartificial livers (Rozga et al., 1993a; LePage et al., 1994; Chen S.C. et al., 1996; Arkadopoulos et al., 1998; Brusse and Gerlach, 1999; MacDonald et al., 1999; Patzer et al., 1999; Watanabe et al., 1999) all consist of: (1) a bioreactor – a support structure with a cell compartment that is connected to channels such as hollow fibers for supplying essential nutrients and gases; (2) cells that form tissue within the cell compartment; and (3) a microenvironment comprising a nutrient medium, extracellular matrix components (or artificial scaffoldings), and serum supplements and/or purified hormones and growth factors. Of the more than 40 designs for bioreactors (Gerlach, 1996; Brusse and Gerlach, 1999; MacDonald et al., 1999), the most common are variants of three types. Flat Plate (or Flat Bed) Bioreactors The flat plate (FP) design in Figure 53.3a is the simplest and allows rapid analyses of cells in monolayer configurations. A bottom matrix layer can be used when matrices are known to induce extended times

Bioartificial Livers 931

(a)

(c)

Flat plate

Spheroid = matrix

Hollow fiber (b)

(d)

Microcarrier

= cells & matrix

Figure 53.3 Four bioartificial liver designs. The FP design consists of hepatocytes cultured on top of a supporting matrix such as type I collagen (a). The hollow fiber design is a concentric cylinder containing a cell compartment traversed by hollow fibers of varying chemistry and used as supply lines create (b). The spheroid design consists of cells and matrix components within the core of a container (c). The microcarrier design consists of cells bound to extracellular matrix components attached to the surface of a microcarrier (d).

(7–10 days) of culture viabilities (Dunn et al., 1991). A second matrix layer may be placed above the cells to “sandwich” them and initiate top and bottom cell-matrix attachments, a design applicable for liver cells that exist in vivo between two matrix layers. This sandwich design is yet another way to further extend (2 months) the viable and functional responses of cultured cells (Dunn et al., 1991). FP designs are relatively simple and have been extensively investigated (Dunn et al., 1991; Koike et al., 1996). Both cell morphology and functional responses are considered baseline results and used as standard responses for comparison analyses when investigating cellular effects inside other bioartificial liver designs. There are two subclasses of this category:

•

•

High throughput designs include “cells-on-a-chip” (e.g. those developed by Dr. Linda Griffith or by Dr. Sangeeta Bhatia and their associates at MIT) that are utilized for small numbers of cells and small volumes of culture medium and reagents. They are ideal for rapid surveys of large numbers of drugs or factors but do not permit tissue in significant amounts, thereby obviating or minimizing the possibility to do extensive biochemistry, cell, or molecular biology (Bhatia et al., 1994; Griffith et al., 1997; Allen et al., 2001; Powers et al., 2002). Large-scale flat bed bioreactors (Bader et al., 1995a, b, 1996; DeBartolo et al., 2000) in which cells are plated as monolayers onto a very large surfaces (culture plastic, synthetic or natural scaffolding, or onto feeder cells) and the nutrient medium is perfused over the cell layer. The flat bed bioreactors can be of sufficient size to accommodate hundreds of millions of cells and offer long-term viability of cells (typically 3–4 weeks or more) with retention of expression of tissue-specific functions as long as the cells are given appropriate matrix substrata, feeder layers, and relevant soluble signals. The limitations for flat bed bioreactors are that they do not permit establishment of 3-dimensionality required for the high densities of cells (billions of cells) required to achieve maximal differentiated functions.

932 THERAPEUTIC APPLICATIONS: CELL THERAPY

Hollow Fiber Bioreactors The hollow fiber design depicted in Figure 53.3b is a concentric cylinder containing a cell compartment and with hollow fibers used as the supply lines. HF bioreactors were first demonstrated by Wolf and Munkelt (Wolf and Munkelt, 1975), who cultured hepatocytes in the extracapillary space (ECS) of a bioreactor derivative of an original design by Knazek and Gullino (Knazek et al., 1972). The hollow fibers are semi-permeable with easily modified attributes (e.g. cell wall thickness, inner and outer diameters, pore sizes, etc.) and prepared by extrusion technologies from various polymers. These are the only bioreactors in which organ tissues achieve full differentiation (Gerlach, 1996; Busse and Gerlach, 1999; MacDonald et al., 1999). Two major advantages of the HF design are: (1) the increased cell-surface area used to seed the cells and (2) the ability to separate both liquid and gaseous nutrient sources such that the cell-core is maximally penetrated by each nutrient phase (e.g. gas or liquid). The benefits of cylindrical geometries are that the number of cultured cells supported by single nutrient sources may expand by 6-fold – as a way to compact the overall design (McClelland and Coger, 2000). Then by employing cylindrical hollow fiber tubes that are easily modified and interchanged, such that various material characteristics and porosity fractions may adjust nutrient mass transfer rates, the nutrient concentrations can be more effectively controlled as they disperse in both radial and longitudinal directions inside the bioartificial liver’s cell-core. The advantage of the hollow fiber bioreactors is that large tissue masses can be achieved providing high levels of tissue-specific functions (if normal cells are used). The limitations are that hollow fiber bioreactor designs shut down quickly, within a few days, due to the natural tendency of adherent cell types to attach to the hollow fibers and deposit cellular materials onto them resulting in “fouling” of them and blockage of the mass transfer of nutrients to the cells (MacDonald et al., 2001a). An especially important result of this problem is the transfer of oxygen (O2). In vivo all cells are within 500 μm from a vascular supply, and liver cells are even closer, being within 50–100 μm of their blood supply. The cell compartments of hollow fiber bioreactors have little or no microvasculature and almost all of them are well beyond the 50–100 μm limit for liver cells to be distant from a source of nutrients. The semi-permeable hollow fibers are minimally sufficient to accommodate reasonable flow of nutrients (MacDonald and Wolfe, 1999; Wolfe et al., 2002). This limitation affects O2 with its diminutive diffusion coefficient through tissues, as compared to other nutrient diffusion coefficients. Convective pathways for nutrient distribution are generally unavailable inside the bioreactor’s compartments, so nutrient diffusion rates limit 3D expansions and restrict geometrical configurations. Other drawbacks to them are the extreme difficulties in retrieving tissue and the lack of visibility during the time when the cells are within the system. Currently, the HF bioartificial livers are the most clinically investigated bioreactors with four different prototypes being utilized in FDA trials (Allen et al., 2001). Some variables from in vivo may be mimicked by modifications to HF designs (Figure 53.4). In this figure, the hollow fiber is displayed with three unique modifications – labeled A, B, and C – which have been segmented (e.g. micropatterned) along the outer HF surface. Area A signifies a Matrigel-coated subdivision with attached hepatocytes. Matrigel causes the liver cells to aggregate into spheroids (Joly et al., 1997; Hamamoto et al., 1998; Funatsu and Nakazawa, 2002). This is beneficial in that the differentiated cell functions are known to increase when cells are in a spheroid format (Joly et al., 1997; Hamamoto et al., 1998). Area B is collagen type I substratum that causes hepatocytes to flatten and spread (Dunn et al., 1991; Koike et al., 1996). Area C combines the borders of two individual surface modifications such that asymmetrical cell responses are mechanically forced to interact. This merged border is beneficial in terms of interconnecting specific amounts of parenchymal and non-parenchymal cell phenotypes and promoting cellular communications. In this way, specific cell types and precise distribution patterns of in vivo organs may be imitated using micropatterning technologies. Thus, the ability to modify the receptor–ligand interactions between artificial surfaces and living tissues is currently one alternative being used to improve bioreactor devices.

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Gas nutrients

Concentric hollow fiber cylinders

Liquid nutrients Area A

Area B

Area C

Figure 53.4 The outer surface of a hollow fiber cylinder with three unique surfaces used to exclusively activate distinctive receptor–ligand cell attachments. Area A indicates spheroids of hepatocytes attached onto Matrigel. Area B indicates hepatocytes that have spread into a monolayer when on type I collagen. Area C shows the interaction between both culture systems as different cell morphologies are brought together.

HF BIOREACTORS SUCCESSFUL WITH LIVER A larger cell mass, as required in bioreactors for organs such as liver, needs integral oxygenation and subsequently a four-compartment (4-C) cell culture space (Xu et al., 2000). However, only two closed bioreactors achieve this in combination with a low shear environment and result in maintenance of differentiated functions: a multicoaxial bioreactor (MCB) developed by Dr. Jeff MacDonald (MacDonald et al., 2001a; Wolfe et al., 2002) and brought to functional reality by Drs. Randall McClelland and Robin Coger (McClelland and Coger, 2000, 2003, 2004) and a woven hollow fiber bioreactor developed by Dr. Joerg Gerlach (Gerlach, 1994; Gerlach et al., 1996, 1997; Brusse and Gerlach, 1999).The MCB design can be easily modified by changing the relative diameters of the two coaxial tubes that are rigorously maintained axially, thereby maintaining control over mass transfer of nutrients. Extensive characterization of rodent liver cells in the MCB system has been done by McClelland and Coger who showed that the reason for this superior performance is due to homogenous and reliable mass transfer, properties that were extensively modeled by McClelland (McClelland and Coger, 2000, 2003, 2004). This design mimics the liver acinus architecture permitting replication of scale (MacDonald and Wolfe, 1999; Wolfe et al., 2002). The only bioreactor that achieves integral oxygenation and decentralized mass exchange and has been engineered to both experimental scale and to clinical scale is that developed by Gerlach and associates (Gerlach et al., 1997; Busse and Gerlach, 1999; Busse et al., 1999; Gerlach et al., 2002, 2003). To enable integral oxygenation and distributed mass exchange with low gradients, Gerlach and associates developed a 4-C bioreactor specific for clinical liver support and that accommodates 400–800 g of normal liver cells. Their approach incorporated the strategy of spontaneous re-assembly of cells into tissues and with synthesis of their own native extracellular matrix after inoculation into 4-C bioreactors. Gerlach and associates have shown that a homogeneous mixture of a suspension of adult human liver cells will reassemble as liver tissue after injection into such a bioreactor to form well-defined liver structures, such as neo-sinusoidal structures and neo-space of Dissé. Two main determinants of this tissue re-assembly were identified: (1) hepatocytes always aggregated between the interwoven artificial capillary beds, if the technology provided low micro-environmental gradients as well as integral oxygenation and CO2 removal and (2) perfusion channels within the aggregates occurred regularly, if the medium flow between the artificial capillaries was enabled and if co-culture with appropriate non-parenchymal cells was employed. As a result, the endothelial cells regularly re-endothelialized

934 THERAPEUTIC APPLICATIONS: CELL THERAPY

perfusion channels between hepatocytes, and stellate cells while extracellular matrix depletion was seen regularly.

METHODS OF INTRODUCTION OF CELLS TO THE CELL COMPARTMENT These HF systems can be seeded with cells alone or in extracellular matrix (HF) (Figure 53.3b) with cells in spheroid encapsulation (SE) format (Figure 53.3c), or with cells bound onto microcarriers (perfusion scaffold (PS) systems) (Figure 53.3d). The spheroid design version shown in Figure 53.3c is an ideal structure; in that spherical geometries take advantage of maximum surface areas in which cells may attach when in a confined space. The microcarrier design in Figure 53.3d takes advantage of spheroid structures by allowing cells to aggregate into small subsets, which are then disseminated around nutrient sources. Straightforward modifications of aggregate size and cell densities distributions help to maximize best-case scenarios when monitoring cell function responsiveness. Given the relative success of HF, further information is given on them. Cells can be introduced on their own, in combination with scaffolds, encapsulated or bound to microcarriers. Liver cells do best when there is recognition of their critical adhesion mechanisms. HF Bioartificial Livers with Cells Encapsulated as Spheroids Encapsulation of cells (e.g. in alginate) to generate spheroids prior to seeding them in HFs is beneficial in that cells are more tightly grouped together and achieve maximal differentiated functions in a format that also shields the aggregates of cells from shear forces (Figure 53.3c) as external liquid nutrients are dynamically circulated throughout the HF. These dynamic flows are utilized to produce concentration baths of equivalent dilutions, such that metabolic support is readily available and not influenced by irregular gradients. HF Bioartificial Livers with Cells on PSs The PS systems are unique in their ability to percolate nutrients throughout 3D constructs. In these systems, cells are supported within porous gels or on microcarriers such that nutrients permeate from source pathways and ultimately fill enclosed culture containers. After filling and saturating the constructs, spent nutrients are released due to overcapacity or induced pressure gradients and recycled back into the source pathways. This process continues for 24-hour periods, at which time fresh nutrients replenish the system. The benefits of PS designs are their simplified scale-up capabilities; where the addition of extra scaffold constructs allows for significantly more cellular add-ons. However, when expanding this type of bioartificial liver design, distinguishable concentration gradients may arise as nutrient sources are further dislodged from cell seed locations. Thus, it is imperative to adjust nutrient concentrations such that nutrient sources offer adequate distributions throughout its constructs.

COMPUTER-REGULATED BIOREACTORS Each of the bioartificial liver designs entails distinctive microenvironments used to stimulate particular cell responses and provides a system in which to sustain millions to billions of hepatocytes in a defined microenvironment. To examine the particulars of cell interactions in each bioreactor, detailed microenvironments have been developed using computer “circuit board” technologies, as shown with Figure 53.5. That is, instead of positioning elaborate pathways of electrical lines across computer “motherboard” surfaces, the lines are replaced with nutrient conduits that are computer actuated to deliver or cut-off life supporting sources. In Figure 53.5, electrical feedback loops interconnect biochips, computers, pump actuators, and monitors such that varying amounts of nutrient metabolites (e.g. media, O2, CO2) may be exclusively dispersed in response to cell sensing nanobots (nanoscaled devices positioned near seeded cells to either obtain information or perform

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Feed back loop

Biochip culture

O2

CO2

Media

Pump

Nutrient input (s) = Cells and matrix

= Cells attachment surfaces

Computer Monitor = Nutrient source inlet/outlet

Figure 53.5 A biochip culture technique where the microenvironment is acutely controlled. The biochip is modified to permit cell attachment while pathways for nutrient support are entwined within the chip structure. Pumps interconnecting nutrient sources, biochips, and computer-controlled software packages are dynamically integrated with feedback loops and monitoring systems. specific duties). In this way, by modifying the motherboard surfaces such that the hepatocyte receptor–ligand chemistries cause surface binding – creating a biochip, a cells microenvironment can be acutely controlled while its reactions are instantaneously monitored. This biochip technology may revolutionize how bioartificial livers are designed, how they are loaded with cultures, and how nutrients perfuse throughout their structures, such that cell proliferation and differentiation may be purposely activated.

MODELING FOR FUTURE DESIGN IMPROVEMENTS Modeling of cells and tissues within specific bioreactors can now be done using super computers (Storaasli et al., 1993; Marcin et al., 2000), commercial software programs such as Fluent, Gambit, and Fidap (Fluent, Inc.; Lebanon, New Hampshire), mathematical solution codes such as Maple (Maplesoft; Waterloo, Canada), Matlab, and Mathcad, and computer design packages (e.g. Pro E, Autocad). The ability to computationally predict cellular responses within diverse bioartificial liver designs can make use of design standards during development phases. In this way, both 2D and 3D computer designs of original tissue organs, cell-based tissue cultures, and inert support structures may be organized into composite grid subdivisions such that intermingled reactions of computational systems occur simultaneously. In this way, a natural liver can be segmented via superimposed finite difference grids, where these subdivisions are then designated to function as specific tissue tasks. These tasks may then be appended into discrete locations of other bioartificial liver structures – to promote similar in-vivo and in-vitro modeling techniques and results. To visualize this modeling process, Figure 53.6a illustrates a segmented natural liver using a meshing scheme consistent with finite difference analysis. To contrast the geometrical shape of this in vivo organ, a bioartificial liver prototype donated by Dr. Jeff MacDonald of UNC Chapel Hill is visualized in Figure 53.6b. In this bioartificial liver structure, the four input channels on the left extremity that are connected with the four output channels on the right extremity help to illustrate a bioreactor that contains four distinctive interior annular spaces. These spaces, which are concealed by the bioartificial liver’s opaque shell, are detached from each other such that one represents the cell-core space while others contain supporting nutrients for metabolic stability. For both the natural liver and its bioartificial liver complement designs, the goals are to synthesize proteins, metabolize drugs, have enzymatic activity, and conjugate bilirubin as ways to maintain body homeostasis. Therefore for modeling purposes, the tasks assigned to each grid subdivision are representative of cellular reactions and/or nutrient flux distributions. This common assignment of tasks between the natural liver and bioartificial liver

936 THERAPEUTIC APPLICATIONS: CELL THERAPY

Protein synthesis

Drug metabolism

Liver

(a) Enzyme activity

Bilirubin conjugation (b)

Bioartifical liver

(BAL)

Figure 53.6 Comparisons between the liver in vivo (a) and a bioartificial liver (bioartificial liver) prototype developed by Dr. Jeffrey MacDonald and modified by McClelland and Coger and (b). Liver is segmented with finite difference meshing schemes to visualize how computational fragments may be interconnected. By labeling the nodal points as black dots along the intersecting meshes, finite difference modeling may be used to replicate liver activities. For the bioartificial liver designs the cellular functions would mimic liver, but nutrient sources are uniquely different as shown by eight access ports at the reactor’s extremities. In both systems, the functional goals are to have normal tissue activities handing synthesis, metabolism, enzymatic digestion, and conjugation.

models can be labeled via nodal positions to distinguish the locality and interconnectedness of rendered functions, shown as an array of black dots (nodal points) in Figures 53.6a and 53.7a. By interconnecting the grids and applying input, output, generation, and uptake conditions, then dynamic responses of nutrient flow through the bioartificial liver’s organic and inorganic components may be graphically illustrated and numerically monitored by solving matrix-derived linear algebra models as shown in Figure 53.7c. Thus by predicting concentration variances, then cell viabilities and functional relationships may be correlated within the system modeling predictions. Subsequently, innate research benefits arise from the ability to computationally modify all parts of the computer structure and to analyze each component prior to beginning experimental investigations. Here, analyzed computer modifications of cell density, cylinder thickness, cylinder lengths, material properties, nutrient concentrations, flow characteristics, etc. are ways to decrease both capital costs and large experimental lab times. Finally, the project details are pre-confirmed with computation analysis such that only trivial experimental changes are necessary to produce functional, improved, and novel bioartificial liver systems.

BIOLOGICAL AND CELL SOURCING ISSUES The biological issues consist of those associated with epithelial–mesenchymal relationships or with stem cells and maturational lineages and are discussed at length in a separate chapter (Cheng et al., 2007) and in a recent review on stem cells (Schmelzer et al., 2006). Here will be presented a summary of the cell types utilized in bioartificial livers being used clinically (Allen et al., 2001; Sauer et al., 2003). The cell types utilized in liver assist devices are hepatoma cell lines (Sussman and Kelly, 1993), porcine hepatocytes (Demetriou et al., 1995; Chen S. et al., 1996), and human hepatocytes (Kamlot et al., 1995; Bornemann et al., 1996; Brusse and Gerlach, 1999; Gerlach and Zeilinger, 2002). Human hepatoma cell lines are easy to use in the bioreactors but are so muted in their differentiated functions as to be of minimal use for support of patients (Sussman and

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(b) Cross-sectional view

(a) “Meshed” hollow fiber conical cylinders

Initial condition Finite difference Nodes of resistance networks

*

Ce ac llula tiv r ity

=

Dynamic concentration responses

Initial condition (c)) Linear algebra model to predict nutrient transport

Figure 53.7 A finite difference meshing scheme applied to the exterior and interior bioartificial liver components (a). Included on this mesh is nodal positions labeled as arrays of black dots. By cross-sectioning the device (b), the grid mesh is displayed as intertwined subsections of concentric hollow fibers. Using a linear algebra technique and matrix multiplication (c), the nodal resistance networks are combined with cell activity and initial conditions to predict dynamic nutrient responses inside the bioartificial liver.

Kelly, 1993; Gerlach, 1996). Embryonic stem cell lines could, theoretically, be lineage restricted to fully mature liver tissue, but this has not yet been accomplished (O’Shea, 1999; Levenberg et al., 2004). Porcine hepatocytes have proven partially effective (Rozga et al., 1993a; Demetriou et al., 1995; Watanabe et al., 1999) but have been found inadequate because of distinct metabolic responses (Gerlach et al., 2002) and because of ongoing concerns about viruses (Irgang et al., 2003). Immunological concerns have been by use of cells from genetically transformed porcine animals (Bornemann et al., 1996). This minimized response also occurs by means of membrane separations in the direct contact between species are eliminated but metabolic reactions remain active as metabolites are transported through membrane pores. The best clinical results have occurred with use of fresh or cryopreserved human liver cells (Gerlach et al., 2002). Unfortunately, it is impossible at present to convert to this option given the difficulties in obtaining high-quality, mature human liver cells. Sourcing of human liver cells is discussed at length in a separate review in the book (Cheng et al., 2007). This severe sourcing problem is now presumed to be overcome in the near future by the use of human hepatic stem cells (Schmelzer et al., 2006; Sicklick et al., 2006). Their expansion potential is sufficient to supply the number of human liver cells needed; their usefulness, therefore, is dependent on the ability to differentiate them to fully mature hepatocytes, a goal that is the focus of various studies.

FUNCTIONAL ANALYSIS The ability to determine stability conditions for bioreactor cultures is necessary to understand functional “timelines,” whereas these timelines are integrated with metabolic activities inside bioartificial liver devices (Demetriou et al., 1995; Gerlach, 1996; Brusse and Gerlach, 1999; Watanabe et al., 1999; Sauer et al., 2003).

938 THERAPEUTIC APPLICATIONS: CELL THERAPY

All of the devices used clinically are opaque obviating visibility of the cells and tissues and imposing analyses of the cells to derive from media stream analyses or on magnetic resonance imaging (MRI) or nuclear magnetic resonance spectroscopy (NMRS). Albumin and urea concentrations are analyzed initially to confirm cellular bioreactor effectiveness, since they have been used for all past studies enabling comparisons to be made. Then, more detailed assays such as gluconeogenesis, glycolysis, cytochrome P450, and tyrosine kinase help to assess bioreactor functional similarities when compared with normal findings from livers in vivo. Thus, multiple function assays are able to detect bulk tissue responses within the bioartificial liver cell-cores.

STATIC VERSUS DYNAMIC NUTRIENT INPUTS Existing data on liver cells maintained ex vivo are based largely on static culture techniques such as monolayer cultures. The introduction of flow perfusion of cultures leads to dynamic nutrient exchange yielding cell responses that more closely resembles that from liver in vivo. A qualifier is that dynamic flow produces shear stresses that can lead to adverse cell effects; the liver in vivo is perfused under very low shear conditions. HF systems do not produce shear as the convective nutrient pathways and cell-core spaces are separated by hollow fiber wall membranes that quench shear forces. Thus, nutrients must permeate into the cell-core spaces by diffusion or dispersion transports; where dispersion concurrently harnesses diffusion and convective coefficients such that their transport magnitudes are jointly beneficial to the cultures needs. For these systems, the use of dynamic flow systems has illustrated better culture performance based upon the increased function levels exhibited by the cells (McClelland and Coger, 2003), such that the bioartificial liver efficacy is improved. MRI AND NMRS ANALYSES OF BIOARTIFICIAL LIVERS Although media stream analysis is valuable, it does not give information on cell morphology, proliferation, or differentiation except with respect to secreted products. To examine all aspect of cultures within bioreactors, image analyses or metabolomic analyses are necessary. Since the outer casings of bioreactors are opaque, direct investigations with microscope are not possible. The alternative is to use MRI or NMRS to investigate the cell responses in the interiors of bioreactor systems. MRI MRI enables millimeter structures to be visualized to show united tissue elements. Using this technique, the radial cell locations “in reference to” the nutrient sources are analyzed using cross-sectional image formats. As the cross-sections are accumulated through the longitudinal span of the bioreactors, then 3D reconstructions of both cells and tissues and of the bioreactor infrastructure can be visualized. In this way, the cell masses may be monitored as they assemble from isolated seeded cells at day 0 and into interconnected tissue structures during the culture periods. Also, since gas and liquid nutrient sources are separated, then the migration of the cells throughout the cell-core space in response to heterogeneous nutrient concentrations is a way to demonstrate positions of healthier microenvironments. By including bioreactor structures (i.e. fiber walls) within these images, then geometrical patterns of nutrient pathways may also be analyzed. In this way, the bioartificial liver structures and arrangements may be scrutinized following the cell loading process and after sustaining dynamic forces, where displacements of both organic and inorganic components may occur from fluid pressures, shear trends, and transformed materials properties. These changed properties may develop as functions of capillary flows, wetting properties, and porous foulings (MacDonald et al., 2001a) in ways that modify the mass transport coefficients of the system. Additionally, material compositions may be affected such that yielding, buckling, and concentric cylinder shifts alter design parameters and initiate heterogeneous activities. Using this technique, the entire HF bioartificial liver is inserted into a miniaturized MRI apparatus such that intact

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bioartificial livers may be intermittently investigated throughout the length of its culture stability. In this way, precisely imaged culture data may be correlated with tissue function responses. NMRS NMR is a technique used to study physical, chemical, and biological properties of matter. It is a method that harnesses the reverberation characteristics of particle movement as non-invasive options that may be used to examine organic components of the bioartificial liver cell-cores. It occurs when the nuclei of certain atoms are initially immersed in static magnetic fields and are then exposed to secondary oscillating magnetic fields to induce resonance (www.rit.edu/htbooks/nmr/inside.htm). In these cases, the resonance frequency and particle displacements associated with intermolecular bonding can be calibrated to distinguish between molecular subcategories. Thus, the calibration may be extrapolated to label expressions of viable or lifeless cells as the molecules within the cell-core cultures transform. For this method, bioartificial livers are positioned inside orthogonally aligned electrical coils (MacDonald et al., 1998) such that each coil may be “pre-tuned” to distinguish between solution or cellular components – where tuning regulates the penetrating powers based upon sample densities and electrical fluxes. In this way, the organization and placement of a particles’ chemical bonds may be analyzed to index for specific cellular activities. Thus, as particle concentrations (e.g. sodium, carbon13, phosphate, fluoride, etc.) change or as cell activities (e.g. ATP, NADH, glycine production, etc.) vary, then spectroscopy analysis is utilized to classify the cultures viable and functional activities.

CRYOSECTIONING Another method of image investigations is to dismantle the bioreactor to facilitate traditional microscopy studies. For this process the bioreactors must be frozen in ways that all bioartificial liver components remain positionally constrained. To accomplish this, the bioreactors nutrient pathways are filled with freezing solutions (e.g. Tissue Tek OCT) after disconnecting them from recirculating nutrient support networks (McClelland and Coger, 2003). Then the bioartificial livers are frozen in liquid nitrogen in order to maintain alignment of geometrical structures and of cells and tissues. Next, after coarse sectioning and micron cryostat slicing to mechanically subdivide the bioreactor, the cross-sectional slices are affixed to microslides and imaged via bright light microscopy, as shown by the 4 and 20 micrographs in Figure 53.8a and d, respectively. In Figure 53.8a, the centralized hollow fiber cylinder is used to transport liquid nutrients throughout the interior of the bioartificial liver. This hollow fiber is positioned such that liquid nutrients diffuse radially outward, through the inner fiber wall, to support the cells within the cell-core space. Additionally, a hollow fiber wall that borders the periphery of the cell-core space is utilized as O2 input diffuses radially inward – through the outer fiber wall – from its nutrient channel such that O2 input also assists the cultures metabolic needs. As shown by the outer hollow fiber wall “detail” in Figure 53.8d, changing fiber material compositions may modify the mass transport rate at which nutrients enter the cell-core space. In this figure, the outer fiber wall is displayed as a woven substrate composed of pore sizes and densities that may be altered to amend fiber wall conductiveness. Also shown is the extruded inner fiber wall in which diffusion is the overriding transport mechanism. Since the hollow fibers are easily interchanged within the bioreactor design, then the rates, concentrations, and routes of nutrient input into the cell-core spaces are ways to control stability of the cultures. Extending this image analysis beyond the bright light technique is possible by labeling the cells with fluorescent probes. This cell labeling is accomplished prior to bioartificial liver freezing by modifying the recirculation media to include particular molar concentrations of fluorescent probes. In most cases, a dual probe setup is utilized as ways to distinguish between functional and non-functional cells. After labeling, freezing, and cross-sectioning the bioartificial livers, then fluorescent or laser excitation and emission microscopy may

940 THERAPEUTIC APPLICATIONS: CELL THERAPY

Detail

Outer fiber wall

O2

(a)

(b) Hepatocytes

Media

O2

Cell-core

Inner fiber wall

Cell-core O2

(c)

(d)

Media

Figure 53.8 Four micrographs of the interior concentric cylinders of a bioartificial liver. The 4 micrograph of image (a) displays the nutrient media annulus with radial diffusion in the outward direction; the cellcore space containing matrix and cells; the O2 input sources diffusing radially inward; and the hollow fiber walls as black concentric rings. Image (b) displays a 20 micrograph hepatocytes labeled with probe, calcein AM. As shown, the labeled cells extend from the outer to the inner fiber walls. Image (c) displays a similar setup as image (b); however, the cells have been labeled with ethidium homodimer-1 to indicate compromised membranes. Image (d) is a 20 micrograph displaying two uniquely different hollow fiber cylinder materials. The outer cylinder is a woven constructs in which the “detail” illustrates large pores while the inner fiber wall is densely packed and limits transport to diffusion.

be accomplished, as displayed with the confocal images in Figure 53.8b and c. In this figure, both the outer and inner hollow fiber walls are included to demonstrate locations of cultured cells in response to gaseous or liquid nutrient inputs. In Figure 53.8b, hepatocytes are labeled with calcein AM (Molecular Probes, Eugene Oregon) to tag polarized and functional mitochondrial membranes. In Figure 53.8c, hepatocytes are labeled with ethidium homodimer-1 (Molecular Probes, Eugene Oregon) to demonstrate compromised cell membranes. In this way, merged images of identically imaged cultures are able to reveal limitation parameters associated with seeding distances and nutrient source locations; thus, the annulus thickness of the cell-core space can be tailored for each hollow fiber design. Numerous fluorescent options to visually analyze the HF bioartificial liver cell-cores are available such that patterns of cell viability, cell proliferation, and cell differentiation may be linked to nutrient source concentrations and viable cell locations within cell-core spaces. These locations help to specify a cell’s status when cultured at specific distances from both liquid and gaseous nutrient sources. In this way, boundary conditions for future bioartificial liver designs will be standardized and tabulated as ways to predict cell viabilities when analyzing new structures as technology enhances material properties.

DISCUSSION The infrastructure designs of bioreactors used for bioartificial livers have remained somewhat constant over the past 25 years. These designs are limited because they are dependent for mass transfer of nutrients on supply lines, for example, the hollow fibers, that cannot reproduce the intricacies of capillary networks found inside tissues. Several modifications are used currently to enhance nutrient transport techniques by lowering transport resistances of material properties within the bioreactors inert structures. One modification is a composition rearrangement of structural molecules. By reorganizing material molecules or altering techniques of matter solidifications, then homogeneous and heterogeneous porous substrates may be exploited to control nutrient transfers. In this way, small molecules may have free reign throughout the device while larger molecules are limited in their direction. Additionally, as the pore sizes change, then transport coefficients are modified such that diffusion responses may impede or facilitate culture stabilities. A second modification is to chemically alter the inert

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structures such that material properties are conducive to the culture environment. In this way cellular activity may respond by aggregating next to the material, by attaching to the material, or by embedding within the material in ways to prolong viable states. A third modification is the arrangement of nutrient pathways such that culture metabolism is maximized. By characterizing transport effects through membranes and matrices that make up the cultures environment, then efficient designs may be computationally determined and graphically designed such that development of new structures are both adequately supported and simple to construct. Biological variables have been addressed in diverse ways that are helping to maintain the cells in a more functional state and that are helping to overcome the severe shortages of human tissue. Wholly defined media, supplements, and extracellular matrix scaffolds are now available for expansion or differentiation of liver cells. The sourcing problems for human tissues are not fully solved though stem cell and maturational lineage biology offer the greatest potential for facilitating the development of bioartificial human livers in the near future. Embryonic stem cells that can be lineage restricted to liver or hepatic stem cells have become sourcing options that are currently being explored. Clinical programs with bioartificial livers are ongoing in FDA I, II, and III trials (Mazariegos et al., 2002). At present there are limitations regarding scale-up into larger tissue masses, but as intricacies in microtechnology improve (e.g. nanobots) and material property components advance, our ability to replicate in vivo tissue masses is quickly becoming a reality. This reality is industry approved, as companies focus these innovations to alter medical treatment techniques or improve diagnostic tools applications as ways to better medical technology.

SUMMARY Bioartificial livers consist of both acellular and cellular components. Acellular inert materials support viable cell compartments used for tissue seeding, extracellular matrix and scaffold incorporation, and pathways for supplying gas and medium nutrients essential for cell regulatory signals. Bioartificial livers are created to imitate liverspecific functions, to provide a system in which to grow tissue-specific pathogens such as viruses (Bader et al., 1998; Nagamori et al., 2000), to analyze the effects of genetic alterations on tissue functioning (Parens, 1995; Frankel and Chapman, 2001), to provide highly differentiated tissue that can be used as a model for screening of drugs or treatments, or to provide a device that will assist or temporarily replace host organs that are diseased (Dixit, 1995; Bornemann et al., 1996; Nagamori et al., 2000). The biological variables governing bioartificial livers comprise the use of specific extracellular matrices (Reid et al., 1992; Zern and Reid, 1993; LeCluyse, 2000; Brill et al., 2002), media (MacDonald et al., 2001a), and regulatory factors inducing cell proliferation or differentiation (Dickson and Salomon, 1998). The inert bioreactor variables comprise infrastructure design, material surface chemistries (Gerlach et al., 1996; Mayer et al., 2000; Catapano et al., 2001; Naruse et al., 2001), and the detailed arrangements of supply lines that mimic capillary networks (McClelland and Coger, 2000, 2003; Wolfe et al., 2002). All combined variables determine the longevity of the tissue in the bioreactor’s cell compartment and the extent of growth, and/or differentiation of the cells. Ongoing challenges include sourcing of human liver cells and identifying methods for improved nutrient mass transfer supporting bioreactor cell environments. Evolving these concepts will enable differentiated tissues to survive and stably function for weeks. This survival offers expanded opportunities for academic, clinical, and industrial investigators studying liver biology or needing patient assist devices that support liver functions. ACKNOWLEDGEMENTS Funding derived from NIH grants (DK52851, AA014243, IP30-DK065933), a Department of Energy Grant (DE-FG02-02ER-63477), and by a sponsored research grant from Vesta Therapeutics (Research Triangle Park in Durham, North Carolina).

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MacDonald, J., Wolfe, S., Roy-Chowdhury, I., Kubota, H. and Reid, L. (2001a). Effect of flow configuration and membrane characteristics on membrane fouling in a novel multicoaxial hollow fiber bioartificial liver. Ann. NY Acad. Sci. 944: 334–343. MacDonald, J., Xu, A., Kubota, H., LeCluyse, E., Hamilton, G., Liu, H., Rong, Y., Moss, N., Lodestro, C., Luntz, T., Wolfe, S. and Reid, L. (2001b). Liver cell culture and lineage biology. In: Atala, A. and Lanza, R. (eds.), Methods of Tissue Engineering. San Diego: Academic Press, pp. 151–201. MacDonald, J.M. and Wolfe, S.P. (1999). Bioreactor design and process for engineering tissue from cells. US Patent 113918.200. MacDonald, J.M., Xu, A.S.L., Hiroshi, K., LeCluyse, E., Hamillton, G., Liu, H., Rong, Y.W., Moss, N., Lodestro, C., Luntz, T., Wolfe, S.P. and Reid, L. (2002). Ex vivo maintenance of cells from the liver lineage. In: Lanza, W.L., Langer, R. and Vacanti, J. (eds.), Methods of Tissue Engineering. San Diego: Academic Press, pp. 151–201. Malchesky, P. (1994). Nonbiological liver support: historic overview. Artif. Organs 18(5): 342–347. Marcin, P., Dent, E. and Kucaba-Pietal, A. (2000). Recent advances in solvers for nonlinear equations. Comput. Assist. Mech. Eng. Sci. 7: 493–505. Mayer, J., Karamuk, E., Akaike, T. and Wintermantel, E. (2000). Matrices for tissue engineering-scaffold structure for a bioartificial liver support system. J. Control. Release 64: 81–90. Mazariegos, G., Patzer II, J.F., Lopez, R., Giraldo, M., deVera, M., Grogran, T., Zhu, Y., Fulmer, M., Amoit, B. and Kramer, D. (2002). First clinical use of a novel bioartificial liver support system (BLSS). Am. J. Transplant. 2: 260–266. McClelland, R. and Coger, R. (2000). Use of micropathways to improve oxygen transport in a hepatic system. J. Biomech. Eng. 122: 268–273. McClelland, R. and Coger, R. (2003). Modeling O2 transport within engineered hepatic devices. Biotechnol. Bioeng. 82(1): 12–27. McClelland, R. and Coger, R. (2004). Effects of enhanced O2 transport on hepatocytes packed within a bioartificial liver device. Tissue Eng. 10(1/2): 253–266. Nagamori, S., Hasumura, S., Matsuura, T., Aizaki, H. and Kawada, M. (2000). Developments in bioartificial liver research: concepts, performance, and applications. J. Gastroenterol. 35(7): 493–503. Naruse, K., Sakai, Y., Lei, G., Sakamoto, Y., Kobayashi, T., Puliatti, C., Aronica, G., Morale, W., Leone, F., Qiang, S., Ming, S., Ming, S., Li, Z., Chang, S., Suzuki, M. and Makuuchi, M. (2001). Efficacy of nonwoven fabric bioreactor immobilized with porcine hepatocytes for ex vivo xenogeneic perfusion treatment of liver failure in dogs. Artif. Organs 25(4): 273–280. NIDDK (2002). National Institute of Diabetes & Digestive & Kidney Diseases. from www.niddk.nih/gov. O’Shea, K.S. (1999). Embryonic stem cell models of development. Anat. Rec. 257(1): 32–41. Parens, E. (1995). The goodness of fragility: on the prospect of genetic technologies aimed at the enhancement of human capacities. Kennedy Inst. Ethic. J. 5(2): 141–153. Patzer, J.F., Mazariegos, G.V., Lopez, R., Molmenti, E., Gerber, D., Riddervold, F., Khanna, A., Yin, W.Y., Chen, Y., Scott, V.L., Aggarwal, S., Kramer, D.J., Wagner, R.A., Zhu, Y., Fulmer, M.L., Block, G.D. and Amiot, B.P. (1999). Novel bioartificial liver support system: preclinical evaluation. Ann. NY Acad. Sci. 875: 340–352. Powers, M., Janigian, D., KE, W., Baker, C., Beer Stolz, D. and Griffith, L. (2002). Functional behavior of primary rat liver cells in a three-dimensional perfused microarray bioreactor. Tissue Eng. 8(3): 499–513. Reid, L.M. (1990). Defining hormone and matrix requirements for differentiated epithelia. In: Pollard, J.W. and Walker, J.M. (eds.), Basic Cell Culture Protocols. Totowa, NJ: Humana Press, Inc. 75: Chapter 21, pp. 237–262. Reid, L.M., Fiorino, A.S., Sigal, S.H., Brill, S. and Holst, P.A. (1992). Extracellular matrix gradients in the space of Disse: relevance to liver biology (editorial). Hepatology 15(6): 1198–1203. Rozga, J., Holzman, M.D., Ro, M.S., Griffin, D.W., Neuzil, D.F., Giorgio, T., Moscioni, A.D. and Demetriou, A.A. (1993a). Development of a hybrid bioartificial liver. Ann. Surg. 217(5): 502–509; discussion 509–511. Rozga, J., Williams, F., Ro, M.S., Neuzil, D.F., Giorgio, T.D., Backfisch, G., Moscioni, A.D., Hakim, R. and Demetriou, A.A. (1993b). Development of a bioartificial liver: properties and function of a hollow-fiber module inoculated with liver cells. Hepatology 17(2): 258–265. Sauer, I., Zeilinger, K., Pless, G., Kardassis, D., Theruvath, T., Pascher, A., Mueller, A., Steinmueller, T., Neuhaus, P. and Gerlach, J. (2003). Extracorporeal liver support based on primary human liver cells and albumin dialysis – treatment of a patient with primary graft nonfunction. J. Hepatol. 39(4): 649–653.

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Schmelzer, E., McClelland, R., Melhem, A., Zhang, L., Yao, H., Wauthier, E., Turner, W., Furth, M., Gerber, D., Gupta, S. and Reid, L. (2006). Hepatic stem cells and the liver’s maturational lineages: implications for liver biology, gene expression and cell therapies. In: Potten, C. (ed.), Tissue Stem Cells. London: Dekker, pp. 161–214. Sicklick, J., Li, Y., Melhem, A., Schmelzer, E., Zdanowicz, M., Huang, J., Caballero, M., Fair, J., Ludlow, J., McClelland, R., Reid, L. and Diehl, A. (2006). Hedgehog signaling maintains resident hepatic progenitors throughout life. Am. J. Gastroint. Liver Physiol. 290(5): G859–G870. Soulsby, C. (1999). Dietary management of hepatic encephalophathy in cirrhotic patients: survey of current practice in the United Kingdom. BMJ 318: 1391. Storaasli, O., Nguyen, D., Baddourah, M. and Qin, J. (1993). Computational mechanics analysis for parallel-vector supercomputers. Comput. Syst. Eng. 4(4–6): 349–354. Sussman, N.L. and Kelly, J.H. (1993). Liver assist devices (LADs) will not be used to treat fulminant hepatic failure (FHF), but its consequences, namely hepatic encephalopathy (HE) (letter). Artif. Organs 17(1): 43–45. Sussman, N.L. and Kelly, J.H. (1996). Artificial liver support. Clin. Invest. Med. – Medecine Clinique et Experimentale 19(5): 393–399. UNOS (2002). Liver Data Annual Reports. United Network for Organ Sharing. from www.unos.org. Watanabe, F.D., Arnaout, W.S., Ting, P., Navarro, A., Khalili, T., Kamohara, Y., Kahaku, E., Rozga, J. and Demetriou, A.A. (1999). Artificial liver. Transplant. Proc. 31(1–): 371–373. Wolf, C. and Munkelt, B. (1975). Bilirubin conjugation by an artificial liver composed of cultured cells and synthetic capillaries. Trans. Am. Soc. Artif. Intern. Organs 21: 16–27. Wolfe, S.P., Hsu, E., Reid, L.M. and Macdonald, J.M. (2002). A novel multicoaxial hollow fiber bioreactor for adherent cell types. Part I: Hydrodynamic studies. Biotechnol. Bioeng. 77: 83–90. Wright, J., Chilcott, J., Holmes, M. and Brewer, N. (2002). The Clinical and Cost Effectiveness of Pulsatile Machine Perfusion vs. Cold Storage of Kidneys for Transplantation Retrieved From Heat-Beating and Non-Heart-Beating Donors. Sheffield, England, School of Health and Related Research, University of Sheffield: (Report). Xu, A., Luntz, T., Macdonald, J., Kubota, H., Hsu, E., London, R. and Reid, L.M. (2000). Liver stem cells and lineage biology. In: Lanza, R., Langer, R. and Vacanti, J. (eds.), Principles of Tissue Engineering. New York: Lands Press, pp. 559–598. Zern, M. and Reid, L. (1993). Extracellular Matrix: Its Chemistry, Biology, and Pathobiology. New York: Marcel Dekker, Inc.

54 Neuronal Transplantation for Stroke Douglas Kondziolka and Lawrence Wechsler

INTRODUCTION In the United States, stroke is the third leading cause of death and the most common cause of serious adult disability. Approximately 30% of patients become severely and permanently disabled and many others have permanent impairment. The economic burden for stroke is huge. Stroke prevention and early intervention to minimize the damage caused by stroke have received great attention. Rehabilitation therapy is important to maximize functional recovery in the early period after stroke. However, once recovery has plateaued and the neurologic deficits are fixed, there is no known treatment (Zivin and Choi, 1991). The role of cellular therapy as one avenue of regenerative medicine has been explored. Preclinical studies first established the potential for cultured neuronal cells derived from a teratocarcinoma cell line to be tested for safety and efficacy in the treatment of human stroke. In an animal model of stroke that caused reproducible learning and motor deficits, injection of neuronal cells resulted in a return of learning behavior retention time and motor function (Borlongan et al., 1998). Studies in monkey, rat, and mouse of up to 14 months demonstrated no clinically important toxicity and no tumor formation (Kleppner et al., 1995). In experimental animals, these neuronal cells integrated with the host brain, sent out axonal processes, released neurotransmitters, and demonstrated typical neuronal proteins (Trojanowski et al., 1993; Kleppner et al., 1995). A phase 1 study evaluated 12 patients with acceptable safety and clinical improvement in some patients that correlated with changes on positron emission tomography (PET) (Kondziolka et al., 2000; Meltzer et al., 2001). A second twocenter clinical trial further evaluated the safety and effectiveness of neuronal cell transplantation in patients with substantial functional motor deficits following cerebral infarction using higher cell numbers and a comparison to observational controls (Kondziolka et al., 2004). This research tested the hypothesis that implantation of neuronal cells would be safe, feasible, and lead to improvement of motor neurologic deficits resulting from basal ganglia cerebral infarction. THE BASIS OF CELLULAR TRANSPLANTATION Transplantation of human neuronal cells is one approach for ameliorating functional deficits caused by central nervous system (CNS) disease or injury. Several investigators have evaluated the effects of transplanted fetal tissue, rat striatum, or cellular implants into small animal stroke models (Nishino et al., 1993; Johansson and Grabowski, 1994). Although transplanting primary human fetal neurons into patients with neurodegenerative disease continues to be evaluated, the widespread clinical use of primary human tissue is likely to be limited due to the ethical and logistical difficulties inherent in obtaining large quantities of fetal neurons (Thompson et al., 1999). For this reason, much effort has been devoted to developing alternate sources of

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human neurons for use in transplantation. One alternate source is the Ntera 2/cl.D1 (NT2) human embryonic carcinoma-derived cell line. These cells proliferate in culture and differentiate into pure, post-mitotic human neuronal cells (LBS-Neurons) upon treatment with retinoic acid (RetA) (Andrews et al., 1984; Pleasure and Lee, 1993). Thus, NT2 precursor cells appear to function as CNS progenitor cells with the capacity to develop diverse mature neuronal phenotypes. When transplanted, these neuronal cells survive, extend processes, express neurotransmitters, form functional synapses, and integrate with the host (Kleppner et al., 1995; Trojanowski et al., 1997). During the retinoic acid induction process, the LBS-Neuronal precursor cells – which share many characteristics of neuroepithelial precursor cells –undergo significant changes resulting in the loss of neuroepithelial markers and the appearance of neuronal markers. The final product is a 95% pure population of human neuronal cells that appear virtually indistinguishable from terminally differentiated, postmitotic neurons (Andrews et al., 1984; Pleasure and Lee, 1993). The cells are capable of differentiation to express a variety of neuronal markers characteristic of mature neurons, including all three neurofilament proteins (NFL, NFM, and NFH); microtubule associated protein 2 (MAP2), the somal/dendritic protein; and tau, the axonal protein. Their neuronal phenotype makes them a promising candidate for replacement in CNS disorders, as a virtually unlimited supply of pure, post-mitotic, terminally differentiated human neuronal cells. In patients disabled by stroke, the concept of restoring function by transplanting human neuronal cells into the brain is innovative and only recently conceived (Bonn, 1998; Thompson et al., 1999). Research in a rat model of transient focal cerebral ischemia demonstrated that transplantation of fetal tissue restored both cognitive and motor functions (Nishino et al., 1993; Borlongan et al., 1997, Borlongan et al., 1998a). Sanberg, Borlongan and colleagues showed that transplants of LBS-Neurons could also reverse the deficits caused by stroke (Borlongan et al., 1998b). The preclinical studies of LBS-Neurons were carried out in a model of transient focal, rather than global, ischemia in order to maximize the chances of functional recovery. In several studies, animals received ischemic insults to the striatum and were tested for behavioral deficits 1 month later. Behavioral testing was conducted using a passive avoidance learning and retention task and a motor asymmetry measure. Animals that showed significant behavioral deficits received neuronal transplantation, and then were periodically re-evaluated during the 6-month post-transplantation period. Animals that received transplants of LBS-Neurons and cyclosporine treatment showed amelioration of ischemia-induced behavioral deficits throughout the 6-month observation period. They demonstrated complete recovery in the passive avoidance test, as well as normalization of motor function in the elevated body swing test. In comparison, control groups receiving transplants of rat fetal cerebellar cells, medium alone, or cyclosporine failed to show significant behavioral improvement (Borlongan and Sanberg, 1995; Borlongan et al., 1995; Saporta et al., 1999). Subsequent studies showed that these cells released glial-derived neurotrophic factor (GDNF) after transplantation into ischemic rats. A second study that evaluated response in comparison to the number of cells transplanted, confirmed the efficacy of transplanted LBS-Neurons in reversing the behavioral deficits resulting from transient ischemia in an MCA occlusion rat model (Saporta et al., 1999).

POTENTIAL MECHANISMS OF CELL TRANSPLANTATION In addition to a humoral mechanism of action, some evidence suggests a direct action of surviving implanted neuronal cells. Animal transplantation studies of LBS-Neurons reveal graft survival, mature neuronal phenotype, and integration into host brain in vivo (Trojanowski et al., 1993; Kleppner et al., 1995; Trojanowski et al., 1997). LBS-Neurons grafted into different regions of the CNS of nude mice showed viable cells in 90% of recipients, with some grafts surviving for up to 14 months. Grafted LBS-Neurons initially remained similar to their in vitro counterparts, but then progressively acquired the phenotype of fully mature neurons in vivo (Borlongan et al., 1995; Kleppner et al., 1995). Transplanted neurons formed synapse-like structures and elaborated dendrites and axons,

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and have been tested electrophysiologically. Thus, transplanted LBS-Neurons demonstrated survival for at least 14 months post-implantation, a fully mature neuronal phenotype in vivo, and integration with the host CNS. In our first human trial, fluorodeoxyglucose-18 PET showed increased uptake at the target site that correlated with the clinical response, and an autopsy evaluation of one graft 27 months after surgery showed surviving donor cells (Meltzer et al., 2001; Nelson et al., 2002). The neuronal cells could improve neurologic function through a number of different mechanisms. These include provision of neurotrophic support (acting as local pumps to support cell function), provision of neurotransmitters, reestablishment of local interneuronal connections, cell differentiation and integration, and improvement of regional oxygen tension (Jakeman and Reier, 1991; Bregman et al., 1993, Himes et al., 1994, Snyder et al., 1997; Tessler et al., 1997). Transplanted cells may also act to limit the reactive glial response and to limit retrograde degeneration, although this may not contribute to repair in a chronic injury (Bregman and Reier, 1986). We believe that axonal reconnections through the grafted cells (serving as a “bridge”) over large distances is less likely, although this phenomenon has been observed in spinal cord injury models.

PRODUCTION OF NEURONAL CELLS FOR HUMAN USE In two clinical trials we used LBS-Neurons (Layton BioScience, Inc., Gilroy, CA) produced using antibiotic free conditions in a class 10,000 clean room, according to cGMP protocols. The NT2/D1 human precursor cell line was plated in culture from a well-characterized working cell bank. This stock culture was passaged 2 times a week in DMEM/F-12 growth media. The NT2/D1 cells were induced to differentiate into neurons by the addition of 10 μM RetA. After 6 weeks of RetA treatment the cultures were harvested with trypsin/EDTA and replated at lower cell densities. These cultures were maintained in DMEM/F-12 media containing 5% FBS and a mitotic inhibitor mixture (10 μM FUdR, 10 μM uridine, and 1 μM AraC) for total of 6 days. The cells were selectively harvested, purified, and extensively tested. The LBS-Neurons were cryopreserved in freezing media, and stored in the vapor phase of liquid nitrogen (Kondziolka et al., 2004).

INSTITUTIONAL PREPARATION OF NEURONAL CELLS On the day of surgery 1 h prior to implantation, vials were thawed, gently washed twice in Isolyte S (McGaw Inc., Irvine, CA) and centrifuged at 200 g for 7 min at room temperature. The cell pellet was gently resuspended in Isolyte S. The viable cell count was determined with a sample of the LBS-Neuron suspension using 0.4% trypan blue, and the cells were resuspended to a final concentration of 3.3  107 cells/ml in Isolyte S and aliquoted at 120 l per sterile 1.0 ml vial. An aliquot was considered acceptable only if more than 50% of cells were viable. Depending upon the dose to be administered, one or more vials were prepared. Vials were loaded into a closed holder and carried by hand in an upright position to the operating room for immediate use. In our first study, doses of 2 and 6 million cells, and in a second trial, 5 and 10 million cells were cleared for evaluation by the United States Food and Drug Administration. The trial was approved by the Institutional Review Boards at the University of Pittsburgh (studies 1 and 2), and Stanford University (study 2), and was reviewed by a separate Data and Safety Monitoring Board. Stopping rules were established during the initial meeting of the committee.

CLINICAL TRIAL DESIGN The second, larger study was an open-label trial with observer-blinded neurologic evaluation of patients with stroke who received stereotactic implants of human neuronal cells (Kondziolka et al., 2004). The first nine

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patients were randomized to either surgery plus rehabilitation (n  7) or rehabilitation alone (n  2) (surgery consisted of 5 million cells divided into 25 implants along five trajectories; 10 μl per implant). The next nine patients were randomized to receive either surgery with 10 million cells plus rehabilitation (n  7) or rehabilitation alone (n  2). Patients were evaluated for safety and efficacy at visits occurring at day 3, and at frequent intervals during the first year. They continued to be seen at 1-year intervals after surgery. Cyclosporine-A (6 mg/kg ideal body weight per day administered orally twice daily) was administered 1 week prior to surgery and continued for 6 months. Thirteen of the 18 patients were men and five were women. Patient age varied from 24 to 70 years. The mean age was 59 years for the 5 million cell patients, 58 for the 10 million cell patients, and 46 for the control patients. Nine strokes were right sided and nine were left sided. The mean time since the onset of the stroke was 3.5 years (range, 1–5) with no difference between groups. Nine strokes were ischemic and nine hemorrhagic. In the control group, three patients had ischemic strokes and one had a hemorrhagic stroke. Study evaluations consisted of complete neurologic examinations, National Institutes of Health Stroke Scale (NIHSS) and European Stroke Scale (ESS) performed at baseline and repeated at all follow-up visits (Brott et al., 1989; Hanston et al., 1994). Stroke Impact Scales and Everyday Memory Questionnaires were performed at baseline and weeks 4, 8, 12, 18, 24, 26, 28, 36, and 52. Neurological Function Questionnaires were performed at weeks 12, 24, and 52. All measures were completed by trained and blinded observers. Patients wore hats to prevent detection of evidence of prior surgery and they were instructed not to reveal their status. Fugl-Meyer scores, Gait tests, Action Research Arm testing, and Grooved Pegboard testing were conducted by the blinded physical therapist. Magnetic resonance imaging (MRI) scans and different serologic tests were obtained at baseline and after surgery. Statistical analyzes were performed on an intent-to-treat basis. The 6-month evaluation was designed a priori as the timepoint for the efficacy assessment. All statistics and data analysis were performed by an independent biostatistician. Sample size was determined (alpha  0.05, power  0.8) to show a significant change from baseline of 5 points on the ESS motor score. Two subjects randomized to rehabilitation (no surgery) per dose group were not factored into the power analysis, but were selected in order to collect data on non-surgical outcomes. Statistical testing utilized a two-tailed t-test and significance determined at the 0.05 level.

NEURONAL TRANSPLANTATION TECHNIQUE One week prior to surgery, all anticoagulant medications were discontinued. After stereotactic frame application, a contrast-enhanced computed tomography (CT) scan was performed for targeting. Coronal and sagittal views were used to define a safe trajectory that entered a cortical gyrus and spared a sulcus. Stereotactic coordinates were obtained for each instrument placement. We determined a point in the basal ganglia inferior to the center of the stroke, and four other targets inferior to the stroke (anterior, posterior, medial, and lateral to the central target, usually spaced by 5 mm) (Kondziolka et al., 2000). For each of the five planned trajectories, the patient was to receive five cell implants spaced equally across a distance of 20–25 mm. A burr hole was created, the dura opened, and a 1.8 mm outer diameter 15 cm long stabilizing probe was inserted to a point 4 cm proximal to the final target. A 0.9 mm outer diameter cannula was then inserted down to the deepest target point for the first implantation (Kondziolka et al., 2000). A 10 μl volume of cells was injected slowly at each target site over 2 min. The instrument was then withdrawn to the more proximal targets along each trajectory. The total time for all implantations was approximately 150 min (Kondziolka et al., 2004). All patients were discharged home the morning after surgery. No cell-related serious adverse events occurred. No clinically significant laboratory, radiographic, or electrocardiographic abnormalities that could be attributed to the neuronal-cell implantation were seen.

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FUNCTIONAL OUTCOMES We used the ESS score as a primary outcome measure because it collected detailed, functional assessments of motor function, even though its validated use was in acute stroke. The ESS scores were collected for each patient pre-operatively on the day of surgery (baseline) and at predetermined intervals through 24 weeks post-implantation. On the ESS, higher scores indicated better neurologic performance, and a significant change was noted to be 3 points. Between the pre-operative evaluation and the day of surgery, 11 of 12 patients had no change in their ESS score and 1 patient improved by 1 point. The mean baseline ESS total score was 68 for the 5 million cell group, 70 for the 10 million cell group, and 70 for the controls. The mean baseline ESS motor score was 28 for the 5 million cell group, 28 for the 10 million cell group, and 30 for the controls. The 6-month evaluation was calculated as the average score between the three evaluations at weeks 24, 26, and 28. This was done to average out the scores related to any increased or decreased patient effort obtained at any one visit. At the 6-month follow-up evaluation post-implantation of LBS-Neurons, four of the seven 5 million cell patients had improved scores on the ESS (range, 5.3–15 points), two patients were unchanged and one patient deteriorated (4.5 points) compared to their baseline scores. In the 10 million cell group, two of seven improved (6.5 and 14.5 points), three deteriorated (4.5–5.5 points), and two were unchanged. In the control group, one of four improved by 3.5 points, and the other three remained unchanged. The mean change in ESS score from baseline to 6 months for all implanted patients was 2.7 points in comparison to 0.75 points in the control group (p  0.148). In patients who received 5 million cells, the mean change at 6 months was 4.74 points. In an evaluation where patients served as their own controls (pre- versus post-surgery), the mean total ESS increased from 69.3 to 74.4 at 6 months (p  0.146). Motor elements of the ESS accounted for much of the change noted in patients treated with LBSNeurons. The ESS-Motor is a validated subscore of the ESS that is a composite of the individual scores for Facial movement, Arm outstretched, Arm raising, Wrist extension, Fingers, Leg maintain position, Leg flex, Foot dorsiflexion, and Gait (Bregman et al., 1993). Mean ESS-Motor score at baseline for all implanted patients was 28.0 and at 6 months, 30.7. The mean change in ESS-Motor score from baseline to 6 months for all implanted patients was 2.6 points in comparison to 1.0 point in the control group (p  0.756). In patients who received 5 million cells, the mean change at 6 months was 3.74 points. In an evaluation where patients served as their own controls (pre- versus post-surgery), the mean motor ESS increased from 28.0 to 32.2 at 6 months (mean change  4.12, 95% C.I.  –0.3–8.5; p  0.066). No difference was identified in the NIHSS scores when the control patients were compared to all surgical patients, or to the separate 5 and 10 million cell groups. The Stroke Impact Scale was used to measure the degree of disability caused by the stroke, and the effects of surgery on different elements of patient performance (Bregman and Reier, 1986). Implanted patients had higher daily activity scores at 6 months than control patients (p  0.056), although a significant change was only noted when 6-month scores for implanted patients were compared to their baseline (p  0.045). Compared to controls, scores for communication (p  0.199), feelings/mood (p  0.413), percent recovery (p  0.426), and meaningful activity (p  0.417) did not change. Everyday memory scores in implanted patients improved compared to control (p  0.012), and compared to their own baseline (mean change  13, 95% C.I.  4.9–21.2; p  0.004). The Fugl-Meyer Assessment of Motor Recovery After Stroke has been used extensively in studies focusing on functional recovery following stroke (Brott et al., 1989). Assessment with the Fugl-Meyer includes items for evaluating motor function in upper and lower extremities, as well as items assessing balance, sensation, range of motion and pain (Hanston et al., 1994). When implanted patients were compared to control, no difference was identified in the pre-surgery versus 6 months scores in upper or lower extremity function, balance, or sensation. When scores from the implanted patients alone were compared, a trend to improvement in hand movement (mean  1.15, 95% C.I.  0.07–2.4; p  0.06) and wrist movement (mean  0.92, 95% C.I.  0.05–1.9; p  0.06) was found.

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ACTION RESEARCH ARM TEST The four subtests of grasp, grip, pinch, and gross movement were evaluated. At 6 months, implanted patients had improved gross movement scores compared to controls (mean  1.64; p  0.017). The mean score in patients with 5 or 10 million cells was 1.57 (p  0.043) and 1.71 (p  0.051). In comparing pre- versus postsurgery scores within the same patient, improvements were noted in gross movement (p  0.001) and grasp (p  0.037), but not in grip or pinch movements. COGNITIVE TESTING Neuropsychological battery was performed before and after surgery in patients treated at the University of Pittsburgh (Stilley et al., 2004). IMAGING Serial MRI at 6 and 12 months showed no anatomic or structural changes in the brain after surgery when compared to baseline. There was no evidence of edema, contrast-enhancement, mass effect, or change in the contours of the infarction. One patient developed a chronic subdural hematoma on the side of surgery within 1 month of surgery. This was evacuated without side effects. The mean infarction size, measured in three dimensions using calipers, at baseline was 13 mm (left–right), 22 mm (anterior–posterior), and 19 mm (superior– inferior). There was no difference in infarct size between groups and no change in infarct size after surgery. SAFETY AND FEASIBILITY IN STROKE PATIENTS A major objective of this study was to demonstrate safety and feasibility of the neuronal-cell implantation procedure. This goal was met, in that implantation was carried out successfully in all 26 patients in two clinical trials. Although no new neurological deficits were identified from surgery, two new neurological events occurred (seizure and chronic subdural hematoma). A small seizure risk should be expected given the limited cortical transgression that occurs at surgery, the spinal fluid loss, and the accumulation of intracranial air around a brain with some degree of atrophy. Since many stroke patients take antiplatelet medications or other anticoagulants, some risk for delayed intracranial hemorrhage should be expected. A second objective was to demonstrate the longer-term safety of neuronal cell implantation. This goal was also met, in that no adverse events related to the implantation have occurred during 24–36 months of follow-up in these patients or in patients from our first clinical trial (52–60 months follow-up) (Kondziolka et al., 2004). No patient sustained any permanent morbidity related to cyclosporine-A use, although the drug led to variable degrees of fatigue. Review of the laboratory data listings reveals no consistent and clinically significant changes in hematology, chemistry, or urinalysis values. This study was also intended to provide some information on the efficacy of neuronal-cell implantation in improving stroke-related neurologic deficits. Study limitations included the paucity of information known regarding optimum patient criteria (age, stroke age, size, type, or location), adequate cell number, location and number of the brain implantation sites, use of immunosuppression, lack of larger control or study groups, and the best way to evaluate the patient response. We did not control for any effect of cyclosporine-A in the treatment groups. The study designs can be criticized for not controlling for a placebo effect, or any effect of cyclosporine-A, but these were randomized trials designed to address those issues. For the ESS, the increases tended to be larger in the group of patients receiving 5 million cells, both in the total scores and in the composite motor subscale scores. These patients had a higher incidence of ischemic rather than hemorrhagic strokes, which were more frequent in the 10 million cell group. However, the small

952 THERAPEUTIC APPLICATIONS: CELL THERAPY

number of patients in the 5 million or 10 million cell groups limit the ability to compare results between groups. In addition, the ESS was developed for use in acute stroke management and may not be best suited to evaluate changes in the chronic stroke patient. For that reason, we also included measures of disability and more chronic deficits such as the Stroke Impact Scale and the Fugl-Meyer scores. A recent study of individuals admitted to inpatient rehabilitation following stroke found Fugl-Meyer motor impairment scores on admission to be a predictor of motor impairment at discharge as well as activities of daily living and mobility functional outcome (Shelton et al., 2001). These indications of efficacy must be tempered by the fact that signs of improvement were not consistent. Some patients had worse stroke scale and disability scores at the end of 6 months than they had at the time of implantation, although these changes were modest. In our first trial, an equal number of patients had no improvement or worsening in stroke or disability scales as had those with improvement (Kondziolka et al., 2005). In that trial, clinical improvement correlated with change on fluorodeoxyglucose PET imaging at both the implant site and in the contralateral cerebellum (remote diaschisis effect) (Kondziolka et al., 2000).

GOING FORWARD After completion of two clinical trials in cellular transplantation for motor stroke, we believe that further research should focus on the development of new cell lines as well as refining clinical inclusion criteria. Because of the wide variety of patients and clinical factors evaluated in the first two studies (age, degree of deficit, spectrum of neurologic symptoms, stroke size, stroke type (hemorrhagic or ischemic), length of immunosuppression) it is difficult to make firm comments regarding candidacy. We believe that patients with younger strokes may have more potential to improve since their motor deficits are less likely to be fixed at the level of the distal musculature. On the other hand, both motor and cognitive improvements were measured in patients who were several years out from their stroke. We found some evidence to suggest that ischemic stroke may be more suited to cell therapy than hemorrhagic stroke, although this difference was not significant. We propose in a later study that an earlier stroke age be evaluated (3–12 months). Eventually, the concept of a placebo effect would need to be tested, if a reasonable and consistent level of clinical improvement was identified. The development of additional cell lines, either including neuronal precursor of stem cell lines will foster additional basic research and hopefully, further clinical studies.

REFERENCES Andrews, P., Damjanov, I., et al. (1984). Pluripotent embryonal carcinoma clones derived from the human teratocarcinoma cell line Tera-2. Differentiation in vivo and in vitro. Lab. Invest. 50: 147–162. Bonn, D. (1998). First cell transplant aimed to reverse stroke damage. Lancet 352: 119. Borlongan, C., Cahill, D., et al. (1995). Locomotor and passive avoidance deficits following occlusion of the middle cerebral artery. Physiol. Behav. 58: 909–917. Borlongan, C., Koutouzis, T., et al. (1997). Neural transplantation as an experimental treatment modality for cerebral ischemia. Neurosci. Biobehav. Rev. 21: 79–90. Borlongan, C. and Sanberg, P. (1995). Elevated body swing test: a new behavioral parameter for rats with 6-hydroxydopamine-induced hemiparkinsonism. J. Neurosci. 15: 5372–5378. Borlongan, C., Saporta, S., et al. (1998a). Viability and survival of hNT neurons determine degree of functional recovery in grafted ischemic rats. Neuroreport 9: 2837–2842. Borlongan, C., Tajima, Y., et al. (1998b). Transplantation of cryopreserved human embryonal carcinoma-derived neurons (NT2N cells) promotes functional recovery in ischemic rats. Exp. Neurol. 149: 310–321. Bregman, B. and Reier, P. (1986). Neural tissue transplants rescue axotomized rubrospinal cells from retrograde death. J. Comp. Neurol. 244: 86–95.

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Bregman, B., Kunkel-Bagden, E., et al. (1993). Recovery of function after spinal cord injury: mechanisms underlying transplant-mediated recovery of function differ after spinal cord injury in newborn and adult rats. Exp. Neurol. 123: 3–16. Brott, T., Adams Jr., H., et al. (1989). Measurements of acute cerebral infarction: a clinical examination scale. Stroke 20: 864–870. Hantson, L., De Weerdt, W., et al. (1994). The European stroke scale. Stroke 25: 2215–2219. Himes, B., Goldberger, M., et al. (1994). Grafts of fetal central nervous system tissue rescue axotomized Clarke’s nucleus neurons in adult and neonatal operates. J. Comp. Neurol. 339: 117–131. Jakeman, L. and Reier, P. (1991). Axonal projections between fetal spinal cord transplants and the adult rat spinal cord: a neuroanatomical tracing study of local interactions. J. Comp. Neurol. 307: 311–334. Johansson, B. and Grabowski, M. (1994). Functional recovery after brain infarction: plasticity and neural transplantation. Brain Pathol. 4: 85–95. Kleppner, S., Robinson, K., et al. (1995). Transplanted human neurons derived from a teratocarcinoma cell line (NTera-2) mature, integrate, and survive for over 1 year in the nude mouse brain. J. Comp. Neurol. 357: 618–632. Kondziolka, D., Steinberg, G., et al. (2004). Evaluation of surgical techniques for neuronal cell transplantation used in patients with stroke. Cell Transplant. 13: 749–754. Kondziolka, D., Steinberg, G., et al. (2005). Neurotransplantation for patients with subcortical motor stroke: a phase 2 randomized trial. J. Neurosurg. 103: 38–45. Kondziolka, D., Wechsler, L., et al. (2000). Transplantation of cultured human neuronal cells for patients with stroke. Neurology 55: 565–569. Meltzer, C., Kondziolka, D., et al. (2001). Serial [18F] fluorodeoxyglucose positron emission tomography after human neuronal implantation for stroke. Neurosurgery 49: 586–592. Nelson, P., Kondziolka, D., et al. (2002). Clonal human (hNT) neuron grafts for stroke therapy: neuropathology in a patient 27 months post-implantation. Am. J. Neuropath. 160: 1201–1206. Nishino, H., Koide, K., et al. (1993). Striatal grafts in the ischemic striatum improve pallidal GABA release and passive avoidance. Brain Res. Bull. 32: 517–520. Pleasure, S. and Lee, V. (1993). NTera 2 cells: a human cell line which displays characteristics expected of a human committed neuronal progenitor cell. J. Neurosci. Res. 35: 585–602. Saporta, S., Borlongan, C., et al. (1999). Neural transplantation of human neuroteratocarcinoma (hNT) neurons into ischemic rats. A quantitative dose–response analysis of cell survival and behavioral recovery. Neuroscience 91: 519–525. Shelton, F., Volpe, B., et al. (2001). Motor impairment as a predictor of functional recovery and guide to rehabilitation treatment after stroke. Neurorehabil. Neural Repair 15: 229–237. Snyder, E., Park, K., et al. (1997). Potential of neural “stem-like” cells for gene therapy and repair of the degenerating central nervous system. Adv. Neurol. 72: 121–132. Stilley, C., Ryan, C., et al. (2004). Changes in cognitive function after neuronal cell transplantation for basal ganglia stroke. Neurology 63: 1320–1322. Tessler, A., Fischer, I., et al. (1997). Embryonic spinal cord transplants enhance locomotor performance in spinalized newborn rats. Adv. Neurol. 72: 291–303. Thompson, T., Lunsford, L., et al. (1999). Restorative neurosurgery: opportunities for restoration of function in acquired, degenerative, and idiopathic neurological diseases. Neurosurgery 45: 741–752. Trojanowski, J., Kleppner, S., et al. (1997). Transfectable and transplantable postmitotic human neurons: a potential “platform” for gene therapy of nervous system diseases. Exp. Neurol. 144: 92–97. Trojanowski, J., Mantione, J., et al. (1993). Neurons derived from a human teratocarcinoma cell line establish molecular and structural polarity following transplantation into the rodent brain. Exp. Neurol. 122. Zivin, J. and Choi, D. (1991). Stroke therapy. Sci. Am. 265: 56–63.

55 Cell-Based Drug Delivery Grace J. Lim, Sang Jin Lee, and Anthony Atala

INTRODUCTION Cell-based drug delivery can be defined as delivery of biological products from living cells for therapy. The use of cells to deliver therapeutic molecules in response to biological need is a physiologically favorable venue in drug delivery systems. Most biopharmaceuticals such as proteins, antibodies, hormones, growth factors, and enzymes are expensive, difficult to manufacture and require frequent administration because the body quickly degrades them. Most common approaches to drug delivery involve polymeric drug formulation where drugs are delivered from polymer-based implants in which the rate of release is controlled by the diffusion of the drug from the delivery system or by the timed degradation of the drug depot (Langer, 1990; Chen and Mooney, 2003; Lee and Kim, 2005; Stayton et al., 2005). Novel approaches using the cell’s capability to produce biological therapeutics are being developed for clinical applications in diabetes treatment, wound healing, pain control, and cancer therapy (Aebischer et al., 1991; Sun et al., 1996; Gappa, 2001; Sakiyama-Elbert et al., 2001; Xu et al., 2002; Kim et al., 2004). A major benefit of a cell-based drug delivery system lies in improving the patient’s compliance since this system would provide more concentrations of therapeutic products steadily at a localized site in a manner that is triggered by cellular activity. Therefore, instead of taking pills or injections frequently, injection or implantation of cells can deliver desirable therapeutics for as long as the cells function. This system would also permit the rate of drug release to be varied as a function of regeneration of damaged surrounding tissues since the drug release is regulated by biological feedback. The cell-based delivery system could be particularly useful when long-term protein delivery such as growth factor is required, where one would like to vary the rate of drug release spatially as a function of tissue remodeling. This chapter will provide an overview of cell-based protein delivery approaches related to tissue regeneration and restoration of normal tissue function and will describe cell sources and cell encapsulation systems associated with avoiding rejection and improving cell function.

CELLS AND CELL PRODUCTS AS DRUG SOURCES A cell-based drug delivery system requires an appropriate source of functional cells. The ability to source, cultivate, and manipulate proper cell types often limits what can be accomplished in cell-based therapy. The simplest sources of cells are primary cells from human (autologenic and allogenic origins) and animals (xenogeneic origin). Pancreatic islets, hepatocytes, kidney cells, parathyroid cells, chondrocytes, and adrenal chromaffin cells are important examples of primary cells which have been used for cell-based therapeutic delivery systems (Aebischer et al., 1991; Koo and Chang, 1993; Iwata et al., 1994; Sun et al., 1996; Hasse et al., 1997; Sefton et al., 1997; Wang et al., 1997; Calafiore et al., 1999; Chandy et al., 1999; Gappa et al., 2001; Sakai et al., 2001; Orive et al., 2003; Kim et al., 2004; Haque et al., 2005). The major advantage of using primary cells as a drug source is their simple application because they are fully differentiated cells. Therefore the biological therapeutics produced by primary cells can be readily used without further processing such as viral design for efficient gene transfection, differentiation, production, and purification

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in cases such as using stem cells and engineered cells. However, the disadvantages of using primary cells are variability in cell numbers, cell quality, dissection, and tissues that may arrive in many states or conditions. The most straightforward application of cell-based therapies is the local delivery of therapeutic compounds by engineered cells at the site of transplantation. The advantage of genetically engineered cells is that a steady and potentially more physiologic concentration of a therapeutic compound may be achieved without the complication of systemic side effects. For instance, baby hamster kidney (BHK) cells were transfected with a human nerve growth factor (hNGF) fusion gene and were encapsulated in a semipermeable polymeric membrane and were transplanted into rat brains. The engineered cells successfully continued to release hNGF in vivo (Emerich et al., 1994). A similar approach was taken to release chemotherapeutic molecule from cells. To inhibit the growth of blood vessels in tumors, BHK cells were transfected with human endostatin (hES) expression vector and encapsulated with alginate and poly(L-lysine). The endostatin, an inhibitor of angiogenesis, was continuously released from microencapsulated engineered cells and more effectively reversed the growth of blood vessels feeding a tumor compared to discrete injections of the same molecule (Joki et al., 2001). The attributes of engineered cells for cell sources are their lower immunogenicity and higher capacity for in vivo survival. Problems associated with engineering cells involve gene transfection efficiency, risk of viral vectors, related safety, and multiple purification processing. However, with advances in genetic engineering techniques, application of genetically modified cells for therapeutic delivery is improving and promising. Stem cells and their derivatives have emerged as a promising source for cell-based drug delivery because of their ability to differentiate into various somatic cell types, the virtually unlimited donor source for transplantation, and the advantage of being flexible to a wide spectrum of genetic manipulations. For example, the antiepileptic potential of adenosine was exploited by intracerebral implants of cells engineered to release adenosine. The local release of adenosine from these encapsulated engineered cells was demonstrated to suppress seizures in kindled rats (Huber et al., 2001). However, long-term studies were precluded by the limited viability of encapsulated fibroblasts. To achieve long-term cell survival and potentially direct integration of therapeutic cells into the affected host tissue, stem cell-derived brain implants should constitute a superior source for cell grafting. Encapsulated embryonic stem (ES) cell-derived embryoid bodies and glial precursor cells released paracrine adenosine when the capsules were grafted into the lateral brain ventricles of kindled rats and successfully suppressed the seizure (Guttinger et al., 2005). Bone marrow stem cells (BMSCs) are another representative cell source for cell-based therapy since BMSCs transplantations are performed in thousands of patients as part of cancer treatments each year (Liu and Chang, 2002). Bone marrow transplants allow cancer patients to survive potentially lethal doses of chemotherapy and radiation since high doses of cytotoxic drugs and radiation destroy hematopoietic stem cells. These are the bone marrow cells that give rise to all blood cell types, leaving patients prone to life-threatening infections and anemia. The quantity of stem cells that can be harvested for large-scale use is still limited at present, and immune protection is required when using allogeneic stem cells. Because ES cells can be maintained and expanded in an undifferentiated state, it is possible to generate virtually unlimited numbers of cells for transplantation. However, direct grafting of undifferentiated ES cells is restricted by the formation of teratomas associated with tumor growth and low graft survival (Lindvall et al., 2004). Therefore, a protocol has been established that permits the efficient in vitro generation of precursors for oligodendrocytes and astrocytes. Each cell source has advantages and disadvantages depending on applications. It is necessary to select proper cell sources in consideration of different diseases, therapeutic efficacy, and long-term safety.

CELL ENCAPSULATION FOR THERAPEUTIC DELIVERY MACHINERY Cell encapsulation has been the primary machinery for cell-based therapeutic delivery systems. Cell microencapsulation is probably the preferable system for cell transplantation and can be used in both organ replacement

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956 THERAPEUTIC APPLICATIONS: CELL THERAPY

Figure 55.1 Schematic diagram of cell encapsulation. Nutrients and oxygen diffuse across the membrane, whereas inflammatory cells, antibodies, and immune cells are excluded.

and the continuous and controlled delivery of drugs. This technique consists of enclosing the biologically active material within a polymeric matrix surrounded by a semipermeable membrane that is designed to circumvent immune rejection. The capsule membrane allows the bi-directional diffusion of nutrients, oxygen, and waste and the secretion of the therapeutic product. It has the advantage of preventing immune cells and antibodies, which might destroy the enclosed cells, from entering the capsule (Figure 55.1). The capsules deliver large molecular weight proteins like insulin through routes other than conventional injection. In this case, insulin is being manufactured not in the drug company’s facilities, but in the transplanted cells, and insulin is delivered directly to the patient in response to glucose levels in the blood. The rising concept of cell-based therapeutics requires advances in cell encapsulation technology, and there have been successful efforts in applying this technology for the treatment of human diseases including renal failure, neurological disorders, cancers, and liver diseases (Aebischer et al., 1991; Emerich et al., 1994; Hasse et al., 1997; Liu and Chang, 2002; Brodie and Humes, 2005). Parameters for Cell Encapsulation Since Chang proposed the idea of using ultrathin polymer membrane microcapsules of the immunoprotection of transplanted cells in 1964 (Chang et al., 1964), a great number of techniques of cell encapsulation have been developed. Such increased interest in this field started when Lim and Sun published their results on the pancreatic islet encapsulation in the way of mild electrostatic cross-linking of sodium alginate and its complexation by poly(L-lysine), which is now the most commonly used cell encapsulation technique (Lim et al., 1980). The study showed that microencapsulated islets implanted in rats corrected the diabetic state for several weeks by producing insulin. Following this technique, several polymeric encapsulation systems have been developed and are currently being tested in clinical trials. For example, Novocell, Inc. has developed a photopolymerizable poly(ethylene glycol) polymer to encapsulate individual cells or cell clusters. Although much effort has been focused on identifying alternative systems to alginate/poly(L-lysine) chemistry, none have overcome all of the disadvantages of the poly(L-lysine). Therefore, although the polycation is required for xenotransplantation, and this renders the solution problematic, the allografts are likely to revert to the uncoated alginate beads. There are a variety of cell encapsulation methods using polymeric materials for treatment of disease and relevant methods and these are listed in Table 55.1.

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Table 55.1 Cells and materials used for cell encapsulation Cells

Functions

Materials

Human and animal primary cells Pancreatic islets

Diabetes

Alginate–poly(L-lysine)–alginate (Sun et al., 1996) Alginate–aminopropylsilicate–alginate (Sakai et al., 2001) Alginate–poly(L-ornithine) (Calafiore et al., 1999) Alginate–cellulose sulfate-poly(methylene-co-guanidine) (Wang et al., 1997) Agarose–poly(styrene sulfonic acid) (Iwata et al., 1994) Poly(N-isopropylacrylamide-co-acrylic acid) (Gappa et al., 2001) Aginate–poly(L-lysine)–poly(ethyleneimine)–protamine–heparin (Tatarkiewicz et al., 1994) Alginate–chitosan–polyethylene glycol (Chandy et al., 1999)

Hepatocytes

Liver transplantation

Alginate–chitosan (Haque et al., 2005) Hydroxyethyl methacrylate–methyl methacrylate (Sefton et al., 1997)

Kidney cells

Erythropoietin

Alginate–poly(L-lysine)–alginate (Koo et al., 1993)

Parathyroid cells

Parathyroid hormone

Alginate (Hasse et al., 1997)

Chromaffin cells

Catecholamines

Alginate–poly(L-lysine)–alginate (Aebischer et al., 1991; Kim et al., 2004)

Chondrocytes

Chondrocyte transplantation

Alginate (Grunder et al., 2004)

Hybridomas

Antibody production

Alginate–agarose (Orive et al., 2003)

Stem cells BMSC

Improve hepatocyte survival

Alginate–poly(L-lysine)–alginate (Liu and Chang, 2002, 2005)

Embryonic cells

Epilepsy

Polyethersulfone hollow fiber (Lindvall et al., 2004)

Mesenchymal stem cells

Tissue repair

Collagen–agarose (Batorsky et al., 2005)

Genetically engineered cells BHK cells

hNGF

Poly(acrylonitrile–vinyl chloride) (PAN–PVC) (Winn et al., 1994)

BHK cells

VEGF

Polysulfone hollow fiber (Yano et al., 2005)

BHK cells

Human ciliary neurotrophic factor (hCNTF)

Polyethersulfone (Bachoud-Levi et al., 2000)

Myoblasts

Mouse growth hormone (GH)

Alginate–poly(L-lysine)–alginate (Al-Hendy et al., 1996)

Mouse C2C12 myoblasts

Adenosine

Polyethersulfone (Guttinger et al., 2005)

SK2 hybridoma cells

Anti-human interleukin-6 (hIL-6) monoclonal antibodies

Alginate–poly(L-lysine)–alginate (Okada et al., 1997)

Mouse Ltk fibroblast

Human growth hormone (hGH)

Alginate–poly(L-lysine)–alginate (Basic et al., 1996)

Xenogeneic tumor cells (Neuro2A)

Beta-endorphin

Polyethersulfone hollow fiber (Saitoh et al., 1995)

iNOS-expressing cells

Inducible nitric oxide synthase gene (iNOS) for tumor suppression

Alginate–poly(L-lysine) (Xu et al., 2002)

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Membrane permeability is a function of both transport and thermodynamic properties, which are dependent upon the molecular characteristics of both the membrane and solute population. Thus materials for cell encapsulation should be selected or designed for each specific therapeutic device, as one may engineer several different membranes with required membrane properties for a desired application. The use of different membranes allows for variations in permeability, mass transfer, mechanical stability, buffering capability, biocompatibility, and other characteristics. A balance, however, has to be maintained among the physical properties of capsule membranes so as to support the entrapped cells’ survival. The mass transport properties of a membrane are critical since the influx rate of molecules, essential for cell survival, and the outflow rate of metabolic waste ultimately determine the viability of entrapped cells. Ordinarily the desired capsule permeability is determined by the molecular weight cutoff (MWCO) and is applicationdependent. The MWCO is the maximum molecular weight of a molecule that is allowed passage through the pores of the capsule membrane. For transplantation, the MWCO must be high enough to allow passage of nutrients but low enough to reject antibodies and other immune molecules (Uludag et al., 2000). Table 55.2 summarizes the MWCO of membranes and related therapeutic molecules. Recent efforts at defining the membrane permeability of biologically relevant proteins, rather than the use of arbitrary markers of varying molecular size, will likely have greater predictive capacity with respect to in vivo performance. Cell encapsulation technology plays important roles in not only by providing immune protection by isolating encapsulated cells from host tissue but by maintaining the phenotype of cells by providing a proper 3D environment and subsequently enhancing the production of therapeutic biologics from cells. For example, when autologous chondrocytes were expanded in capsule in vitro, these cells did not dedifferentiate and maintained their phenotype by high expression of type I collagen and a decrease in type II collagen expression (Grunder et al., 2004). A co-encapsulation approach is widely used to increase the duration of viability and function of cells. For example, co-encapsulated hepatocytes with BMSCs resulted in increased viability of the hepatocytes in vitro and in vivo, and also significantly prolonged the lowering of high systemic bilirubin levels in congenital Gunn rats with defects in the liver enzyme uridine diphosphate glucuronosyltransferase (UDPGT) (Liu and Chang, 2002). Challenges in Cell Capsule Technology In spite of a great promise of cell encapsulation concepts, there have been continuous challenges in cell capsulebased therapeutic delivery. The major challenge is long-term cell survival or prolonged cell viability in capsules. However, the encapsulated cells have a limitation due to the supply of nutrients and oxygen. Nutrients typically include low molecular weight solutes such as glucose, macromolecules such as albumin, and transferrin for iron uptake. Growth factors may also be required. Although the transport limitations for macromolecules have not yet been quantified, it is likely that oxygen supply limitations are the most serious. A class of microporous membranes that induce neovascularization membrane is in direct contact with the bloodstream at an arterial pO2 of 100 mmHg. By contrast, extravascular devices implanted intraperitoneally or in subcutaneous tissue are exposed to the average pO2 of the microvasculature (40 mmHg). Implantation in soft tissue is further disadvantaged if a foreign-body response occurs, in which an avascular layer typically 100 μm thick is produced adjacent to the membrane. This fibrotic tissue increases the distance between blood vessels and the implant, and the fibroblasts in the avascular layer consume oxygen. Researchers dealing with the limitations in oxygen transport attempted using cross-linked hemoglobin (Hb-C), and inclusion of materials that induce neovascularization in the vicinity of the implant use (Chae et al., 2004). Prolonged glucose normalization of streptozotocin-induced diabetic mice was observed by transplantation of rat islets co-encapsulated with cross-linked hemoglobin, while the mice that received the conventional control islet microcapsule (without Hb-C) transplant showed graft failure in 4 weeks, exhibited by hyperglycemia, weight loss, and deteriorated glucose tolerance.

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Table 55.2 MWCOs of semipermeable cell capsule membranes (Prakash et al., 2005) MWCOs of semipermeable membranes

Molecules (molecular weight) Leukocytes IgM (950 kd) Urease (482.7 kd) C19 (410 kd) Fibrinogen (339 kd) Phenylalanine NH3 lyase (320 kd) Catalase (247 kd) C4 (210 kd)

Hollow fiber (200 kd)

C5 (195 kd) IgE (190 kd) Human leukocytes antigen (180–210 kd) C3 (185 kd) IgA (170–720 kd) C2 (170 kd) C8 (163 kd) IgD (160 kd) IgG (150 kd) Tyrosinase (128 kd) C6 (110 kd) C7 (100 kd) Transferrin (81 kd) C9 (79 kd)

Alginate–Poly(L-lysine)–Alginate (60–70 kd)

Albumin (66.248 kd) Hemoglobin (64 kd) FactorX (55 kd) Tumor necrosis factor (TNF) (51 kd) Platelet-derived growth factor-C (46/30 kd) Superoxide dismutase (31.187 kd)

Cellulose nitrate or polyamide (30 kd)

Fibroblast growth factor-7 (28 kd) Vascular endothelial growth factor (21/42 kd) Bone morphogenic proteins-4 (18/33 kd) Interleukin-beta (17 kd) Fibroblast growth factor-2 (17 kd) Insulin-like growth factor-1 (17 kd) Fibroblast growth factor-1 (15.5 kd) Platelet-derived growth factor-B (14/33 kd) Platelet-derived growth factor-A (14 kd) Nerve growth factor (13 kd) C3a (9000 d) Epidermal growth factor (6 kd) Insulin (5.7 kd) Beta-endorphin (3.4 kd)

Lipid-complexed polymer (100–200 d)

Glucose (180 d) Tyrosine (163 d) Phenylalanine (147 d) Glutamine (128 d) Aspargine (114 d) Creatinine (113 d)

Lipid vesicles (lipophilic)

Urea (60 d) Carbon dioxide (44 d) Ammonia (17 d) Oxygen (16 d)

960 THERAPEUTIC APPLICATIONS: CELL THERAPY

Device geometry also critically affects the local pO2 to which cells are exposed. A spherical geometry is known to be most advantageous because of the high surface area to volume ratio. Thus, an islet (150 μm in diameter) microencapsulated in an alginate bead (600–800 μm in diameter) was shown to be less susceptible to oxygen mass transfer than a tubular or planar diffusion chamber (Colton, 1995).

CELL-BASED PROTEIN FACTORY Cells can be manufactured to produce a therapeutic protein as a protein factory in vivo for controlled release. One of these approaches include manipulating cells to deliver growth factor, that would allow the stable incorporation of growth factors within a cell in-growth matrix in a manner such that local enzymatic activity associated with tissue regeneration could trigger growth factor release. A research group (Sakiyama-Elbert et al., 2001) investigated this approach in the context of peripheral nerve regeneration by designing modified beta-nerve growth factor (NGF) fusion proteins and testing their ability to promote neurite extension. They selected fibrin as the cell in-growth matrix, and the transglutaminase activity of factor XIIIa to covalently incorporate NGF fusion proteins within fibrin matrices as shown in Figure 55.3. Novel NGF fusion proteins, which contained an exogenous factor XIIIa substrate to allow incorporation into fibrin matrices, were expressed recombinantly. An intervening plasmin substrate domain

(a)

(b)

Figure 55.2 Confocal images of microcapsulated chromaffin cells. (a) Before implantation of microencapsulated chromaffin cells. (b) After retrieval of the microcapsules from the subarachnoid space 30 days after implantation. Images were captured with a confocal laser scanning microscope (100 magnification) (Kim et al., 2004).

Factor IIIa substrate

Active or inactive plasmindegradable substrate

Human -NGF

Degraded plasminsubstrate Plasmin cleavage

Human -NGF

Fibrin Plasmin Fibrin

Figure 55.3 Cell triggered growth factor delivery. β-NGF fusion proteins with exogenous domains for growth factor immobilization via the transglutaminase factor XIIIa and cell-triggered release via the proteolytic activity of plasmin (Sakiyama-Elbert et al., 2001).

Cell-Based Drug Delivery 961

was placed between the factor XIIIa substrate and the NGF domain to allow cell-mediated growth factor release in response to plasmin, which is generated by invading cells. The results showed that by placing an enzymatically degradable linker between the cross-linking substrate and the growth factor domain in the fusion protein, growth factors can be delivered in an active form in response to cell-regulated processes. It further suggested that the release of immobilized growth factors in a manner that can be temporally and spatially regulated by cell-associated enzymatic processes may be important in the context of wound healing. Thus, delivery systems that allow drug release to be regulated by the progress of wound healing through a cell in-growth matrix could prove to be more effective in promoting successful tissue regeneration. As another example, implantable protein factory (ImPACT) products have been created by Cell Based Delivery (CBD), Inc., using muscle cells. CBD’s ImPACT™ products deliver predictable, therapeutic levels of proteins, such as vascular endothelial growth factor (VEGF) and fibroblast growth factor (FGF) to stimulate rapid, sustained angiogenesis. It was reported that the growth factors were shown to be functional for an extended period in the body, and the use of these products for various cardiovascular diseases, hormonal growth deficiency, musculoskeletal disease, and solid tumors are in preclinical trials with animals. Rugged and compact, ImPACT™ tissue implants measure only 20 mm long and 1–2 mm wide. Using standard catheter or minimally invasive techniques, interventional cardiologists and cardiothoracic surgeons can position ImPACT™ at selected sites with fibrin glues or stents. ImPACT products have survived in animals for up to 6 months with no sign of tissue loss, indicating that long-lasting drug delivery can be achieved with just a single procedure. Currently available protein-delivery systems, such as biodegradable microspheres, provide several weeks of continuous dosing. Moreover, because muscle cells live for years, autologous ImPACT products could ultimately last a year or longer. For example, products placed below the skin could deliver proteins systemically for the treatment of chronic diseases such as hemophilia or anemia. For production of therapeutic substances for suppressing cancer, cell-based delivery system is currently in use. hES secreting cells were engineered using BHK-21 for cancer therapy by a research group (Joki et al., 2001). It was found that cell-based delivery of endostatin, an inhibitor of angiogenesis, more effectively reverses the growth of blood vessels feeding a tumor rather than do discrete injections of the same protein. Therefore rather than expressing a therapeutic protein in cultured cells, then purifying it into the patient, it would make it easier for patients by implanting cells directly once or a few times a year rather than taking a pill or injection daily. Protein delivery based on cells is promising and attracts many scientists as an alternative therapeutic administration.

DRUG-LOADED TUMOR CELL SYSTEM Cell-based drug delivery system does not necessarily need living cells only. Dead cells can be used for controlled delivery of therapeutic molecules. A drug-loaded tumor cell (DLTC) system has been developed for lung metastasis-targeting drug delivery. Doxorubicin was loaded into B16-F10 murine melanoma, and the loading process led to the death of all the carrier cells. The diameter of DLTC was approximately 15 μm (Shao et al., 2001). The amount and rate of doxorubicin being released from the DLTC mainly depended on the drug loading and carrier cell concentration. Over 6 month storage in phosphate buffered saline (PBS) at 4°C, the decrease in intracellular drug concentration and the carrier cell numbers were less than 25% and 5%, respectively. After a bolus injection of 30 μg doxorubicin in either DLTC form or free solution into the mice tail veins, drug deposit in the lung from DLTC was about 4-fold of that achieved by free drug solution. In spite of potential problems associated with using dead cells as drug carriers, the finding from extensive research strongly suggested the DLTC system possessed a lung-targeting activity that may be partially attributed to its specific surface characteristics.

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Table 55.3 Clinical applications of cell-based drug delivery Year

Cell types

Formulation

Target diseases

References

1994

Pancreas islets

Diabetes

Soon-Shiong et al.

1994

Amyotrophic lateral sclerosis

Aebischer et al.

1996

Encapsulated xenogeneic cells BHK cells

Amyotrophic lateral sclerosis

Aebischer et al.

1999

Xenogeneic cells

2001

BHK cells Allogeneic CYP2B1expressing cells Parathyroid cells

Amyotrophic lateral sclerosis and chronic cancer pain Amyotrophic lateral sclerosis Pancreatic tumor

Abicht and Lochmuller

2000 2001

Alginate high in guluronic acid Alginate–poly(L-lysine)– alginate Alginate–poly(L-lysine)– alginate Polymer-based hollow fibers Hollow fiber Cellulose sulfate

Chronic hypoparathyroidism

Tibell et al.

2004

BHK cells

Polytetrafluroethylene (PTFE) membrane Thermoplastic polyethersulfone

Huntington’s disease

Bloch et al.

Zurn et al. Lohr et al.

Table 55.4 Selected companies working on cell-based therapies Company

Cell-based drug delivery BioHybrid Technologies (Shrewsbury, MA) Islet Sheet Medical (San Francisco, CA) Ixion Biotechnology (Alachua, FL) Neurotech (Evry, France) Novocell (Irvine, CA) Oxford BioMedica (Oxford, UK) Layton Biosciences (Sunnyvale, CA) Cell Based Delivery (Providence, RI) Sertoli Technologies (Cranston, RI)

Cell-based immunotherapy Aastrom Bioscience (Ann Arbor, MI) CellExSys (Seattle, WA) Geron (Menlo Park, CA)

Technology

Major disease focus

Encapsulation system for allografts

Encapsulated pancreatic islet cell allografts; therapeutic protein delivery Retrievable bioartificial pancreas for diabetes Unencapsulated islet cell allografts for diabetes Therapeutic protein delivery to eye and brain Encapsulated islet cell allografts for diabetes Prodrug activating enzyme (CYP2B6) for treating cancers Central nervous system (CNS) disorders (stroke, tumors, Parkinson’s disease, Alzheimer’s disease)

Encapsulated pancreatic islet cells Pancreatic islet-producing human stem cells Encapsulation system for allografts Individually polymer-coated pancreatic islet cells MacroGen (macrophages as gene delivery systems) Human neuronal stem cells

Implantable protein-expressing muscle tissue allografts Sertoli cells to protect implanted allografts

Therapeutic protein delivery for chronic diseases Pancreatic islet cell allografts, therapeutic protein delivery, and cell-based gene therapy

Autologous cell processing system, bone marrow, and cord blood stem cells Ex vivo production of cytotoxic T lymphocytes Human ES cells; dendritic cell vaccines

Dendritic cell-based cancer vaccine, solid tissue, and blood regeneration with stem cells Cell-based treatments for hepatitis B and C, cancer Cell-based treatments for cancer, diabetes, osteoarthritis

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Table 55.4 (Continued) Company

Technology

Major disease focus

Xcyte Therapies (Seattle, WA) Immuno-Designed Molecules (Paris, France)

Ex vivo expansion of T-cells Immunotherapeutics, and dendritic cells used as vaccines; cell processor technologies

Autologous cytotoxic T-cell generation to treat cancer, infectious diseases Cell drugs, to enhance immunity to treat cancer

Cell retrieval and expansion Gamida-Cell (Jerusalem, Israel) Nexell Therapeutics (Irvine, CA)

System for expanding stem cell populations ex vivo System for isolating hematopoietic stem cells

TEI Biosciences (Boston, MA) Progenitor Cell Therapy (Saddle Brook, NJ)

Signaling molecules to induce stem cell differentiation Cell therapy manufacturing services

Cytomatrix (Woburn, MA)

A 3D matrix, an artificial thymus, for the growth and maturation of T-cells A 3D matrix for growing cells ex vivo

Select Therapeutics (Woburn, MA)

Hematopoietic stem cells-derived umbilical cord blood for use in high dose chemotherapy Stem cell therapy for chronic granulomatous disease and other hereditary blood disorders; cancer vaccines Tissue engineering using derived cell types Good manufacture practice (GMP) factory and distribution system to grow and deliver autologous therapies nationwide (USA) Cell culture devices for bench research and bioreactors for commercial production of cells Expansion of hematopoietic stem cells for bone marrow transplants, cytotoxic T-cells to treat cancer

SUMMARY Delivery of biological products from living cells in response to biological need is a physiologically attractive approach. There have been successful business and clinical attempts for producing therapeutic proteins from various types of cells, as is summarized in Tables 55 3 and 55.4. Cell-based delivery system might allow a lower total drug dose to be incorporated within the delivery system, and spatial regulation of release could permit a greater percentage of the drug to be released at the time and place of greatest cellular activity. The significance of cell-based drug delivery is using biological feedback control in drug release, which could overcome the limit of a polymer-based drug delivery system. However, issues on long-term viability, risk of immune development, related safety, and retrieval of the unwanted cells should be addressed to further explore their possible clinical applications. Many experimental applications of drug delivery systems are easing their way into the clinic, and the hope that cells may be used for therapeutics seems increasingly likely to be realized.

REFERENCES Abicht, A. and Lochmuller, H. (1999). Technology evaluation: CRIB (CNTF delivery) CytoTherapeutics Inc. Curr. Opin. Mol. Ther. 1: 645–650. Aebischer, P., Tresco, P.A., Sangen, J. and Winn, S.R. (1991). Transplantation of microencapsulated bovine chromaffin cells reduces lesion-induced rotational asymmetry in rats. Brain Res. 560: 43–49. Aebischer, P., Buchser, E., Joseph, J.M., Favre, J., de Tribolet, N., Lysaght, M., Rudnick, S. and Goddard, M. (1994). Transplantation in humans of encapsulated xenogeneic cells without immunosuppression. A preliminary report. Transplantation 58: 1275–1277.

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Aebischer, P., Schluep, M., Deglon, N., Joseph, J.M., Hirt, L., Heyd, B., Goddard, M., Hammang, J.P., Zurn, A.D., Kato, A.C., Regli, F. and Baetge, E.E. (1996). Intrathecal delivery of CNTF using encapsulated genetically modified xenogeneic cells in amyotrophic lateral sclerosis patients. Nat. Med. 2: 696–699. Al-Hendy, A., Hortelano, G., Tannenbaum, G.S. and Chang, P.L. (1996). Growth retardation – an unexpected outcome from growth hormone gene therapy in normal mice with microencapsulated myoblasts. Hum. Gene Ther. 7: 61–70. Bachoud-Levi, A.C., Deglon, N., Nguyen, J.P., Bloch, J., Bourdet, C., Winkel, L., Remy, P., Goddard, M., Lefaucheur, J.P., Brugieres, P., Baudic, S., Cesaro, P., Peschanski, M. and Aebischer, P. (2000). Neuroprotective gene therapy for Huntington’s disease using a polymer encapsulated BHK cell line engineered to secrete human CNTF. Hum. Gene Ther. 11: 1723–1729. Basic, D., Vacek, I. and Sun, A.M. (1996). Microencapsulation and transplantation of genetically engineered cells: a new approach to somatic gene therapy. Arif. Cells Blood Substit. Immobil. Biotechnol. 24: 219–255. Batorsky, A., Liao, J., Lund, A.W., Plopper, G.E. and Stegemann, J.P. (2005). Encapsulation of adult human mesenchymal stem cells within collagen–agarose microenvironments. Biotech. Bioeng. 92: 492–500. Bloch, J., Bachoud-Levi, A.C., Deglon, N., Lefaucheur, J.P., Winkel, L., Palfi, S., Nguyen, J.P., Bourdet, C., Gaura, V., Remy, P., Brugieres, P., Boisse, M.F., Baudic, S., Cesaro, P., Hantraye, P., Aebischer, P. and Peschanski, M. (2004). Neuroprotective gene therapy for Huntington’s disease, using polymer-encapsulated cells engineered to secrete human ciliary neurotrophic factor: results of a phase I study. Hum. Gene Ther. 15: 968–975. Brodie, J.C. and Humes, H.D. (2005). Stem cell approaches for the treatment of renal failure. Pharmacol. Rev. 57(3): 299–313. Chae, S.Y., Kim, Y.Y., Kim, S.W., and Bae, Y.H. (2004). Prolonged glucose normalization of streptozotocin-induced diabetic mice by transplantation of rat islets coencapsulated with crosslinked hemoglobin. Transplantation. 78: 392–397. Chandy, T., Mooradian, D.L. and Rao, G.H. (1999). Evaluation of modified alginate–chitosan–polyethylene glycol microcapsules for cell encapsulation. Artif. Organs 23: 894–903. Chang, T.M.S. (1964). Semipermeable microcapsules. Science. 146: 524–525. Chen Calafiore, R., Basta, G., Luca, G., Boselli, C., Bufalari, A., Cassarani, M.P., Giustozzi, G.M. and Brunetti, P. (1999). Transplantation of pancreatic islets contained in minimal volume microcapsules in diabetic high mammalians. Ann. NY Acad. Sci. 875: 219–232. Colton, C.K. (1995). Implantable biohybrid artificial organs. Cell Transplant. 4: 415–436. Chen, R.R. and Mooney, D.J. (2003). Polymeric growth factor delivery strategies for tissue engineering. Pharm. Res. 20: 1103–1112. Emerich, D.F., Winn, S.R., Harper, J., Hammang, J.P., Baetge, E.E. and Kordower, J.H. (1994). Implants of polymerencapsulated human NGF-secreting cells in the nonhuman primate: rescue and sprouting of degenerating cholinergic basal forebrain neurons. J. Comp. Neurol. 349: 148–164. Gappa, H., Baudys, M., Koh, J.J., Kim, S.W. and Bae, Y.H. (2001). The effect of zinc-crystallized glucagon-like peptide-1 on insulin secretion of macroencapsulated pancreatic islets. Tissue Eng. 7: 35–44.Grunder, T., Gaissmaier, C., Fritz, J., Stoop, R., Hortschansky, P., Mollenhauer, J. and Aicher, W.K. (2004). Bone morphogenetic protein-2 enhances the expression of type II collagen and aggrecan in chondrocytes embedded in alginate beads. Osteoarthritis Cartilage 12: 559–567. Guttinger, M., Padrun, V., Pralong, W.F. and Boison, D. (2005). Seizure suppression and lack of adenosine A1 receptor desensitization after focal long-term delivery of adenosine by encapsulated myoblasts. Exp. Neurol. 193: 53–54. Haque, T., Chen, H., Ouyang, W., Martoni, C., Lawuyi, B., Urbanska, A.M. and Prakash, S. (2005). In vitro study of alginate–chitosan microcapsules: an alternative to liver cell transplants for the treatment of liver failure. Biotechnol. Lett. 27: 317–322. Hasse, C., Klock, G., Schlosser, A., Zimmermann, U. and Rothmund, M. (1997). Parathyroid allotransplantation without immunosuppression. Lancet 350: 1296–1297. Huber, A., Padrun, V., Deglon, N., Aebischer, P., Mohler, H., and Boison, D. (2001). Grafts or adenosine-releasing cells suppress seizures in kindling epilepsy. Proc Natl Acad Sci U. S. A. 98: 7611–7616 Iwata, H., Takai, T., Kobayashi, K., Oka, T., Tsuji, T. and Ito, F. (1994). Strategy for developing microbeads applicable to islet xenotransplantation into a spontaneous diabetic NOD mouse. J. Biomed. Mater. Res. 28: 1201–1207.

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Joki, T., Machluf, M., Atala, A., Zhu, J., Seyfried, N., Dunn, I., Abe, T., Carroll, R. and Black, P. (2001). Continuous release of endostatin from microencapsulated engineered cells for tumor therapy. Nat. Biotechnol. 19: 35–39. Kim, Y.M., Jeon, Y.H., Jin, G.C., Lim, J.O. and Baek, W.Y. (2004). Immunoisolated chromaffin cells implanted into the subarachnoid space of rats reduce cold allodynia in a model of neuropathic pain: a novel application of microencapsulation technology. Artif. Organs 28: 1059–1066. Koo, J. and Chang, T.M. (1993). Secretion erythropoietin from microencapsulated rat kidney cells: preliminary results. Int. J. Artif. Organs 16: 557–560. Langer, R. (1990). New methods of drug delivery. Science 249: 1527–1533. Lee, M. and Kim, S.W. (2005). Polyethylene glycol-conjugated copolymers for plasmid DNA delivery. Pharm. Res. 22: 1–10. Lim, F. and Sun, A.M. (1980). Microencapsulated islets as bioartificial endocrine pancreas. Science. 210: 908–910. Liu, Z.C. and Chang, T.M. (2002). Increased viability of transplanted hepatocytes when hepatocytes are co-encapsulated with bone marrow stem cells using a novel method. Artif. Cells Blood Substit. Immobil. Biotechnol. 30: 99–112. Liu, Z.C. and Chang, T.M. (2005). Transplantation of bioencapsulated bone marrow stem cells improves hepatic regeneration and survival of 90% hepatectomized rats: a preliminary report. Artif. Cells Blood Substit. Immobil. Biotechnol. 33: 405–410. Lohr, M., Hoffmeyer, A., Kroger, J., Freund, M., Hain, J., Holle, A., Karle, P., Knofel, W.T., Liebe, S., Muller, P., Nizze, H., Renner, M., Saller, R.M., Wagner, T., Hauenstein, K., Gunzburg, W.H. and Salmons, B. (2001). Microencapsulated cellmediated treatment of inoperable pancreatic carcinoma. Lancet 357: 1591–1592. Okada, N., Miyamoto, H., Yoshioka, T., Katsume, A., Saito, H., Yorozu, K., Ueda, O., Itoh, N., Mizuguchi, H., Nakagawa, S., Ohsugi, Y. and Mayumi, T. (1997). Cytomedical therapy for IgG1 plasmacytosis in human interleukin-6 transgenic mice using hybridoma cells microencapsulated in alginate-poly(L)lysine–alginate membrane.Biochim. Biophys. Acta. 1360: 53–63. Orive, G., Hernandez, R.M., Gascon, A.R., Calafiore, R., Chang, T.M., De Vos, P., Hortelano, G., Hunkeler, D., Lacik, I., Shapiro, A.M. and Pedraz, J.L. (2003). Cell encapsulation: promise and progress. Nat. Med. 9: 104–107. Prakash, S. and Jones, M.L. (2005). Artificial cell therapy: new strategies for the therapeutic delivery of live bacteria. J Biomed Biotech. 1: 44–56. Saitoh, Y., Taki, T., Arita, N., Ohnishi, T. and Hayakawa, T. (1995). Cell therapy with encapsulated xenogenic tumor cells secreting beta-endorphin for treatment of peripheral pain. Cell Transplant. 1: S13–S17. Sakai, S., Ono, T., Ijima, H. and Kawakami, K. (2001). Synthesis and transport characterization of alginate/aminopropylsilicate/alginate microcapsule: application to bioartificial pancreas. Biomaterials 22: 2827–2834. Sakiyama-Elbert, S.E., Panitch, A. and Hubbell, J.A. (2001). Development of growth factor fusion proteins for cell triggered drug delivery. FASEB J. 15: 1300–1302. Sefton, M.V., Hwang, J.R. and Babensee, J.E. (1997). Selected aspects of the microencapsulation of mammalian cells in HEMA–MMA. Ann. NY Acad. Sci. 831: 260–270. Shao, J., DeHaven, J., Lamm, D., Weissman, D.N., Runyan, K., Malanga, C.J., Rojanasakul, Y. and Ma, J.K. (2001). A cellbased drug delivery system for lung targeting: I. Preparation and pharmacokinetics. Drug Deliv. 8: 61–69. Shao, J., DeHaven, J., Lamm, D., Weissman, D.N., Malanga, C.J., Rojanasakul, Y. and Ma, J.K. (2001). A cell-based drug delivery system for lung targeting: II. Therapeutic activities on B16-F10 melanoma in mouse lungs. Drug Deliv. 8: 71–76. Soon-Shiong, P., Heintz, R.E., Merideth, N., Yao, Q.X., Yao, Z., Zheng, T., Murphy, M., Moloney, M.K., Schmehl, M. and Harris, M. (1994). Insulin independence in a type 1 diabetic patient after encapsulated islet transplantation. Lancet 343: 950–951. Stayton, P.S., El-Sayed, M.E., Murthy, N., Bulmus, V., Lackey, C., Cheung, C. and Hoffman, A.S. (2005). “Smart” delivery systems for biomolecular therapeutics. Orthod. Craniofac. Res. 8: 219–225. Sun, Y., Ma, X., Zhou, D., Vacek, I. and Sun, A.M. (1996). Normalization of diabetes in spontaneously diabetic cynomologus monkeys by xenografts of microencapsulated porcine islets without immunosuppression. J. Clin. Invest. 98: 1417–1422. Tatarkiewicz, K., Sitarek, E., Fiedor, P., Sabat, M. and Orlowski, T. (1994). In vitro and in vivo evaluation of protamine–heparin membrane for microencapsulation of rat Langerhans islets. Artif. Organs 18: 736–739.

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Tibell, A., Rafael, E., Wennberg, L., Nordenstrom, J., Bergstrom, M., Geller, R.L., Loudovaris, T., Johnson, R.C., Brauker, J.H., Neuenfeldt, S. and Wernerson, A. (2001). Survival of macroencapsulated allogeneic parathyroid tissue one year after transplantation in nonimmunosuppressed humans. Cell Transplant. 10: 591–599. Uludag, H., De Vos, P., and Tresco, P.A. (2000). Technology of mammalian cell encapsulation. Adv Drug Deliv Rev. 42: 29–64. Wang, T., Lacik, I., Brissova, M., Anikumar, A.V., Prokop, A., Hunkeler, D., Green, R., Shahrokhi, K. and Powers, A.C. (1997). An encapsulation system for the immunoisolation of pancreatic islets. Nat. Biotechnol. 15: 362–385. Winn, S.R., Hammang, J.P., Emerich, D.F., Lee, A., Palmiter, R.D. and Baetge, E.E. (1994). Polymer-encapsulated cells genetically modified to secrete human nerve growth factor promote the survival of axotomized septal cholinergic neurons. Proc. Natl Acad. Sci. USA 91: 2324–2328. Xu, W., Liu, L. and Charles, I.G. (2002). Microencapsulated iNOS-expressing cells cause tumor suppression in mice. FASEB J. 16: 213–215. Yano, A., Shingo, T., Takeuchi, A., Yasuhara, T., Kobayashi, K., Takahashi, K., Muraoka, K., Matsui, T., Miyoshi, Y., Hamada, H. and Date, I. (2005). Encapsulated vascular endothelial growth factor-secreting cell grafts have neuroprotective and angiogenic effects on focal cerebral ischemia. J. Neurosurg. 103: 104–114. Zurn, A.D., Henry, H., Schluep, M., Aubert, V., Winkel, L., Eilers, B., Bachmann, C. and Aebischer, P. (2000). Evaluation of an intrathecal immune response in amyotrophic lateral sclerosis patients implanted with encapsulated genetically engineered xenogeneic cells. Cell Transplant. 9: 471–484.

Part VI Therapeutic Applications: Tissue Therapy

56 Fetal Tissues Seyung Chung and Chester J. Koh

INTRODUCTION The field of regenerative medicine aims to replace damaged, diseased, or malformed tissue through the development of biological substitutes which can restore and maintain normal function. In following the principles of cell biology and transplantation, materials science and engineering, many current strategies for regenerative medicine depend upon a sample of autologous cells from the diseased organ of the host. Usually, a small piece of donor tissue is dissociated into individual cells, and either implanted directly into the host, or are expanded in culture, attached to a support matrix, and then reimplanted into the host after expansion (Oberpenning et al., 1999). The use of autologous cells avoids immunological rejection, and thus the deleterious side effects of immunosuppressive medications can be avoided. Ideally, both structural and functional tissue replacement will occur with minimal complications. However, for many patients with extensive end-stage organ failure, a tissue biopsy may not yield enough normal cells for expansion and transplantation. In other instances, primary autologous human cells cannot be expanded from a particular organ, such as the pancreas. In these situations, stem cells are envisioned as a viable source of cells, as they can serve as an alternative source of cells from which the desired tissue can be derived. Combining the regenerative medicine techniques discovered over the past few decades with this potentially endless source of versatile cells could lead to novel sources of replacement organs to replace diseased, damaged, or malformed tissue. Embryonic stem cells exhibit two remarkable properties: the ability to proliferate in a undifferentiated, but pluripotent state (self-renew), and the ability to differentiate into many specialized cell types (Brivanlou et al., 2003). They can be isolated by immunosurgery from the inner cell mass of the embryo during the blastocyst stage (5 days post-fertilization), and are usually grown on feeder layers consisting of mouse embryonic fibroblasts or human feeder cells (Richards et al., 2002). More recent reports have shown that these cells can be grown without the use of a feeder layer (Amit et al., 2003), and thus avoid the exposure of these human cells to mouse viruses and proteins. These cells have demonstrated longevity in culture by maintaining their undifferentiated state for at least 80 passages when grown using current published protocols (Thomson et al., 1998; Reubinoff et al., 2000). Human embryonic stem cells have been shown to differentiate into cells from all three embryonic germ layers in vitro. Skin and neurons have been formed, indicating ectodermal differentiation (Schuldiner et al., 2000, 2001; Reubinoff et al., 2001; Zhang et al., 2001). Blood, cardiac cells, cartilage, endothelial cells, and muscle have been formed, indicating mesodermal differentiation (Kaufman et al., 2001; Kehat et al., 2001; Levenberg et al., 2002). And pancreatic cells have been formed, indicating endodermal differentiation (Assady et al., 2001). In addition, as further evidence of their pluripotency, embryonic stem cells can form embryoid bodies, which are cell aggregations that contain all three embryonic germ layers, while in culture, and can form teratomas in vivo (Itskovitz-Eldor et al., 2000). However, the harvesting of human embryonic stem cells requires the destruction of human embryos, which has raised significant ethical and political concerns in the United States. In August, 2001, in a compromise

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between the stem cell research advocates and critics, the federal government ordered that only previously generated human embryonic stem cell lines could be approved for federal research funding, and over 70 different cell lines met this criterion at that time. However, as stated in National Institutes of Health (NIH) testimony before Congress in April, 2003, only 11 stem cell lines are currently available, which has had deleterious effects on the progress of stem cell research in the United States (Kennedy, 2003). In addition, most of the approved cell lines were grown in the presence of mouse cells (feeder cells), which can supply many needed growth factors, but which also expose the human cells to potential contamination from mouse viruses or proteins. This may render the current cell lines unsuitable for human therapeutic purposes. These barriers to progress in embryonic stem cell research have spawned the search for alternate sources of stem cells, and the use of fetal tissue as a source of stem cells may overcome some of the political and ethical controversies surrounding embryonic stem cells.

STEM CELLS DERIVED FROM FETAL TISSUES Fetal stem cells are not a new concept and in fact they have been in clinical use over the past 10 to 20 years. For example, umbilical cord blood (UCB) stem cells are widely used in the treatment of hematological disorders (Watt and Contreras, 2005), and fetal neural tissue has been associated with some clinical improvement in the treatment of Parkinson’s disease (Lindvall and Bjorklund, 2004). Several sources of stem cells derived from fetal tissues have been investigated, and a select few are listed below. For hematopoiesis, fetal bone marrow, as opposed to adult bone marrow, cord blood, or peripheral blood, appears to be the ideal source of stem cells for engraftment and therapeutic reconstitution, as they have a very high proliferative capacity, low immunogenicity, and the highest number of primitive stem/progenitor cells (Michejda, 2004). As transplanted mature hepatocytes have enormous repopulating capacity under conditions of continuous liver injury, progenitor cells from fetal liver cells have been isolated and transplanted, where up to 10% of a normal liver can be repopulated (Dabeva and Shafritz, 2003; Rollini et al., 2004). However, further studies are necessary to determine the regenerative capabilities of these cells for both liver regeneration as well as for other mesenchymal tissues. Mesenchymal stem cells (MSCs) have been isolated from human fetal blood, liver, and bone marrow, where they exhibited clonogenicity and were able to differentiate into the adipogenic, osteogenic, and chondrogenic lineages (Campagnoli et al., 2001). The fetal kidney may also be a potential source of MSCs, since the metanephric mesenchyme may represent a pluripotent population with a predilection toward the epithelial and stromal cell lineages (Al-Awqati and Oliver, 2002). Second-trimester fetal lung was also noted to be a source of MSCs, especially for the osteogenic and adipogenic lineages (in ‘t Anker et al., 2003). IMMUNOLOGICAL CONSIDERATIONS Fetal cells have long been known to exist in a microchimerism state in females during pregnancy, and it appears that microchimerism persists until decades later (O’Donoghue et al., 2004). Since fetal MSCs do not express human leukocyte antigen (HLA) class II antigens and may not express HLA class I antigens as well, this may help to explain this phenomenon. Since these early fetal stem cells appear to have a pre-immune status, this may be ideal for allogenic transplantation/mismatch situations, as both undifferentiated and differentiated fetal MSCs do not elicit alloreactive lymphocyte proliferation (Gotherstrom et al., 2004). ETHICAL CONSIDERATIONS: FETUS AND OOCYTES The ethics and politics of transplantation of human fetal tissue have long been debated, and opposition to this type of medical therapy still exists (Annas and Elias, 1989). Although relatively fewer individuals would object

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to the use of diseased or anencephalic fetal tissue for transplantation purposes, usually unsuitable fetal tissue is available from this source because of the associated pathology, such as chromosomal anomalies and infection (Abouna, 2003). In addition, tissue samples that otherwise would be discarded also tend to generate less controversy in its usage. However, for other tissue sources from viable fetuses, different beliefs about the beginning of life from opposing political and ethical groups have created vocal opposition to significant federal funding for this type of research, which has stymied progress in the field. A similar debate on the ethics of donated tissue has arisen in the area of therapeutic cloning and oocyte donation, and perhaps lessons can be learned from the resulting discussions (Magnus and Cho, 2005). Some areas for discussion include ethical oversight of collaborations between scientists working in countries with different standards, the protection of tissue donors, and the avoidance of unrealistic expectations arising from the research.

REGENERATIVE MEDICINE APPLICATIONS OF FETAL TISSUES Stem cells derived from fetal tissues have been utilized in regenerative medicine applications for many organ systems, and some of the recent studies are highlighted below. Neural Tissue Neural tissue regeneration is a complex biological phenomenon for which many laboratories have investigated time and resources in the quest for novel solutions to diseases such as Parkinson’s and Alzheimer’s. For peripheral nervous system injuries, these nerves can generally regenerate on their own if the injuries are small. Larger injuries, however, must be surgically treated, typically with nerve grafts harvested from elsewhere in the body (Schmidt and Leach, 2003). Mahoney and Anseth reported on their experience with neural tissue regeneration, where neural cells within degradable hydrogels were photoencapsulated and then monitored by confocal microscopy to study the key cellular functions over time (Mahoney and Anseth, 2006). Holecko et al. presented some immunohistochemical strategies for assessing the interactions at the immediate interface between microscale implanted devices and the surrounding brain tissue during inflammatory astrogliotic reactions (Holecko et al., 2005). In addition, a novel implantable device was reported by Phillips et al. that delivers a tethered aligned collagen guidance conduit containing Schwann cells into a peripheral nerve injury site (Phillips et al., 2005). Brain tissue fragments in amniotic fluid of rats with NTD (neural tube defects) were examined by Mendonca et al. (2005) and were found to reflect the evolution from exencephaly to anencephaly, and thus could support the aspiration hypothesis of how brain tissue nodules are found in the lungs of subject with NTD. Hopefully these findings can help lead to novel therapies for those with neurodegenerative disorders. Heart In cardiac research, current experimental efforts have focused on cellular cardiomyoplasty, myocardial tissue engineering, and myocardial regeneration as alternative approaches to whole organ transplantation (Krupnick et al., 2004). Such strategies may offer novel forms of therapy to patients with end-stage heart failure in the near future. With regards to fetal heart tissue, several preliminary studies have been reported. Embryonic cardiomyocytes were shown to have the ability to remodel the abdominal aorta into a spontaneous pulsating apparatus and to function as a vascular pump (Okamura et al., 2002). In addition, Fukuda et al. has suggested that cell transplantation therapy for the patients with heart failure might possibly be achieved using the regenerated cardiomyocytes from autologous bone marrow cells (Fukuda, 2003). Gardiner’s group has suggested that the treatment of fetal arrhythmias may be rationalized by the use of fetal electrocardiography and magnetocardiography,

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and that further information may be derived by defining the natural history of complete heart block and mechanisms of tachyarrhythmia (Gardiner, 2005). Another research area of interest in cardiac research is valvular interstitial cells. Most valvular interstitial cells in normal valves are quiescent with a fibroblast-like phenotype. However, valvular interstitial cells in developing, diseased, adapting, and engineered valves are adjusted to a dynamic environment through the activation of these cells and secretion of proteolytic enzymes. They appear to be mediating extracellular matrix remodeling (“developing/remodeling/activated” phenotype), which is then followed by a normalization of phenotype (Rabkin-Aikawa et al., 2004). Lung For lung tissue engineering, lung cell replacement therapeutics is being studied for treating respiratory-specific diseases such as cystic fibrosis and idiopathic pulmonary fibrosis. Directing the differentiation of mouse embryonic stem cells into the respiratory cell lineages has been tried extensively in the recent past (Denham et al., 2006). One of the novel findings is that all of the lung cell-seeded scaffolds undergo cell-mediated contraction that appeared to coincide with the findings by immunohistochemistry of smooth muscle actin expression in some cells. These results seem to demonstrate the capability of dissociated lung cells to form lung histotypic structures in collagen–GAG (glycosaminoglycan) tissue-engineering scaffolds in vitro (Chen et al., 2005). In other fetal lung research, vitamin D3-upregulated protein 1 (VDUP1) was suggested to be an important mediator of expansion-induced lung cell proliferation and alveolar epithelial cell (AEC) differentiation in the developing lung (Filby et al., 2006). Their studies revealed reduced fetal lung expansion and increased VDUP1 mRNA levels in experimental fetal subjects after 7 days. VDUP1 was found to be localized to airway epithelium in small bronchioles, AECs, and some mesenchymal cells. Skin Skin tissue engineering was one of the early organ systems to which regenerative medicine techniques were applied. Autologous skin grafting is the gold standard for treatment of deep second and third degree burns. However, the majority of current researches in skin tissue engineering focuses on the synthesis of complex threedimensional (3D) polymer scaffolds containing functional biomolecules to which cells are introduced. Hohlfeld and associates developed fetal skin constructs to improve healing of such severe burns (Hohlfeld et al., 2005). Their simple techniques provided complete treatment without auto grafting, showing that fetal skin cells might have great potential to treat burns and eventually acute and chronic wounds of other types. Sun et al. showed that co-culture with fibroblasts enables keratinocytes and endothelial cells to proliferate without serum, and that keratinocytes and endothelial cells appear to self-organize according to the native epidermal–dermal structure given the symmetry-breaking field of an air–liquid interface (Sun et al., 2005). In order to gain insight into the biology of fetal skin during culture, Vuadens and colleagues have investigated the cellular proteins during four culture passages (P00, P01, P04 as well as P10) using high-resolution two-dimensional (2D) gel electrophoresis and mass spectrometry (MS) (Vuadens et al., 2003). Bioinformatic analyses were focused on a region of each gel corresponding to pI between 4 and 8 and M(r) from 8,000 to 35,000. In this area, 373 42 spots were detected (N  18). Twenty-six spots presented an integrated intensity that increased as the passage number increased, whereas five spots showed a progressively lower intensity in later passages. MS analysis was performed on those spots that were unambiguously identified on preparative 2D gels. These observations were interpreted as reflecting either an oxidative stress related to cell culture, or, alternatively, maturation, differentiation, and the aging of the cells. Kaviani et al. consistently isolated subpopulations of fetal mesenchymal cells from human amniotic fluid and rapidly expanded them in vitro (Kaviani et al., 2003). These human mesenchymal amniocytes attach firmly to both polyglycolic acid polymer and acellular human dermis, and thus it was hypothesized that amniotic fluid may be a valuable and practical cell source for fetal tissue engineering.

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Muscle Recently, Vickers and associates have demonstrated that several types of musculoskeletal connective tissue cells, including chondrocytes, fibrochondrocytes, ligament fibroblasts, osteoblasts, and MSCs can express the gene for the contractile actin isoform, alpha-smooth muscle actin (α-SMA) and can contract analogs of extracellular matrix in vitro (Vickers et al., 2004). These findings appear to indicate that control of α-SMA-enabled contraction may be important when employing synovial cells for cartilage repair procedures and warrant further investigation into the physiological role of α-SMA expression in synovial cells. In another muscle-based study, Danielsson et al. implanted smooth muscle cell–scaffold constructs in the dorsal subcutaneous space of athymic mice (Danielsson et al., 2006). Their in vivo studies identified the presence of fluorescent-labeled transplanted smooth muscle cells until day 3. Thereafter angiogenesis was induced and infiltration of mouse fibroblasts and polymorphonuclear cells were observed. The polymorphonuclear cells were noted to have completely disappeared after 3 weeks. Human UCB has been regarded as a possible alternative source for cell transplantation and cell therapy for muscle tissue engineering because of its hematopoietic and non-hematopoietic (mesenchymal) potential. Gang and associates demonstrated that UCB-derived MSCs possessed a potential of skeletal myogenic differentiation and that these cells could be a suitable source for skeletal muscle repair and muscle tissue engineering (Gang et al., 2004). In addition, skeletal muscle has been well characterized as a reservoir of myogenic precursors or satellite cells with the potential to participate in cellular repopulation therapies for muscle replacement purposes. However, recent evidence suggests that the post-natal muscle compartment can be considered an alternative to bone marrow as a source of multipotent stem cells. These cells have also been called muscle-derived stem cells (MDSCs). These MDSCs, when primed with appropriate environmental cues, can differentiate into a variety of non-muscle cells as well. Sinanan et al. showed that the CD56 subpopulation within adult human skeletal muscle is heterogeneous and is composed of both lineage-committed myogenic cells as well as multipotent stem cells (the candidate MDSCs) (Sinanan et al., 2004). These multipotent stem cells were able to form nonmuscle tissue such as fat and bone. Bone Human fetal bone cells have been envisioned for use in bone tissue engineering and in the regeneration of adult skeletal tissue (Montjovent et al., 2004). To construct bioengineered bone structures, vascularized bone grafts theoretically have many advantages over non-vascularized free grafts. However, the availability of these grafts can be extremely limited. Kim and colleagues sought to determine whether new vascularized bone could be engineered by the transplantation of osteoblasts around existing vascular pedicles using biodegradable, synthetic polymer as a cell delivery vehicle (Kim and Kim, 2005). They found that the degree of osteoid and bone formation progressed over time as blood vessels invaded the growing tissue. This tissue appeared to ultimately undergo morphogenesis to become organized trabeculated bone with a vascular pedicle. To date, there has been no description of human primary fetal bone cells successfully treated with differentiation factors. The characterization of these fetal bone cells is particularly important as the pattern of secreted proteins from osteoblasts has been shown to change as the bone tissue ages (Montjovent et al., 2004). Regardless of the source of chondrocytes, all fetal cartilage constructs appear to resemble hyaline cartilage, both grossly and histologically, in vitro. Fuchs et al. showed that in vivo, engineered implants retained their hyaline characteristics for up to 10 weeks after implantation, but then remodeled into fibrocartilage by 12 weeks post-operatively (Fuchs et al., 2003). Mononuclear inflammatory infiltrates surrounding residual polymer fibers were noted in all of the specimens but they were most prominently in the acellular controls. As a result, Fuchs et al. concluded that fetal bone tissue engineering may have some utility in the treatment of severe congenital chest wall defects at birth.

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Hematopoietic Cells Human UCB has been regarded as an alternative source for cell transplantation and cell therapy based on its hematopoietic and non-hematopoietic (mesenchymal) potentials. A number of trials have been undertaken to isolate MSCs from UCB because MSCs from bone marrow have been regarded as suitable material for cell/gene therapeutics and because they have been noted to be multipotent. Recently, Lee and colleagues have shown that cryopreserved human UCB fractions can be used as an alternative source of MSCs for experimental and clinical application (Lee et al., 2005). In another study, Leveen and associates have generated a mouse model using the Cre/lox system that exhibits an inducible deficiency of transforming growth factor beta receptor II (TbetaRII) (Leveen et al., 2005). With this approach, transforming growth factor beta (TGF-beta) signaling deficiency was restricted to the hematopoietic system by bone marrow transplantation. No increase of thymocyte apoptosis was observed, but TbetaRII-deficient CD8+ thymocytes displayed a 2-fold increase in proliferation rate, as determined by bromodeoxyuridine (BrdU) incorporation in vivo. They suggested that TGF-beta functions as an immune regulator critical for T-cell function (Leveen et al., 2005). As mentioned in the previous section, MSCs were isolated from human UCB and were induced to differentiate into skeletal muscle cells for muscle tissue engineering purposes (Gang et al., 2004). During cell culture expansion, the UCB-derived mononuclear cells gave rise to adherent layers of fibroblast-like cells expressing MSC-related antigens such as SH2, SH3, alpha-smooth muscle actin, CD13, CD29, and CD49e. When these UCB-derived MSCs were incubated in promyogenic conditions for up to 6 weeks, they began to express myogenic markers by both flow cytometry and reverse transcriptase-polymerase reaction analyses, where two early myogenic markers, MyoD and myogenin, were expressed after 3 days of incubation but not after 2 weeks. Pancreatic Islet Cells Fetal pancreatic tissue has been suggested as a possible cell source for islet replacement therapy in type 1 diabetes mellitus. While this tissue usually consists of a small amount of β-cells, a raft of immature and/or progenitor cells may have the potential to proliferate and differentiate into functional insulin-producing cells. Suen and associates showed that freshly isolated fetal islet-like cell clusters were poorly responsive to glucose challenge as compared with adult islets (Suen et al., 2005). They showed that both the expansion and differentiation of fetal islet-like cell clusters could be enhanced with the exposure of appropriate growth factors and microenvironments. Their data indicated that in vivo exendin-4 treatment may have enhanced the growth and differentiation of fetal mice islet-like cell clusters, and thus promoted the functional maturation of the graft after transplantation. Recently, Zhang et al. showed that monoclonal side population (SP) progenitors were isolated from human fetal pancreas, which may be a novel method of purifying pancreatic progenitor cells from human tissues (Zhang et al., 2005). For insulin gene expression and islet cell survival, integrin receptors are known to play a major role, as they are involved in tissue morphogenesis and homeostasis by regulating cell interactions with extracellular matrix proteins (Wang et al., 2005). Wang and colleagues examined the expression pattern of integrin subunits in human fetal pancreas specimens (8–20 weeks fetal age) and investigated the relevance of beta1 integrin function for insulin gene expression and islet cell survival. The alpha3, alpha5, and alpha6 beta1 integrins were expressed in ductal cells at 8 weeks, before glucagon- and insulin-immunoreactive cells budded off, and their levels gradually increased in both the ductal cells and the islet clusters for up to 20 weeks. This provided important evidence of a molecular basis for cell–matrix interactions during islet development and suggested that beta1 integrin plays a vital role in regulating islet cell adhesion, gene expression, and survival. Zhang and colleagues isolated nestin-positive cells isolated from human fetal pancreas and discovered that these cells possess the characteristics of pancreatic progenitor cells since they have highly proliferative potential and the capability of differentiation into insulin-producing cells in vitro (Zhang et al., 2005). Interestingly, the nestin-positive pancreatic progenitor cells shared many of the same phenotypic markers as bone-marrow-derived MSCs.

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Kidney Several interesting studies on embryonic kidney tissue have been published. Dakovic-Bjelakovic’s group studied nephrons in the kidney’s cortex during gestation (Dakovic-Bjelakovic et al., 2005), and found that nephron polymorphism was the main characteristic of the human fetal kidney during development. In younger fetuses, there was a wide nephrogenic zone just below the renal capsule, which contained the condensed mesenchyme and terminal ends of the ureteric bud. Nephrons were noted to be in different stages of development and were located around the ureteric bud, but they branched in the cortical nephrogenic zone and induced nephrogenesis. COX-2 is known as an important regulator of fetal renal growth and function, and its inhibition may lead to congenital oligonephropathy. Hartleroad and associates investigated whether maternal administration of a selective COX-2 inhibitor would adversely affect fetal renal growth. They found that fetal kidney size was unaffected and concluded that maternal administration of therapeutic doses of celecoxib did not adversely affect fetal renal growth after analyzing for vascular endothelial growth factor (VEGF) and its soluble receptors, matrix metalloproteinase (MMP)-2 and -9, tissue inhibitor of metalloproteinase (TIMP)-1 and -2, COX-2, and total cellular protein levels (Hartleroad et al., 2005). The orchestration of kidney organogenesis is complex and requires the interaction of many morphoregulatory molecules that lead to coordinated organ development. For example, chemokines can induce cell motility during embryogenesis by activating specific receptors. Lu and colleagues examined CXCR-1-4 and SDF-1 mRNA in various fetal tissues including kidney (Lu et al., 2005). They discovered that the expression of CXCR-3 in kidney, liver, and brain was dependent upon gestational age, and that CXCR-1-4 protein was expressed in non-hematopoietic cells in the brain, heart, intestine, and kidney. They concluded that CXCR-1-4 and SDF-1 genes are widely expressed in the normal human fetus and that these gene products could influence kidney fetal development. Bladder The bladder serves as a reservoir for the storage of urine and it maintains a low intraluminal pressure as it fills under normal conditions. Bladder reconstruction has been attempted with both natural materials and synthetic polymers. For bladder regeneration, Ram-Liebig et al. investigated the optimum conditions for the proliferation of urothelial cells, in order to obtain confluent coverage of large surfaces of biocompatible membranes, and for their terminal differentiation (Ram-Liebig et al., 2004). They concluded that the mitogenic effects of the extracellular matrix content of biological membranes and fibroblastic inductive factors were synergistic with each other, and may be able to compensate for a low fetal calf serum concentration and the absence of other additives. They found that lowering the fetal calf serum concentration to 1% in the culture medium inhibited the proliferation of urothelial cells and permitted their terminal differentiation. Several congenital and acquired diseases of the bladder may need, due to lack or destruction of functional tissue, mechanically stable biomaterials as cell carriers for the engineering of these tissues. Collagen scaffolds have some advantageous characteristics for tissue engineering purposes because of their capacity to induce tissue regeneration and their biocompatibility. Recently, Danielsson and colleagues evaluated cell growth by WST-1 proliferation assay and showed improved growth of bladder cells when modified collagen scaffolds were used (Danielsson et al., 2006). The cell penetration assessed by histology showed similar results on both modified and native scaffolds. In vivo studies in athymic mice showed the presence of the fluorescent-labeled transplanted smooth muscle cells in the cell–scaffold constructs until day 3. Thereafter angiogenesis was noted and infiltration of mouse fibroblasts and polymorphonuclear cells were observed. Nyirady and colleagues characterized the developmental changes to the normal bladder by examining the in vitro contractile properties of the fetal sheep detrusor smooth muscle bladder at different gestational ages (Nyirady et al., 2005). They found that fetal development between 65 and 140 days in the sheep was associated with increased contractile activation, which correlated with an increase of muscle development in the earlier stages (65–110

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days). In later stages (110–140 days), muscle development appeared to be complete but functional innervation of the tissue was still noted.

CONCLUSIONS The use of fetal tissue for regenerative medicine purposes has been investigated for essentially every organ systems, and some applications, especially in the area of hematopoietic cells, have been in use for several years. There are currently unresolved ethical and moral issues regarding the use of some fetal tissues. However, fetal tissues in conjunction with regenerative medicine techniques may offer novel methods to treat many diseases that currently have suboptimal available treatments.

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in ‘t Anker, P.S., Noort, W.A., et al. (2003). Mesenchymal stem cells in human second-trimester bone marrow, liver, lung, and spleen exhibit a similar immunophenotype but a heterogeneous multilineage differentiation potential. Haematologica 88(8): 845–852. Itskovitz-Eldor, J., Schuldiner, M., et al. (2000). Differentiation of human embryonic stem cells into embryoid bodies compromising the three embryonic germ layers. Mol. Med. 6(2): 88–95. Kaufman, D.S., Hanson, E.T., et al. (2001). Hematopoietic colony-forming cells derived from human embryonic stem cells. Proc. Natl Acad. Sci. USA 98(19): 10716–10721. Kaviani, A., Guleserian, K., et al. (2003). Fetal tissue engineering from amniotic fluid. J. Am. Coll. Surg. 196(4): 592–597. Kehat, I., Kenyagin-Karsenti, D., et al. (2001). Human embryonic stem cells can differentiate into myocytes with structural and functional properties of cardiomyocytes (comment). J. Clin. Investig. 108(3): 407–414. Kennedy, D. (2003). Stem cells: still here, still waiting (comment). Science 300(5621): 865. Kim, W.S. and Kim, H.K. (2005). Tissue engineered vascularized bone formation using in vivo implanted osteoblast–polyglycolic acid scaffold. J. Korean Med. Sci. 20(3): 479–482. Krupnick, A.S., Kreisel, D., et al. (2004). Myocardial tissue engineering and regeneration as a therapeutic alternative to transplantation. Curr. Top. Microbiol. Immunol. 280: 139–164. Lee, M.W., Yang, M.S., et al. (2005). Isolation of mesenchymal stem cells from cryopreserved human umbilical cord blood. Int. J. Hematol. 81(2): 126–130. Leveen, P., Carlsen, M., et al. (2005). TGF-beta type II receptor-deficient thymocytes develop normally but demonstrate increased CD8+ proliferation in vivo. Blood 106(13): 4234–4240. Levenberg, S., Golub, J.S., et al. (2002). Endothelial cells derived from human embryonic stem cells. Proc. Natl Acad. Sci. USA 99(7): 4391–4396. Lindvall, O. and Bjorklund, A. (2004). Cell therapy in Parkinson’s disease. NeuroRx 1(4): 382–393. Lu, W., Gersting, J.A., et al. (2005). Developmental expression of chemokine receptor genes in the human fetus. Early Hum. Dev. 81(6): 489–496. Magnus, D. and Cho, M.K. (2005). Ethics. Issues in oocyte donation for stem cell research. Science 308(5729): 1747–1748. Mahoney, M.J. and Anseth, K.S. (2006). Three-dimensional growth and function of neural tissue in degradable polyethylene glycol hydrogels. Biomaterials 27(10): 2265–2274. Mendonca, E.D., Gutierrez, C.M., et al. (2005). Brain tissue fragments in the amniotic fluid of rats with neural tube defect. Pathology 37(2): 152–156. Michejda, M. (2004). Which stem cells should be used for transplantation? Fetal Diagn. Ther. 19(1): 2–8. Montjovent, M.O., Burri, N., et al. (2004). Fetal bone cells for tissue engineering. Bone 35(6): 1323–1333. Nyirady, P., Thiruchelvam, N., et al. (2005). Contractile properties of the developing fetal sheep bladder. Neurourol. Urodyn. 24(3): 276–281. Oberpenning, F., Meng, J., et al. (1999). De novo reconstitution of a functional mammalian urinary bladder by tissue engineering (see comment). Nat. Biotechnol. 17(2): 149–155. O’Donoghue, K., Chan, J., et al. (2004). Microchimerism in female bone marrow and bone decades after fetal mesenchymal stem-cell trafficking in pregnancy (see comment). Lancet 364(9429): 179–182. Okamura, S., Suzuki, A., et al. (2002). Formation of the biopulsatile vascular pump by cardiomyocyte transplants circumvallating the abdominal aorta. Tissue Eng. 8(2): 201–211. Phillips, J.B., Bunting, S.C., et al. (2005). Neural tissue engineering: a self-organizing collagen guidance conduit. Tissue Eng. 11(9–10): 1611–1617. Rabkin-Aikawa, E., Farber, M., et al. (2004). Dynamic and reversible changes of interstitial cell phenotype during remodeling of cardiac valves. J. Heart Valve Dis. 13(5): 841–847. Ram-Liebig, G., Meye, A., et al. (2004). Induction of proliferation and differentiation of cultured urothelial cells on acellular biomaterials. BJU Int. 94(6): 922–927. Reubinoff, B.E., Pera, M.F., et al. (2000). Embryonic stem cell lines from human blastocysts: somatic differentiation in vitro (comment). (Erratum appears in Nat. Biotechnol. 2000 May; 18(5): 559). Nat. Biotechnol. 18(4): 399–404. Reubinoff, B.E., Itsykson, P., et al. (2001). Neural progenitors from human embryonic stem cells (comment). Nat. Biotechnol. 19(12): 1134–1140.

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Richards, M., Fong, C.Y., et al. (2002). Human feeders support prolonged undifferentiated growth of human inner cell masses and embryonic stem cells (comment). Nat. Biotechnol. 20(9): 933–936. Rollini, P., Kaiser, S., et al. (2004). Long-term expansion of transplantable human fetal liver hematopoietic stem cells. Blood 103(3): 1166–1170. Schmidt, C.E. and Leach, J.B. (2003). Neural tissue engineering: strategies for repair and regeneration. Annu. Rev. Biomed. Eng. 5: 293–347. Schuldiner, M., Yanuka, O., et al. (2000). Effects of eight growth factors on the differentiation of cells derived from human embryonic stem cells. Proc. Natl Acad. Sci. USA 97(21): 11307–11312. Schuldiner, M., Eiges, R., et al. (2001). Induced neuronal differentiation of human embryonic stem cells. Brain Res. 913(2): 201–205. Sinanan, A.C., Hunt, N.P., et al. (2004). Human adult craniofacial muscle-derived cells: neural-cell adhesion-molecule (NCAM; CD56)-expressing cells appear to contain multipotential stem cells. Biotechnol. Appl. Biochem. 40(Pt 1): 25–34. Suen, P.M., Li, K., et al. (2005). In vivo treatment with glucagon-like peptide 1 promotes the graft function of fetal isletlike cell clusters in transplanted mice. Int. J. Biochem. Cell Biol. 38: 951–960. Sun, T., Mai, S., et al. (2005). Self-organization of skin cells in three-dimensional electrospun polystyrene scaffolds. Tissue Eng. 11(7–8): 1023–1033. Thomson, J.A., Itskovitz-Eldor, J., et al. (1998). Embryonic stem cell lines derived from human blastocysts (comment). (Erratum appears in Science 1998 December 4; 282(5395):1827). Science 282(5391): 1145–1147. Vickers, S.M., Johnson, L.L., et al. (2004). Expression of alpha-smooth muscle actin by and contraction of cells derived from synovium. Tissue Eng. 10(7–8): 1214–1223. Vuadens, F., Crettaz, D., et al. (2003). Plasticity of protein expression during culture of fetal skin cells. Electrophoresis 24(7–8): 1281–1291. Wang, R., Li, J., et al. (2005). Role for beta1 integrin and its associated alpha3, alpha5, and alpha6 subunits in development of the human fetal pancreas. Diabetes 54(7): 2080–2089. Watt, S.M. and Contreras, M. (2005). Stem cell medicine: umbilical cord blood and its stem cell potential. Semin. Fetal Neonatal Med. 10(3): 209–220. Zhang, L., Hong, T.P., et al. (2005a). Nestin-positive progenitor cells isolated from human fetal pancreas have phenotypic markers identical to mesenchymal stem cells. World J. Gastroenterol. 11(19): 2906–2911. Zhang, L., Hu, J., et al. (2005b). Monoclonal side population progenitors isolated from human fetal pancreas. Biochem. Biophys. Res. Comm. 333(2): 603–608. Zhang, S.C., Wernig, M., et al. (2001). In vitro differentiation of transplantable neural precursors from human embryonic stem cells (comment). Nat. Biotechnol. 19(12): 1129–1133.

57 Engineering of Large Diameter Vessels Saami K. Yazdani and George J. Christ INTRODUCTION Vascular disease affects millions of people worldwide, occurs at all levels of the vascular tree, and represents a major cause of morbidity and mortality. When the extent of vascular disease is severe and requires vessel bypass or replacement, the available supply of healthy native collateral vessels is frequently inadequate. The only currently available clinical alternative is the use of synthetic vascular grafts. While synthetic grafts have been reasonably successful for larger diameter vessels (i.e. 6 mm), they have faired poorly in smaller caliber vessels. Tissue-engineered blood vessels (TEBV) have been forwarded as a viable clinical alternative for both indications, with a majority of the clinical success limited to large caliber vessels, while the preclinical work to date has focused primarily on smaller caliber vessels. However, the virtually epidemic increase in end-stage renal disease (ESRD) has highlighted the deficiencies of synthetic grafts, even when used for relatively large caliber vessels. In this scenario, a huge demand for improved dialysis vascular access is anticipated (Figure 57.1). Certainly this clinical indication requires larger caliber TEBV (6 mm). The goal of this report, therefore, is to briefly review the status of TEBV research, and moreover, to describe the challenges and opportunities associated with creating large caliber TEBV, such as those that might be used for improved dialysis vascular access. In so doing, we will pay special attention to the importance of the vascular smooth muscle cell (SMC) to TEBV. To date, relatively little attention has been paid to the importance of the medial SMC layer to both vessel function and accelerated vessel maturation (both in vitro and in vivo). Both of these beneficial properties of smooth muscle have important implications for the further development and clinical translation of vascular tissue engineering. As such, the creation of TEBV for dialysis vascular access provides an extraordinary opportunity to further examine the role of the SMC in TEBV. To this end, we will address how the presence of the SMC can help meet the physiological characteristics/demands of the bioengineered vessels that would be required for such clinical success, and finally, outline one currently envisioned strategy for achieving this end. PREVALENCE AND IMPACT OF VASCULAR DISEASE Vascular diseases are the second leading cause of morbidity and mortality in the United States. (www.american heart.org). Abnormal vascular function contributes to coronary artery disease, stroke, peripheral arterial disease, renal insufficiency, and diabetic neuropathy. In 2003 alone, nearly 500,000 coronary artery bypass graft surgeries were performed and over 100,000 lower extremity bypass procedures are performed (www.americanheart.org, Birkmeyer et al., 2002). Important risk factors for vascular disease include older age, hypertension, hyperlipidemia, smoking, diabetes, and chronic renal insufficiency (Collins et al., 2003a). Population trends are unfavorable with respect to vascular disease, as the US population is ageing, diabetes is reaching epidemic proportions, and chronic renal disease, especially ESRD, is now epidemic

978

From dialysis machine

Looped graft Artery

To dialysis machine

Vein

Figure 57.1 Schematic illustration of dialysis vascular access. The common two scenarios of dialysis vascular access are demonstrated here (adopted from the National Kidney and Urological Disease Information Clearing house, http://kidney.niddk.nih.gov/).

(McClellan, 1994; Gilbertson et al., 2005). With respect to ESRD, there is a significant unmet medical need for autologous dialysis vascular access graft. Such grafts are clearly of relatively large diameter when mature (6 mm) and thus, represent an important target for the relatively large diameter TEBVs that are the subject of this report.

THE NEED FOR IMPROVED DIALYSIS VASCULAR ACCESS In the United States, 297,928 individuals received chronic dialysis therapy in December 2002 (Rafii and Lyden, 2003). This number is projected to increase to 712,290 by 2015, more than a doubling of the dialysis population in just over one decade (Gilbertson et al., 2005). Presently only a mature native radial artery to cephalic vein fistula achieves the ideal access route of blood circulation for hemodialysis. A close alternative is another site of native arteriovenous fistula (AVF) within the upper extremity, for example, an upper arm brachial artery to cephalic or basilic vein fistulas. Regardless, only 33% of hemodialysis patients in the United States achieve dialysis via a native AVF while the majority requires a prosthetic polytetrafluoroethylene (PTFE) artery to vein bypass grafts (AVBG, 41%) or chronic indwelling central venous catheters (McClellan, 1994; National Kidney Foundation, 2000, 2001; Hsu et al., 2004). A detailed discussion of the limitations of PTFE is well beyond the scope of this report. Suffice it to say, that stenosis is the most common problem, and moreover, the presence of the PTFE creates a foreign body response (Kohler and Kirkman, 1999; Huber et al., 2003, 2004). In addition, endothelialization occurs only within the first 1–2 cm at anastomoses, and furthermore, prosthetic materials are prone to infection. In fact, chronic cannulation with needles inserted through the skin and left in place for hours during dialysis predisposes to frequent graft infection (National Kidney Foundation, 2002; Basaran et al., 2003; Huber et al., 2004; Neville et al., 2004). Failure of the lumen surface to heal in PTFE grafts may also predispose to hematogenous seeding. Finally, as PTFE does not regenerate, the graft wall deteriorates over time from

979

980 THERAPEUTIC APPLICATIONS: TISSUE THERAPY

Table 57.1 Properties of the IDEAL vascular graft for hemodialysis access Anti-thrombogenic Anti-inflammatory Resistant to injury and intimal proliferation (i.e. stenosis) Resists degradation of the scaffold Maintains structural integrity and adverse remodeling under a wide range of pressure conditions Pharmacological/physiological and mechanical similarities to native vessels Durability Rapid replacement (i.e. increased maturation rate) Easy accessibility Durability and structural integrity in the face of repeated punction over a prolonged period of time (years) Suitable geometry (diameter and length) to achieve high-volume blood flow

chronic puncture, predisposing to pseudoaneurysm formation, skin breakdown, cannulation site bleeding, and graft infection. For all of the aforementioned reasons, creating an autologous blood vessel of the appropriate geometry for AVBG directly addresses many of the limitations of the PTFE grafts currently used for dialysis access. Certainly, a cellularized vessel wall with luminal endothelial coverage is likely to be more resistant to thrombosis and infection. Furthermore, the cellularized wall of a mature bioengineered vessel should allow healing at puncture sites to prevent vessel wall deterioration and provide resistance to infection superior to PTFE. Moreover, the engineered blood vessel is likely to have a compliance profile better matched to the outflow vein than PTFE, which in turn should reduce the extent of outflow venous stenosis (Kohler and Kirkman, 1999). All of these properties are prerequisites for the next generation of dialysis vascular access, and are summarized in Table 57.1.

VASCULAR PHYSIOLOGY RELEVANT TO TEBVs Blood is carried from the heart to the capillaries by the arteries, and then returned via the venous circulation. The magnitude of the cardiovascular problems described above has certainly served to focus most TEBV research on the arterial side of the vascular tree, which will also remain the subject of this report. In that regard, the arterial vascular tree can be subdivided into three general types of arteries based both on their location in the vascular tree, as well as the functions they serve. As blood is moved away from the heart, it moves from large elastic arteries that have a strictly conduit function (e.g. aorta) to more medium-sized muscular arteries that have a distributive function, and eventually to small muscular arteries and arterioles, which provide the majority of the resistive function. The lumen to wall ratio decreases as one moves down the vascular tree away from the heart, and similarly, so does the ratio of the elastic component versus the smooth muscle component (Boulpaep, 2003). Regardless of the considerable differences in function, the vessel wall in all three types of arteries possess three distinct layers (tunics) which are the intima, media, and adventitia (Figure 57.2). The innermost layer encountered traversing the vessel wall from the luminal side is the tunica intima, which is in direct contact with moving blood. The intima is covered by the endothelium, which in turn, resides on a thick basement membrane referred to as the internal elastic lamina. The endothelium provides the anti-thrombogenic surface that ensures continuous laminar blood flow. The middle layer in the vessel wall is the tunica media. The media is composed of SMCs embedded in a matrix of collagen, elastin, and proteoglycans, the ratio and composition of which varies along the vascular tree (see below). The media resides between the internal elastic lamina and the tunica externa (i.e. adventitia). The adventitia represents the outermost portion of the vessel wall, and is primarily comprised of loose connective tissue, fibroblasts, and small nerve fibers. Of note, nerve fibers rarely penetrate the adventitial–medial SMC border.

Tu nic aa dv en titi a

Tu n

ica

m

Tu

ni

ed

ia

ca

in

tim

a

Engineering of Large Diameter Vessels 981

Arteries Elastic

8–12 mm Internal diameter 1–2 mm Wall thickness Elastic fibers  Smooth muscle cells  Collagen fibers 

Medium

Small

2–5 mm 0.5–1 mm   

0.1–1 mm 0.1–0.25 mm   

Figure 57.2 Structure and composition of the arterial wall. Representative H&E staining illustrating the major components of the vessel wall (Boulpaep, 2003).

The physiological characteristics of each vessel depend on their location in the vascular tree. Of note, there is no native vessel that mimics the physiological characteristics of the proposed dialysis vascular access graft (i.e. AVF). While arteriovenous anastomoses are quite common in the circulation (e.g. for rapid shunting of blood in the skin for heat exchange), categorizing the behavior of the AVF as proposed herein is somewhat unique. In fact, arteriovenous anastomoses naturally occur between small muscular arteries and venules to bypass the capillary network and provide rapid shunting of blood. The proposed bioengineered AVF described herein would be a much larger vessel (6 mm), and therefore, has some unique characteristics. Thus, the ideal AVF must possess some hybrid characteristics, for example, the compliance of large elastic arteries and perhaps the tone of large- to medium-sized muscular arteries. The main goal of these bioengineered vessels is to maintain a non-thrombogenic and non-proliferative surface, while retaining the ability to adapt and remodel to external stimuli, and moreover, be able to heal in response to repetitive puncture wounds (i.e. 3/week). Clearly, to incorporate all of these features will require the presence of both SMCs and endothelial cells (ECs). A brief review of the phenotypic and functional characteristics of these two vascular wall cell types most directly pertinent to TEBV is provided below. ECs There are many excellent reviews on ECs and the reader is referred to a few of these for more details (Cines et al., 1998; Michiels, 2003; Aird, 2006). ECs line the entire vascular tree and provide a functional barrier between blood and the vascular wall cells and tissue parenchyma. Perhaps more importantly, they serve as a

982 THERAPEUTIC APPLICATIONS: TISSUE THERAPY

biologically active lining of the blood vessels and play a critical role in the control of vascular tone. Regulation of vascular tone is accomplished via a variety of endothelium-derived vasoactive substances. Some important endothelium-derived vasorelaxants include nitric oxide (NO), prostacyclin (PGI2), and endothelium-derived hyperpolarizing factor (EDHF). The endothelium also provides an important source of constrictor substances such as endothelin-1, superoxide anions/radicals, angiotensin II, thromboxane A2, and endoperoxides. These are synthesized and released in response to a wide variety of environmental and mechanical stimuli. In addition to regulation of vascular tone, the endothelium is also responsible for the maintenance of vessel wall permeability (i.e. regulating the flow of nutrients, biological molecules), as well as the balance between coagulation and fibrinolysis, the composition of the subendothelial matrix, the adhesion and extravasation of leukocytes, and mediation of inflammatory processes in the vascular wall. Prevention of thrombotic events is accomplished by maintaining a healthy monolayer of ECs that retain the ability to secret anti-thrombotic agents such as NO, PGI2, tissue plasminogen activator (tPA), and thrombomodulin. All of these EC functions are controlled via membrane bound proteins, junctional proteins, and a variety of cell surface receptors, and are critical to circulatory homeostasis, and thus, normal organ function. Smooth Muscle Cells Vascular myocytes are interposed between the variable autonomic innervation on one side of the vessel (adventitial or abluminal side), and the endothelium on the other. This anatomical arrangement has important mechanistic implications for coordinated vessel function, as the size of the medial SMC layer varies from a single cell in the terminal arteriole to numerous relatively concentric layers of muscle such as those that encircle the large elastic and muscular arteries. Nonetheless, the role of myocytes in most vessels is similar, that is, to maintain vessel tone at some partial level of contractility, with the ability to become further constricted, or relaxed, as the physiological necessities of the vessels dictate. More importantly, contraction and relaxation of individual myocytes in the vessel wall must be coordinated both across the width of the muscle layer (i.e. perpendicular axis to the vessel wall), as well as along the length (i.e. longitudinal axis) of the blood vessel. The exact mechanism(s) that endow the vascular myocyte with the ability to accomplish this task differs throughout the vascular tree, and the details of such are well beyond the scope of this report. Those mechanisms pertinent to the conduit-type bioengineered vessels that are the subject of this report are described briefly below. It is hard to overestimate the importance of the vascular SMC to circulatory homeostasis and function. In this regard, vascular SMCs make at least two major contributions to TEBV function: (1) contractility/tone and (2) accelerated tissue maturation/formation. Both of these properties are illustrated in Figure 57.3 and are discussed in more detail below. The “tone” or contractility provided by the presence of SMCs in the vessel wall ensures that the TEBV will not be passively dilated in the presence of increased systemic pressure. In fact, a direct contribution of SMC tone to vascular diameter and/or compliance has been demonstrated in vitro (Figure 57.3) and in vivo in both human vessels and animal models at all levels of the vascular tree (Barra et al., 1993; Bank et al., 1995; Kuecherer et al., 2000; Safar et al., 2000; Moosmang et al., 2003; Jarajapu and Knot, 2005). Examples include modulation of pulse pressure and compliance in large elastic conduit vessels such as the aorta, as well as autoregulation of blood flow in specialized circulations (i.e. cerebral arterioles). Control of medial SMC tone is modulated by intravascular pressure and filling (myogenic response in muscular arteries and arterioles), circulating neurotransmitters and hormones (neurogenic response), as well as factors released from surrounding tissues (metabolic response). There are a variety of neurotransmitters known to regulate vasoconstriction (e.g. neuropeptide Y (NPY), norepinephrine (NE), and ATP (i.e. purinergic signaling)) as well vasorelaxation (e.g. vasoactive intestinal polypeptide (VIP), calcitonin gene related peptide (CGRP), and NO

(a)

(b)

H&E

Movat

(c)

(d)

(e) 175

Passive diameter Diameter with tone

Diameter ( M)

150

125

100

75

WKY

50 0

50

100

150

200

250

Pressure (mmHg)

Figure 57.3 Vascular SMC infiltration and function. (a) The EC and SMC cell seeded engineered grafts 2 weeks after implantation in sheep contained uniform cellularity throughout the vascular walls. (b) Abundant elastin fibers were observed in the entire arterial wall with a prominent distribution in the serosa and luminal surface. (c) In the EC-only graft 15 days post-implantation the vessel showed a poorly organized thrombotic deposit (arrow). (d) In the EC-only seeded graft 130 days post-implantation, the vessel architecture looked relatively normal. The vessel lumen in the lower panels is in the center (direction arrows point). (e) In the presence of SMC contraction, the heightened contractile response of the muscle cell resist the passive dilation due to the increase of intramural pressure, resulting in constant diameter within the 50–150 mmHg range. However, calcium depletion ablates SMC contraction and leads to passive vessel dilation over the same pressure range. This phenomenon clearly documents the importance of vascular muscle tone to vascular function. (The authors would like to thank Dr. Yagna P.R. Jarajapu for providing Figure 57.3e).

984 THERAPEUTIC APPLICATIONS: TISSUE THERAPY

(Christ and Barr, 2000; Christ and Wingard, 2005; del Valle-Rodriguez et al., 2006)). Furthermore, as noted above, ECs also release both relaxing and contracting factors (see section above). As described in detail elsewhere, all of these processes can be further integrated via intercellular communication through gap junctions (Christ et al., 1996, 1999; Brink et al., 2000; Wang et al., 2001; Lagaud et al., 2002a, b; Haefliger et al., 2004; Haddock and Hill, 2005). In fact, gap junction (Cx40, Cx43, and Cx45) appear to play a role in the control of vascular tone in a variety of ways. First, they can help coordinate locally restricted signals arising across the vessel wall (i.e. integrating neural and endothelial signals that originate on opposite sides of the vessel). Second, they can help orchestrate responses along the length of the blood vessel (up- and down-stream vasodilation or constriction). Third, they can provide a safety factor ensuring syncytial SMC responses, even when not all cells in the vessel wall can respond to any given stimulus (i.e. cellular heterogeneity). However, in addition to tone/compliance as discussed above, the presence of SMCs in TEBV also appears to accelerate vascular tissue maturation/formation. This point is also illustrated in Figure 57.3, where the anatomy/histology of the blood vessel appears almost “normal” only 2 weeks after implantation; as opposed to the relatively immature looking vessel observed at the same time point in an endothelial-only seeded implant. In this scenario, the presence of the SMC layer may confer a third advantage of special significance to the TEBV for dialysis vascular access. That is, the repeated puncturing of the vessel wall (i.e. typically 3/week) may require the presence of the additional cell type for tissue/wound healing. Finally, it would seem that the presence of the SMC and the commensurate cell-to-cell interactions with the endothelium would be required to confer the full range of phenotype(s) and function(s) characteristic of the native vessel wall.

TISSUE-ENGINEERED VASCULAR GRAFTS: A BRIEF REVIEW OF THE LITERATURE From the aforementioned discussion it is clear that tissue engineering of vascular conduits directly addresses a critical need for improved treatment options for vascular disease. There have been a significant number of review publications on the topic of TEBV (Ratcliffe, 2000; Tiwari et al., 2001; Rabkin and Schoen, 2002; Teebken and Haverich, 2002; Sales et al., 2005; Vara et al., 2005; Isenberg et al., 2006) and many of the primary studies are summarized in Table 57.2. Provided below are brief summaries of some of the seminal research findings. Small diameter TEBV (4 mm): Clinical studies have indicated that EC-seeded synthetic grafts have high patency rates in human coronary artery bypass grafts and in lower extremity artery bypass grafts (Deutsch et al., 1999; Laube et al., 2000). In the last two decades many attempts have been made to engineer endotheliallined, patent 4–6 mm arterial substitutes. Weinberg and Bell (1986) were the first to engineer blood vessel substitutes by seeding ECs, SMCs, and fibroblasts on preformed collagen gels. However, mechanical and burst strengths were poor, precluding in vivo implantation. A similar approach was taken by L’Heureux et al. (1998), who used SMCs, fibroblasts, and EC to engineer a polymer-free blood vessel that had better mechanical properties and performed reasonably well in vivo (three out of six implanted grafts remained patent after 7 days). In addition, Niklason et al. (1999) described seeding SMCs and ECs on biodegradable polymers made of polyglycolic acid (PGA) and the implanted grafts remained patent up to 24 days. Most recently, L’Heureux et al. (2006) have implanted autologous TEBV extracted from fibroblasts for up to 8 months in rats, canine, and primate models. These grafts demonstrated tissue integration, suitable mechanical properties, and formation of vasa vasorum. Badylak et al. (1989) introduced the concept of a native collagen-rich matrix (small intestinal submucosa) as a vascular graft, and Huynh et al. (1999) showed that these grafts, in the absence of cells, were fully endothelialized within 3 months and impregnated with SMC, improving long-term patency. Kaushal et al. (2001) showed similar results by maintaining vascular graft patency for greater than 4 months by seeding decellularized porcine arterial segments with ECs from circulating progenitor cells. In fact, the explanted grafts exhibited contractile activity and NO-mediated vascular relaxation similar to the native

Engineering of Large Diameter Vessels 985

carotid artery (Kaushal et al., 2001). These early studies demonstrated the capabilities of collagen matrices to mature and remodel via cell infiltration of SMC in the vessel wall and EC coverage of the lumen, leading to development of vasomotor tone and responsiveness. Large diameter TEBV (6 mm): There is significantly less information available concerning the investigation and development of large diameter TEBV. However, as noted above, with increased longevity worldwide, and the rapidly expanding number of patients with diabetes and renal disease that will require dialysis, a need to create a functional, patent, autologous large diameter graft that can remodel and regenerate is clearly emerging. Shin’oka et al. (2005) have developed a tissue-engineered graft from a PGA/PLLA or poly (L-lactide) scaffold seeded with autologous bone marrow cells on the luminal surface to treat pediatric patients with congenital heart defects. The performance of these grafts was first evaluated in animal models (Watanabe et al., 2001). The results of the animal studies revealed that seeded TEBV remain patent for up to 6 months with no sign of stenosis or dilation. Moreover, when retrieved, the endothelium of the vessel stained positive for functional endothelial-specific surface marker (Factor VIII). In a groundbreaking clinical study, the peripheral pulmonary arteries of 23 patients were replaced with large diameter autologous seeded biodegradable scaffolds (PGA/PLLA  autologous bone marrow cells). Long-term follow-up of these seminal clinical studies (32 months) have shown no complications such as thrombosis, stenosis, or obstruction associated with the implants. Importantly, these results demonstrate the potential of TEBV to remodel, grow, and remain patent in a growing patient. Opitz et al. (2004a) investigated the development of a tissue-engineered graft for aortic replacement. The challenges of a bioengineered aorta clearly present a significant departure from the TEBV investigations that have been conducted elsewhere in the vascular system. The scaffold for these studies was constructed from poly-4-hydroxybutyrate (P-4-HB, Tepha Inc., Cambridge, MA) and endothelialized and impregnated with SMC within a bioreactor system. Dynamic preconditioning of the scaffold for 2 weeks resulted in a TEBV with a rupture force of approximately 80% of the native ovine aorta, the target replacement arterial segment. In vivo experiments of the TEBV demonstrated that the implanted grafts remained patent up to 3 months, followed by significant dilation and thrombus formation of the graft likely due to insufficient elastic fiber synthesis. The development of large diameter vessels within our group has blossomed from the knowledge gained from past experiences in developing small diameter TEBVs (Kaushal et al., 2001; Amiel et al., 2006; Stitzel et al., 2006). Despite the obvious differences in the clinical application, the approach in developing both large and small TEBV share many common features. One strategy for so doing is outlined below.

TISSUE-ENGINEERED VASCULAR GRAFTS: A BRIEF REVIEW OF THE PROCESS An approach to the construction of relatively large diameter tissue-engineered vessels is illustrated in Figure 57.4, and reflects the general approach taken by several groups for the development of both large and small diameter TEBV. This over-simplified conceptual framework does not depict the numerous complexities associated with this process. In fact, each step in the TEBV process, from selection of the scaffold, to cell source (i.e. isolation of progenitor cells, etc.), cell seeding conditions and bioreactor TEBV preconditioning protocols, to selection of the appropriate animal model for implantation needs to be thoroughly evaluated. Each of these steps has a critical impact on the TEBV process that will likely vary with each TEBV for each indication. Certainly, with respect to the best “recipe” for TEBV, the devil is in the details. Below we provide some basic concepts, features, and requirements for each step in the process. Scaffolds Various synthetic and naturally derived biomaterials have been used in constructing vascular grafts but none have proven entirely satisfactory. The goal is always the same, that is, to develop a reproducible, biocompatible

986 THERAPEUTIC APPLICATIONS: TISSUE THERAPY

Table 57.2 Summary of TEBV studies Authors

Year

Graft type

Diameter (mm)

Cell(s)

Methods Static

Dynamic

Weinberg et al.

1986

Collagen

6

Bovine aortic SMC, EC, fibroblasts

1 week rotation @ 1RPM

n/a

Badylak et al.

1989

Small intestine submucosa

10

None

n/a

n/a

L’Heureux et al.

1998

VSMC and fibroblast sheet

3

SMC (human umbilical cord)

3 months of total maturation (SMC, fibroblast, EC)

n/a

Campbell et al.

1999

Myofibroblast tube

3,5

Mesothellum, myofibroblast (autologous)

n/a

n/a

Huynh et al.

1999

Intestinal collagen layer (small intestine submucosa)

4

None

n/a

n/a

Shum-Tim et al.

1999

PGA–PHA

7

Ovine carotid SMC, EC, fibroblasts

1 week of incubation (mixed cell population)

n/a

Niklason et al.

1999

PGA

3.1

EC, SMC (bovine aorta, porcine carotid artery (in vivo))

30 min of static seeding

8 weeks of pulsatile conditions

Teebken et al.

2000

Acellular porcine artery

n/a

EC, myofibroblasts (both human saphenous vein)

60 min of incubation

4 day pulsatile condition

Hoerstrup et al.

2001

PGA/P-4-HB

5

EC, myofibroblast (ovine carotid artery)

4 days of static seeding

Up to 28 days of pulsatile conditions

Niklason et al.

2001

PGA

3.1

EC, SMC (bovine aorta)

Rotation for 30 min

Up to 8 weeks of pulsatile conditions

Kaushal et al.

2001

Acellular porcine artery

4

EPC (ovine peripheral blood)

Rotation for 6 h

Steady flow

Teebken et al.

2001

Acellular porcine artery

4

EC (porcine external jugular vein)

60 min of incubation

n/a

Watanabe et al.

2001

PGA-CL/LA

10

Canine femoral vein SMC, fibroblasts

1 week of incubation (mixed cell population)

n/a

Mckee et al.

2002

PGA

3.1

EC (HUVEC), SMC (Human aortic)

16 h of static seeding

7 weeks of pulsatile condition

Berglund et al.

2003

Hybrid collagen

3

EC (human coronary EC)

60 min of incubation

n/a

Nasseri et al.

2003

PGA/P-4-HB

5,12

Myofibroblast (ovine carotid artery)

Rotation @ 5 RPM 5–10 days

n/a

Yu et al.

2003

PTFE

4

EC, SMC (both Rabbit Jugular vein)

Rotation at 1 RPM for 2 h

n/a

Shirota et al.

2003

Polyurethane

1.5

EPC (human peripheral

Rotation @ 120 degrees each

n/a

Engineering of Large Diameter Vessels 987

In vivo model

Outcome

In vitro Outcome

n/a

n/a

EC and SMC were seeded with success

Canine (infrarenal aorta)

100% patency for up to 52 week (n  9)

n/a

Canine (femoral artery)

Three out of six grafts remained patent at 7 days

EC and SMC were organized successfully to mimic the structure of native artery

Rat (aorta), rabbit (carotid)

Rats: 67% patency at 4 months (n  30), Rabbit: 70% patency at 4 months (n  20)

n/a

Rabbit (carotid artery)

100% patency at 28 days (n  9), 53 days (n  4), and 90 days (n  4)

n/a

Ovine (aortic replacement)

100% patency at 10 days (n  1), 84 days (n  3), 150 days (n  3)

n/a

Porcine (saphenous artery)

100% patency at 4 weeks for preconditioned graft (n  1), nonpreconditioned vessels occluded at 3 weeks (n  2)

Endothelium layer was achieved, SMC impregnation of the scaffold was achieved

n/a

n/a

EC were seeded with success

n/a

n/a

Endothelium layer was achieved

n/a

n/a

EC and SMC were seeded with success

Sheep (carotid artery)

100% patency at 15 days and 130 days after implantation (n  7)

Endothelium layer was achieved

Sheep (carotid artery)

54% patency at 1 week (n  8) and 71% patency at 4 months (n  8) for seeded graft

n/a

Canine (Inferior vena cava)

100% patency at 3 (n  1), 4 (n  1), 5 (n  1), 6 months (n  1)

n/a

n/a

n/a

EC and SMC were seeded with success

n/a

n/a

EC were seeded with success dynamic rotation seeding can culture myofibroblasts onto tubular polymer scaffold

n/a

n/a

Rabbit (aorta shunt)

Retention rate of EC at 1 h is 65% and 1 day (51%), EC/SMC at 1 h (98%), and 1 day (90%)

EC were seeded with success

n/a

n/a

EPCs were seeded with success

(Continued)

988 THERAPEUTIC APPLICATIONS: TISSUE THERAPY

Table 57.2 (Continued) Authors

Year

Graft type

Diameter (mm)

Cell(s)

Methods Static

blood)

hour followed with 4 days of static seeding

Dynamic

Naito et al.

2003

PLLA/PGA

20

EC, SMC (peripheral vein)

10 days of seeding

n/a

Williams et al.

2004

PGA

4.5

EC, SMC (bovine thoracic aorta)

24 h syringe pump cell seeding

4–16 days of pulsatile conditions

Baguneid et al.

2004

Polyester

n/a

EC, SMC (porcine aorta)

1 h of slow rotation

Pulsatile conditions

McFetridge et al.

2004

Acellular porcine artery

5–12

EC, SMC (human umbilical vein)

Rotation for 2 h

Steady flow

Opitz et al.

2004

P-4-HB

15

EC, SMC (ovine carotid artery)

4 days of rotation

14 days of pulsatile conditions

Opitz et al.

2004

P-4-HB

4

EC, SMC (ovine carotid artery)

Rollar mixer

Pulsatile conditions

Hibino et al.

2004

PLLA/PGA

8

EC, SMC (femoral vein), BMC (Iliac bone)

1 week for vein cells, 1 h for BMC

n/a

Shin’oka et al.

2005

PLLA/PGA

12–24

BMC (anterior superior Iliac spine)

2–4 h

n/a

Poh et al.

2005

PGA

3

EC, SMC (human saphenous vein)

5 days of static seeding

Up to 7 weeks of pulsatile condition

Jeong et al.

2005

PLCL

4

SMC (rabbit aorta)

2 days of static seeding

8 weeks of pulsatile conditions

Laflamme et al.

2005

VSMC sheet

3

EC, SMC (human umbilical vein)

1 week of maturation

n/a

Williams et al.

2005

PGA

4.5

EC, SMC (bovine aorta)

Syringe pump of for 24 h

24 days of pulsatile conditions

Borschel et al.

2005

Acellular rat femoral artery

1

EC (rat heart)

Over night Incubation

n/a

Xu et al.

2005

Acellular carotid

n/a

SMC (canine saphenous vein)

24 h of static seeding after

Dual syringe pump over night

artery

dynamic

Yang et al.

2005

Poly (diol citrate)

3

EC, SMC (human aortic)

2 day of static seeding up to 8 weeks

n/a

L’Heureux et al.

2006

Fibroblast sheet

4.2

EC (saphenous vein)

3 h of static seeding

3 day pulsatile (from 3 to 150 ml/min)

Laflamme et al.

2006

VSMC sheet

3

EC, SMC (human umbilical vein)

3 weeks of maturation

n/a

Hoerstrup et al.

2006

PGA

18

EC, myoflbroblast (ovine carotid artery and jugular vein)

7 days of static seeding

2 weeks of pulsatile (from 50 to 550 ml/min)

Leyh et al.

2006

Acellular ovine pulmonary artery

n/a

EC (ovine carotid artery)

(4 h (static)  12 h (0.1 RPM)) 3

n/a

Engineering of Large Diameter Vessels 989

In vivo model

Outcome

In vitro Outcome

Human (pulmonary artery)

100% patency at 4 months (n  1)

n/a

n/a

n/a

Endothelium layer and SMC impregnation of the scaffold was achieved

n/a

n/a

Endothelium layer and SMC impregnation of the scaffold was achieved

n/a

n/a

Endothelium layer was achieved, SMC impregnation had limited success

n/a

n/a

Endothelium layer was achieved, SMC impregnation of the scaffold was achieved

Sheep (descending aorta)

100% patency at 1, 3, 6, 12 weeks (n  4), Thrombus formation and dilation at 24 weeks but still patent

n/a

Canine (inferior vena cava)

100% patency at 4 weeks (n  8)

n/a

Human (pulmonary artery)

100% patency at 1–32 months (n  23)

n/a

n/a

n/a

Endothelium layer and SMC impregnation of the scaffold was achieved

n/a

n/a

SMC Impregnation of the scaffold was achieved

n/a

n/a

Contraction could be induced via endothelin

n/a

n/a

EC and SMC were seeded with success

Rat (femoral artery)

89% patency at 4 weeks (n  9)

EC were seeded with success

n/a

n/a

Mechanical strength increases with preconditioning

n/a

n/a

EC and SMC were seeded with success

Rats (abdominal aorta), primate

86% patency at 90–225 days (n  12, rats), (for primate) 100 patency at 6 weeks (n  1) and 8 weeks (n  2)

Cellular TEBV was achieved

n/a

n/a

Similar contraction in the TEBV could be induced via endothelin as compared to native artery

Ovine (pulmonary artery)

100% patency at all time points which included 20 weeks (n  3), 50 weeks (n  2), 80 weeks (n  3), and 100 weeks (n  4)

EC and SMC were seeded with success

Ovine (pulmonary artery)

100% patency at 6 months (n  5), increase in diameter was observed

n/a

990 THERAPEUTIC APPLICATIONS: TISSUE THERAPY

1. Autologous cell harvest from circulating blood.

2. Progenitor cell isolation and expansion.

3. Decellularization and static seeding of vascular scaffolds.

4. Bioreactor preconditioning.

5. Surgical implant: sheep model, jugular-carotid A-V fistual 6. Surgical implant: dialysis vascular access graft.

Figure 57.4 Schematic depiction of the TEBV process.

scaffold similar to that characteristic of native vasculature. With respect to the synthetic constructs, polymers and electrospun scaffolds are both very attractive options due to the control one has over composition, architecture, and the reproducibility of the manufacturing process. The current generations of polymers are mostly biodegradable and include polylactic acid (PLA), PGA, polyhydroxyalkanoate (PHA), and polydioxanone (PDS). These polymers can be used singly, or in combination to optimize the desired mechanical performance and biocompatibility of the graft. Similar to polymers, electrospinning techniques can take advantage of a variety of materials to create scaffolds. Electrospinning involves creation of an electromagnetic field with a high-voltage source. Exposure to high voltage causes polymers in volatile solvents to elongate and splay into small fibers and be drawn/sprayed onto a grounded surface (i.e. a mandrel) where they can be spun into tubular structures. By controlling the characteristics of individual fiber formation during the electrospinning process, as well as the rotational speed of the mandrel (see Stitzel et al., 2006) structural characteristics such as porosity and geometry can be precisely controlled. Thus, from a commercial perspective, synthetic scaffolds are very attractive for the clinical translation of TEBV. However, from a biological perspective, decellularized vessels (i.e. natural scaffolds), possess a biochemical composition, ultrastructural architecture, and biomechanics similar to native vessels. Not surprisingly, decellularized collagen-based vascular scaffolds derived from porcine blood vessels have been successfully used for TEBV in vivo (Kaushal et al., 2001). Similar approaches have been used in a variety of clinical applications for developing tissue-engineered vascular patches (Cho et al., 2005), heart valves (Lichtenberg et al.,

Engineering of Large Diameter Vessels 991

(a)

(c)

(b)

(d)

Percent Collagen type I

60.20%

Collagen type II

5.30%

Collagen type III

14.80%

Elastin

19.70%

Figure 57.5 Natural scaffolds derived from porcine arterial segments. (a) H&E of native porcine carotid artery. (b) H&E of decellularized porcine carotid artery. (c) Segment of native porcine carotid artery. (d) Segment of retrieved TEBV following in vivo implantation. The collagen and elastin composition of the decellularized porcine carotid artery are provided in the table below.

2006), and bladders (Gabouev et al., 2003). To summarize, while synthetic scaffolds will undoubtedly provide an important source of “off the shelf ” scaffold material for clinical TEBV, the natural scaffold still provide the ultimate “gold” standard with respect to the biological requirements and characteristics of native vessels required to guide the development of the TEBV in vivo. The TEBV strategy outlined below utilizes the decellularized scaffold. Step 1: Removal of cells from mature arteries produces a collagen-based scaffold that is amenable for seeding and growth of vascular cells. Prior work has established a working protocol for preparation of scaffolds from animal arteries using a multi-step decellularization process. Details of the procedure can be found in previous literature that shows the overall concept (Kaushal et al., 2001; Amiel et al., 2006). As shown in Figure 57.5, decellularized scaffolds preserve their extracellular matrix architecture, including internal and external elastin layers and several layers of collagen. Moreover, the decellularization process removed all cellular components, maintaining only collagen and elastin components. The quantity and distribution of collagen and elastin in a vascular scaffold is vital information in consideration for scaffold material in developing TEBV. Mechanical characteristics of vascular grafts play a significant influence in long-term patency of the implant. In fact, compliance mismatch is thought to be one of the most important factors predisposing prosthetic vascular

992 THERAPEUTIC APPLICATIONS: TISSUE THERAPY

(b)

120 100 60 40 Native Decellularized

20 0 1

1.05

1.1

1.15

1.2

2000 1500 1000

0 (d)

3.5 Decellularized Native

3.0 2.5

Stress (MPa)

Stress (MPa)

Native Decellularized

2500

500

Diameter strain [mm/mm] (c)

3500 3000

80

Pressure (mmHg)

Pressure (mmHg)

(a)

2.0 1.5 1.0 0.5 0.0 0

50

100

Strain (mm/mm)100

150

3.5 3.0 2.5 2.0 1.5 1.0 0.5 0.0

Decellularized Native

0

20

60 40 Strain (mm/mm)  100

80

Figure 57.6 Mechanical behavior of native and decellularized grafts. (a) Pressure versus diameter measurements of the decellularized vessel compared to native. (b) Burst pressure measurements of native and decellularized vessels. (c) Stress versus strain measurements in the axial direction. (d) Stress versus strain measurements in the circumferential direction.

grafts to intimal hyperplasia, thrombosis, and occlusion. If the TEBV is stiff then flow disturbances and tissue vibration may predispose to hyperplasia. Conversely, a TEBV that is too compliant may result in the formation of an aneurysm. As such, we have rigorously analyzed the biomechanical characteristics of the decellularized scaffolds. To measure compliance, decellularized vascular scaffolds were immersed in a water bath, cannulated at one end, and pressurized, while recording the diameter change using a digital camera. Figure 57.6 summarized the data and demonstrates that the decellularized scaffolds are similar to that of the native artery. Moreover, burst strength testing and stress–strain measurements, demonstrate that the decellularization process does not disturb the mechanical integrity to the extent that failure might occur in vivo (Figure 57.6). Cell Source Step 2: There are numerous potential cell sources available for cellularizing the synthetic or naturally derived scaffolds. The strategy that we are currently pursuing is to isolate progenitor cells from circulating blood and expand them to obtain the EC and SMC that are required for TEBV, as outlined in Figure 57.7. The overall concept is to utilize cell-selective markers to isolate and expand the progenitor cells prior to differentiation and further proliferation for seeding purposes. This process is well characterized with respect to differentiation of ECs from endothelial progenitor cells, but further research is required for obtaining similar procedures for derivation of SMCs from circulating muscle progenitor cells. The latter work is ongoing in our group. Cell Seeding and Preconditioning Steps 3 and 4: The final steps in creating TEBV involve the development of a bioreactor system for cell seeding and preconditioning; that is to expose TEBV to in vivo conditions they will face upon implantation. Seeding TEBV consists of depositing cells (EC and/or SMC) onto a three-dimensional scaffold to achieve a confluent

Engineering of Large Diameter Vessels 993

Tissue engineered blood vessel cell source Smooth Muscle Cells

Endothelial Cells

Primary Veins Arteries

Progenitors Blood Bone Marrow

Primary Veins Arteries

EC MSC

EPC

CD133

MPC

CD133

VE-cadherin

CD31

Desmin

VE-cadherin

Vimentin

Desmin

SMC CD34

Figure 57.7 Identification of progenitor-derived EC and SMCs. As illustrated, mesenchymal cells (MS) are collected from sheep blood and separated into endothelial progenitor (EP) and muscle progenitor (MP) cell populations. The cells are then subcultured into differentiated SMC and EC types.

monolayer of EC at the inner surface and/or SMC on the outside. A variety of approaches have been attempted in seeding both the endothelium and SMCs, and recent published studies have demonstrated highly evolved bioreactor systems to produce and monitor the mechanical forces required for cell seeding and/or preconditioning (Thompson et al., 2002; Barron et al., 2003; Mironov et al., 2003; McCulloch et al., 2004; Narita et al., 2004; Williams and Wick, 2004; Portner et al., 2005; Soletti et al., 2006). The theory behind the use of bioreactors for TEBV derives from studies demonstrating that mechanical stress accelerated cell and tissue growth and phenotypic differentiation (Braddon et al., 2002; Nerem, 2003; Jeong et al., 2005; Kurpinski et al., 2006). In this regard, a properly designed bioreactor system provides physiologically relevant stress in a three-dimensional tissue, accelerating tissue maturation, and development functional properties. While we are unaware of any published studies documenting that bioreactor preconditioning per se is capable of producing a relatively mature and fully functional vessel in vitro, this certainly seems an area worthy of further investigation. It corresponds to intuition that implantation of a more mature functional TEBV would accelerate tissue formation and maturation in vivo; thereby providing for quicker restoration of function, and presumably, promoting more widespread clinical applications. Regardless of the precise operational concept, a bioreactor system for development of TEBV should be capable of the following functions:

• • • • • • • •

Permitting static and/or dynamic seeding. Providing and monitoring physiological flow rate and pressure. Capable of dynamic data display and recording (archival). Providing physiological axial and circumferential stress. Providing an external bath. Maintaining desired concentration of gases and nutrients in the culture medium. Maintaining temperature and sterility. Be easily portable and accessible for transportation and use in surgical procedure.

994 THERAPEUTIC APPLICATIONS: TISSUE THERAPY

(a)

Shear stress (dynes/cm2)

(b)

Pressure transducer

25

Pulsatile flow

15 10 5 0

Pump

Steady flow

20

0

24

48

72 96 Time (h)

120

1.5

2

3.5

Flow meter Bioreactor

Bypass External media bath TEBV

Shear stress (dynes/cm2)

(c)

144

168

25 20 15 10 5 0

0

0.5

1

2.5

3

4

4.5

5

Time (s)

Figure 57.8 Bioreactor system. (a) Image of the bioreactor flow system. The bioreactor provides an external media bath, optical access, a bypass system, control over flow and pressure conditions, and the ability to maintain sterility. (b) Summary of the 7 day preconditioning protocol of the TEBV. (c) Pulsatile shear conditions during the final 48 h of preconditioning.

(a)

(b)

(c)

(d)

Figure 57.9 Cell seeding of decellularized scaffolds. (a) H&E staining of the decellularized vessel after static EC seeding. (b) H&E staining illustrating the presence of a confluent monolayer of EC within the lumen of the decellularized vessel after 7 days in the bioreactor. (c) Static seeding of vascular SMCs after 48 h. (d) One week bioreactor preconditioned decellularized scaffold seeded with vascular SMCs.

Obviously, the optimal preconditioning protocol(s) required to seed and mature TEBV are still being developed. However, Figure 57.8 shows the general features of a bioreactor system, while Figure 57.9 shows some preliminary results with both EC and SMC seeding on decellularized scaffolds. We are currently investigating the impact of various bioreactor protocols on the efficiency of cell seeding and the phenotypic differentiation

Engineering of Large Diameter Vessels 995

of ECs and SMCs. Major parameters of interest include rotational speed of scaffold during seeding, optimal cell seeding density and time course of cell seeding protocol, and duration of bioreactor preconditioning period (i.e. days or weeks). Clearly further development and refinement of the bioreactor system is required, but unequivocally, such development holds intrinsic scientific value, and moreover, will likely be required to ensure the widespread clinical application of TEBV.

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Opitz, F., Schenke-Layland, K., Richter, W., Martin, D.P., Degenkolbe, I., Wahlers, T. and Stock, U.A. (2004a). Tissue engineering of ovine aortic blood vessel substitutes using applied shear stress and enzymatically derived vascular smooth muscle cells. Ann. Biomed. Eng. 32: 212–222. Opitz, F., Schenke-Layland, K., Cohnert, T.U., Halbhuber, K.J., Martin, D.P. and Stock, U.A. (2004b). Tissue engineering of aortic tissue: dire consequence of suboptimal elastic fiber synthesis in vivo. Cardiovasc. Res. 63: 719–730. Poh, M., Boyer, M., Solan, A., Dahl, S.L., Pedrotty, D., Banik, S.S., McKee, J.A., Klinger, R.Y., Counter, C.M. and Niklason, L.E. (2005). Blood vessels engineered from human cells. Lancet 364: 2122–2124. Portner, R., Nagel-Heyer, S., Goepfert, C., Adamietz, P. and Meenen, N.M. (2005). Bioreactor design for tissue engineering. J. Biosci. Bioeng. 100: 235–245. Rabkin, E. and Schoen, F.J. (2002). Cardiovascular tissue engineering. Cardiovasc. Pathol. 11: 305–317. Rafii, S. and Lyden, D. (2003). Therapeutic stem and progenitor cell transplantation for organ vascularization and regeneration. Nat. Med. 9: 702–712. Ratcliffe, A. (2000). Tissue engineering of vascular grafts. Matrix Biol. 19: 353–357. Safar, M.E., Blacher, J., Mourad J.J. and London, G.M. (2000). Stiffness of carotid artery wall material and blood pressure in humans, application to antihypertensive therapy and stroke prevention. Stroke 31: 782–790. Sales, K.M., Salacinski, H.J., Alobaid, N., Mikhail, M., Balakrishnan, V. and Seifalian, A.M. (2005). Advancing vascular tissue engineering: the role of stem cell technology. Trends Biotechnol. 23: 461–467. Shin’oka, T., Matsumura, G., Hibino, N., Naito, Y., Watanabe, M., Konuma, T., Sakamoto, T., Nagatsu, M. and Kurosawa, H. (2005). Midterm clinical result of tissue-engineered vascular autografts seeded with autologous bone marrow cells. J. Thorac. Cardiovasc. Surg. 129: 1330–1338. Shirota, T., He, H., Yasui, H. and Matsuda, T. (2003). Human endothelial progenitor cell-seeded hybrid graft: proliferative and antithrombogenic potentials in vitro and fabrication processing. Tissue Eng. 9: 127–136. Shum-Tim, D., Stock, U., Hrkach, J., Shin’oka, T., Lien, J., Moses, J., Stamp, A., Taylor, G., Moran, A.M., Landis, W., Langer, R., Vacanti, J.P. and Mayer Jr., J.E. (1999). Tissue engineering of autologous aorta using a new biodegradable polymer. Ann. Thorac. Surg. 68: 2298–2305. Soletti, L., Niepnice, A., Guan, J., Stankus, J.J., Wanger, W.R. and Vorp, D.A. (2006). A seeding device for tissue engineered tubular structures. Biomaterials 27: 4863–4870. Stitzel, J., Liu, J., Lee, S.J., Komua, M., Berry, J., Soker, S., Lim, G., Van Dyke, M., Czerw, R., Yoo, J.J. and Atala, A. (2006). Controlled fabrication of a biological vascular substitute. Biomaterials 27: 1008–1094. Teebken, O.E. and Haverich, A. (2002). Tissue engineering of small diameter vascular grafts. Eur. J. Endovasc. Surg. 23: 475–485. Teebken, O.E., Bader, A., Steinhoff, G. and Haverich, A. (2000). Tissue engineering of vascular grafts: human cell seeding of decellularized porcine matrix. Eur. J. Vasc. Endovasc. Surg. 19: 381–386. Teebken, O.E., Pichlmaier, A.M. and Haverich, A. (2001). Cell seeded decellularized allogeneic matrix grafts and biodegradable polydioxanone-prostheses compared with arterial autografts in a porcine model. Eur. J. Vasc. Endovasc. Surg. 22: 139–145. Thompson, C.A., Colon-Hernandez, P., Pomerantseva, I., MacNeil, B.D., Nasseri, B., Vacanti, J.P. and Oesterle, S.N. (2002). A novel pulsatile, laminar flow bioreactor for the development of tissue-engineered vascular structures. Tissue Eng. 8: 1083–1088. Tiwari, A., Salacinski, H.J., Hamilton, G. and Seifalian, A.M. (2001). Tissue engineering of vascular bypass grafts: role of endothelial cell extraction. Eur. J. Vasc. Endovasc. Surg. 21: 193–201. Vara, D.S., Salacinski, H.J., Kanna, R.Y., Bordenave, L., Hamilton, G. and Seifalian, A.M. (2005). Cardiovascular tissue engineering: state of the art. Pathol. Biol. 53: 599–612. Wang, H.Z., Day, N., Valcic, M., Hsieh, K., Serels, S., Brink, P.R. and Christ, G.J. (2001). Intracellular communication in cultured human vascular smooth muscle cells. Am. J. Physiol. Cell Physiol. 281: C75–C88. Watanabe, M., Shin’oka, T., Tohyama, S., Hibino, N., Konuma, T., Matsumura, G., Kosaka, Y., Ishida, T., Imai, Y., Yamakawa, M., Ikada, Y. and Morita, S. (2001). Tissue engineered vascular autograft: inferior vena cava replacement in a dog model. Tissue Eng. 7: 429–439. Weinberg, C.B. and Bell, E. (1986). A blood vessel model constructed from collagen and cultured vascular cells. Science 231: 397–400. Williams, C. and Wick, T.M. (2004). Perfusion bioreactor for small diameter tissue-engineered arteries. Tissue Eng. 10: 930–941.

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58 Engineering of Small Diameter Vessels Chrysanthi Williams and Robert T. Tranquillo

INTRODUCTION Since 1918, cardiovascular disease has been the number 1 killer in the United States, claiming over 1.4 million persons in 2002 alone, with coronary heart disease being the single largest killer of American males and females (American Heart Association, 2003). Coronary arteries supply blood to the myocardium of the heart, and coronary heart disease occurs due to cholesterol–lipid–calcium deposits on the inner walls that narrow the vessel lumen and prevent adequate blood supply. Atherosclerosis, a form of arteriosclerosis (hardening of the arteries), is a multifactorial disease and is influenced by diet, cigarette smoking, diabetes, high blood pressure, and exercise (Burke et al., 1997). Several theories have been formulated to explain the localized nature of atherosclerosis. Fluid mechanical theories predict that atherogenesis occurs in areas that have a relatively complex geometry, a fairly large Reynolds number, and a lower than average wall shear stress throughout the pulsatile cycle (Ku et al., 1985; Giddens et al., 1990; Lieber and Giddens, 1990). Flow in these areas is complex, unsteady, and sometimes turbulent. Solid mechanical views blame sites of high stress, such as bifurcations, and constricted or dilated areas. A blood vessel that is under internal pressure and longitudinal stretch experiences stress concentration under the following conditions: increased radius of curvature, saddle shape, areas in the neighborhood of a small side branch, and bending of the wall (Fung, 1996). The disease begins with the focal eccentric accumulation of lipid in the intima with intracellular lipid visible mainly in macrophages and smooth muscle cells (SMCs) with time. This leads to the formation of a fatty streak, which is composed of SMCs, matrix fibers, and lipids. At a later stage, the fatty streak becomes the preatheroma, which contains multiple extracellular lipid pools, as well as collagen and elastin fibers accumulated beneath the endothelium. The subendothelial zone may subsequently become more organized to form the fibrous cap, which resembles the normal media layer in structure and thickness, and does not contain any macrophages or lipids. As the lipid pools coalesce into lipid cores, the intima becomes disorganized, and this lesion type is termed an atheroma. As the disease develops, the lesion becomes stratified due to the increasing amount of fibrous tissue in deep and superficial layers, and the localization of lipid cores between the fibrous regions, forming fibroatheromas (Glagov et al., 1995). However, as the lesion enlarges, the artery also enlarges by an outward bulging of the wall beneath the growing plaque to compensate for the narrowing that has occurred. Lumen stenosis becomes evident when the plaque takes up approximately 40% or more of the lumen area (Bassiouny et al., 1997). Intimal thickening or hyperplasia could also be a response to vascular intimal injury. When the vascular wall is injured, SMCs proliferate in and migrate from the media to the intima and synthesize extracellular matrix (ECM) proteins. SMCs undergo dedifferentiation, lose their ability to contract, gain the capacity to divide, and increase ECM synthesis (Assoian and Marcantonio, 1996). SMCs residing in the intima lose their

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thick myosin-containing filaments and greatly increase the amount of organelles involved in protein synthesis, such as rough endoplasmic reticulum and Golgi apparatus. The migratory and proliferative activity of SMCs is regulated by both growth promoters, such as platelet-derived growth factor, basic fibroblast growth factor, and interleukin 1, and inhibitors, such as heparan sulfate, nitric oxide, and transforming growth factor-β. Intimal SMCs may return to a non-proliferative state when either the overlying endothelial layer is reestablished or the abnormal chronic endothelial stimulation ceases. Atheromas in advanced disease almost always undergo patchy or massive calcification, and atherosclerotic lesions cause clinical disease by one of the following mechanisms: slow narrowing of the intima that results in ischemia of the tissues perfused by the involved vessels; sudden occlusion of the lumen by superimposed thrombosis or hemorrhage into an atheroma; thrombosis followed by embolism; weakening of the wall of a vessel, causing an aneurysm or rupture (Schoen, 1994). Several approaches are taken to treat atherosclerotic cardiovascular disease of small caliber arteries (6 mm), and the most common ones are briefly described next. Balloon angioplasty or percutaneous transluminal coronary angioplasty is a procedure used to dilate narrowed arteries. A catheter is inserted with a deflated balloon at its tip into the narrowed vessel, the balloon is inflated, compressing the plaque and enlarging the inner diameter of the artery, and then the balloon is deflated and the catheter removed. About 70–90% of these procedures also involve the placement of a stent, which is a wire mesh tube that is initially collapsed to a small diameter, placed over a balloon catheter and moved into the area of the blockage. When the balloon is inflated, the stent expands, locks in place, and holds the artery open. Concerns with stents include injury to the vessel wall during insertion, and acute thrombosis or intimal hyperplasia as consequences of injury (Didisheim and Watson, 1996). Drug-eluting stents that slowly release a drug around the stent to prevent restenosis have been more successful than bare metal stents, but their long-term advantage and mortality in multivessel coronary artery disease are still being assessed (Guyton, 2006; Kivela and Hartikainen, 2006). Coronary artery bypass graft operation, first performed in 1964, is an invasive procedure, which is done to reroute, or “bypass,” blood around occluded arteries and improve the supply of blood and oxygen to the heart. Grafts commonly used include the great saphenous vein from the leg, internal mammary artery from the chest, radial artery from the arms, and sometimes arteries from the stomach. Although 70–82% of the saphenous vein substitutes remain patent in 5 years (Lytle et al., 1985) and 61% after 10 years (Goldman et al., 2004), stenosis due to intimal hyperplasia, which is a flow-restricting lesion, and limited availability are important limitations. Internal mammary and radial arteries are often used in the coronary circulation and are preferred for key artery branches because they tend to remain open longer, but also have limited availability (Cameron et al., 1996; Conte, 1998). Although bypass surgery is a common procedure, it carries some serious risks, such as heart attack, stroke, or even death. Native vessels that are used as substitutes have limited availability, and synthetic grafts used to replace small diameter arteries induce clotting and fail. Significant work has been performed toward the design of biomaterials to serve as small diameter conduits. Since the blood-contacting surface of biomaterials often induces clotting, researchers have incorporated heparin to prevent coagulation, seeded the inner surfaces with endothelial cells (ECs) that among other functions provide the anti-thrombogenic properties of native vessels, and/or modified the surface otherwise. Although some of these approaches have been successful shortterm, complications such as thrombosis, infection, and graft failure arise with time. The properties of synthetic grafts, their surface characteristics in particular, have been modified to make them more suitable for small diameter vessel applications. The graft surface has been modified to prevent platelet adhesion either by coating with chemicals or by attaching ECs. Therefore, researchers have modified the surface properties of synthetic materials to reduce their thrombogenicity by modifying surface chemical groups, grafting peptides (Mann et al., 1999; Ko and Iwata, 2002), proteins (Ye et al., 2000; Laredo et al., 2003),

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growth factors (Greisler et al., 1996), heparin, or treating the surface with plasma (Marois et al., 1999; Lee et al., 2000; Chevallier et al., 2005). Since ECs confer blood vessels their anti-thrombogenic properties in vivo, considerable effort has been expended toward the endothelialization of synthetic non-resorbable grafts. Researchers very quickly realized that ECs do not readily adhere to the surface of these grafts, and when they do adhere, they do not remain adhered over time and upon exposure to blood flow. The shortcomings of the therapies described above have led researchers to the development of tissueengineered grafts. An increasing number of researchers support the idea of developing a living blood vessel substitute that closely mimics the native arteries. A living vascular graft will have the ability to respond to hemodynamic changes and other stimuli, remodel, and self-heal. Among the desired properties of a blood vessel substitute are adequate mechanical strength, controllable adaptation to changing hemodynamics, compliant elasticity, zero tolerance for failure, long-term fatigue strength and durability, low thrombogenicity, biocompatibility, suturability, easy handling, and low cost. Although some tissue engineers aim at reproducing the arterial wall architecture and function, most are striving toward restoring function with an arterial replacement possessing a composition and architecture that confers the required properties, not necessarily a replica of the native artery. Small diameter tissue-engineered grafts are typically composed of cells and a scaffold. The scaffold serves as a template that provides the required geometry and has such properties to allow the cells to remodel the scaffold and deposit their own ECM proteins (Figure 58.1). Since the scaffold is resorbable, it is, in principle, degraded

Post-cellularized approach

Pre-cellularized approaches Cultured tissue cells

Harvested tissue





Preformed synthetic polymer

Biopolymer/hydrogel

De Novo synthesis (“self-assembly”) Decellularize chemical treat

Cell Ingrowth

Fibrillogenesis with cell entrappment

Artificial tissue

Cell ingrowth (before or after implantation)

In vitro remodeling (chemical and mechanical signaling)

Implantation in vivo remodeling

Figure 58.1 The general approach to tissue engineering. A scaffold is combined with tissue cells, which subsequently remodel the scaffold by synthesizing ECM to form an engineered tissue.

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over time resulting in tissue formation without any foreign material present. The choice of cells and scaffold is paramount, and the optimal combination is unclear as of yet. However, the in vitro development of vascular grafts is not feasible without the appropriate environmental cues. The following sections address three major components of tissue engineering a small diameter vascular graft: cells, scaffolds, and environmental stimuli.

CELL SOURCING CONSIDERATIONS Cell sourcing is a critical component of a tissue-engineered graft. In order to select the appropriate source, several questions must be answered. What cell type, species, passage, differentiation stage, and donor age should be chosen? How many cell types and which should be used? In what spatial and temporal fashion should the different cell types be introduced to the scaffold? The cell types residing in native blood vessels (i.e. ECs, SMCs, and fibroblasts, or their progenitors) are the obvious candidates as cell types. ECs ECs form a monolayer that lines the entire vascular system and have a remarkable capacity to adjust their number and arrangement to suit local requirements. ECs play an important role in tissue homeostasis, fibrinolysis, and coagulation (thrombogenicity), vasotone regulation, growth regulation of other cell types, and blood cell activation, and migration during physiological and pathological processes (Risau, 1995; Shireman and Pearce, 1996; Aird, 2006; Liebner et al., 2006). Thrombogenicity is an imprecisely defined vascular property, but it implies the qualitative and quantitative assessment of platelet and fibrin deposition on the vascular luminal surface (endothelium). A vessel that is thrombogenic may be so for a variety of reasons, many or most of which are likely related to endothelial dysfunction. Normal quiescent endothelium exhibits limited or absent expression of secreted and cell-associated procoagulant proteins, including platelet adhesogens (e.g. P-selectin) and activators of thrombin generation (tissue factor). Membrane phospholipid asymmetry is maintained in healthy endothelium in order to prevent exposure of the highly pro-thrombotic aminophospholipids that support the assembly of coagulation enzymatic complexes. Conversely, the loss of normal anti-thrombotic or anti-fibrinolytic mechanisms, or the loss or inhibition of mechanisms that prevent platelet adhesion, may also induce thrombogenicity. In tissueengineered vessels, the physical detachment of ECs resulting in exposure of the procoagulant sub-endothelial surface may be at least as important as EC activation as a mechanism for thrombogenicity. Due to the critical role ECs play in determining the thrombogenicity of a vascular graft, EC sourcing dictates to a large extent the patency of a graft (Heyligers et al., 2005). There is a high demand for ECs that can be isolated from the patient requiring bypass surgery to eliminate the need for long-term anti-coagulation therapy and graft rejection. The main EC sources currently explored are umbilical vein ECs (L’Heureux et al., 1998; McKee et al., 2003), venous ECs, mesothelial cells as an alternative to ECs, and progenitor ECs. Venous ECs have been successfully isolated from a single saphenous vein biopsy and used to engineer vascular grafts (Grenier et al., 2003). However, EC isolation from short (2.5 cm long) vein segments without any SMC contamination has proven difficult, and vein biopsies are invasive procedures. Mesothelial cells line serosal cavities and most internal organs in the body and share many characteristics and functions with ECs (Herrick and Mutsaers, 2004). Although this EC source is promising, more work is ongoing to fully characterize the phenotype of these cells and their potential of providing an alternate source of EC-like cells (Campbell et al., 1999). An attractive EC source and the subject of much research have been circulating EC endothelial progenitor cells (EPCs) (Matsumura et al., 2003; Cho et al., 2004). EPCs are found in postnatal bone marrow, have high proliferative capacity, and can differentiate into mature ECs (Hristov et al., 2003). In contrast to mature

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ECs that have low proliferative capacity and limited ability to replace damaged ECs, EPCs are active participants in the repair of denuded endothelium. EPCs have been successfully used to line engineered vascular grafts that remained patent post-implantation in animal models (Kaushal et al., 2001; Matsumura et al., 2003; Matsuda, 2004; Cho et al., 2005). The main drawback of using EPCs is low yield (Matsuda, 2004), which becomes a critical factor when multiple grafts are needed for bypass surgery. Blood outgrowth ECs have also been collected from human peripheral blood through the outgrowth of a marrow-derived, transplantable circulating, putative endothelial progenitor (Lin et al., 2000, 2002; Sieminski et al., 2005). These cells have a cobblestone morphology, take up acetylated LDL, have Weibel–Palade bodies, express multiple endothelial markers (CD34, VE-cadherin, CD31, P1H12, αvβ3 integrin, β1 integrin, thrombomodulin, von Willebrand factor (vWF), flk-1), have a quiescent phenotype, and hold promise as an alternate EC source. Smooth Muscle Cells Vascular SMCs perform many functions, including vasoconstriction and dilation in response to normal or pharmacological stimuli; synthesis of various types of collagen, elastin, and proteoglycans; elaboration of growth factors and cytokines; and migration and proliferation. SMCs are capable of expressing a range of phenotypes or alterations in character (Chamley-Campbell and Campbell, 1981; Thyberg et al., 1990; Owens, 1995). Modulations in cell phenotype may occur as a result of cell–cell interactions, alterations of ECM, or in response to other signals such as hormones (Stegemann and Nerem, 2003). At one end of this phenotypic spectrum are SMCs in contractile state with 80–90% of the cytoplasmic volume occupied with contractile apparatus (Tagami et al., 1986). Organelles such as rough endoplasmic reticulum, Golgi, and free ribosomes are few in number and located in the perinuclear region. SMCs in the contractile phenotypic state exhibit reduced proliferation and matrix production. At the other end of the spectrum is the synthetic state, which is seen in development, repair, and pathological conditions. SMCs with synthetic phenotype proliferate and actively produce ECM proteins, and their cytoplasm contains few filament bundles, but large amounts of rough endoplasmic reticulum, Golgi, and free ribosomes (Ross, 1971). Aortic SMCs in synthetic state synthesize four-fold the amount of collagen and five-fold the amount of sulfated glycosaminoglycans compared to contractile SMCs, whereas the amount of non-collagen protein synthesized doubles under the same conditions. These increases are not related to cell proliferation since synthetic state cells are maintained in a quiescent growth state in these experiments (Campbell, 1985). Contractile SMCs switch to a more synthetic phenotype in vitro (Thyberg et al., 1985), and this change may be irreversible depending on the culture conditions (Chamley-Campbell and Campbell, 1981; Stadler et al., 1989). Vascular SMCs have been successfully used to engineer small diameter vascular grafts (see sections below), although a non-invasive method of harvesting these cells from patients is impossible. Fibroblasts Fibroblasts and SMCs share several functions such as collagen, elastin, and proteoglycan synthesis and contractile behavior. In the normal adult, some of these functions are specifically exerted by the fibroblast (e.g. collagen synthesis) or the SMC (e.g. contractility), but during development or pathological conditions this can change. SMCs secrete collagen during development and during the formation of an atheromatous plaque, whereas contractility may be exerted by fibroblasts during wound healing, resulting in wound contraction (Desmouliere and Gabbiani, 1995). Therefore, fibroblasts (termed myofibroblasts) can acquire SMC-like features during wound contraction and disease and express SMC-specific markers in particular situations. Fibroblasts can be easily harvested from patients through a simple skin biopsy (Normand and Karasek, 1995) rendering these cells a preferred cell type for vascular graft tissue engineering if they prove to provide sufficient function.

Engineering of Small Diameter Vessels

Other Cell Sources Human ECs and SMCs that are used in vascular tissue engineering are oftentimes harvested from young donors. However, the majority of patients requiring bypass surgery are elderly whose SMCs become senescent and do not produce mechanically robust arteries in vitro (McKee et al., 2003). McKee et al. introduced ectopic expression of the human telomerase reverse transcriptase subunit into human SMCs to extend their lifespan and were able to engineer mechanically stronger grafts using infected vascular cells from elderly men compared to cells non-infected with the human telomerase reverse transcriptase subunit (Poh et al., 2005). While autologous ECs must be used to avert an unacceptable inflammatory/immune response, it is unclear whether that is the case for SMCs. Allogeneic SMCs (or other matrix-producing cells that are used to fabricate the tissue-engineered artery) may prove acceptable based on precedents like the tissue-engineered skin Apligraf®, which is fabricated from allogeneic fibroblasts. Bone marrow-derived progenitor cells have been used as an alternate source of SMCs. Progenitor cells were either seeded into polylactic-co-glycolic acid (PLGA) scaffolds and implanted in the peritoneal cavity of athymic mice (Cho et al., 2004) or exposed to 10% cyclic strain at 1 Hz for 7 days (Hamilton et al., 2004); in both cases, cells expressed markers of SMC differentiation. Simper et al. isolated circulating smooth muscle progenitor cells from human blood and showed evidence of SMC outgrowth through the expression of smooth muscle α-actin, myosin heavy chain, and calponin (Simper et al., 2002). Riha et al. have presented a thorough overview of stem cell sourcing for vascular tissue engineering applications (Riha et al., 2005).

SCAFFOLDS FOR SMALL DIAMETER TISSUE-ENGINEERED VESSELS The choice of scaffold in fabricating a cellularized tubular construct dictates the method of fabrication, so the main methods of fabrication are briefly summarized first (Figure 58.1). The use of synthetic polymers typically involves first synthesizing the polymer and processing it into a tube, and then cellularizing the tube, as the polymer synthesis conditions are typically cytotoxic (there are some exceptions, such as PEG-based scaffolds (Seliktar et al., 2004; DeLong et al., 2005). The use of certain biopolymers that self-assemble under physiological conditions allow for cellularization as the biopolymer is formed into a tube. In the scaffold-free approach, cells are cultured so as to produce a sheet of tissue that is subsequently formed into a tube. These in vitro fabrication approaches are distinct from in vivo fabrication approaches wherein a decellularized vessel (or decellularized tissue formed into a tube), or a polymeric rod or tube, is implanted and cellularized by host cells with subsequent tissue growth and remodeling. Synthetic Scaffolds Much of the work in the engineering of vascular grafts has focused on the use of biodegradable synthetic polymer scaffolds. The main advantage of these scaffolds is that they can provide the initial strength necessary for implantation while being biodegradable and can be readily processed into tubes. The obvious disadvantage is that they are synthetic biomaterials that may elicit an immune response. Also, cellularization may be difficult. Semicrystalline polymers such as polyglycolic acid (PGA) and poly-L-lactic acid (PLLA) degrade by bulk hydrolysis. Degradation occurs first in the amorphous domains, which are more accessible to water, and crystallinity gradually increases resulting in a highly crystalline material that is much more resistant to hydrolysis than the starting material. The increase in crystallinity is believed to occur due to an increased mobility of the partially degraded polymer chains, which enables a realignment of the polymer chains into a more ordered crystalline state (Anderson, 1995). PLGA copolymers degrade via non-specific hydrolytic scission of their ester bonds to reform the monomers lactic acid and glycolic acid. Other factors such as pH, heat, and carboxypeptidases may also contribute to the degradation process. In vivo and in vitro experiments with PLGA copolymers have studied their degradation and biocompatibility (Lu et al., 2000). PLLA is thought to

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degrade via simple hydrolysis, whereas PGA may also be subjected to enzyme-mediated hydrolysis. PGA absorption takes 6–17 weeks, and the tensile strength falls to about 10% after 3 weeks, depending on the molecular weight. Only mild inflammatory responses have been caused by PLGA polymers and in some instances phagocytic and giant cells have been observed. There is also concern that the local acidity following degradation can induce cell dedifferentiation (Niklason et al., 1999). PLGA polymers have been approved by the Food and Drug Administration (FDA) for suture materials, bone plates and screws, and cardiovascular woven meshes. Thus, these polymers have been used in various tissue engineering applications. Biodegradable polymeric scaffolds have been used successfully in vascular tissue engineering. Niklason et al. seeded PGA scaffolds with bovine aortic SMCs in a bioreactor system for 8 weeks under pulsatile conditions and subsequently applied bovine aortic ECs for 3 days with continuous flow (Niklason et al., 1999). The resulting endothelium stained positive for vWF and platelet endothelial cell adhesion molecule (PECAM), and the SMCs expressed smooth muscle α-actin and calponin. The grafts showed high SMC density and collagen production, had burst pressure of over 2000 mmHg, and contracted in response to serotonin, endothelin-1, and prostaglandin F2a. Pulsatility increased wall thickness, collagen production, and suture retention strength. Culture medium was supplemented with ascorbic acid, copper ion, and amino acids to support matrix production, which resulted in vessels with higher burst strength. Implantation of scaffolds seeded with autologous ECs and SMCs and cultured in vitro under pulsatile flow into the right saphenous artery of miniature swine resulted in patent grafts after 4 weeks although decreased flow was observed. Limitations of this approach were low elastin production compared to that of native vessels and the presence of dedifferentiated SMCs around residual polymer fragments. Hoerstrup et al. coated PGA meshes with a thin layer of poly-4-hydroxybutyrate, seeded them with ovine myofibroblasts under static conditions for 4 days, seeded them subsequently with ovine ECs, and cultured them in a pulse duplicator bioreactor for up to 28 days (Hoerstrup et al., 2001). DNA and collagen content increased continuously for 21 days but small amounts of matrix were produced which resulted in low mechanical strength. In a separate study with human cells, culture conditions were optimized with ascorbic acid and basic fibroblast growth factor for increased collagen production (Hoerstrup et al., 2000). Shin’oka et al. implanted a polycaprolactone–polylactic acid copolymer reinforced with woven PGA that was cultured with autologous cells for 10 days in vitro in the right pulmonary artery of a 4-year-old girl in Japan (Shin’oka et al., 2001). Seven months later there was no evidence of graft occlusion or aneurysm formation. Although the demands for a blood vessel substitute of the pulmonary circulation are not as high as those for the systemic circulation, the successful implantation of a completely tissue-engineered graft in humans is still a very exciting accomplishment. Biologic Scaffolds Biopolymers, typically a reconstituted type I collagen gel or fibrin gel, are formed with and compacted by tissue cells, where an appropriately applied mechanical constraint to the compaction yields circumferential alignment of fibrils and cells characteristic of the arterial media (L’Heureux et al., 1993; Barocas et al., 1998; Seliktar et al., 2000). It is this last feature that is most attractive about a biopolymer-based tissue-engineered artery. This follows from two axioms (i) that native artery function, particularly mechanical function, depends on structure (particularly alignment of the SMCs and collagen fibers in the medial layer) as much as it depends on composition, and (ii) that the tissue-engineered artery should serve as a functional remodeling template, so that while providing function during the remodeling, the artificial tissue also provides a template for the alignment of the growing tissue. Cells entrapped in a tube of forming biopolymer gel exert traction on the network of fibrils. When the gel contracts around a non-adhesive mandrel, typically over 1–2 weeks, fibrils and cells become circumferentially aligned. Collagen gels have been previously used to engineer small diameter grafts (Weinberg and Bell, 1986; Seliktar et al., 2000) but possessed insufficient mechanical strength for arterial replacements. Attempts at improving the mechanical strength of collagen gels have been moderately successful (Tranquillo et al., 1996;

Engineering of Small Diameter Vessels

Girton et al., 1999; Seliktar et al., 2000). Huynh et al. used submucosal collagen isolated from porcine small intestine coated with type I bovine collagen and treated the inner surface with heparin–benzalkonium to prevent coagulation (Huynh et al., 1999). The collagen layers were cross-linked to increase the mechanical strength, and the grafts remained patent and thrombi-free for up to 13 weeks when implanted in rabbits. However, the response of the human cardiovascular system to animal collagens remains unknown. When a fibrin gel is used, fibrin is replaced by cell-produced ECM over longer times. Fibrin is the major structural protein of a blood clot and can be readily obtained from plasma (Gilbert et al., 2001). Cells entrapped in fibrin gel are able to proliferate and deposit collagen (Figure 58.2) and elastic fibers to a greater extent compared to cells entrapped in collagen gel (Grassl et al., 2002; Long and Tranquillo, 2003; Ross and Tranquillo, 2003) resulting in stronger and stiffer tissues (Grassl et al., 2003). SMCs in fibrin produce around 3–5 times more collagen than SMCs in collagen depending on the concentration of an inhibitor used to control fibrinolysis (Grassl et al., 2002). Collagen fibrils produced by the cells adopt the alignment of the contracted fibrin fibrils, that is, when “media-equivalents” are fabricated such that the SMC induced fibrin contraction results in circumferential alignment, the collagen subsequently produced by the SMC is also circumferentially aligned (Grassl et al., 2003). Ross and Tranquillo (2003) performed a study characterizing the tissue growth and development process that occurs in vitro with this system (Figure 58.3). Following fibrin gel contraction during week 1, peak rates of SMC proliferation, collagen production, and tropoelastin production occur between weeks 1 and 4.

Figure 58.2 Remodeling of fibrin disks revealed by sections stained with Masson’s trichrome stain: fibrin is pink-red, collagen is green, cell nuclei are purple. Top surface of disk is up and adherent surface of disk is down. Thickness (vertical dimension) is 150 μm. 100 Elastin

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Figure 58.3 Time-course of fibrin remodeling into tissue by neonatal SMCs. Gene expression (curves with labels on left), SMC and collagen content, and mechanical properties are plotted together as a percentage of peak value during the 5-week incubation period.

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Organized, cross-linked collagen and elastic fibers replace the degrading fibrin over weeks 3–5 and are manifested as increased mechanical strength. The peak rate of SMC proliferation (weeks 1–2) precedes that for maximum collagen production (weeks 2–4), which is consistent with the 3-week time point of maximum expression of collagen type I and III from quantitative reverse transcription-polymerase chain reaction (qRTPCR). Insoluble elastin quantification reveals that the majority of elastic fibers are produced by week 4, which is also consistent with the qRT-PCR data showing a dramatic downregulation of tropoelastin expression by week 4, indicating elastogenesis occurs during the early stages of tissue growth and development. There is a strong upregulation of lysyl oxidase expression during weeks 1–3 with a peak in expression at week 3, correlating with the phases of collagen and tropoelastin production. Mechanical strength doubles over weeks 4–5 when production of collagen and elastic fibers and expression of lysyl oxidase are subsiding. This may be due in part to the more organized collagen fibrils evident from the histological sections in weeks 3–5. Fibrin-based grafts were successfully implanted into ovine jugular veins for 15 weeks (Swartz et al., 2005; Yao et al., 2005). Explanted grafts were populated with a monolayer of ECs, circumferentially aligned SMCs, significant amounts of collagen, and elastin fibers. Although pre-implantation grafts possessed significant amounts of residual fibrin and were not mechanically strong for the arterial circulation, this in vivo study holds great promise for fibrin gel-based vascular grafts. Decellularized native vessels have also been used as scaffolds, with (Hodde et al., 2002; Amiel et al., 2006) or without recellularization (Hiles et al., 1995; Martin et al., 2005) prior to implantation. Although decellularized grafts may have a shorter route to the clinic by avoiding graft rejection and immune response due to the presence of cells, recellularization often improves patency. Indeed, 4-week patency rates of recellularized rat arteries with ECs were 89% compared to 29% for acellular controls (Borschel et al., 2005). In a hybrid approach, decellularized, porcine small intestinal submucosa was seeded with human umbilical vein ECs for 2 weeks to allow the deposition of a basement membrane (conditioning), and ECs were subsequently removed (Woods et al., 2004). ECs were then reseeded and showed enhanced organization of adherens cell junctions, increased metabolic activity, downregulation of pro-inflammatory prostaglandin PGI2, and decreased adhesion of resting or activated human platelets compared to a non-conditioned graft. Cell Seeding Techniques A confluent and quiescent endothelium is one important characteristic that many current methods have not yet addressed. Grafts are typically endothelialized at later stages in culture and only very few days prior to in vivo experiments (Pawlowski et al., 2004). This results in the formation of an immature endothelium that is not firmly adhered to its substrate. Other roadblocks are the EC expression of an activated phenotype, increased vascular permeability compared to healthy vessels, and distress signals from ECs to underlying cells. ECs are typically stained for EC-specific markers to identify their presence and location. More recently, other functional assays have been used to test whether ECs are able to transmit signals to underlying cells, and whether they upregulate pro-inflammatory and pro-thrombogenic markers in response to drugs such as tumor necrosis factor alpha (TNF-α) and thrombin (Remy-Zolghadri et al., 2004). Although these steps provide some insight, more systematic research is needed to characterize the endothelium of tissue-engineered vascular grafts and ensure that it has reached maturity prior to implantation. Kim et al. compared three different methods of seeding rat aortic SMCs on PGA matrices: static (culture dishes), stirred (spinner flasks), and agitated (50 ml tubes on orbital shaker) (Kim et al., 1998). The dynamic seeding methods yielded constructs with higher cell density, more uniformly distributed cell population, and greater elastin deposition compared to static seeding. Burg et al. also compared static (petri dish), dynamic (spinner flask), and perfusion bioreactor seeding of PGA meshes with rat aortic ECs, by seeding under static or dynamic conditions for 24 h, followed by

Engineering of Small Diameter Vessels

a maturation phase of 6 days in either a static, dynamic, or bioreactor system (Burg et al., 2000). This bioreactor design allows simultaneous testing of replicates but exposes each scaffold to a slightly different environment, and it is therefore not optimal. Nevertheless, it was found that dynamic seeding followed by a bioreactor maturation phase yielded scaffolds with the best cellular attachment and distribution and highest metabolic activity. Centrifugation of SMCs into PLGA-coated PGA scaffolds at 2500 rpm for 10 min was found superior to static or spinner flask seeding terms of seeding efficiency and cellular distribution within scaffolds, especially when spins were broken into 1 min segments (Godbey et al., 2004). Systematic mechanical regimens have also been used to optimize seeding, especially for EC. Baguneid et al. developed a bioreactor in which SMC-seeded scaffolds were exposed to pulsatile shear stress (9.32 dynes/cm2 on average with a maximum of 32.1 dynes/cm2 and 120 mmHg systolic pressure) for 7 days prior to EC seeding (Baguneid et al., 2004). Endothelialized grafts were exposed to 1 h of physiological shear stress before harvesting. During that time, EC retention decreased from 100% to approximately 75% (Baguneid et al., 2004). These results show that a more gradual increase of shear stress over a longer period of time is needed to maximize EC retention. Niklason et al. gradually increased flow rate in the lumen of endothelialized grafts from 1.98 to 6 ml/min or 0.1 to 0.3 dynes/cm2 (Niklason et al., 1999). Clearly, mechanical stimulation of EC was minimal, and ECs were not exposed to a physiological-like environment prior to implantation. It is difficult to assess what phenotype ECs were expressing because only vWF and PECAM staining were reported (Niklason et al., 1999). Finally, basement membrane formation has been shown to have a significant effect on EC adhesion and retention (Baguneid et al., 2004). Fibrin-based media-equivalents were seeded with EC (achieving 99% EC surface coverage in as little as 2 days post-seeding) and placed in a pulsatile flow bioreactor (Isenberg et al., 2006a). The media-equivalents were exposed to steady and pulsatile flow, and EC elongation and alignment were observed in the flow direction, but only when the flow was in the laminar regime (Re  2100). EC surface coverage remained high (95%) in the presence of pulsatile flow up to (at least) 10 dynes/cm2 for 48 h, indicating that ECs were highly adherent to the grafts. Both static and flow conditioned media-equivalents expressed vWF, a marker of properly functioning ECs, suggesting that ECs exposed to flow in the bioreactor were normal. In summary, there is strong evidence that supports the hypothesis that when direct cell entrapment (as for the biopolymers) is not an option, dynamic seeding produces constructs with higher cell densities, uniform cell distribution, higher metabolic activity, and superior mechanical strength. Scaffold-Free A unique approach was taken by L’Heureux et al. (1998), who created a tissue-engineered blood vessel made exclusively of cultured human cells and their matrix without any synthetic biomaterials (Figure 58.4). Human vascular SMCs were cultured to produce a cohesive cellular sheet (media layer) that was wrapped around a mandrel, and a similar sheet of human skin fibroblasts was subsequently wrapped around the media to produce the adventitia. After a minimum of 8 weeks maturation, the mandrel was removed and the lumen was seeded with human umbilical vein ECs (intima). The overall culture period was 3 months, excluding cell expansion. The resulting graft had burst strength of over 2,000 mmHg and good handling and suturability when implanted in a canine model. The media layer was also tested with pharmacological stimuli and showed contractile/relaxation responses (L’Heureux et al., 2001). In another study where human blood vessels were engineered with a similar approach, abdominal interpositional graft patency was 85% in nude rats with time points up to 225 days (L’Heureux et al., 2006). Although the cell sheet-based tissue engineering technique has been very successful, the overall development time is long, and scale-up capabilities to meet patient demand do not appear straightforward.

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Figure 58.4 Cell sheet assembly method for the development of three-layered vascular grafts. CELLULAR INTERACTIONS ECs are believed to recruit SMC-pericytes to newly formed blood vessels during vasculogenesis (Hungerford and Little, 1999). A recent study suggests that a small percentage of primary ECs may give rise to SMCs via transdifferentiation in vitro (Frid et al., 2002). However, the origin of these SMC-like cells remains unclear, and bone marrow-derived progenitor cells may be implicated. Although further work is required to establish whether transdifferentiation of ECs into SMCs is possible, these two cell types clearly interact significantly in vivo. ECs used to be regarded solely as a lining of all blood vessels that form a barrier to blood. However, it is now known that ECs are involved in numerous cell signaling pathways and communicate with SMCs via heterocellular junctions and other mediators (Davies, 1986). Intermittent fenestrations in the internal elastic lamina are 0.5–1.5 μm in large vessels and 0.1–0.45 μm in capillaries, thereby allowing direct contact between the two cell types (Saunders and D’Amore, 1992). EC–SMC co-culture experiments have revealed an EC effect on SMC proliferation (Vouyouka et al., 2004), migration (Casscells, 1992), phenotype (Brown et al., 2005), and ECM production. Heterotypic interactions between SMCs and ECs have been investigated in a number of in vitro assays. Typical results in “transfilter” assays, where cells are cultured on opposite sides of a membrane (e.g. Dacron), are that confluent ECs inhibit SMC proliferation while subconfluent ECs stimulate SMC proliferation (Axel et al., 1996; Yoshida et al., 1996), perhaps via PDGF-AB from the ECs (Axel et al., 1997), although there are species differences (Imegwu et al., 2001) and effects depend on the ECM coating the substrate. Likewise, EC proliferation can be inhibited by the presence of SMCs (Imegwu et al., 2001). SMC proliferation depends on the EC proliferative state, and EC can be either stimulators or inhibitors (Campbell and Campbell, 1986; Casscells, 1992). Synthetic SMCs in the presence of proliferating endothelium have an increased proliferation rate, whereas confluent, quiescent endothelium inhibits SMC proliferation. ECs co-cultured with SMCs across a 10 μm thick porous polycarbonate membrane induce SMC proliferation (Waybill et al., 1997; Waybill and Hopkins, 1999), but in another study, where a 13 μm thick polyethylene terephthalate membrane was used, this effect was decreased with time and led to growth inhibition by day 4 (Fillinger et al., 1993). Respectively, SMCs inhibit EC proliferation in vitro when the two cell types are in direct contact but not when they are separated by 1–2 mm (Orlidge and D’Amore, 1987). Conditioned medium from EC–SMC cocultures also inhibits EC proliferation to the same degree as direct contact co-culture itself, due to activated transforming growth factor-β production (Antonelli-Orlidge et al., 1989). Three-dimensional co-culture studies of ECs and SMCs in collagen gels show that SMCs affect not only EC proliferation, but also EC alignment and elongation (Imberti et al., 2002; Ziegler et al., 1995). ECs seeded onto SMC-contracted collagen gel were growth-inhibited relative to tissue culture plastic, and proliferation was further reduced in the presence of physiological shear flow for 24 h (Ziegler et al., 1995). In a co-culture

Engineering of Small Diameter Vessels

study where ECs were seeded directly on SMCs, a lower (and decreasing over time) attachment efficiency of ECs was observed when they were seeded on proliferating compared to quiescent SMCs (Lavender et al., 2005). Therefore, cell signaling between ECs and SMCs is a two-way communication. Also, ECs affect SMC matrix deposition but these effects appear less studied. Culturing rabbit aortic SMC monolayers in conditioned medium from confluent bovine aortic EC stimulates glycosaminoglycans synthesis up to 120% within 24 h by the SMC (Merrilees et al., 1990) and keeps the SMC in a differentiated, contractile phenotype. In a different study with pig and rat aortic ECs and SMCs, hyaluronic acid and sulfated glycosaminoglycans production increases in co-culture compared to separate cultures of the two cell types (Merrilees and Scott, 1981). In contrast, collagen synthesis and collagen type I expression decrease in EC–SMC in vitro co-culture models (Powell et al., 1997). ECs have also been shown to increase gene expression of vascular endothelial growth factor, platelet-derived growth factors AA and BB, and transforming growth factor β, and decrease expression of basic fibroblast growth factor in co-cultured SMCs (Heydarkhan-Hagvall et al., 2003). In a more recent study where ECs and SMCs were seeded into a tubular PGA scaffold and co-cultured in a perfusion bioreactor, 15-day co-culture increased cell proliferation, decreased collagen and proteoglycan deposition, and led to higher SMC expression of contractile proteins compared to 2-day co-cultures (Williams and Wick, 2005). Lavender et al. investigated the effects of various substrates and medium formulations on their ability to successfully co-culture SMCs and ECs under steady, laminar flow conditions (Lavender et al., 2005). They found that medium conditions that yielded a quiescent SMC population allowed for the direct culture of a distinct, confluent, and adherent EC monolayer on top of the SMC layer for up to 10 days. Furthermore, under these conditions, ECs increased their rate of acetylated–LDL uptake, and PECAM expression in EC borders was decreased. These co-culture studies reveal strong interactions between ECs and SMCs that are dependent on co-culture method and cell densities in vitro. This communication occurs through mediators released into the surrounding medium or through direct cell–cell contact in vivo, and is an important factor in the control of blood vessel growth, remodeling, and function. The studies described above clearly show that SMC–EC interactions can greatly affect development of tissue-engineered vascular grafts. However, these interactions will not only determine the grafts’ in vitro properties but may also determine the grafts’ in vivo patency through SMC effects on EC thrombogenicity. Human umbilical vein ECs co-cultured in vitro with human umbilical cord SMCs in the presence of TNF-α supported significantly higher levels of platelet adhesion compared to ECs cultured with TNF-α but in the absence of SMCs (Tull et al., 2006). It has also been shown that when human umbilical cord ECs are co-cultured with human umbilical cord SMCs under disturbed flow conditions in a vertical-step flow chamber, neutrophil, peripheral blood lymphocyte, and monocyte adhesion to ECs and transmigration through the EC monolayers are significantly increased compared to controls in the absence of SMCs (Chen et al., 2006).

BIOREACTOR CULTURES Tissue engineering studies have elucidated the importance of bioreactors for improving cell seeding, ECM production, and tissue architecture and differentiation compared to static culture techniques (Carrier et al., 2002; Davisson et al., 2002; Pei et al., 2002). Several bioreactor systems are currently available for the development of vascular grafts. One of these systems engineered by Niklason et al. is composed of four chambers assembled in parallel, a medium reservoir, a pulsatile pump, and a compliance chamber (Niklason et al., 1999, 2001). In each chamber, a highly distensible silicone tubing is inserted through the lumen of the polymer scaffold. Mixing is achieved through stirrer bars and magnetic stirplates, each scaffold is seeded initially with an SMC suspension for 30 min, each chamber is filled with culture medium and pulsatile flow is introduced through the silicone tubing. After 8 weeks the

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silicone tubing is removed, and medium flow is applied directly through the construct. Subsequently, EC suspension is injected into the lumen, and cells are allowed to adhere for 90 min under static conditions. Flow rates are then increased over a 3-day culture period from 0.033 to 0.1 ml/s. This bioreactor has several advantages such as continuous perfusion and pulsatile flow capabilities and has been used to generate mechanically robust vascular grafts. However, SMC and EC seeding involves opening each bioreactor chamber which may compromise asepsis. The significance of culture in perfusion bioreactors has been since shown in several studies (McFetridge et al., 2004; Opitz et al., 2004; Williams and Wick, 2004; Engbers-Buijtenhuijs et al., 2006). Rabbit SMCs seeded into poly(lactide-co-caprolactone) scaffolds and exposed to 5% radial distension and 25 mmHg pressure at 1 Hz exhibited higher proliferation and collagen deposition, significant cell alignment in the radial direction and upregulation of smooth muscle α-actin compared to static controls (Jeong et al., 2005). Commercially available spinner flasks (Bellco Glass, Inc.) have also been used for long-term (7 weeks) culture of rat SMCs into PGA matrices bonded with PLLA (Kim and Mooney, 1998). SMCs were seeded into the scaffolds by placing them on an orbital shaker at 100 rpm for 24 h, and the seeded constructs were then transferred to the spinner flasks and cultured at 40 rpm. SMCs proliferated in the bonded scaffolds and produced elastin over the 7-week culture time. Thus, cell seeding and construct growth took place in two different systems unlike in a single system as with Niklason et al. (1999). Seliktar developed a dynamic mechanical conditioning bioreactor, in which up to four constructs can be mounted on an expandable silicone sleeve (Seliktar, 2000). The intraluminal pressure was regulated to produce a 10% cyclic change in the outer diameter of each silicone tube, and it was shown that mechanical preconditioning of collagen gels seeded with rat aortic SMCs improves their mechanical strength and histological organization (Seliktar et al., 2000). Although this system is not used for the seeding phase of vascular graft development, it has the potential to be used for growth.

IN VIVO CONSTRUCT FABRICATION Harvesting cells and expanding them in culture oftentimes result in a phenotypic change and may prevent robust tissue formation in vitro. Furthermore, an “optimized” cell culture medium for a given cell type and tissue has not yet been developed. To overcome these limitations, researchers have attempted recruiting cells in vivo and using the body as the bioreactor. Campbell et al. developed grafts in the recipient’s peritoneal cavity by inserting silastic tubing as a mandrel in the peritoneal cavity of rats or rabbits. The intima of the grafts became populated with mesothelial cells and the media with myofibroblasts within 2 weeks (Campbell et al., 1999). Although this innovative method of graft development needs to be further evaluated in long-term in vivo experiments to establish patency, it bypasses complicated cell and scaffold sourcing-related issues. In a different study, six different kinds of polymeric rods were inserted subcutaneously into rabbits for up to 3 months (Nakayama et al., 2004). The tissue formed on these biotubes was mostly collagen-rich ECM and fibroblasts, and endothelialization of the resulting grafts was not addressed. Another approach is the implantation of a cell-free scaffold to recruit cells from the host. Hyaluronan-based grafts implanted in rat abdominal aortas were populated with ECs and SMCs within 21 days and contained collagen and elastic fibers (Lepidi et al., 2006). CONCLUSION Significant advances have been made toward the development of a small diameter vascular graft, although the challenges remain substantial. Development of vascular substitutes is time-consuming and in most cases one graft is produced at a time. This approach raises the issue of just how efficient and cost-effective the process can be, and also how reproducibility can be ensured (Ratcliffe and Niklason, 2002). A functional small diameter

Engineering of Small Diameter Vessels

vascular graft possesses appropriate mechanical properties, including physiological compliance and viscoelasticity and, critically, adequate burst strength, without any propensity for permanent creep that leads to aneurysm. It also possesses transport properties, such as appropriate permeability to plasma and proteins. Finally, it exhibits physiological properties, such as vasoconstriction/relaxation responses, insofar as these responses indicate a physiological SMC phenotype. From a practical standpoint, suturability and simplicity of handling are necessary, and from a commercial standpoint, it must be fabricated in a process that scales well with quantity and be a product that can be shipped and stored. Meeting all criteria simultaneously remains a challenge. For example, high burst strength is often associated with compliance mismatch (L’Heureux et al., 1998), which can lead to intimal hyperplasia at the suture line. Conversely, collagen-based constructs that possess physiological compliance have lacked high burst strength (Girton et al., 2000). Fibrin-based constructs yield higher burst strengths and physiological compliance (Isenberg et al., 2006b), although there is no accepted standard for what constitutes a minimum burst pressure at implantation. Notably, no approach has yet resulted in all the key features of the media layer, namely circumferential alignment of SMCs, collagen fibers, and elastin lamellae. In fact, mature (i.e. crosslinked) elastin fibers have only been reported in the self-assembly approach, and in association with fibroblasts, not SMCs (L’Heureux et al., 1998). The developmental downregulation of elastogenesis in SMCs creates a major hurdle (McMahon et al., 1985; Johnson et al., 1995). Indeed, elastic recoil is critical to abolish permanent creep and is conferred by elastin lamellae in the large elastic arteries (Opitz et al., 2004, Patel et al., 2006), whereas lamellae are less prominent in smaller diameter muscular arteries, which are the targets of vascular tissue engineering. It remains to be seen whether other ECM can confer both elasticity and physiological compliance in the absence of elastin lamellae. This question is related to a broader challenge for the field of tissue engineering: the need for a predictive basis for the optimal combination of cell source/scaffold/stimulation/bioreactor. This will hinge on a more complete understanding of how the cell integrates the various signals at the cellular and molecular level. This understanding will translate into biophysical models that relate cell cycle regulation and the production and assembly of ECM components in response to these integrated signals, and ultimately into multiscale mechanical models that relate the evolving ECM at the molecular level to macroscopic mechanical and functional properties. There are recent continuum mechanical models of vascular growth and remodeling that are aimed in this direction (Taber, 2001; Humphrey and Rajagopal, 2003; Gleason and Humphrey, 2004). Ultimately, the growth and remodeling that occur following implantation in response to signals that the tissue engineer has little or no control over will determine the success of tissue-engineered vessels. There is scant information about how growth and remodeling depend on the properties at implantation. Furthermore, there is no imminent solution to the extreme immunogenicity of non-autologous ECs. Even if a construct could be pre-fabricated from non-autologous SMCs, it would still take many days to weeks to isolate and expand the patient’s ECs to the numbers required for seeding a construct of useful length, for example, with circulating EC progenitor cells (Hristov et al., 2003; Matsumura et al., 2003; Cho et al., 2005) or blood outgrowth ECs (Lin et al., 2000, 2002), both of which possess high proliferative capacity and can differentiate into mature ECs. The associated time lag, however, might limit the applicability of vascular grafts fabricated with these cell sources to patients with anticipated repeat procedures. The optimal sources for SMCs and ECs remain to be determined, but economic and regulatory considerations would favor prefabrication of small diameter vascular grafts from non-autologous, genetically unmodified cells.

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Merrilees, M.J. and Scott, L. (1981). Interaction of aortic endothelial and smooth muscle cells in culture. Effect on glycosaminoglycan levels. Atherosclerosis 39(2): 147–161. Merrilees, M.J., Campbell, J.H., Spanidis, E. and Campbell, G.R. (1990). Glycosaminoglycan synthesis by smooth muscle cells of differing phenotype and their response to endothelial cell conditioned medium. Atherosclerosis 81(3): 245–254. Nakayama, Y., Ishibashi-Ueda, H. and Takamizawa, K. (2004). In vivo tissue-engineered small-caliber arterial graft prosthesis consisting of autologous tissue (biotube). Cell Transplant. 13(4): 439–449. Niklason, L.E., Gao, J., Abbott, W.M., Hirschi, K.K., Houser, S., Marini, R. and Langer, R. (1999). Functional arteries grown in vitro. Science 284(5413): 489–493. Niklason, L.E., Abbott, W., Gao, J., Klagges, B., Hirschi, K.K., Ulubayram, K., Conroy, N., Jones, R., Vasanawala, A., Sanzgiri, S. et al. (2001). Morphologic and mechanical characteristics of engineered bovine arteries. J. Vasc. Surg. 33(3): 628–638. Normand, J. and Karasek, M.A. (1995). A method for the isolation and serial propagation of keratinocytes, endothelial cells, and fibroblasts from a single punch biopsy of human skin. In Vitro Cell. Dev. Biol. Anim. 31(6): 447–455. Opitz, F., Schenke-Layland, K., Cohnert, T.U., Starcher, B., Halbhuber, K.J., Martin, D.P. and Stock, U.A. (2004). Tissue engineering of aortic tissue: dire consequence of suboptimal elastic fiber synthesis in vivo. Cardiovasc. Res. 63(4): 719–730. Opitz, F., Schenke-Layland, K., Richter, W., Martin, D.P., Degenkolbe, I., Wahlers, T. and Stock, U.A. (2004). Tissue engineering of ovine aortic blood vessel substitutes using applied shear stress and enzymatically derived vascular smooth muscle cells. Ann. Biomed. Eng. 32(2): 212–222. Orlidge, A. and D’Amore, P.A. (1987). Inhibition of capillary endothelial cell growth by pericytes and smooth muscle cells. J. Cell. Biol. 105(3): 1455–1462. Owens, G.K. (1995). Regulation of differentiation of vascular smooth muscle cells. Physiol. Rev. 75(3): 487–517. Patel, A., Fine, B., Sandig, M. and Mequanint, K. (2006). Elastin biosynthesis: The missing link in tissue-engineered blood vessels. Cardiovasc. Res. 71(1): 40–49. Pawlowski, K.J., Rittgers, S.E., Schmidt, S.P. and Bowlin, G.L. (2004). Endothelial cell seeding of polymeric vascular grafts. Front. Biosci. 9:1412–1421. Pei, M., Solchaga, L.A., Seidel, J., Zeng, L., Vunjak-Novakovic, G., Caplan, A.I. and Freed, L.E. (2002). Bioreactors mediate the effectiveness of tissue engineering scaffolds. FASEB J. 16(12): 1691–1694. Poh, M., Boyer, M., Solan, A., Dahl, S.L., Pedrotty, D., Banik, S.S., McKee, J.A., Klinger, R.Y., Counter, C.M. and Niklason, L.E. (2005). Blood vessels engineered from human cells. Lancet 365(9477): 2122–2124. Powell, R.J., Hydowski, J., Frank, O., Bhargava, J. and Sumpio, B.E. (1997). Endothelial cell effect on smooth muscle cell collagen synthesis. J. Surg. Res. 69(1): 113–118. Ratcliffe, A. and Niklason, L.E. (2002). Bioreactors and bioprocessing for tissue engineering. Ann. NY Acad. Sci. 961: 210–215. Remy-Zolghadri, M., Laganiere, J., Oligny, J.F., Germain, L. and Auger, F.A. (2004). Endothelium properties of a tissueengineered blood vessel for small-diameter vascular reconstruction. J. Vasc. Surg. 39(3): 613–620. Riha, G.M., Lin, P.H., Lumsden, A.B., Yao, Q. and Chen, C. (2005). Review: application of stem cells for vascular tissue engineering. Tissue Eng. 11(9–10): 1535–1552. Risau, W. (1995). Differentiation of endothelium. FASEB J. 9(10): 926–933. Ross, R. (1971). The smooth muscle cell. II. Growth of smooth muscle in culture and formation of elastic fibers. J. Cell. Biol. 50(1): 172–186. Ross, J.J. and Tranquillo, R.T. (2003). ECM gene expression correlates with in vitro tissue growth and development in fibrin gel remodeled by neonatal smooth muscle cells. Matrix Biol. 22(6): 477–490. Saunders, K.B. and D’Amore, P.A. (1992). An in vitro model for cell–cell interactions. In Vitro Cell. Dev. Biol. 28A(7–8): 521–528. Schoen, F.J. (1994). Blood Vessels. In: Cotran, R.S., Kumar, V. and Robbins, S.L. (eds.), Robbins Pathologic Basis of Disease. W.B. Saunders Company. Philadelphia, PA. Seliktar, D., (2000). Dynamic Mechanical Conditioning Regulates the Development of Cell-Seeded Collagen Constructs In Vitro: Implications for Tissue-Engineered Blood Vessels. Atlanta: Georgia Institute of Technology. Seliktar, D., Black, R.A., Vito, R.P. and Nerem, R.M. (2000). Dynamic mechanical conditioning of collagen-gel blood vessel constructs induces remodeling in vitro. Ann. Biomed. Eng. 28(4): 351–362.

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Seliktar, D., Zisch, A.H., Lutolf, M.P., Wrana, J.L. and Hubbell, J.A. (2004). MMP-2 sensitive, VEGF-bearing bioactive hydrogels for promotion of vascular healing. J. Biomed. Mater. Res A. 68(4): 704–716. Shin’oka, T., Imai, Y. and Ikada, Y. (2001). Transplantation of a tissue-engineered pulmonary artery. N. Engl. J. Med. 344(7): 532–533. Shireman, P.K. and Pearce, W.H. (1996). Endothelial cell function: biologic and physiologic functions in health and disease. AJR Am J Roentgenol. 166(1): 7–13. Sieminski, A.L., Hebbel, R.P. and Gooch, K.J. (2005). Improved microvascular network in vitro by human blood outgrowth endothelial cells relative to vessel-derived endothelial cells. Tissue Eng. 11(9–10): 1332–1345. Simper, D., Stalboerger, P.G., Panetta, C.J., Wang, S. and Caplice, N.M. (2002). Smooth muscle progenitor cells in human blood. Circulation 106(10): 1199–1204. Stadler, E., Campbell, J.H. and Campbell, G.R. (1989). Do cultured vascular smooth muscle cells resemble those of the artery wall? If not, why not? J. Cardiovasc. Pharmacol. 14(Suppl 6):S1–S8. Stegemann, J.P. and Nerem, R.M. (2003). Altered response of vascular smooth muscle cells to exogenous biochemical stimulation in two- and three-dimensional culture. Exp. Cell Res. 283(2): 146–155. Swartz, D.D., Russell, J.A. and Andreadis, S.T. (2005). Engineering of fibrin-based functional and implantable smalldiameter blood vessels. Am. J. Physiol. Heart Circ. Physiol. 288(3): H1451–H1460. Taber, L.A. (2001). Biomechanics of cardiovascular development. Annu. Rev. Biomed. Eng. 3:1–25. Tagami, M., Nara, Y., Kubota, A., Sunaga, T., Maezawa, H., Fujino, H. and Yamori, Y. (1986). Morphological and functional differentiation of cultured vascular smooth-muscle cells. Cell Tissue Res. 245(2): 261–266. Thyberg, J., Nilsson, J., Palmberg, L. and Sjolund, M. (1985). Adult human arterial smooth muscle cells in primary culture. Modulation from contractile to synthetic phenotype. Cell Tissue Res. 239(1): 69–74. Thyberg, J., Hedin, U., Sjolund, M., Palmberg, L. and Bottger, B.A. (1990). Regulation of differentiated properties and proliferation of arterial smooth muscle cells. Arteriosclerosis 10(6): 966–990. Tranquillo, R.T., Girton, T.S., Bromberek, B.A., Triebes, T.G. and Mooradian, D.L. (1996). Magnetically orientated tissueequivalent tubes: application to a circumferentially orientated media-equivalent. Biomaterials 17(3): 349–357. Tull, S.P., Anderson, S.I., Hughan, S.C., Watson, S.P., Nash, G.B. and Rainger, G.E. (2006). Cellular pathology of atherosclerosis: smooth muscle cells promote adhesion of platelets to cocultured endothelial cells. Circ. Res. 98(1): 98–104. Vouyouka, A.G., Jiang, Y. and Basson, M.D. (2004). Pressure alters endothelial effects upon vascular smooth muscle cells by decreasing smooth muscle cell proliferation and increasing smooth muscle cell apoptosis. Surgery 136(2): 282–290. Waybill, P.N., Chinchilli, V.M. and Ballermann, B.J. (1997). Smooth muscle cell proliferation in response to co-culture with venous and arterial endothelial cells. J. Vasc. Interv. Radiol. 8(3): 375–381. Waybill, P.N. and Hopkins, L.J. (1999). Arterial and venous smooth muscle cell proliferation in response to co-culture with arterial and venous endothelial cells. J. Vasc. Interv. Radiol. 10(8): 1051–1057. Weinberg, C.B. and Bell, E. (1986). A blood vessel model constructed from collagen and cultured vascular cells. Science 231(4736): 397–400. Williams, C. and Wick, T.M. (2004). Perfusion bioreactor for small diameter tissue-engineered arteries. Tissue Eng. 10(5–6): 930–941. Williams, C. and Wick, T.M. (2005). Endothelial cell-smooth muscle cell co-culture in a perfusion bioreactor system. Ann. Biomed. Eng. 33(7): 920–928. Woods, A.M., Rodenberg, E.J., Hiles, M.C. and Pavalko, F.M. (2004). Improved biocompatibility of small intestinal submucosa (SIS) following conditioning by human endothelial cells. Biomaterials 25(3): 515–525. Yao, L., Swartz, D.D., Gugino, S.F., Russell, J.A. and Andreadis, S.T. (2005). Fibrin-based tissue-engineered blood vessels: differential effects of biomaterial and culture parameters on mechanical strength and vascular reactivity. Tissue Eng. 11(7–8): 991–1003. Ye, Q., Zund, G., Jockenhoevel, S., Schoeberlein, A., Hoerstrup, S.P., Grunenfelder, J., Benedikt, P. and Turina, M. (2000). Scaffold precoating with human autologous extracellular matrix for improved cell attachment in cardiovascular tissue engineering. Asaio J. 46(6): 730–733. Yoshida, H., Nakamura, M., Makita, S. and Hiramori, K. (1996). Paracrine effect of human vascular endothelial cells on human vascular smooth muscle cell proliferation: transmembrane co-culture method. Heart Vessels 11(5): 229–233. Ziegler, T., Alexander, R.W. and Nerem, R.M. (1995). An endothelial cell-smooth muscle cell co-culture model for use in the investigation of flow effects on vascular biology. Ann. Biomed. Eng. 23(3): 216–225.

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59 Vascular Assembly in Engineered and Natural Tissues Eric M. Brey and Larry V. McIntire INTRODUCTION Mass transfer limitations are an important obstacle to clinical application of engineered tissues (McIntire, 2002). The majority of cells need to be within 100–200 μm of a blood supply to receive adequate oxygen and nutrients for survival (Carmeliet and Jain, 2000). Engineering a tissue of clinically relevant size requires the formation of extensive, stable microvascular networks in the tissue (Brey et al., 2005). Clinical trials have shown some promise when treating tissue ischemia with the genes or proteins of naturally occurring angiogenic factors, but efficacy has not been proven in randomized, double blind, placebo controlled trials (Epstein et al., 2001). The situation in tissue engineering can be even more challenging. Ischemic tissues have a pre-existing, but diseased, microvasculature that requires remodeling or regrowth, while engineered tissues may require the de novo generation of a complete microvasculature. A number of issues need to be addressed in order to successfully vascularize engineered tissues for clinical use, including: control of protein or gene levels both spatially and temporally, long-term stability of the microvasculature formed, and the structure and phenotype of the vessels formed. In addition, an important barrier to clinical success is the decreased sensitivity of many within the potential patient population, including older patients, diabetics, and patients with heart disease to vascular interventions (Brey and Greisler, 2005; Poh et al., 2005). In this chapter, techniques under investigation for stimulating vascular assembly within engineered tissues will be presented identify some of the limitations of each approach identified. VASCULAR ASSEMBLY: BASIC MECHANISMS Angiogenesis/Vasculogenesis Neovascularization is the process by which new vascular structures assemble. Under normal adult physiological conditions vascular networks are relatively quiescent, with neovascularization primarily limited to wound healing and steps in the female reproductive processes. However, abnormal neovascularization can result from a number of diseases. It is believed that under normal physiological conditions tissues contain an equal balance of positive and negative regulators of neovascularization. Neovascularization is initiated when some environmental stimulus tilts this balance toward a higher relative level of positive factors, a time known as the “angiogenic switch” (Carmeliet and Jain, 2000). A number of factors have been identified that can turn on this “switch,” including hypoxia, hypoglycemia, mechanical stresses, inflammation, and genetic mutations related to cancer or disease. There are two primary mechanisms by which neovascularization can occur: vasculogenesis and angiogenesis (Carmeliet, 2005). Vasculogenesis is the in situ assembly of endothelial progenitors into capillaries, while angiogenesis is the formation of new capillaries from pre-existing vessels. Angiogenesis and vasculogenesis were initially

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Figure 59.1 Immunostain showing CD31 positive capillaries (blue) interacting with desmin positive pericytes (brown). The pericytes exhibit the classic structure, coating a portion of the capillary surface and extending cytoplasmic processes along the vessel surface.

considered independent events, with vasculogenesis occurring exclusively during embryogenesis and angiogenesis in adults. It is now recognized that both mechanisms can contribute to neovascularization in a single microenvironment (Augustin, 2001; Hirschi and Goodell, 2001).Adult neovascularization is thought to proceed primarily via angiogenesis, with the contribution of vasculogenesis ranging between 3.5% and 25% (Hirschi and Goodell, 2001). The early steps of neovascularization are often treated as an endothelial cell (EC) only phenomenon, with ECs invading the extracellular matrix (ECM), migrating toward some stimulus, and then assembling into network structures. However, other cells are present during these times and may play an important role in guiding the formation of the initial structures (Gerhardt and Betsholtz, 2003; Ponce and Price, 2003; Brey et al., 2004). Pericytes are vascular mural cells that interact with ECs on microvessels and play a role in vessel stability, mechanical function, and response to physiological changes. Unlike smooth muscle cells (SMCs) in larger vessels, pericytes do not coat entire capillary surfaces but extend cytoplasmic processes that encompass only a fraction of the vessel surface (Figure 59.1). Pericytes can be present on early capillary sprouts (Gerhardt and Betsholtz, 2003; Ponce and Price, 2003; Brey et al., 2004), but it is not clear what role they play in neovascularization. Remodeling/Stabilization Neovascularization leads to an initial excess of vessels formed. The networks are then remodeled and stabilized to meet the specific metabolic demands of the tissue (Darland and D’Amore, 2001). Prior to stabilization the vessels may be in a growth factor dependent phase (Benjamin et al., 1998, 1999), a time in which growth factors are thought to provide survival signals that inhibit vessel regression while awaiting stabilizing interactions (Benjamin et al., 1998). It is at this time that pericyte recruitment to, and proliferation on, the immature vessels increases. The vessels are then considered “mature” suggesting that they are quiescent, independent of angiogenic survival factors (Abramsson et al., 2002), and less responsive than immature vessels to anti-angiogenic signals (Jain, 2001; Gee et al., 2003). Pericyte growth on, and recruitment to, new vessels is mediated in part by platelet-derived growth factorBB (PDGF-BB), TGF-β1, and EC-pericyte contact (Hirschi et al., 1998, 1999). However, the source of the pericytes recruited to the vessels has not been determined. Evidence suggests that they may result from the proliferation and migration of mural cells from pre-existing vessels or arise from differentiation of mesenchymal cells in the surrounding tissue. It is possible that both sources contribute to the resulting mural cells.

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The presence of mural cells on the surface of ECs is often treated as an indication of vessel stability; however, this is not entirely true. As discussed previously, pericytes are often present during initial vessel sprout formation. In addition, stable capillaries in some tissues are devoid of pericytes and research has shown that in some cases mural cell-coated vessels regress following loss of neovascular stimuli (Gee et al., 2003; Brey et al., 2004). It is likely that additional signals (produced by the pericyte and/or ECs), such as transforming growth factor-β1 (TGF-β1) (Hirschi et al., 1999), angiopoietin-1 (Ang-1) (Uemura et al., 2002), or basement membrane proteins, may be required for vessel stabilization. In fact, Ang-1 alone can provide vessel integrity in the absence of pericytes (Uemura et al., 2002). Patterning/Structure The microvasculature has a distinct hierarchy and is divided into defined arterial and venous regions and the overall architecture and capillary structure varies between different tissues. Capillaries may be continuous, discontinuous, or fenestrated depending on the particular tissue bed. These structures and vessel organization are vital to the function of the resultant microvasculature. The mechanisms causing EC organization into distinct patterns have not been elucidated, but recent advances are providing greater insight into the control of this process. The eph family of receptors and their ephrin ligands play a role in arterial/venous specification (Hayashi et al., 2005). Ephrin-B2 has been shown to be exclusively expressed on arterial ECs, while its receptor (Eph-B4) is expressed by venous ECs (Yancopoulos et al., 1998). ECs acquire an arterial or venous designation at the earliest stages of development, prior to the formation of networks and establishing flow (Wang et al., 1998). These ECs continue to communicate with one another as networks develop during embryogenesis. While it is not entirely clear how specification occurs in the adult, it seems possible that this arterial–venous crosstalk would also be present during adult neovascularization. Alternative splicing of vascular endothelial growth factor (VEGF), a potent stimulant of neovascularization, results in six different isoforms that contribute in defined ways to the process of neovascularization. Isoforms of VEGF are expressed with and without ECM binding regions. The ECM binding isoforms result in steep interstitial VEGF gradients that guide vessel sprouting and influence branching patterns (Gerhardt et al., 2003). Without ECM binding isoforms the gradient is more gradual, resulting in sprouts with altered spatial orientation. While VEGF clearly plays an important role, it is not the only factor that modulates microvascular patterning. Recent evidence suggests that nerves and vessels have common guidance signals. VEGF, which was once thought to be an EC-specific factor, is now known to have neuroprotective effects and guide neuronal patterning through direct action on neuronal cells (Zachary, 2005). In addition, known axonal guidance factors, such as members of the slit and semaphorin families, are also involved in guiding blood vessel sprouts (Autiero et al., 2005). In the inner retinal vascular plexus, astrocytes were seen guiding sprouts as they migrated along the VEGF gradient (Gerhardt et al., 2003). Other proteins without known axonal guidance roles have been shown to contribute to vascular patterning. Transgenic overexpression of fibroblast growth factor-1 (FGF-1 or acidic FGF) significantly increased the number of branching structures in the heart without altering capillary density (Fernandez et al., 2000). While controlling concentration gradients and relative levels of pleiotrophic factors is likely to effect pattern formation, tissue-specific angiogenic factors may also play a role in the structure of vessels formed. Endocrine gland-derived VEGF (EG–VEGF) selectively stimulates endocrine gland ECs and can induce ECs to the fenestrated phenotype often seen in endocrine glands (LeCouter et al., 2002). The existence of tissuespecific factors such as EG–VEGF suggests that the specific proteins used, in addition to their spatial and temporal levels, play a role in controlling microvascular structure. Arteriogenesis Distinct from angiogenesis or vasculogenesis, arteriogenesis may be the mechanism of vascular remodeling that proves to be the most important for tissue engineering success. Arteriogenesis is not the formation of new

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vessels but is the process by which pre-existing vessels are enlarged in order to increase blood flow. The resulting high conductance vessels rapidly increase blood volumes unlike capillaries formed via angiogenesis or vasculogenesis. Arteriogenesis is a structural remodeling of existing vessels driven in part by changes in vessel shear stress. An increase in shear may activate ECs to release factors that recruit monocytes to the collaterals. These monocytes produce many of the chemical mediators of arteriogenesis, including tumor necrosis factor-α (TNF-α), leading to inflammation (van Royen et al., 2001). This local inflammatory environment plays an important role in providing signals vital to the enlargement of the collaterals.

GENES AND PROTEINS Proteins The process of neovascularization is controlled by a complex spatial and temporal expression of proteins. Many neovascularization strategies have focused on stimulating angiogenesis by injecting these growth factors either systemically or directly into the target tissue. VEGF, FGF-1, and FGF-2 (or basic FGF) were some of the first identified angiogenic growth factors. These growth factors have enjoyed the most popularity as therapies, reaching the point of clinical trials for the treatment of myocardial and/or peripheral limb ischemia (Cao et al., 2005). Many other growth factors also play a role in neovascularization and are under investigation in animal and in vitro models. The angiopoietins (Ang-1 and Ang-2), placental growth factor (PlGF), FGF-4, hepatocyte growth factor (HGF), and platelet-derived growth factor-BB (PDGF-BB) are a few of the many other proteins with the potential to be used for therapeutic neovascularization. Soluble ephrin-B2 has also been shown to induce neovascularization primarily through stimulation of venous angiogenesis (Hayashi et al., 2005). Regardless of the protein used, high doses and repeat injections are required in order to achieve a significant response due to short protein half-lives in vivo and their rapid diffusion out of target tissues. Sustained protein levels are most likely needed in order to form a stable microvasculature, while the high doses have lead to concerns for potential side effects, including hyperpermeable vessels, stimulation of tumor growth, abnormal vascular function, hypotension, and hypervascularity (Epstein et al., 2001). Even if the levels could be controlled, many pathologies in need of tissue engineering interventions have altered ability to respond to growth factors. Ischemic conditions lead to tissues that are already rich in VEGF, FGF’s, etc., but these patients do not have a sufficient angiogenic response (Kiefer et al., 2003; Ruel et al., 2003). Will the use of these soluble factors as therapies be able to stimulate a sufficient response in these patients to vascularize an engineered tissue? Genes The genes of many of the proteins described in the previous sections have also been investigated as neovascularization therapies. The goals of these approaches are typically local overexpression of the protein within the engineered tissue. Gene therapies offer an advantage over protein therapies in that protein levels may be increased for a longer period of time, but the duration and levels of expression are difficult to control. Overexpression of a single protein has shown promise in animal models, but clinical results have fallen short of expectations. This may be overcome by overexpression of multiple proteins for a synergistic effect on neovascularization. PlGF has recently been shown to regulate crosstalk between VEGF receptors (Autiero et al., 2003), and when combined with VEGF and PlGF was significantly more potent than each factor alone in an animal model that was refractory to a single protein (Autiero et al., 2003). Combined therapy may be more effective in treatment of diseased or elderly patients who are known to have an impaired response to a single protein. Combined delivery of adenovirus mediated VEGF and Ang-1 have also been shown to promote greater perfusion and vessel stability in muscle flaps than VEGF alone (Lubiatowski et al., 2002).

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As with proteins, it is not clear that gene therapies alone will be successful in diseased patient populations. In addition, the method of gene delivery must be carefully selected to balance the need for maximum transfection efficiency with minimum risk from the vector used for gene delivery. Molecular Modification of Natural Signals Molecular and recombinant techniques can be used to modify the biological properties of naturally occurring proteins. FGF-1 has been engineered for increased mitogenicity for both ECs and SMCs (Xue et al., 2000), to resist thrombin degradation (Erzurum et al., 2003), for heparin independence (Xue et al., 2001; Brewster et al., 2004), for relative specificity for ECs over SMCs (Xue et al., 2001), and to localize to specific tissue microenvironments. A combined approach can be used to engineer FGF-1 with more than one of these properties (Brewster et al., 2004). These novel forms of FGF-1 have unique properties from wild type FGF-1 in vitro and have been used to promote endothelialization of vessels in animal models of cardiovascular disease (Tassiopoulos and Greisler, 2000). It is hypothesized that the in vivo results occur, at least in part, to the ability of these mutants to increase transmural angiogenesis. Studies have shown that at least some of these mutants are more potent than FGF-1 for stimulating neovascularization (Brey, 2003), but further characterization is required in order to determine their therapeutic potential and detailed mechanism of action. As discussed previously, alternative splicing of VEGF plays an important role in controlling normal microvascular network formation. Use of a single isoform is likely to result in abnormal microvascular structure. Zinc transcription factors have been engineered to stimulate synthesis of VEGF (Rebar et al., 2002). The use of transcription factors allows the cell to dictate splicing of the multiple isoforms of VEGF in the correct stoichiometry, resulting in a more normal vasculature than with the injection of a single VEGF splice variant. Vessels in a mouse model formed in response to engineered transcription factor were less leaky than those stimulated by the VEGF165 protein alone. Hypoxia-inducible factor-1α (HIF-1α) is a transcription factor that increases in concentration with reductions in local oxygen levels and activates mediators of angiogenesis, including VEGF. A constitutively active form of HIF-1α was designed by deleting the C-terminal half of HIF-1α and replacing it with the transactivation domain of herpes simplex virus VP16 protein (HIF-1α/VP16) (Vincent et al., 2000). This HIF-1α/VP16 transcription factor retains the DNA binding of HIF-1α, activates VEGF gene expression independent of hypoxia, and promotes an increase in capillary density and maximal blood flow over VEGF DNA injection in a rabbit model of hindlimb ischemia (Vincent et al., 2000). Molecular engineering techniques have the potential to overcome some of the limitations inherent in using naturally occurring growth factors. However, since these proteins may be more potent or longer lasting than naturally occurring factors, the potential of deleterious side effects is even more significant. Techniques will need to be developed to control their local concentrations.

BIOMATERIALS Presentation of Genes and Proteins As discussed previously, gene and protein therapies have shown promise for promoting neovascularization in animal models, but clinical success has been elusive. The high levels and repeated administration required for a therapeutic response cause a concern for deleterious side effects and the formation of abnormal vessels within the target tissue. In addition, it is the local tissue concentration and not the total dose of protein that determines the threshold between normal and pathological microvascular structure (Ozawa et al., 2004). Methods are needed for delivering sustained levels of active angiogenic proteins to target tissues, thus reducing the dose

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required, localizing the delivery, and sustaining the benefit. Both natural and synthetic protein formulations have been developed for the delivery of angiogenic growth factors, with delivery largely governed by diffusion. The following advances attempt to develop biomaterials that release growth factors in response to stimuli in the tissue milieu. Electrostatic interactions between biomaterials and proteins can be used to prolong protein bioavailability. Basic proteins (i.e. FGF-2 or VEGF) encapsulated in acidic gelatin (Ikada and Tabata, 1998) or alginate hydrogels (Gu et al., 2004) will result in formation of ionic complexes between the protein and the material at physiological pH. Diffusion from the materials is then hindered and the material released more slowly. Basic materials would not have strong interactions with the basic proteins leading to rapid protein release via diffusion. If electrostatic interactions are strong enough release will occur only as the material is degraded; however, in most cases the interactions only delay release via diffusion (Ikada and Tabata, 1998). Natural signals occurring in healing tissue can be exploited for covalent attachment of proteins to fibrin gels (Schense and Hubbell, 1999). Chimeric proteins are synthesized of a factor XIIIa substrate attached to a growth factor or other protein. These chimeras covalently incorporate into fibrin during coagulation through the action of transglutimase factor XIIIa. The covalently bound proteins cannot rapidly diffuse out of the fibrin. Instead, they are released as the fibrin is degraded by proteolytic activity within the tissue microenvironment. VEGF121 covalently bound to fibrin gels in this manner exhibited a dose dependent enhancement in EC growth in vitro (Hahn et al., 2006). By extending the availability of VEGF121 in vivo these gels were able to stimulate formation of an organized vascular network, unlike the chaotic vessels formed in response to diffusible VEGF121 (Hayashi et al., 2005). This approach has also been used to stimulate angiogenic signaling through the presentation of ephrin-B2 (Hirschi and Goodell, 2001). Its presentation by modified fibrin gels stimulated angiogenesis in the chick chroioallontoic membrane (CAM) but did not show a relative specificity for venous formation (Hirschi and Goodell, 2001). In addition to modifying fibrin gels, angiogenic signals have been covalently incorporated into polyethylene glycol (PEG) hydrogels (Seliktar et al., 2004). PEG hydrogels containing both VEGF and RGD cell adhesive sequences demonstrate enhanced EC anchorage in vitro. These matrices can be further modified to include peptide sequences that are sensitive to degradation by matrix metalloproteinases (MMPs). These matrices not only support EC growth but are degraded by proteases released from the cells (Seliktar et al., 2004). This combination of cell adhesiveness, growth factor release, and directed degradation presents a novel example of how many design parameters can be incorporated into a single biomaterial. However, the challenge of multiple parameter optimization is not trivial. While RGD sequences can support EC growth, it is not clear how having only a single adhesion sequence will alter cell function and phenotype. It is possible that single amino acid sequences will not be able to substitute for the complex microenvironment that ECs are exposed to in vivo. Gelatin, alginate, modified PEG, and bioactive fibrin hydrogels react to biochemical signals in the tissue milieu to deliver angiogenic proteins. In addition to biochemical stimuli, tissues are subject to a dynamic mechanical environment (Lee et al., 2000). Alginate hydrogels will respond to mechanical signals to increase release of encapsulated protein. VEGF reversibly binds to alginate at physiological pH through electrostatic interactions. In the absence of mechanical forces, VEGF release from the hydrogels was constant, but increased up to 5 times in the presence of cyclic strain. When gels were strained in vivo they stimulated greater angiogenesis than unstrained gels (Lee et al., 2000). It is not clear how well hydrogel release can be tuned to different mechanical environments. This characteristic is most likely not unique to alginate hydrogels as mechanical forces are likely to increase convection, and hence drug delivery, from most hydrogels. The techniques described above assume that the limitation of growth factor therapies is a function of the way it is presented to the tissues. However, this neglects the complex interplay of multiple growth factors that

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may be essential to the formation of a stable microvasculature is neglected. A single growth factor may not be able to initiate the entire neovascularization cascade. Polymeric systems can be synthesized that release two growth factors each with distinct kinetics (Richardson et al., 2001). PDGF-BB encapsulated into poly(lactideco-glycolide) (PLG) microspheres and mixed with particulate polymer and VEGF165 can be processed to form a PLG scaffold. These scaffolds are able to deliver active levels of VEGF165 and PDGF-BB with distinct kinetics. When these scaffolds were implanted subcutaneously in small animal models, incorporation of VEGF alone increased the number of vessels formed, while PDGF-BB alone increased the coverage of vessels by smooth muscle alpha actin (SMA) positive cells (a marker of mural cells). Controlled delivery of VEGF and PDGF-BB resulted in both a more dense and more mature vasculature. Systems for the delivery of two (and possibly more) growth factors with distinct kinetics may prove to improve the therapeutic control of neovascularization. However, what two growth factors should be delivered and with what kinetics? Sequential delivery of VEGF164 and Ang-1 from alginate beads was able to spatially control vascular patterning (Peirce et al., 2004). VEGF164 beads were implanted subcutaneously into the dorsal subcutaneous tissue of rats. After 7 days, the VEGF164 beads were removed and replaced with Ang-1 beads due to its role in vascular stabilization. This combined approach not only increased the density of SMA positive vessels but also resulted in a more persistent angiogenic response than VEGF164 alone. While these advances have improved the options for delivery of angiogenic growth factors, a number of issues remain. The duration and level of delivery will need to be optimized in order to achieve the maximum therapeutic benefit. Research has largely focused on optimizing a single factor in terms of one or two vascular parameters, primarily microvascular density. Clinical success will require more than just an increase in the number of blood vessels. Increased blood flow, normalized oxygen levels, and proper vessel phenotype have to be achieved in the target tissues. It remains to be seen if under optimal delivery conditions a single factor can stimulate a therapeutic improvement.

CELL THERAPIES Mature ECs can be suspended in engineered tissues to increase neovascularization. These cells are thought to mimic vasculogenesis and assemble into capillary structures. A number of different cell types and biomaterials have been used (Nor et al., 1999, 2000, 2001a, b; Polverini et al., 2003), and results have shown that these cells can fuse with invading host vessels, recruit perivascular cells, and establish flow (Nor et al., 2001b). The incorporation of transplanted ECs into new vessels may be enhanced through paracrine interactions with other cell types. ECs suspended in Matrigel survive longer in vivo with the inclusion of fibroblasts (Sieminski et al., 2002). Similarly, ECs and fibroblasts seeded in a tissue-engineered skin equivalent formed tubular structure and inosculated with the host vasculature, while EC-only equivalents did not (Black et al., 1998). ECs seeded on a cultured skin substitute containing both fibroblasts and keratinocytes formed vascular structures that at times became invested with perivascular cells; however, engraftment was not improved over substitutes without ECs (Supp et al., 2002). SMCs combined with ECs seeded on polyglycolic acid (PGA) scaffolds were able to develop into vascularized structures when implanted, but the importance of each cell type to this process was not clear (Park et al., 1999). Support cells can prolong the existence of transplanted vessels (Koike et al., 2004). ECs seeded alone into fibronectin–collagen type I gels and implanted into mice showed little perfusion and regressed from the gels after 60 days. The addition of mesenchymal precursor cells to ECs resulted in the formation of vessels that

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established flow through connections with the mouse circulatory system. The precursor cells that associated with the vessels expressed the mural cell marker SMA and were still functional and stable 4 months after implantation. Transplanted ECs can increase neovascularization in engineered tissues, and there are many examples indicating that these cells inosculate with host vessels. However, the relative contributions of transplanted EC assembly into new vessels versus the release of angiogenic factors from these cells has not been adequately addressed. The growth factor mixture secreted by these cells may be more important than their incorporation into new vessels. A number of other questions exist before the routine use of transplanted ECs is realized. Determination of the optimal cell source, delivery method, and concentration requires further investigation. If autologous cells have to be used, will we be able to isolate a sufficient number of properly functioning ECs from the diseased patient population that is in most dire need of therapeutic neovascularization? Progenitor and Stem Endothelial progenitor cells (EPCs) (Amrani and Port, 2003; Park et al., 2004; Suh et al., 2005) and ECs derived from embryonic stem cells (Levenberg et al., 2002, 2005) may also be used to increase neovascularization in engineered tissues. These cells may have increased proliferative capacity relative to mature ECs and may be less sensitive to the short-term hypoxic conditions in engineered tissues prior to establishing a blood flow. Clinical trials suggest that EPCs may improve the function of ischemic tissues. Injected EPCs selectively localize in ischemic tissues and increases vascular density (Park et al., 2004). However, the mechanism for this increase is not clear. EPCs may develop into new vascular structures or may increase neovascularization indirectly by recruiting monocytes/macrophages that then secrete angiogenic factors (Suh et al., 2005). ECs derived from human embryonic stem (hES) cell lines can also be used to vascularize engineered scaffolds. When seeded in poly(L-lactic acid)/poly(lactic-co-glycolic acid) (PLLA/PLGA) scaffolds and implanted in mice these ECs formed vascular structures and inosculated with host vasculature (Levenberg et al., 2002). In initial studies the vessels persisted for up to 7 days in vivo. In order to formulate an effective stem, or progenitor, cell therapy a consensus must emerge about how these cells are defined and the parameters for their isolation and culture. In addition, circulating EPCs are known to contribute to tumor angiogenesis. Will transplanted EPCs localize to tumors and enhance their growth and metastases? Genetically Modified ECs can be genetically modified ex vivo in order to improve the response to the transplanted cells and increase the transfection efficiency of the gene therapy. Bcl-2 is an anti-apoptosis protein that is upregulated during angiogenesis. ECs transfected to overexpress Bcl-2 transplanted in mice showed increased vascular density over transplantation of ECs alone. EPCs can also be modified to further enhance their therapeutic function. EPCs transfected to express VEGF stimulate a greater improvement in blood flow and angiogenesis in animal models of ischemia than progenitor cells alone (Iwaguro et al., 2002; Ikeda et al., 2004). While they have a longer life than fully differentiated cells, EPCs isolated from adults have reduced telomerase activity and regenerative capacity relative to embryonic stem cells. EPCs isolated from bone marrow and transfected to express telomerase reverse transcriptase (TERT) are more resistant to apoptosis and drastically increased neovascularization in an animal model of limb ischemia (Murasawa et al., 2002). By combining cell and gene therapies these groups were able to achieve a greater therapeutic response.

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VASCULARIZATION IN VITRO Three-Dimensional Cell Culture Methods Under appropriate culture conditions, many types of ECs are able to spontaneously develop into tubular structure when homogeneously suspended on, or in, biomaterials. Matrigel™, collagen, and fibrin have all been used as the structural support for EC network formation in vitro. While the network structures are different from microvascular networks found in vivo, attempts have been made to use these “vascularized” gels as scaffolds for tissue engineering. The hope is that this scaffold would rapidly establish a blood supply by minimizing the steps of EC recruitment, migration, and proliferation. The ECs derived from hES described in previous sections have been used to fabricate a prevascularized muscle tissue (Levenberg et al., 2005). The ECs were seeded with myoblasts and embryonic fibroblasts onto a polymer scaffold and assembled into vascular structure when cultured under appropriate conditions in vitro. The addition of embryonic fibroblasts to the scaffolds greatly increased the lumen size and total area of vessels formed, possibly due to stabilization induced by the fibroblasts. By prevascularizing the scaffolds, the transplanted muscle had increased blood flow and viability following implantation into animal models. Rather than using a homogeneous EC population to vascularize engineered tissues, microvascular fragments containing both ECs and perivascular cells could be used to rapidly establish blood flow in engineered tissues (Shepherd et al., 2004). Fragments suspended in collagen gels form an extensive network of tubes following 5 days in culture. When implanted, the networks established flow and a mature microvasculature bed was present for 28 days. The original fragments were the source of over 80% of the microvasculature in the gels at this time, suggesting that the prevascularized tissue was not replaced by host vessels, but instead fused with the host vessels. Photolithographic Techniques Photolithographic techniques allow cell, biomaterial, or protein patterning with high spatial precision. Unlike the cell culture methods described in the previous section, photolithographic techniques could, in theory, allow generation of patterns that match structures found in vivo. Kaihara et al. used standard photolithographic techniques to create patterned structures with capillary diameters down to 10 μm similar to the branched architecture of microvascular networks (Kaihara et al., 2000). ECs and hepatocytes cultured on these two-dimensional (2D) structures could be lifted as a cell sheet and transplanted in vivo. While the hepatocytes survived upon implantation, it is not clear that the vascular microstructure was maintained. The ability to induce polymerization of hydrogels using light can be exploited to generate patterned hydrogel structures with photolithography (Liu and Bhatia, 2002; Hahn et al., 2005, 2006;). By polymerizing through prefabricated photomasks, PEG hydrogels can be formed with branching networks of channels similar to those found in the microvasculature (Figure 59.2). As with other photolithographic techniques, complex structures can be generated with high spatial precision in 2D. While methods have been developed for the formation of simple two layer patterns (Liu and Bhatia, 2002; Hahn et al., 2005, 2006), it is yet to be seen if these methods could be applied to generate complex three-dimensional (3D) structures. Other microfabrication techniques could also be used to design prevascularized structures. Under defined pattern geometries, ECs can spontaneously develop into capillary-like structures with continuous lumens. When cultured on 10 μm patterns of fibronectin ECs form tubes, while 30 μm channels resulted in EC monolayers (Dike et al., 1999). Another recent technique allowed the patterning of 10–50 μm channels on chitosan that supported EC growth (Wang and Ho, 2004). The 10 and 20 μm channels were never spanned by more than a single cell but did not form lumens. When combined with the fibronectin studies, it suggests that capillary formation is not only pattern but also substrate dependent.

Vascular Assembly

1000 m

600 m

1000 m

500 m

Figure 59.2 PEG hydrogels formed using non-contact photolithography. Gels have a patterned network of channels similar to those found in microvascular networks. (a, b). The photomasks used to generate the patterns and (c, d) the resultant hydrogels. Prevascularizing tissue scaffolds in vitro suffers from the same cell sourcing questions described in previous sections. The optimal number of cells must be isolated, cultured, and the phenotype controlled in order for the vascular networks to function properly. In addition, it is inevitable that any implanted tissue (natural or engineered) will be remodeled upon implantation (Badylak et al., 2002). Many of these approaches focus entirely on ECs, but mural cells play an important for the vessel stabilization that may assist in maintaining the implanted structure. Microfabrication techniques can allow generation of highly complex structures with high spatial precision in 2D. However, extension of these approaches to the 3D geometries necessary for neovascularization will require new innovations. Cell culture models allow 3D network formation but the structure of the vessels is very different from the organized hierarchy of vessels found in vivo.

VASCULARIZATION IN VIVO It may also be possible to take advantage of the patient’s own regenerative capacity to prevascularize engineered tissues in vivo. Scaffolds can be implanted in a highly vascular location and harvested at a later time as a vascularized tissue. The vascularized tissue could then be transferred to a compromised location. This approach in which the body is used as a “bioreactor” (also termed “in vivo tissue engineering” (Daly et al., 2004)) could be used alone to vascularize tissues or as a supplement to cell, molecular, and/or tissue engineering strategies. Implantation of PLLA substrates in the mesentery was able to prevascularize a tissue scaffold, but the rapid ingrowth of fibrovascular tissue resulted in decreased void space so that future cell seeding was not practical (Wake et al., 1994). However, when a collagen–glycosaminoglycan matrix was allowed to vascularize for 10 days on full-thickness wounds it served as a favorable substrate for cultured epithelial autografts (Orgill et al., 1998). Another approach to fabricating vascularized tissue was to fill a polycarbonate chamber with an ECM component and implant it subcutaneously along a vascular pedicle. The pedicle allows de novo vascularization of the ECM and can be transferred along with the scaffold to the recipient site using conventional microsurgical techniques (Cassell et al., 2001). It has recently been shown that a 3D vascularized bone segment of defined geometry can be generated by implanting a chamber containing a cell/matrix/growth factor mixture against the periosteum (Thomson et al.,

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1999; Cheng et al., 2005; Brey et al., 2006). The natural osteogenic and angiogenic potential of the periosteum acts as an in vivo bioreactor to promote neovascularization and bone formation within the chamber. At some optimal time, the chamber and the bone is then transferred along with the periosteum to the recipient location as a vascularized bone flap. The success of this approach is not entirely due to the in vivo environment but requires optimization of the chamber components and the duration of implantation (Thomson et al., 1999; Cheng et al., 2005; Brey et al., 2006). This technique has been successfully applied to the clinical fabrication of a vascularized bone flap for transfer to a mandible defect (Cheng et al., 2006). The approaches discussed here attempt to take advantage of the natural regenerative capacities of the body to assist in neovascularizing tissue. Success of these strategies requires a careful consideration of the balance between parenchymal tissue function and neovascularization. Regardless of the approach, the prevascularized tissue will likely undergo further remodeling upon transfer to the compromised location (Badylak et al., 2002). Successful neovascularization strategies may require combination of a therapy with a surgical technique that maximizes the body’s natural regenerative capacity.

ANALYSIS AND ASSESSMENT In Vitro techniques Initial screens of neovascularization strategies can be performed using in vitro EC proliferation and migration assays. While these are vital steps in vessel formation, they make up only part of the process. The entire process involves multiple steps, including proliferation, migration, and assembly into 3D networks. For this reason, in vitro models where ECs assemble into complex vascular-like networks have become a vital tool in the development and evaluation of neovascularization strategies. These models allow study of the organization of ECs into vessel structures under controlled environmental conditions, and can be classified into three main categories: spontaneous tube formation (i.e. tubulogenesis), organ, and sprouting models (Figure 59.3). Under appropriate culture conditions, many types of ECs are able to spontaneously develop into tubular structure when homogeneously suspended on, or in, 3D gels. Matrigel™, collagen, and fibrin have all been used as the structural support for models of tubulogenesis. During neovascularization, ECs first form sprouts and then establish a hollow lumen, but in some tubulogenesis models the morphological steps of angiogenesis occur in reverse order, with lumen formation followed by organization into tubes (Davis et al., 2002). In addition, tubulogenesis models mimic vasculogenesis more closely than angiogenesis. A commonly used in vitro organ model is the rat aortic ring assay. A segment of either rat or mouse aorta is dissected and embedded in an ECM gel (Nicosia and Ottinetti, 1990). Upon implantation, sprouts grow

Figure 59.3 (a) Tubulogenesis model of vasculogenesis. ECs cultured on Matrigel will develop a network of tubes in response to angiogenic factors. (b) A sprouting model of angiogenesis. A capillary network forms when an EC aggregate is implanted in a 3D fibrin gel and stimulated by growth factors.

Vascular Assembly

from the ring in a manner analogous to sprouting angiogenesis. Both ECs and mural cells contribute to these sprouts, a phenomenon that also occurs in vivo. Although the formation of sprouting vessels containing both ECs and mural cells more closely approximates angiogenesis than tubulogenesis models, mechanistic information is limited as angiogenesis in vivo is primarily a microvascular event. Moreover, it is difficult to isolate the contribution of each cell type to the formation of branched networks due to the multiple cell types present in the organ explants. Sprouting models allow the study of angiogenesis in a defined cellular and ECM environment (Vernon and Sage, 1999; Korff et al., 2001; Uriel et al., 2006). Aggregates of ECs sandwiched in an ECM gel form a branched network of tubes in response to soluble or mechanical factors. This model recapitulates many of the steps of angiogenesis in their natural sequence: vessel sprouting, lumen formation, and development of a branched network of tubes. The most controlled model to date for studying angiogenesis consists of suspending a defined number of both ECs and SMCs as a co-culture spheroid in collagen gels (Korff et al., 2001). The cells form an aggregate consisting of a surface layer of ECs, with underlying SMCs, and allowed study of the interactions between ECs and SMCs and their roles in angiogenesis and vessel stabilization. Under precise EC:SMC ratios the aggregates were quiescent and resistant to low levels of VEGF, while high levels of VEGF stimulated the formation of vessel sprouts. Although it is not clear to what extent spatially or mechanistically that SMCs contributed, ECs and SMCs were both present during sprout formation. These models continue to be vital tools in the development and screening of neovascularization strategies. Some commonly used ECMs, such as Matrigel, can stimulate capillary sprout formation by cells other than ECs, so these events are not distinctly angiogenic. The biochemical conditions of matrix assembly and polymerization need to be carefully controlled in order to maximize the benefit from these models. While these models should be used, caution must be advised in interpreting data obtained from purely in vitro models. Data obtained should be used as a starting point and expanded upon in vivo. Animal Models The in vitro tests allow evaluation of neovascularization therapies under controlled conditions but are a simplification of the actual in vivo process. A number of models are used to assess these therapies under in vivo conditions. The CAM assay is easy to use and allows non-invasive assessment of neovascularization (Zisch et al., 2003; Ehrbar et al., 2004). Materials containing neovascularization factors are grafted onto the CAM typically between embryonic days 7–9. Neovascularization can then be assessed by counting vessels that grow into the grafted material. While easy to use, there is significant inherent variation in this model and the baseline vasculature in the CAM makes quantitation difficult. In addition, the CAM is an embryonic tissue and may not be appropriate for studying adult neovascularization processes. As initial tests in adult animals, engineered tissues are often implanted subcutaneously in the dorsal flanks of mice or rats and harvested at various time points (Nor et al., 1999, 2000, 2001a, b; Shea et al., 1999; Richardson et al., 2001; Polverini et al., 2003; Brey et al., 2004). Histological and immunohistochemical methods are then used to assess the presence and structure of vessels that form within the implanted tissues. Subcutaneous implants are surgically simple and multiple implants can be placed in a single animal but are limited by the inability to non-invasively monitor vessel formation. Corneal implant and window chamber models were developed to allow non-invasive monitoring of blood vessel formation. In the cornea model scaffolds are implanted into a pocket created in the avascular cornea (Staton et al., 2004). The absence of vessels in the cornea makes determination of new vessels straightforward and its location is easily accessible for non-invasive monitoring of the vasculature (Hayashi et al., 2005). Window chambers are prepared by removing a layer of tissue (the dorsal skin is most common, but ears and skull models have also been used) to expose the underlying tissue. Therapies can then be implanted and the exposed tissue covered with glass that is secured in place. The windows allow continuous monitoring of blood vessel

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formation and can be combined with advanced imaging techniques to perform rigorous quantitative analysis of microvascular structure and blood flow (Brown et al., 2001; Abdul-Karim et al., 2003). However, the window chambers are technically challenging and may introduce a high level of baseline neovascularization. While models are designed for ease of analysis they do not reflect an actual clinical situation. It has been suggested from studies of tumor angiogenesis that the mechanism and process of blood vessel formation is dependent on the tissue microenvironment (Fidler, 2002). In order to accurately assess vascular assembly strategies, the ultimate analysis should be formed in the animal model and tissue location most relevant to the clinical situation where they will be applied. While these techniques may preclude the ability to continuously monitor blood vessel formation they will provide the most important pre-clinical data. Quantitative Techniques Quantification of results of an intervention from animal studies are often over-simplified. Proper microvascular structure and vessel phenotype are essential to meeting a tissue’s specific metabolic needs. Techniques that allow 3D quantitative imaging of microvascular structure from biological and engineered tissue samples can provide a unique ability to assess outcomes. Image processing techniques can be combined with serially immunostained tissue sections for unique insight into microvascular structure (Figure 59.4) (Brey et al., 2002). This technique can image large volumes of tissue, allows imaging of both developing and established microvasculature, and has been used to study the 3D interactions between multiple cell types during neovascularization (Brey et al., 2004). Although this imaging method is invasive, it provides unique, quantitative images of 3D cellular interactions from large volume tissue samples. While using serial tissue sections provides unique cellular detail, it can be time-consuming and the microstructure can be distorted by the sectioning process. Micro-computed tomography (micro-CT) can be used for imaging microvascular structure. Resolution with this technique can reach as high 10 μm with a conventional source, but can be as high as 1 μm when using a synchrotron source (Bentley et al., 2002). However, synchotrons are not commonly available and are extremely expensive. Micro-CT requires that samples are injected with a contrast agent and harvested in order to obtain resolution down to the microcirculatory level.

Figure 59.4 3D image of microvascular networks within a model tissue-engineered construct. The developing vessels (red) are seen growing into a fibrin network (white).

Vascular Assembly

At best, these techniques produce 3D quantitative images of patent microvasculature. They do not allow detail that distinguishes between cell types nor allow imaging of vessels that have yet to establish a blood supply. Only immunohistochemistry based techniques allow imaging of both existing and developing vessels. In addition to serial tissue techniques, laser scanning confocal microscopy (LSCM) also offers a method for imaging immunostained tissues. While limited to imaging tissues of a certain thickness, LSCM offers the distinct advantage of optical instead of physical sectioning, which better preserves the vessel architecture. LSCM has been applied extensively the study of microvascular structure in tumors (Brown et al., 2001). Progress in our quantitative understanding of how vessels form and respond to therapies continues to be an important aspect of neovascularization studies. Advances in technology are constantly being made but more are needed. Ideally, techniques would be available that allow non-invasive quantification of both the structure and function of the microvasculature down to the smallest newly formed vessels. With the current tools, something is always compromised. The invasive techniques required to achieve capillary resolution prohibit our ability to monitor changes in microvascular structure over time, while non-invasive methods lack the resolution and the detail essential for comprehensive analysis.

CONCLUSIONS Clinical success of tissue engineering requires the ability to rapidly establish a blood supply within the tissues. Recent progress has shown promise in this field, but a number of issues still need to be addressed. Therapies need to be designed and evaluated for their ability to stimulate neovascularization in adverse healing conditions. Diabetic, elderly, and hypertensive people are just a few within the potential patient population who may have decreased sensitivity to neovascularization therapies. Neovascularization must be controlled both spatially and temporally. As therapeutics become more potent and sustainable it will be vital to localize the effect through novel drug delivery methods or targeted systems. Finally, the microvasculature formed must meet the specific needs of the target tissue. The microvasculature is tremendously diverse with distinct functional and geometric characteristics in each tissue bed which are vital to proper function. Overcoming these issues will bring us closer to the routine application of engineered tissues.

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60 Cardiac Tissue Milica Radisic and Michael V. Sefton

INTRODUCTION: FROM TISSUES TO ORGANS: KEY GOALS AND ISSUES Nearly 8 million people in the United States have suffered from myocardial infarction, with 800,000 new cases occurring each year (American Heart Association, 2004). Myocardial infarction results in the substantial death of cardiomyocytes in the infarct zone followed by pathological remodeling of the heart. The remodeling process involves cardiac dilation, wall thinning and severe deterioration of contractile function leading to congestive heart failure in more than 500,000 patients in the United States each year (American Heart Association, 2002). Conventional therapies are limited by the inability of myocardium to regenerate after injury (Soonpaa and Field, 1998) and the shortage of organs available for transplantation. This chapter will focus on describing cell and tissue-based therapies that have been considered as novel treatment options (Reinlib and Field, 2000). Regardless of the approach to regenerative medicine or the scope of the application (a vascular graft, a pediatric valve or an entire heart) there are three overlapping therapeutic goals – the three R’s:

• • •

Make tissue and organ replacement safer, more effective and more widely available. Repair tissues and organs without having to replace them. Enable tissues and organs to regenerate so that repair and regeneration become one and the same.

Furthermore, the problems of reaching these goals can be summarized (Table 60.1) in three categories (here largely in the context of tissue engineering) (Sefton, 2002; Sefton et al., 2005):

• • •

Cell number: What is the source of cells to be used and how will large numbers be generated? How will they

be supplied with nutrients and oxygen (and have wastes removed) within a device of reasonable volume? Cell function: How will the scaffold, extracellular matrix, and diffusible factors interact to generate the desired cell phenotype? How will the engineered tissue/organ function integrate with the host to ensure a functional outcome? Cell durability: What will happen over the long term as remodeling and/or the host immune/inflammatory system responds to the new tissue?

In order to replace, repair, or regenerate cardiovascular tissue, these central issues of regenerative medicine will need to be addressed. Some of these issues (Table 60.1) reflect the fundamental nature of how an organ is different from a tissue: the large size and 3-dimensional (3D) structure and the presence of multiple cell types that work in unison. Beyond these largely scientific challenges, there are the no less critical, practical questions of manufacturing, sterilization, storage and distribution, and the regulatory and public policy issues that will need to be addressed before such therapies can be made available to the patients who are expected to benefit. Furthermore we will also need new imaging or other non-invasive strategies to monitor the success (or not) of these therapies (i.e. to enable the translation into clinical practice).

1038

Table 60.1 Critical issues associated with tissue engineering a heart (with permission from Sabiston and Spencer, Sefton et al., 2005)

Cell number Function

Durability

• • • • a

Objective

Critical issues

• • • • • •

300 g of cells (3  1011 cells) 200 mL O2/ha

• • •

Fatigue resistance Hypoxia and disease tolerance Host tolerance

• • • • • • • • • •

Cellular phenotype (multiple cell types) Co-ordinated muscle contraction Pump blood Connect to circulation

Cell source/purity Vascularization Microenvironment (soluble and insoluble factors) Pacemaker and electrical conduction Valves and conduits Biomechanical elasticity and strength Non-thrombogenicity Biocompatibility Remodeling Innate/adaptive immune response

Manufacturing and quality control Ethical, legal and social issues Imaging and non-invasive diagnostics Regulatory and public policy issues Based on moderate activity, (Burton, 1972)

CELL AND GENE THERAPY Cell Therapy Treatment options for heart failure and myocardial infraction (MI) are limited by the inability of adult cardiomyocytes to proliferate and regenerate injured myocardium. Cell injection, has thus emerged as an alternative treatment option. In animal models, injection of fetal or neonatal cardiomyocytes improved left ventricular (LV) function and ventricle thickness, thus attenuating pathological ventricular remodeling (Reinecke et al., 1999; Muller-Ehmsen et al., 2002a, b). Differentiated cardiomyocytes are indeed an ideal cell source for injection or tissue engineering, since they contain a developed contractile apparatus and can integrate through gap junctions and intercalated disks with the host cardiomyocytes. However, large numbers of clinically relevant autologous cardiomyocytes are unavailable. In searching for an appropriate cell source (Table 60.2), regeneration of infarcted myocardium has been attempted in animal models by transplantation of skeletal myoblasts (Dorfman et al., 1998), as well as cardiomyocytes derived from embryonic stem (ES) cells (Klug et al., 1996) and bone-marrow-derived mesenchymal stem cells (MSCs) (Toma et al., 2002b). For a review of cell therapy approaches see Laflamme and Murry (2005). The obvious advantage of skeletal myoblasts is that they can be harvested from the patient and expanded in vitro. However, mature skeletal myoblasts do not express gap junctional proteins, thus they are incapable of functionally integrating with the host myocardium. This was the most likely reason for the occurrence of arrhythmias in four out of ten patients in a Phase 1 clinical trial of autologous skeletal myoblast transplantation (Menasche et al., 2003). For further information on myoblast clinical trails see Laflamme and Murry (2005). Hematopoietic stem (HS) cells from bone marrow were tested in their ability to contribute to the regeneration of infarcted myocardium. The general consensus on the effect of injection or the mechanism of action has not been reached yet. Anversa and colleagues (Orlic et al., 2001) reported that HS cells injected into the

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Table 60.2 Cell sources for cardiac tissue engineering, and some of their advantages and disadvantages (with permission, Sefton et al., 2005) Cell sources

Advantages

Disadvantages

Adult cardiac cells

Target cell source

Fetal cardiac cells

Some proliferative potential, appropriate developmental potential; demonstrated efficacy Some proliferative potential; may elicit in vivo healing through indirect mechanisms

Little proliferative or developmental potential, limited resource Limited resource; ethical considerations

Endothelial progenitor cells

Adult bone-marrowderived cells

Significant in vitro proliferative potential; some demonstration of efficacy

ES cells

Significant in vitro proliferative potential; demonstration of efficacy; appropriate developmental potential; sustainable resource

Appropriate developmental potential yet to be demonstrated; may not be appropriate for larger tissue replacement or in vitro tissue engineering Appropriate developmental potential to be demonstrated; safety tolerance after in vitro culture to be determined In vitro culture may introduce genetic changes; safety tolerance after in vitro culture and differentiation to be determined

peri-infarct zone in mice with acute MI gave rise to cardiomyocytes regenerating 68% of the infarct. These results could not be reproduced by other groups (Balsam et al., 2004; Murry et al., 2004). Instead the studies suggest that HS cells differentiate into blood cells (Murry et al., 2004; Nygren et al., 2004), and occasionally fuse with host cardiomyocytes. The discrepancy may lie in the different techniques used. Bone marrow MSCs have also been considered as a cell source for myocardial repair. When injected directly into the hearts of mice (Toma et al., 2002a) and pigs (Shake et al., 2002) post-infarction, the cells attenuated pathological ventricle remodeling and expressed cardiac markers. Contribution of cell fusion to these events remains to be determined. Bone marrow mononuclear cells (BMNCs) (consisting of both HS and MSC) were evaluated in clinical trials (for a review see Dimmeler et al., 2005). In general, the initial clinical studies indicate that bone marrow transplantation is safe and contributed to the increase in ejection fraction (Chen et al., 2004; Wollert et al., 2004) although the mechanism of the effect is unclear. The main advantage of bone marrow as a cell source is that it can be harvested from the patient; however, the frequency of stem cells is generally low (0.1%). Recent emerging work suggests that the heart may contain resident progenitor cells. This is an exciting possibility, as resident progenitors may be an ideal source of autologous cardiomyocytes. However, it appears that there is more than one heart cell subpopulation that fits the description of a cardiac progenitor. C-kit cells isolated from adult rat hearts and expanded under limited dilution gave rise to cardiomyocytes, smooth muscle, and endothelial cells (ECs) when injected into ischemic myocardium (Beltrami et al., 2003). Oh et al. (2003) reported Sca-1 as a marker of resident cardiac progenitors, and expression of cardiac markers upon treatment with 5-azacytidine. LIM homeodomain islet 1 transcription factor (isl1) was also identified as a marker of resident cardiac progenitor cells (Laugwitz et al., 2005). The isl1 cells from mouse hearts were propagated in culture and they differentiated into functional cardiac myocytes when in contact with terminally differentiated cardiomyocytes. It remains to be determined if the progenitors, regardless of their marker, can be isolated from adult human biopsies and if sufficient numbers of cardiomyocytes (108 cells/patient) can be generated in vitro.

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ES cells have enormous proliferative potential, and in combination with nuclear transfer can generate autologous cells. However, the main technical concern in utilization of ES cells is that the presence of a single undifferentiated cell in vivo can potentially yield teratomas (Laflamme and Murry, 2005). Highly pure populations of cardiomyocytes (99.6%) can be generated using a neomyocin-resistant transgene driven by a cardiac marker promoter (Klug et al., 1996; Zandstra et al., 2003). Upon injection into hearts, the ES cell-derived cardiomyocytes formed stable intracardiac grafts (Klug et al., 1996) and improved contractile function (Etzion et al., 2001). Electromechanical integration of the cardiomyocytes derived from human ES cells with the host myocardium was also reported (Kehat et al., 2001). Besides focusing on restoration of contractile function through injection of myogenic cells, regeneration of infarcted myocardium has also been attempted through injection of EC progenitors (Kocher et al., 2001). The regeneration is based on the improvements in infarct neovasculature that lead to improved perfusion and ultimately improved LV function. In most cases described above, the cells were suspended in an appropriate liquid (saline or culture medium) followed by intramyocardial or coronary injection. The main challenges associated with this procedure are poor survival of the injected cells (Muller-Ehmsen et al., 2002b) and washout from the injection site (Reffelmann and Kloner, 2003). According to some estimates, 90% of the cell delivered through a needle leak out of the injection site (Muller-Ehmsen et al., 2002a, b). In addition, significant number of cells (90%) die within days after injection (Zhang et al., 2001; Muller-Ehmsen et al., 2002b). Thus developing improved delivery and localization methods (e.g. hydrogels) and effective anti-death strategies (e.g. heat shock treatment) could significantly improve effectiveness of cell injection procedures. Gene therapy Gene therapy approaches are based on either delivering exogenous genes capable of expressing therapeutic proteins or on the blockade of genes involved in pathological process. The genes can be delivered using nonviral vectors (such as naked plasmids, liposome formulation, and synthetic peptides) or recombinant viruses. Replication defective recombinant viruses are significantly more effective in gene transfer to myocardium compared to the non-viral vectors that are limited by high degradation rate and low genomic integration (Melo et al., 2004a). However, viruses sometimes lead to immune reaction, and there is a small risk that they may become proliferative. In an early work aimed at converting the non-contractile scar tissue into tissue capable of contraction, Murry et al. (1996) used adenovirus to transfer MyoD, a myogenic determination gene, into granulation tissue of rat myocardium post-infraction. In vitro, gene transfer converted fibroblasts into skeletal muscle cells. Similar results (i.e. expression of MyoD, myogenin, and embryonic isoform of myosin heavy chain) were observed in vivo after transfection with high doses of virus (1010 pfu). Restoration of contractile function has also been attempted by normalization of β-adregenic receptor signaling. In rabbits, intracoronary delivery of β2-adregenic receptor gene led to improvements in LV and hemodynamic function (Maurice et al., 1999). Using similar approach, the β-adregenic receptor signaling was rescued in ventricular myocytes from patients with heart failure. Calcium signaling was another target for gene therapy aimed at restoration of contractile function (review in Hajjar et al., 2000). Intracoronary delivery of SERCA2a genes in a rat model of heart failure improved long-term survival, restored systolic and diastolic function, and improved Ca2 ATP-ase activity (del Monte et al., 2001). Antisense inhibition of phospholamban was shown to improve contractility of cardiomyocytes from end-stage heart failure patients (del Monte et al., 2002). Gene therapies for acute MI were limited by the available delivery techniques. In general, the time it takes for transcription and translation is too long for a successful intervention in acute MI (Melo et al., 2004b). However, individuals at risk may benefit from preventive strategies that protect from ischemia/reperfusion

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injury. In that respect, overexpression of antioxidant enzyme systems (HO-1), heat shock proteins and survival genes (Bcl-2 Akt) was demonstrated to be beneficial in small animal models (Melo et al., 2004b). Most recently, a novel gene therapy approach was reported for treatment of acute MI and chronic ischemia. Intramyocardial injection of naked DNA encoding human sonic hedgehog preserved LV function, enhanced neovascularization, and reduced fibrosis and cardiac apoptosis. Sonic hedgehog is a morphogen and a crucial regulator of organ development during embryogenesis, thus transient reconstruction of embryonic signaling had beneficial effect on tissue repair and neovascularization (Kusano et al., 2005). Gene therapy was also utilized to treat ischeamia in patients with coronary artery disease who were not eligible for standard treatment options such as percutaneous angioblasty or surgical vascularization. A number of pre-clinical and clinical trails focused on overexpression of vascular endothelial growth factor (VEGF), fibroblast growth factor (FGF) and hepatocyte growth factor in an attempt to improve collateral blood vessel formation (Melo et al., 2004b). Although functional improvements were reported in large animals, phase II and III clinical trials failed to conclusively prove efficacy and the long-term therapeutic effect (Yla-Herttuala et al., 2004; Markkanen et al., 2005). Although the safety record was excellent in all of the trails, the following reasons were considered as possible causes for the disappointing results in efficacy: a wrong dose, a less-than-optimal route of administration, an inefficient delivery system, an insufficient duration of the treatment, selection of an appropriate animal model in pre-clinical trials as well as selection of an appropriate patient group. While all of the above-mentioned limitations are technical in nature, targeting a single gene, as most commonly used in gene therapy, may have conceptual limitations as well. Most pathological processes are complex and involve expression or down-regulation of multiple genes. In many instances this genetic complexity is not well understood and thus it is difficult to predict a prior what the ultimate effect of overexpression or blockade of a single gene will be. In this respect combination of gene and cell therapy may be a preferred approach in the treatment of heart diseases. One of the major limitations of cell therapy approaches is low cell survival. Thus, transfecting the injected cells with agents that enhance angiogenesis or cell survival may benefit the cell injection procedure. Once in the appropriate location, the cells may contribute to the contractile function and adjust appropriately to the complex physiological stimuli of the local milieu. Li and colleagues demonstrated that injection of VEGF165 transfected cardiomyocytes into cryoinjured rat myocardium sustained VEGF expression and increased capillary density in the border zone as well as regional blood flow within the scar (Sakai et al., 2001). Most other studies focused on the injection of cardiomyocytes expressing growth factors (for review see Fazel et al., 2005) consistently reported that a combination of cell and gene therapy results in improved angiogenesis and functional properties in comparison to cell therapy alone.

SCAFFOLD-BASED APPROACHES While small infarcts may be treated with cell therapy, larger areas of damaged tissue will require excision and replacement with a cardiac patch. The time post-infarction is critical in the success of any regeneration strategy. Upon myocardial infarction, a vigorous inflammatory response is elicited and dead cells are removed by marrow-derived macrophages. Over the subsequent weeks to months, fibroblasts and ECs proliferate forming granulation tissue and ultimately dense collagenous scar. Formation of scar tissue severely reduces contractile function of the myocardium and leads to ventricle wall thinning and dilatation, remodeling, and ultimately heart failure. The best regeneration strategy thus depends on the time post-infarction, that is, new and old infarcts most likely cannot be treated using the same approach. Cell injection strategies will work best if applied shortly after MI. Application of cells and growth factors within hours and days after MI has a potential of directing the wound repair process so that the minimum amount of scar tissue is formed, the contractile function is maintained in the border zone, and pathological

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remodeling is attenuated. Tissue engineering strategies will work in the acute phase as well, but may be more necessary after scar has formed. In that case, larger areas of heart must be replaced or augmented and this is potentially where a scaffold-based approach may be most useful. Cell-free cardiac patches Patients with large transmural akinetic scars often benefit from the Dor procedure (endoventricular circular patch plasty) (Dor et al., 1989; Di Donato et al., 1997). In some cases, however, the success of this procedure is temporary, thus motivating the need for viable tissue patches. In this procedure the scar tissue is excised and the ventricle is closed using a circular Dacron (polyethylene terephthalate) patch lined with endocardium. Another strategy to address pathological remodeling and prevent heart failure is a CorCap cardiac support device. CorCap is an implant-grade polyethylene terephthalate mesh that is wrapped around the heart ventricle to prevent further dilatation and support contractile function. In clinical trials, it was demonstrated that it results in improved quality of life, as well as improved heart size and shape (Starling and Jessup, 2004). Cell-based cardiac patches Self-assembly: In cardiac tissue engineering approaches, most studies suggest that some type of scaffold, an inductive 3D matrix, is necessary to support assembly of cardiac tissue in vitro. An important scaffold-free approach includes stacking of confluent monolayers of cardiomyocytes (Shimizu et al., 2002). Although cardiac patches obtained in this way generate high active force, engineering patches more than 2–3 cell layer thick remains a problem. Recently, 24-mm long and 100-μm thick contractile cardiac organoids were fabricated by self-organization (Baar et al., 2005). Cardiomyocytes were cultivated on a poly(dimethylsiloxane) (PDMS) surface coated with laminin. As laminin degraded, the confluent monolayer detached from the periphery of the substrate moving towards the center and wrapping around a string placed in the center of the plate until a cylindrical contractile organoid was formed. The scaffold approaches can be divided into: (i) hydrogel approaches where cells are either encapsulated and cultivated in vitro or injected directly into MI without pre-culture and (ii) porous and fibrous 3D scaffold approaches where scaffolds are seeded with cells and in most cases cultivated in vitro prior to the utilization as cardiac patches. Hydrogels The most important example of hydrogel-based cardiac tissue engineering includes the work of Eschenhagen and colleagues. Cardiomyocytes were cast in growth factor supplemented collagen gels and cultivated in the presence of cyclic mechanical stretch (Eschenhagen et al., 1997; Fink et al., 2000; Zimmermann et al., 2000; Zimmermann et al., 2002a, b). The main advantage of the hydrogel approach is the higher active force generated by such cardiac tissues, compared to the force generated by tissues on porous or fibrous 3D scaffolds. In addition, collagen and laminin are the main components of the myocardial extracellular matrix, thus they are supportive of cardiomyocyte attachment and elongation. However, the main remaining challenge is tailoring the shape and dimensions of such tissues. One interesting approach to address this issue is the use of extruded collagen type I tubes (Yost et al., 2004). A technique that can potentially combine the advantages of the hydrogel approach with ease in tailoring tissue shape and size is inkjet printing. Cardiac constructs based on feline cardiomyocytes were created by printing cell solution onto alginate and using calcium as a cross-linking agent. This approach may be particularly useful for co-culture (Tao et al., 2004) as it enables precise control over cell location in the tissue construct. Without pre-culture, hydrogels were utilized to provide structural stability and deliver cells for regeneration of infarcted myocardium. Various cell types were injected into myocardium using a biomaterial that

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crosslinks upon injection such as Matrigel (Balsam et al., 2004), fibrin glue (Christman et al., 2004a, b; Ryu et al., 2005), or self-assembled peptide hydrogels (Davis et al., 2005). In general, the studies report prevention of ventricle dilatation and improvement of fractional shortening as well as improved angiogenesis. Kofidis et al. (Balsam et al., 2004) reported that injection of Matrigel or Matrigel and ES cells into infarcted rat hearts resulted in structural stabilization, prevented ventricular wall thinning and improved fractional shortening. Chirsman et al. (2004a, b) demonstrated that injection of skeletal myoblasts into myocardial infarcts using fibrin glue increased cell localization within the infarct after 5 weeks, reduced infarct size and increased vascularization of the scar without causing a significant inflammatory response or foreign body reaction. Similarly, Ryu et al. (2005) found that injection of BMNCs into cryoinjured rat myocardium using fibrin matrix increased the amount of viable tissue and microvessel formation and reduced the amount of fibrous tissue in comparison to the injection of BMNC in culture medium or culture medium alone. Recently, it was demonstrated that a synthetic material, self-assembling peptide hydrogel, can also be utilized for cell injection into the myocardium (Davis et al., 2005). Upon injection, the peptide formed a nanofibrous structure that promoted recruitment of endogenous cells expressing endothelial markers and supported survival of injected cardiomyocytes. Porous scaffolds Three-dimensional cardiac tissue constructs were successfully cultivated in dishes using a variety of scaffolds amongst which collagen sponges were the most common. In the pioneering approach of Li et al. (1999), fetal rat ventricular cardiac myocytes were expanded after isolation, inoculated into collagen sponges and cultivated in static dishes for up to 4 weeks. The cells proliferated with time in culture and expressed multiple sarcomeres. Adult human ventricular cells were used in a similar system, although they exhibited no proliferation (Li et al., 2000) Fetal cardiac cells were also cultivated on porous alginate scaffolds in static 96-well plates. After 4 days in culture the cells formed spontaneously beating aggregates in the scaffold pores (Leor et al., 2000). Cell seeding densities of the order of 108 cells/cm3 were achieved in the alginate scaffolds using centrifugal forces during seeding (Dar et al., 2002). Neonatal rat cardiomyocytes formed spontaneously contracting constructs when inoculated in collagen sponges (Tissue Fleece) within 36 h after seeding (Kofidis et al., 2003) and maintained their activity for up to 12 weeks. The contractile force increased upon addition of Ca2 and epinephrine. Fibrous scaffolds In a classical tissue engineering approach, fibrous polyglycolic acid (PGA) (Figure 60.1a) scaffolds were combined with neonatal rat cardiomyocytes and cultivated in spinner flasks and rotating vessels (Carrier et al., 1999). The scaffold was 97% porous and consisted of non-woven PGA fibers 14 μm in diameter. This material has advantages from a clinical stand point since it is FDA approved and found in biodegradable sutures. Neonatal rat or embryonic chick ventricular myocytes were seeded onto (PGA) scaffolds by placing a dilute cell suspension in the spinner flasks and mixing for 3 days (50 rpm) (Carrier et al., 1999). Mixing in the spinner flasks (0, 50, or 90 rpm) had a significant effect on the construct metabolism and cellularity. Constructs cultivated in well mixed flasks had significantly higher cellularity index and metabolic activity compared to the constructs cultivated in the static flasks. After 1 week of culture, constructs seeded with neonatal heart cells contained a peripheral tissue-like region (50–70 μm thick) in which cells stained positive for tropomyosin and organized in multiple layers in a 3-D configuration (Bursac et al., 1999) (Figure 60.1a and b). Electrophysiological studies conducted using a linear array of extracellular electrodes showed that the peripheral layer of the constructs exhibited relatively homogeneous electrical properties and sustained macroscopically continuous impulse propagation on a centimeter size scale (Bursac et al., 1999). Constructs based on the cardiomyocytes enriched by preplating exhibited lower excitation threshold (ET), higher

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Figure 60.1 Representative scaffolds used in cardiac tissue engineering. (a) Scanning electron micrograph of a non-woven fibrous PGA scaffold used in a classical approach by Freed and colleagues. (b) Immunohistochemical staining for tropomyosin in constructs based on surface-hydrolyzed PGA seeded with neonatal rat cardiomyocytes and cultivated in rotating vessels for 1 week (with permission from Papadaki et al., 2001). (c) Scanning electron micrograph of a fibrous PLA scaffold obtained by electrospinning followed by uniaxial stretching. (d) Neonatal rat cardiomyocytes cultured on oriented PLA scaffolds exhibited well-developed contractile apparatus (actin – green) (with permission from Bui et al., 2005). (e) Thin polylactide-co-glycolic acid (PLGA) films patterned with laminin using microcontact printing (inset: 15 μm laminin lanes spaced 20 μm apart) and seeded with neonatal rat cardiomyocytes (actin filaments – red, nuclei – blue). (f) Immunohistochemical staining illustrates elements of intercalated disks (N-cadherin-yellow, actin filaments-red) (with permission from McDevitt et al., 2002). (g) Scanning electron micrograph of the knitted Hylonect fabric; arrow indicates the direction of cyclic stretch applied during culture. (h) Cross-section of a construct sampled 2 h after cell seeding, showing the multifilament yarn and immunohistochemical staining for cardiac troponin I. Neonatal rat cardiomyocytes were inoculated into the scaffold using fibrin (with permission from Boublik et al., 2005). (i) Parallel channel array bored in the PGS scaffolds using CO2 laser/scanning engraving system. (j) Neonatal rat heart cells seeded onto channeled PGS scaffolds using Matrigel™ and cultivated in perfusion with 5.4 vol% perfluorocarbon emulsion supplemented culture medium (vimentin stained fibroblasts – red, troponin I stained cardiomyocytes – green, nuclei – blue) (with permission from Radisic et al., 2006).

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conduction velocity, higher maximum capture rate (MCR), and higher maximum and average amplitude of contraction. Laminar flow conditions in rotating bioreactors further improved the PGA-based constructs. The cells in the peripheral layer expressed tropomyosin and had spatial distribution of connexin-43 comparable to the neonatal rat ventricle. The expression levels of cardiac proteins connexin-43, creatine kinase-MM and sarcomeric myosin heavy chain were lower in rotating bioreactors cultivated constructs compared to the neonatal rat ventricle but higher than in the spinner flask cultivated constructs (Papadaki et al., 2001). It is important to note that in both spinner flasks and rotating bioreactors the center of the constructs was mostly acellular due to the oxygen diffusional limitations. Recently, electrospun scaffolds (Figure 60.1c) have gained significant attention as they enable control over structure at sub-micron levels as well as control over mechanical properties, both of which are important for cell attachment and contractile function. Entcheva and colleagues (Zong et al., 2005) used electrospinning to fabricate oriented biodegradable non-woven poly(lactide) (PLA) scaffolds. Neonatal rat cardiomyocytes cultivated on oriented PLA matrices had remarkably well-developed contractile apparatus (Figure 60.1d) and exhibited electrical activity. Thin Films A significant step forward toward a clinically useful cardiac patch was the cultivation of ES cell-derived cardiomyocytes on thin polyurethane films. Cells exhibited cardiac markers (actinin) and were capable of synchronous macroscopic contractions (Alperin et al., 2005). The orientation and cell phenotype could further be improved by microcontact printing of extracellular matrix components (e.g. laminin) as demonstrated for neonatal rat cardiomyocytes cultivated on thin polyurethane and PLA films (Figure 60.1e and f) (McDevitt et al., 2002, 2003). Combination Approaches To combine the benefits of the presence of naturally occurring extracellular matrix (laminin) and the stability of porous scaffolds, neonatal rat cardiomyocytes were inoculated into collagen sponges or synthetic poly(glycerol sebacate) scaffolds (PGS) using Matrigel (Radisic et al., 2006). The main advantage of a collagen sponge is that it supports cell attachment and differentiation. However, the scaffold tends to swell when placed in culture medium, thus creation of a parallel channel array resembling a capillary network is difficult. For that purpose a novel biodegradable elastomer (Wang et al., 2002) with high degree of flexibility was used (Figure 60.1i and j). Freed and colleagues have recently reported that mechanical stimulation of hybrid cardiac grafts is based on knitted hyaluronic-acid-based fabric and fibrin (Boublik et al., 2005) (Figure 60.1g and h). The grafts exhibited mechanical properties comparable to those of native neonatal rat hearts. In a subcutaneous rat implantation model the constructs exhibited the presence of cardiomyocytes and blood vessel ingrowth after 3 weeks.

TISSUE AND ORGAN FUNCTION Successful implantation of engineered tissues requires both maintenance of cellular phenotype and the functional integration of the construct within the host tissue. As progress is made from the state of the art described above to the final goal, it will be necessary to ensure that engineered cardiac cells and tissue not only contract in unison with the surrounding native myocardium to produce the desired force but also that the biograft is electrically integrated with the host to prevent arrhythmogenesis. Underlying such integration and the implicit control of the construct phenotype is the creation of the arborized networks (vessels, lymphatics and nerves) needed to sustain large and complex tissue structures. Then there are the issues associated with blood compatibility, tissue remodeling and, more generally the

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immune and inflammatory responses to the new tissue or cells. Using autologous cells is an approach that is immunologically preferable, but it likely precludes the “off-the-shelf ” concept behind much of the attraction of tissue engineering. Mechanical Elasticity and Strength Development A critical feature of a heart is its mechanical characteristics. Simply speaking the heart must pump blood at a mean pressure of roughly 100 mmHg. Hence heart muscle must stretch in response to capillary filling pressure and eject a volume of blood that varies with demand. The latter requires a uniform and well-coordinated contraction that generates the required power. The mechanical fatigue limitations of a heart that must beat 3  108 times over 10 years must be compared with the flexural fatigue life of synthetic elastomeric materials that are typically much lower. It will be a significant challenge to replicate the complex architecture of the myocardium and its non-linear viscoelastic properties in both resting and activated states (Fung, 1993). While some constructs exhibit a significant burst strength and some groups are very advanced in the use of the tools of biomechanics to advance vascular graft (Nerem, 2003) or heart valve development, this area has received less attention than it deserves (Butler et al., 2000). Tissue Architecture and Electrical Conduction The complexity of the electrical conduction pathways in the heart has received little attention in the tissue engineering literature. The cells need to form the appropriate intercellular connections and matrix arrangements to enable the directed beating of contracting cells to generate the forces required to pump blood (Akins, 2000). The proper formation of the intercalated discs between myocytes are also critical in enabling electrical pulses to be transmitted in the correct direction at normal speeds and in allowing suitable force transmission. The heart also contains specialized cells that participate in the electrical conduction routes found throughout the heart. These specialized cells are crucial to the co-ordination of the heart’s contractile effort, and including them in the proper places in a regenerated substitute may be critical. There are clear differences between the rhythmic twitching of cultured cardiac cells en mass and the organized, efficient, regulated beating of the heart; only the latter will generate the force required to pump blood at systolic pressure levels. It is not difficult to envision the problems yet to be faced. Given the variety of electrical-conduction-related diseases in a normal myocardium, there is good reason to suspect that simple mimicry of heart muscle may fall short of the goal. Thrombogenicity and Endothelialization The need for blood compatibility is another crucial characteristic of cardiovascular constructs. All biomaterials lack the desired non-thrombogenicity and most extracellular matrices initiate thrombosis, endothelialization of the construct is another critical issue. ECs have a reversible plasticity (Augustin-Voss et al., 1991; Lipton et al., 1991; Risau, 1995) and they can become activated (proliferative or adhesive to leukocytes) upon exposure to inflammatory cytokines (e.g. IL1, tumor necrosis factor (TNF)) or to growth factors such as VEGF. Flow and the associated shear stress, normally in the range of 5–20 dyn/cm2, elongate and align cells in the direction of flow (Eskin et al., 1984; Ives et al., 1986) and modify gene expression (McCormick et al., 2001) as well as many other functions including markers of antithrombogenicity. ECs provide a hemocompatible surface by production of molecules that modulate platelet aggregation (e.g. prostacyclin), coagulation (thrombomodulin (Marcum et al., 1984; Esmon, 2000)) and fibrinolysis (Shen, 1998) (e.g. tissue plasminogen activator). They can be transformed into a prothrombotic surface, for example by the action of thrombin or through exposure to some biomaterials (Li et al., 1992; Cenni et al., 1993, 2000; Lu and Sipehia, 2001). Blood compatibility has been a key issue in the development of vascular grafts. Recent

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clinical success (Meinhart et al., 2001) has renewed enthusiasm for seeding grafts with ECs. In some protocols, many of the pre-seeded cells are lost on implantation due to insufficient adhesion (Williams, 1995) and thus the protection from thrombosis provided by the cells is limited due to the incomplete cell coverage. The potential to exploit the presence of circulating EC progenitors has only begun to be explored (Rafii, 2000). It is also worth noting the effects of the endothelium on the neighboring tissue and the corresponding effects on EC phenotype. With vascular smooth muscle cell (VSMC), this bidirectional cross-talk is thought to be a critical regulator of vascular homeostasis (Korff et al., 2001): secretion and expression of molecules such as nitric oxide (Palmer et al., 1988), prostacyclin (Moncada, 1982), and endothelin (Mawji and Marsden, 2003) act on VSMC to regulate vessel tone. Meanwhile, VSMC inhibits EC endothelin 1 (ET-1) production to increase EC NO and eNOS expression (Di Luozzo et al., 2000). Many other relevant systems (e.g. matrix metalloprotease (MMP) secretion and matrix remodeling) are also affected by the interactions between EC and other cell types. Vascularization The intrinsic nature of large cell-based constructs and the corresponding difficulty of supplying cells deep within the construct with nutrients is yet another problem. Diffusion is fine for 100 μm or so and low cell densities can extend this limit, but at the cost of making constructs too large to be useful. Thin or essentially 2D (e.g. a tube) constructs are feasible without an internal blood/nutrient supply. However it is hard to combine cells at tissue densities (108 cells/cm3) into large tissues without some sort of prevascularization or its alternative. Thus, a capillary network (and a lymphatic network) needs to be “engineered” as part of the creation of a larger structure. In a cell-free approach, vascularization and improvement of LV function following MI were achieved by sustained release of basic FGF (bFGF) incorporated into gelatin microspheres (Sakakibara et al., 2003), aFGF from ethylene vinyl acetate copolymer (Sellke and Simons, 1999) and bFGF from heparin-alginate beads (Harada et al., 1994). Mooney and colleagues have incorporated an EC mitogen (VEGF) into three-dimensional porous poly(lactide-co-glycolide) (PLG) scaffolds during fabrication (Sheridan et al., 2000) to promote scaffold vascularization. Sustained delivery of bioactive VEGF translated into a significant increase in blood vessel ingrowth in mice and the vessels appeared to integrate with the host vasculature. We are using microencapsulated VEGF165 secreting cells (prepared by transfection of L929 cells) as a means of exploring this strategy, at least for microcapsules (Vallbacka et al., 2001). Of course VEGF is but one angiogenic factor (Ahrendt et al., 1998) and issues associated with the functional maturity of the vessels and the need for multiple factors may limit this strategy. In a third approach, Vacanti et al., micromachined a hierarchical branched network mimicking the vascular system in 2D. Silicon and Pyrex surfaces were etched with branching channels ranging from 500–10 μm in diameter (Kaihara et al., 2000) that were then seeded with rat hepatocytes and microvascular ECs. Most recently, prevascularized skeletal muscle was created (Levenberg et al., 2005) by co-culturing skeletal muscle cells with ES-cell-derived EC and fibroblasts. It appeared that up to 40% of the engineered blood vessels “connected” to the host vasculature upon implantation, at least in this small animal model. Finally we note that there are initial attempts at adapting endothelial seeding approaches in a modular approach to create scalable and vascularized tissue constructs (Figure 60.2b). ECs were seeded onto sub-mm sized collagen gel cylindrical modules that contained a second cell (e.g. HepG2 or smooth muscle cells or in the future perhaps cardiomyocytes). These modules were packed into a larger tube, thereby creating interconnected channels lined with ECs. These channels permitted the perfusion of whole blood, creating a means of producing uniform, scaleable tissue constructs with an internal vascular supply (McCuigan and Sefton, 2006). Host Response and Biocompatibility Questions related to the immune and inflammatory response to tissue constructs are starting to draw attention. The host response to a tissue engineered construct is manifested by the innate and adaptive immune

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Figure 60.2 Cardiac tissue engineering culture systems focus on achieving adequate oxygen supply for highly metabolically active cells (a, b) and providing appropriate physical cues that lead to differentiated phenotype (c, d). (a) Direct culture medium perfusion of constructs based on neonatal rat cardiomyocytes inoculated into collagen sponges using Matrigel. Medium perfusion resulted in uniform cell distribution and maintenance of cell viability. Immunohistochemical staining illustrated cross-sectional distribution of cells expressing cardiac Troponin I (with permission from Radisic et al., 2004). (b) Modular tissue engineering approach using sub-mm sized EC seeded collagen modules assembled into a larger tube or construct (with permission from McGuigan and Sefton, 2006). (c) Zimmermann and Eschenhagen designed a bioreactor that provides cyclic mechanical stretch to engineered heart tissue based on neonatal rat cardiomyocytes and collagen gel. Mechanical stimulation yielded elongated cardiomyocytes with remarkably well-developed contractile apparatus (with permission from Zimmermann et al., 2002). (d) Cardiac-like electrical field stimulation was applied to collagen sponges inoculated with suspension of neonatal rat cardiomyocytes in Matrigel, resulting in differentiated phenotype and improved tissue assembly (with permission from Radisic et al., 2004).

systems, involving both plasma (e.g. complement) and cellular components (e.g. macrophages, T cells, etc.), that are directed against engineered cells and grafts or the materials used in tissue constructs. This potent immune response is most often mediated by major histocompatability complex mismatches between donor and host tissue in allogeneic transplantations. This response can also be manifested in situations where autologous cells or tissues are engineered to express therapeutic but foreign factors or if these autologous cells are placed in tissue constructs that themselves negatively impact immune consequences (Mikos et al., 1998). Immunosuppressants have enabled the successful transplantation of kidneys, hearts, and other organs. With the advent of tissue engineering, new configurations of tissues and organs (often with an added biomaterial component) are being developed and our understanding of the immune and inflammatory response to these new therapies is being shown to be inadequate. Some xenogeneic cell transplants (mice to rat) survive in situations of cardiac repair despite the species differences (Saito et al., 2002) although this may be specific

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to the animal model or to cardiac repair. The longevity of a transplant is also dependent on the ability of somatic cells to withstand and respond to the stresses of implantation, rejection, and other injuries (Halloran and Melk, 2001). The classic “foreign body reaction” to biomaterials is well known, but the details of the molecular signals (complement regulatory proteins, MMPs) that accompany this phenomenon (in the context of biomaterials) are only beginning to be defined. A variety of approaches have been undertaken or are in development to generate or to improve upon graft acceptance (Rossini et al., 1999). These approaches include methods to block the innate immune response such as by use of drugs or transferred genes to block NFkB signaling pathways, for example. Other methods to block the innate response include the use of antibodies to IL-1 or TNF or the use of anti-adhesion and anti-elastase antibodies. Perhaps nuclear transfer and therapeutic cloning strategies (McLaren, 2000) may be necessary assuming the various ethical issues can be resolved. We must better understand the mechanism of the host response itself so that we can design better biomaterials, select or engineer more suitable cells or devise better strategies for controlling both innate and adaptive immune responses and enable a functional integration of the new tissue with the host.

BIOREACTORS AND CONDITIONING Major efforts in the development of bioreactors for tissue engineering of myocardium focus on (a) providing sufficient oxygen supply for the highly metabolically active cardiomyocytes and (b) providing appropriate physical stimuli necessary to reproduce complex structure at various length scales (subcellular to tissue). The most common culture vessels utilized for tissue engineering of the myocardium include static or mixed dishes, static or mixed flasks, and rotating vessels. These bioreactors offer three distinct flow conditions (static, turbulent, and laminar) and therefore differ significantly in the rate of oxygen supply to the surface of the tissue construct. Oxygen transport is a key factor for myocardial tissue engineering due to the high cell density, very limited cell proliferation, and low tolerance of cardiac myocytes for hypoxia. In all configurations oxygen is supplied only by diffusion from the surface to the interior of the tissue construct, yielding 100 μm thick surface layer of compact tissue capable of electrical signal propagation and an acellular interior (Radisic et al., 2005). Oxygen Supply In an attempt to enhance mass transport within cultured constructs, a perfusion bioreactor that provides interstitial medium flow through the cultured construct at velocities similar to those found in native myocardium (400–500 μm/s) was developed (Radisic et al., 2004b). In such a system oxygen and nutrients were supplied to the construct interior by both diffusion and convection (Figure 60.2a). Interstitial flow of culture medium through the central 5 mm diameter  1.5 mm thick region resulted in physiologic density of viable and differentiated, aerobically metabolizing cells. In response to electrical stimulation, perfused constructs contracted synchronously had lower ET and recovered their baseline function levels of ET and MCR following treatment with a gap junctional blocker; dish-grown constructs exhibited arrhythmic contractile patterns and failed to recover their baseline MCR levels. These studies suggested that the immediate establishment and maintenance of interstitial medium flow markedly enhanced the control of oxygen supply to the cells and thereby enabled engineering of compact constructs. However, most cells in perfused constructs were round and mononucleated, indicating that some of the regulatory signals, either molecular or physical, were not present in the culture environment. In another approach, a separate compartment for medium flow was created by perfusing channeled scaffolds in a configuration resembling the capillary network in vivo. Neonatal rat heart cells were inoculated into the pores of an elastic, highly porous scaffold (PGS) with a parallel channel array and perfused with a synthetic

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oxygen carrier (Oxygent™ in culture medium, PFC emulsion) (Radisic et al., 2006). Constructs cultivated with PFC emulsion had significantly higher DNA content, significantly lower ET and higher relative presence of cardiac markers troponin I and connexin-43 (Western blot) compared to the culture medium alone. Cells were present throughout the construct volume. In this configuration, the presence of PFC emulsion further enhanced the oxygen supply to the cells by improving both axial (convective term) and radial (effective diffusivity) transport properties (Radisic et al., 2005). Differentiation Mechanical Stimulation One significant approach to cardiac tissue engineering, established by Eschenhagen and colleagues (Eschenhagen et al., 1997; Fink et al., 2000; Zimmermann et al., 2000, 2002b) involves the cultivation of neonatal rat heart cells in collagen gel or Matrigel, in the presence of growth factors. The cultured tissues are subjected to sustained mechanical strain. Under these conditions, cardiomyocytes and non-myocytes form 3D cardiac organoids, consisting of a well-organized and highly differentiated cardiac muscle syncytium, that exhibit contractile and electrophysiological properties of working myocardium. First implantation experiments in healthy rats showed survival, strong vascularization, and signs of terminal differentiation of cardiac tissue grafts (Zimmermann et al., 2002a). In the state of the art approach by Eschenhagen and colleagues neonatal rat cardiac cells were suspended in the collagen/Matrigel mix and cast into circular molds (Zimmermann et al., 2002b). After 7 days of static culture, the strips of cardiac tissue were placed around two rods of a custom made mechanical stretcher and subjected to either unidirectional or cyclic stretch (Figure 60.2c). Histology and immunohistochemistry revealed the formation of intensively interconnected, longitudinally oriented cardiac muscle bundles with morphological features resembling adult rather than immature native tissue. Primitive capillary structures were also detected. Cardiomyocytes exhibited well-developed ultrastructural features: sarcomeres arranged in myofibrils, with well-developed Z, I, A, H, and M bands, specialized cell–cell junctions, T-tubules as well as well-developed basement membrane. Contractile properties were similar to those measured for native tissue, with a high ratio of twitch to resting tension and strong β-adrenegenic response. Action potentials characteristic of rat ventricular myocytes were recorded. Electrical Stimulation In native heart, mechanical stretch is induced by electrical signals. Contraction of the cardiac muscle is driven by the waves of electrical excitation (generated by pacing cells) that spread rapidly along the membranes of adjoining cardiac myocytes and trigger release of calcium, which in turn stimulates contraction of the myofibrils. Electromechanical coupling of the myocytes is crucial for their synchronous response to electrical pacing signals, resulting in contractile function and pumping of blood (Severs, 2000). In a recent study, (Radisic et al., 2004a) cardiac constructs prepared by seeding collagen sponges with neonatal rat ventricular cells were electrically stimulated using suprathreshold square biphasic pulses (2 ms duration, 1 Hz, 5 V). The stimulation was initiated after 1–5 days of scaffold seeding (3-day period was optimal) and applied for up to 8 days. Over only 8 days in vitro, electrical field stimulation induced cell alignment and coupling, increased the amplitude of synchronous construct contractions by a factor of 7 and resulted in a remarkable level of ultrastructural organization. Development of conductive and contractile properties of cardiac constructs was concurrent, with strong dependence on the initiation and duration of electrical stimulation. Aligned myofibers expressing cardiac markers were present in stimulated samples and neonatal heart (Figure 60.2d). Stimulated samples had sarcomeres with clearly visible M, Z lines, H, I and A bands. In most cells, Z lines were aligned, and the intercalated disks were positioned between two Z lines. Mitochondria

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(between myofibrils) and abundant glycogen were detected. In contrast, non-stimulated constructs had poorly developed cardiac-specific organelles and poor organization of ultrastructural features. Hence the in vitro application of a single, but key in vivo factor, progressively enhanced the functional tissue assembly and improved the properties of engineered myocardium at the cellular, ultrastructural, and tissue levels.

IMPLANTATION OF CARDIAC PATCHES While significant progress has been made in constructing in vitro cultivation systems and biomaterial scaffolds, very few studies have focused on implantation of cell-based cardiac patches onto viable or injured myocardium (Figure 60.3). In a pioneering study, Li et al. (1999) implanted a construct based on neonatal rat cardiomyocytes and collagen sponges onto the surface of the cryoinjured myocardium of Lewis rats (Figure 60.3). The grafts were implanted 3 weeks post-infraction. After 5 weeks in vivo, the cells survived, supported by the blood vessel ingrowth and integrated with the surrounding tissue. However, the graft did not improve LV function. Attenuation of pathological remodeling (i.e. prevention of ventricle dilatation and maintenance of contractile function) was observed in a study by Leor et al. (2000), where cardiac constructs based on neonatal rat cardiomyocytes and porous alginate scaffolds were implanted onto myocardium of Sprague-Dawley rats that underwent permanent main coronary artery occlusion (Figure 60.3). The grafts were implanted 7 days after MI. After 9 weeks of implantation, the grafts demonstrated integration with host myocardium at the anchorage sites as well as inflammatory infiltrates and presence of fibrous collagen. Zimmerman et al. (2002a) placed cardiac tissue rings cultivated in the presence of mechanical stimulation onto uninjured hearts of Fisher 344 rats for 14 days (Figure 60.3). They noticed that although both cells and collagen were isolated from Fisher rats, immunosuppression was required for maintenance of heart tissue upon implantation. In the absence of immunosuppression, even in the syngeneic approach, cardiac

1 cm Cryoinjured myocardium of Lewis rats

Coronary artery occlusion in Sprague-Dawley rat myocardium

Uninjured heart of Fisher 344 rats

Constructs implanted after 3 weeks Constructs implanted after 1 week Fetal (Sprague-Dawley) rat Fetal (Lewis) rat cardiomyocytes in cardyomyocytes in porous collagen sponge aglinate scaffolds Cultivated under static condition for Cultivated under static conditions 7 days for 4 days

Neonatal (Fisher 344) rat cardiomyocytes in collagen type I gel Cultivation with mechanical stimulation for 12 days

After 5 weeks in vivo no significant After 9 weeks in vivo attenuation After 14 weeks in vivo the implant difference compared to the controls of LV function and maintenance of vascularized and improved the contractile function in comparison level of maturation. to controls without construct Immunosupression was required Li et al. (1999)

Leor et al. (2000)

Zimmermann et al. (2002)

Figure 60.3 Representative studies investigating the effect of implantation of the cardiomyocyte-based constructs on the function of injured or viable hearts.

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constructs completely degraded after only 2 weeks in vivo. It is unknown what exactly caused the response; it is possible that it was the remainder of serum or chick extract. Regardless, the finding has significant implications in the potential implantation of cardiac patches in clinical settings. Limitations related to the source of autologous cardiomyocytes motivated the studies that utilized nonmyocyte-based patches for MI repair. Smooth muscle cells seeded with poly (ε-caprolactone-co-L-lactide) sponge reinforced with poly-L-actide fabric were used in a modified endoventricular circular patch plasty procedure (Dor procedure). Cell seeded grafts resulted in improved LV function (as assessed by echocardiography) compared to cell-free controls (Matsubayashi et al., 2003). A patch made of dermal fibroblasts seeded onto knitted Vicryl mesh (Dermagraft) was used in an attempt to increase angiogenesis upon MI. When placed over the infracted regions on the hearts of severe combined immunodeficient (SCID) mice, the grafts improved microvessel density within the damaged myocardium (Kellar et al., 2001). These studies demonstrated feasibility of cardiac patch implantation, but further studies are necessary to estimate the effect of culture conditions and scaffold type on the in vivo outcome. Although, significant progress has been made in the area of biomaterials and bioreactors, it is currently unknown which cultivation conditions and what biomaterial will best preserve contractile function and prevent pathological remodeling upon implantation. Thus studies that investigate this in a systematic fashion and correlate in vitro parameters (e.g. force of contraction) to in vivo outcomes (e.g. fractional shortening) are required.

SUMMARY Overall, the field of cardiac tissue engineering is very much in its infancy. Although the results to date are exceedingly encouraging, much remains to be done in order to develop clinically relevant approaches, let alone move towards a whole heart. Not surprisingly a NIH task force (National Institutes of Health, 1999) has emphasized the development of heart components such as a cardiac patch or a valve before “graduating” to whole heart engineering. However, significant progress has been made ever since the LIFE initiative embarked on the creation of the artificial heart in 1999. Functional viable cardiac patches have been engineered based on neonatal rat cardiomyocytes and more recently based on ES-cell-derived cardiomyocytes. Various biomaterials have been tested for this purpose and in vitro culture systems have been developed that enhance cardiac construct differentiation (mechanical and electrical stimulation) as well as improve cardiomyocyte survival at high density (medium perfusion). Exciting new findings on resident progenitor cells have also emerged. While the completely artificial heart will remain a dream, the near future will bring a clinically relevant autologous cardiac patch.

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Menasche, P., Hagege, A.A., Vilquin, J.T., Desnos, M., Abergel, E., Pouzet, B., Bel, A., Sarateanu, S., Scorsin, M., Schwartz, K., Bruneval, P., Benbunan, M., Marolleau, J.P. and Duboc, D. (2003). Autologous skeletal myoblast transplantation for severe postinfarction left ventricular dysfunction. J. Am. Coll. Cardiol. 41: 1078–1083. Mikos, A.G., McIntire, L.V., Anderson, J.M. and Babensee, J.E. (1998). Host response to tissue engineered devices. Adv. Drug Deliv. Rev. 33: 111–139. Moncada, S. (1982). Eighth Gaddum memorial lecture. University of London Institute of Education, December 1980. Biological importance of prostacyclin. Br. J. Pharmacol. 76: 3–31. Muller-Ehmsen, J., Peterson, K.L., Kedes, L., Whittaker, P., Dow, J.S., Long, T.I., Laird, P.W. and Kloner, R.A. (2002a). Rebuilding a damaged heart: long-term survival of transplanted neonatal rat cardiomyocytes after myocardial infarction and effect on cardiac function. Circulation 105: 1720–1726. Muller-Ehmsen, J., Whittaker, P., Kloner, R.A., Dow, J.S., Sakoda, T., Long, T.I., Laird, P.W. and Kedes, L. (2002b). Survival and development of neonatal rat cardiomyocytes transplanted into adult myocardium. J. Mol. Cell. Cardiol. 34: 107–116. Murry, C.E., Kay, M.A., Bartosek, T., Hauschka, S.D. and Schwartz, S.M. (1996). Muscle differentiation during repair of myocardial necrosis in rats via gene transfer with MyoD. J. Clin. Invest. 98: 2209–2217. Murry, C.E., Soonpaa, M.H., Reinecke, H., Nakajima, H., Nakajima, H.O., Rubart, M., Pasumarthi, K.B., Virag, J.I., Bartelmez, S.H., Poppa, V., Bradford, G., Dowell, J.D., Williams, D.A. and Field, L.J. (2004). Haematopoietic stem cells do not transdifferentiate into cardiac myocytes in myocardial infarcts. Nature 428: 664–668. National Institutes of Health (NIH) (1999). Working Group on Tissuegenesis and organogenesis for Heart, Lung and Blood Applications. Nerem, R.M. (2003). Role of mechanics in vascular tissue engineering. Biorheology 40: 281–287. Nygren, J.M., Jovinge, S., Breitbach, M., Sawen, P., Roll, W., Hescheler, J., Taneera, J., Fleischmann, B.K. and Jacobsen, S.E. (2004). Bone marrow-derived hematopoietic cells generate cardiomyocytes at a low frequency through cell fusion, but not transdifferentiation. Nat. Med. 10: 494–501. Oh, H., Bradfute, S.B., Gallardo, T.D., Nakamura, T., Gaussin, V., Mishina, Y., Pocius, J., Michael, L.H., Behringer, R.R., Garry, D.J., Entman, M.L. and Schneider, M.D. (2003). Cardiac progenitor cells from adult myocardium: homing, differentiation, and fusion after infarction. Proc. Natl Acad. Sci. USA 100: 12313–12318. Orlic, D., Kajstura, J., Chimenti, S., Jakoniuk, I., Anderson, S.M., Li, B., Pickel, J., McKay, R., Nadal-Ginard, B., Bodine, D.M., Leri, A. and Anversa, P. (2001). Bone marrow cells regenerate infarcted myocardium. Nature 410: 701–705. Palmer, R.M., Ashton, D.S. and Moncada, S. (1988). Vascular endothelial cells synthesize nitric oxide from L-arginine. Nature 333: 664–666. Papadaki, M., Bursac, N., Langer, R., Merok, J., Vunjak-Novakovic, G. and Freed, L.E. (2001). Tissue engineering of functional cardiac muscle: molecular, structural and electrophysiological studies. Am. J. Physiol. Heart Circ. Physiol. 280: H168–H178. Radisic, M., Park, H., Shing, H., Consi, T., Schoen, F.J., Langer, R., Freed, L.E. and Vunjak-Novakovic, G. (2004a). Functional assembly of engineered myocardium by electrical stimulation of cardiac myocytes cultured on scaffolds. Proc. Natl Acad. Sci. USA 101: 18129–18134. Radisic, M., Yang, L., Boublik, J., Cohen, R.J., Langer, R., Freed, L.E. and Vunjak-Novakovic, G. (2004b). Medium perfusion enables engineering of compact and contractile cardiac tissue. Am. J. Physiol. Heart Circ. Physiol. 286: H507–H516. Radisic, M., Malda, J., Epping, E., Geng, W., Langer, R. and Vunjak-Novakovic, G. (2005). Oxygen gradients correlate with cell density and cell viability in engineered cardiac tissue. Biotechnol. Bioeng. 93: 332–343. Radisic, M., Park, H., Chen, F., Salazar-Lazzaro, J.E., Wang, Y., Dennis, R.G., Langer, R., Freed, L.E. and Vunjak-Novakovic, G. (2006). Biomimetic approach to cardiac tissue engineering: oxygen carriers and channeled scaffolds. Tissue Eng. 12: 2077–2091. Rafii, S. (2000). Circulating endothelial precursors: mystery, reality, and promise. J. Clin. Invest. 105: 17–19. Reffelmann, T. and Kloner, R.A. (2003). Cellular cardiomyoplasty – cardiomyocytes, skeletal myoblasts, or stem cells for regenerating myocardium and treatment of heart failure? Cardiovasc. Res. 58: 358–368. Reinecke, H., Zhang, M., Bartosek, T. and Murry, C.E. (1999). Survival, integration, and differentiation of cardiomyocyte grafts: a study in normal and injured rat hearts. Circulation 100: 193–202. Reinlib, L. and Field, L. (2000). Cell transplantation as future therapy for cardiovascular disease?: a workshop of the National Heart, Lung, and Blood Institute. Circulation 101: e182–e187. Risau, W. (1995). Differentiation of endothelium. FASEB J. 9: 926–933.

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61 Regenerative Medicine in the Cornea Heather Sheardown and May Griffith

INTRODUCTION: THE NEED FOR REGENERATIVE MEDICINE IN THE CORNEA The cornea, acting as both the primary refractive element of the eye and its main defense against injury from the external environment, consists of three cellular layers organized into a specialized structure. The outermost stratified, non-keratinized epithelium protects the ocular structures from external insult. The middle, primarily structural, stromal layer is comprised of more than 300 highly ordered layers of primarily Type I collagen interspersed with corneal stromal cells, and gives the cornea both its strength and transparency. The single cell thick posterior endothelial layer is essential for the maintenance of stromal hydration and hence corneal transparency. Any damage or failure of corneal cells due to injury or disease can lead to vision loss, and where irreversible, results in blindness. In most cases, the treatment for such vision loss is transplantation. Currently, transplantation of human corneas usually involves a full-thickness replacement by a surgical technique called penetrating keratoplasty (PK). Lamellar keratoplasty (LKP) is an alternative surgical procedure that requires removal of only the damaged or diseased epithelium and stroma, leaving the endothelium intact, in cases where only the more superficial layers are damaged. Non-penetration of the aqueous humor reduces the rate of rejection and post-operative complications such as leakage, improving long-term graft stability (Johnson et al., 2000; Aucoin et al., 2002). According to the World Health Organization, diseases of the cornea are a major cause of vision loss, second only to cataracts as the leading cause of blindness (Whitcher et al., 2001). Corneal ulceration and ocular trauma are estimated to result in between 1.5 and 2 million new cases of blindness worldwide on an annual basis; corneal scarring resulting from measles is a leading cause of blindness in children. Cornea-induced blindness, affecting more than 10 million individuals worldwide (estimates from the Vision Share Consortium of Eye Banks, USA) is generally related to a loss of transparency and can be managed by replacement by all or part of the host cornea with, most commonly, human donor tissue or with synthetic devices. It has been estimated that approximately 45,000 transplants, either PK or lamellar procedures, are performed each year in the United States. Under ideal circumstances, traditional allograft cornea transplantation has a quite high success rate, with an estimated 80% of grafts remaining clear after 2 years, a number which drops to approximately 65% 5 years post surgery (Beekhuis, 1995; Sit et al., 2001; Carlsson et al., 2003). Reported acute rejection rates range from 13.3% to 65% within 4 months of keratoplasty, and rejection can occur many years later (Smolin and Goodman, 1988). With many diseases, including inactive central corneal scars and keratoconnus, the prognosis is excellent. Other diseases including alkali burns, severe dry eye, immunological disorders, stem cell deficiency, vascularization, or ocular diseases such as Stevens–Johnson syndrome (SJS), ocular citracial pemphigoid, and neurotropic scars secondary to herpes zoster ophthalmicus often result in the eye not being able to support corneal transplants (Trinkaus-Randall, 2000; Khan et al., 2001). In these cases, reported success

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rates are much lower; in cases of repeated graft rejection for example, the success rate of future transplantation drops to near zero (Khan et al., 2001). As is the case with many other organs, worldwide demand for donor corneas exceeds the supply, an imbalance that is projected to worsen. Waiting lists exceeding 2 years are now common in North America (Griffith et al., 2002; Carlsson et al., 2003). Wait times are expected to further increase with the aging population, as older corneas are less suitable for transplantation and these patients are more likely to require transplants. A further decrease in the availability of acceptable donor tissue is expected with the increasing incidence of infectious diseases, including HIV and hepatitis, as well as the growing popularity of laser in situ keratomileusis (LASIK) for correcting refractive errors. These surgically treated corneas are unacceptable donor tissue. Another serious disadvantage of cornea allograft transplantation is the possibility for transmission of infection. Person-to-person transmission of the rabies virus (Houff et al., 1979) and at least one case of Creutzfeldt–Jakob disease (Duffy et al., 1974) have been reported. Hepatitis B and C and HIV can be isolated from tears and there is concern about their possible transmission. Given the knowledge that infectious transmission is possible in prion form, it is conceivable that transmission of as yet unknown pathogens could also occur. These issues are compounded in third world countries, where instances of corneal blindness are rising, yet the skills and resources to perform transplant surgeries are limited (Chirila, 2001; Griffith et al., 2002). It is clearly beneficial to seek corneal replacements. Corneal substitutes designed to replace part of or the full thickness of damaged or diseased corneas range from prosthetic devices that solely address replacement of the cornea’s function through to tissue engineered hydrogels that allow and in fact depend on some regeneration of the host tissues. At present, however, widely accepted corneal substitutes are not available (Chirila, 2001). The prostheses developed are gaining acceptance but none integrates seamlessly into the host tissue. Recent developments in many areas of bioengineered corneas, including clinical trials of an artificial cornea designed as a prosthesis, development of completely natural corneal replacements, and development of biosynthetic matrices that permit host tissue regeneration hold promise for the future. Therefore, the focus of the current chapter will be scaffolding for transplantable, engineered corneal replacements, and on methods to improve the regenerative capacity of either the local cellular components or seeded components.

DESIGN REQUIREMENTS FOR HUMAN CORNEAL REPLACEMENTS Mechanically human corneas are extremely tough and tear resistant, but perfect reproduction of these properties may not be necessary for a substitute or replacement to be functional, provided that the device can maintain optical clarity, survive handling and implantation stresses, post-operative wear and tear, and can protect the more delicate inner parts of the eye. As well, the device should be functionally comparable in other key areas (Chapekar, 2000). High optical transparency with minimal scatter for vision is essential. The human cornea comprises some 300 layers of collagen fibrils arranged in offset sheets, parallel to the plane of the cornea. This precise, ordered structure has been proposed to be essential for the high optical clarity of the cornea (Maurice, 1957). However, most recent evidence suggests that clarity results from a combination of refractive index matching and, more importantly, that the collagen fibrils have diameters less than the shortest wavelength of visible light (ca. 380 nm) (Benedek, 1971; Freegard, 1997). Optical clarity (freedom from absorption and scatter in the visible region) requires a synthetic polymer or biopolymer that is free of chromophores absorbing in the visible region and also free of subunits including crystalline domains or large aggregates of microfibrils that will scatter light. Amorphous synthetic polymers, free of crystallinity, such as poly(methyl methacrylate) (PMMA), are widely used when transparency is required. Biopolymers, such as collagen I and fibrin, are inherently fibrous in nature because of the ready association of their nanofibrillar subunits. However, by careful control of cross-linking conditions, both biopolymers can be locked in a form where their microfibrils are below approximately 300 nm

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diameter, which is well below the lowest visible wavelengths (Li et al., unpublished observations). However, because these clear materials have a disordered array of microfibrillar structures, they are less tough than the composite structure of the natural cornea that is based on thin plies of roughly parallel collagen filaments, with glycosaminoglycans and water contributing to inter filament bonding. Tissue engineered corneal materials that are in use, or being developed, are based on both amorphous synthetic polymers and microfibrillar biopolymers to incorporate both the requirement for transparency as well as the toughness necessary to protect the more delicate inner parts of the eye. Beyond the need for vision restoration, there are other critical questions of bonding or integration with the host tissue and epithelial overgrowth to restore the cornea’s protective surface layer. Even more demanding, but very desirable, is the regeneration of a functional, sensitive nerve network that functions both as a highly effective warning of potential injury, and as a key link in eye humidification through signaling when blinking must occur to prevent the potentially dangerous situation of “dry eye” that can occasionally lead to ulceration and vision loss (Stern et al., 1988).

KERATOPROSTHESES Currently Available Keratoprostheses Keratoprostheses (KPro) (commonly referred to as artificial corneas) are usually completely synthetic constructs designed to replace the central portion of an opaque cornea. Early devices with rigid components required complex surgery and led to high incidences of such complications as extrusion, melting, aqueous leakage, infection, retroprosthetic membrane formation, and glaucoma. Rigidity, lack of free flow of nutrients (oxygen, glucose), and lack of biointegration have also been shown to contribute to device failure. Therefore, in addition to transparency, appropriate refractive index, and sufficient strength to fulfill the protective barrier function of the cornea, it is evident that properties such as oxygen and nutrient permeability, necessary for survival of the surrounding cells and the ability to be colonized by host stromal cells for integration and minimization of inflammation are key design considerations (Hicks et al., 2000; Griffith et al., 2002). While vascularization has been suggested to improve healing and survival of adjacent corneal tissue by provision of nutrients and proteolytic enzyme inhibitors in some studies (Kim et al., 2002; Stoiber et al., 2004), more research into the effects of the degree of vascularization of a host cornea on a device is needed as vascularization is often an indication of increased inflammation as excessive toxic, immune, or inflammatory responses can result in biodegradation, calcification, or tissue melting (Hicks et al., 2000; Griffith et al., 2002). In light of these requirements, the “core and skirt” concept has now been widely adopted (Chirila, 2001). This design is based upon a porous, flexible, biointegratable “skirt” (containing interconnected pores in the 10–30 μm diameter range) that surrounds a clear, central optic. The porous nature of the skirt enables fibroblast ingrowth and extracellular matrix (ECM) deposition in a similar manner to the wound healing process, to anchor the device in the eye (Trinkaus-Randall, 2000; Griffith et al., 2002, 2005). Tissue breakdown around the anchoring skirt and extrusion of the KPro is still a major cause of its failure. In some cases, the implant is then covered by transplanted autologous tissue or eyelid skin (Khan et al., 2001). The posterior surface of the implant should inhibit cellular attachment in order to prevent development of opaque fibrous retroprosthetic membranes, another complication (Griffith et al., 2002). Less studied is the need for coverage of the anterior implant surface by a confluent layer of corneal epithelial cells which is thought to improve long-term device stability, allow for tear film spreading, remove the need for external coverage of the skirt material, and prevent bacterial infection and epithelial downgrowth (Trinkaus-Randall, 2000; Griffith et al., 2002, 2005). The central optic, aside from being transparent, should provide a refractive power similar to that of the normal cornea and should have a diameter sufficient to allow posterior segment visualization and a reasonable field of view.

Regenerative Medicine in the Cornea

The introduction of highly hydrophilic constructs, based on poly(2-hydroxyethyl methacrylate) (pHEMA) hydrogels for both the transparent core and the microporous skirt, has resolved many of the initial problems of extrusion, inflammatory, and immune reactions, although calcification was a problem in the earlier iterations (Vijayasekaran et al., 2000). Recently, multicenter clinical trials with the AlphaCor KPro (based on the pHEMA hydrogel skirt core construction, previously known as the Chirila KPro) (Crawford et al., 2002) showed that a full-thickness, synthetic device can be maintained in the cornea through anchorage via fibroblast ingrowth into the peripheral portion. However, epithelialization of the pHEMA anterior surface did not occur, even though it is realized that epithelialization would be ideal. Nerve regeneration in these prostheses has not been reported or addressed, but would be essential to more normal corneal function. The AlphaCor KPro device has been used as an alternative to donor corneal tissue in patients who would be at high risk of conventional corneal graft failure, a very demanding situation for KPro retention. A number of studies have described the implantation of AlphaCor KPro in patients with various pathologies (Hicks et al., 2002, 2003a, 2004; Crawford et al., 2005). Surgery involved a two-stage procedure in which the device was placed within an intrastromal pocket closed by suturing a conjunctival flap over the anterior surface of the cornea. After a period of 12 weeks, the device optic was exposed by removing the conjunctival flap. The esthetics of this optically functional Kpro have raised some problems (Crawford et al., 2002). Recent results suggest that retention to 1 year was achieved in greater than 80% of cases with an overall increase in survival rate and visual acuity compared to repeated donor grafts (Hicks et al., 2002). Complications included stromal melting, retroprosthetic membrane formation, optic damage, and poor biointegration (Hicks et al., 2005, 2006). A comprehensive and unique program of data collection has allowed for ongoing review of complications and risk and protective factors. In early studies, active ocular simplex virus was found to be a contraindication (Hicks et al., 2003b). More recent results suggest that with appropriate therapies, herpes simplex virus (HSV) does not exclude patients from AlphaCor treatment. It was concluded that a history of HSV should be an exclusion factor for AlphaCor surgery. Additionally, in approximately 20% of clinical trial cases, deposits either on the surface or within the hydrogel optic resulted in diminished vision, and are thought to be related to smoking, or adsorption of certain combinations of medications to the exposed hydrogel leading to calcium deposition (Hicks et al., 2003a, b, 2004). However, studies indicate that elimination of the implicated medications effectively prevented calcium deposit formation in more recently implanted devices (Hicks et al., 2004). Other core skirt KPro based on various materials including a porous semitransparent poly tetrafluoroethylene (PTFE) skirt and a central optic of poly vinyl pyrrolidone (PVP) coated silicone rubber (poly(dimethyl siloxane) or PDMS) (Legeais et al., 1997; Legeais and Renard, 1998), or on poly(butyl methacrylate), hexaethyleneglycolmethacrylate with a dimethacrylate cross-linker (Bruining et al., 2002) have been proposed. Regenerative Medicine Applied to Keratoprosthesis Development A major goal in current keratoprosthesis research is to improve understanding of cellular interactions with the implant materials and to develop polymers which elicit a suitable biological response. The ingrowth of stromal cells has been widely demonstrated in a variety of materials; therefore, growth of a corneal epithelial cell layer, thought to improve the stability of the implants by preventing stromal exposure to the tear film, proteinases, and inflammatory cells, over the anterior device surface (George and Pitt, 2002) is the subject of considerable current investigation. The epithelial cells must be able to migrate from the remaining corneal tissue over the implant surface, attach to the material, and proliferate to restore complete coverage. Factors including surface hydrophilicity, porosity, topography, adhesiveness, and permeability to nutrients have been shown to affect epithelialization of a synthetic implant (George and Pitt, 2002; Sweeney et al., 2003). Generally, materials must be modified to enable epithelialization (Trinkaus-Randall, 2002), although other factors including pore size and surface topography (Evans et al., 2003) can impact device epithelialization.

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Passive coating with ECM proteins, including collagen and laminin, to mimic the epithelial basement membrane and fibronectin, which forms a provisional matrix for cell migration during wound healing, has been shown to promote cell adhesion and outgrowth in vitro (Evans et al., 2000; Griffith et al., 2002; Sweeney et al., 2003). The presence of these matrix proteins on an implant surface is thought to trigger migrating cells to reform a basement membrane by ECM protein secretion and formation of adhesion complexes at the surface (Sweeney et al., 2003). However, preliminary in vivo results have been less conclusive and long-term data are generally lacking. In particular, the effects of proteolytic activity on the surface must be further investigated. It was noted that more rapid wound closure did not lead to the most persistent epithelial coverage, indicating that cell adhesion may be more critical than migration rate for effective wound healing. More recent in vitro work suggests that corneal epithelial cell growth and adhesion were significantly enhanced by tethering of laminin or fibronectin adhesion promoting peptide (FAP) via flexible polyethylene glycol (PEG) chains, more so than by tethering of fibronectin or simple coating of the surface with matrix proteins (Jacob et al., 2005; Wallace et al., 2005). In several other studies (Kobayashi and Ikada, 1991; Merrett et al., 2001), modification with fibronectin-based (RGD(S)) (Kobayashi and Ikada, 1991; Merrett et al., 2001; Aucoin et al., 2002) laminin-based (YIGSR) (Merrett et al., 2001; Aucoin et al., 2002), and a novel collagen-based peptide Gly–Pro–Nleu (Johnson et al., 2000) has been observed to improve epithelial cell adhesion to various surfaces in vitro. Surface modification with combinations of peptides, including the cell adhesion peptides RGDS and YIGSR as well as synergistic counterparts PHSRN and PDSGR, demonstrated that corneal epithelial cell adhesion is greatly improved on surfaces with the cell adhesion peptides and at least one of the counterparts (Aucoin et al., 2002). Another strategy to improve epithelialization is through the use of growth factors. In particular, epidermal growth factor (EGF) is a potent stimulator of corneal epithelial cell proliferation and migration and is active in the wound healing process. Recent work indicates that covalent binding of EGF to PDMS substrates via a PEG tether can significantly improve cell coverage of the polymer in vitro (Klenkler et al., 2005). Furthermore, EGF attached surfaces show significantly greater production of various ECM proteins which are necessary for cell adhesion to occur as shown in Figure 61.1. However, the interactions between the growth factor-modified polymer and the cells are clearly complex and require further study. Underlying surface modifications appear to play a role in the extent of cell coverage as well as the density of the EGF on the surface and the presence of EGF in the cell culture medium. While good coverage with corneal epithelial cells was observed at 4 days with intermediate EGF concentrations on PDMS surfaces resulting in better cell coverage as shown in Figure 61.2. High PEG densities were also found to be undesirable, presumably due to the nonfouling nature of this surface inhibiting the adsorption of adhesion proteins produced by the cells. In contrast to stimulatory effects, epithelial cell attachment to certain parts of the keratoprosthesis must be inhibited to prevent epithelial downgrowth and retroprosthetic membrane formation. Transforming growth factor β (TGFβ) was investigated due to its previously demonstrated ability to inhibit epithelial growth and promote stromal keratocyte proliferation, and hence could potentially be useful for modification of the stromal implant surface. However, the results observed on TGFβ-modified PDMS surfaces in vitro were opposite to those expected; keratocyte adhesion was inhibited and epithelial cell growth enhanced by the surface treatment, indicating the complex nature of growth factor–cell interactions (Merrett et al., 2003). Grafting of PEG to PMMA implants, which typically exhibit high protein deposition and cell adhesion associated with retroprosthetic membrane formation, was investigated (Kim et al., 2001). The modification resulted in decreased keratocyte and inflammatory cell adhesion on the polymer surface in vitro and in rabbit experiments. Permeability to oxygen and nutrients, also a key parameter for survival of cells adjacent to a polymeric implant, is the basis for the development of novel materials. In one study, interpenetrating networks of PDMS and hydrogels were found to glucose permeability levels similar to those of the native cornea (Liu and Sheardown, 2005) and to support corneal epithelial cell adhesion (unpublished data). Novel perfluoropolyether-based

Regenerative Medicine in the Cornea

(a)

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Figure 61.1 Production of ECM proteins by corneal epithelial cells on EGF-modified surfaces. (a) Fibronectin production by cells on EGF-modified PDMS surfaces at 4 days. (b) Fibronectin production by cells plated on unmodified PDMS in the presence of exogenous EGF. Similar results were obtained for laminin and fibronectin. 100 magnification.

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Figure 61.2 Effect of EGF surface density on corneal cell adhesion to EGF-modified surfaces. Clearly, as the EGF density increases, cell interactions with the surfaces decrease with high PEG densities decrease the effect of the tethered EGF. materials with both oxygen and nutrient permeability have shown good success in corneal onlay applications. To enhance epithelial overgrowth, a 5–10 nm layer of collagen I was covalently immobilized on the anterior surface of each lenticule. In a clinical comparison of implanted lenticules and sham wounds, epithelium completely covered the feline, corneal wound bed (sham) by days 3–9, and the exposed lenticule surface (implanted) by days 5–11 in six of the seven implanted corneas. Overall, the corneas in both series were quiet with no signs of thinning, remained transparent by slit lamp examination, maintained multilayered epithelial cover, and supported a stable tear film during the observation period. Light microscopy of the sham-wounded corneas revealed six to seven layers of cells constituting the epithelium on the central portion of the original wound bed at 4 weeks that had increased to eight to ten layers by 8 weeks. This was slightly less than the 12 layers of cells in normal feline

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corneal epithelium. An overall examination of the stromal tissue in the implanted corneas at both time points revealed that the stroma was relatively normal in appearance with no evidence of thinning (stromal melting) (Evans et al., 2002). While intended for onlay applications, this material may also have potential as a substrate material for a keratoprosthesis.

CELL-BASED HUMAN CORNEAL EQUIVALENTS More recently, tissue engineering of functional corneal equivalents has been used to develop constructs to mimic the structure of the native cornea, incorporating the collagenous structural component and the main cell types present in the native cornea. It is believed that this cell-based, tissue engineered approach will lead to the development of corneal substitutes that will play a more important role in patient treatment and in fundamental studies of the cellular and molecular mechanisms during corneal wound healing in the long term (Trinkaus-Randall, 2000). Several groups have been developing corneal equivalents using completely natural materials as potentially implantable replacements. The model developed by the Laboratoire d’Organogenese Experimentale (LOEX) (Germain et al., 1999) uses a self-assembly approach whereby stromal cells are provided with the nutrients and appropriate factors such as ascorbic acid to induce production of sheets of collagen and other ECM macromolecules (Gaudreault et al., 2003). These sheets are stacked together and subsequently seeded with epithelial cells; the endothelial cell layer was not included in initial reconstructions although more recent work has focused on the optimization of the culture conditions for endothelial cells for the inclusion of this layer in the construct (Gagnon et al., 2005). Previous work with tissue engineered blood vessels demonstrated that high tensile strength could be achieved by this method (Auger et al., 2002) suggesting that this might eventually be achieved in the corneal models as well. However, no optical data from this corneal model was reported. In a different approach, Han et al. (2002) have prepared a bioengineered ocular surface tissue replacement consisting of human limbal cells, believed to be corneal epithelial stem cells in a cross-linked, human, fibrin gel. The cells were suspended in a human fibronectin/fibrin gel cross-linked by human thrombin and Factor XIII, derived from a fibrinogen rich cryoprecipitate of human plasma and proliferated in the fibrin gel to near confluence over the 15 days. This bioengineered corneal surface tissue created a transportable, pliable, and stable tissue replacement. Because both the cells and the plasma components of the fibrin gel are of human origin (potentially derived from the patient), this tissue replacement represents a totally autologous bioengineered replacement tissue. However, production of the normal, stratified, epithelial architecture and adequate optical clarity was not discussed. Hybrid Collagen–Synthetic Polymer Matrix Replacements as Scaffolds Recent corneal models are based on the premise that the ideal biomaterial scaffold for achieving regeneration should duplicate the environmental conditions that direct the development of the original tissue. The “tissue template” properties of different ECM macromolecules allow for specification of different cell attractive environments. Creation of engineered tissues for regeneration of specific tissues and organs depends on the exploitation of these properties. However, in the cornea, as in many other tissues adequate tensile strength and toughness are required. An additional need in the cornea is optical transparency. To achieve these requirements, enhancement of the properties of the natural polymers with synthetic components is necessary. While cells growth in two dimensions has been shown on the surfaces of many synthetic polymers, ingrowth or encapsulation (three-dimensional growth) of living cells has only been demonstrated in a few, fully synthetic polymers, particularly polyethylene oxide, polypropylene oxide, and poly(N-isopropylacrylamide) (PNiPAAm) (Lee and Mooney, 2001; Hoffman, 2002). In contrast, many natural biopolymer hydrogels, such as those based

Regenerative Medicine in the Cornea

on alginate, fibrinogen–fibrin, chitosan, agarose, albumin, collagens, and their derivatives, are widely used to encapsulate living cells. Hydrogels of collagen I, the dominant biopolymer in the human cornea, are particularly attractive as matrix replacement type scaffolds, partly because of their strength at relatively low concentrations, resulting from the virtually rigid rod properties of the collagen Type I triple helix (Amis et al., 1985). In addition, collagen brings the cell attachment motif arginine–glycine–glutamic acid (RGD) (Pierschbacher and Ruoslahti, 1987). However, both the biodegradation resistance of collagen I and the strength of hydrogels in general at low concentrations (10 wt/vol.%) need to be enhanced by chemical cross-linking (Hoffman, 2002). A novel NiPAAm-based polymer [poly(N-isopropylacrylamide-co-acrylic acid-co-acryloxysuccinimide] or its YIGSR-modified analog (co-polymers abbreviated to Terpolymer (TERP) and TERP5, respectively), was copolymerized with Type I bovine atelocollagen to give composite hydrogels that could be molded to the curvature and dimensions of a cornea (Li et al., 2003). These hydrogels also showed high optical clarity (Figure 61.3a) with direct transmission and backscatter of visible light comparable to that of human corneas as measured by the same optical method. These collagen–co-polymer matrices had a glucose diffusion permeability coefficient (2.7  10–6 cm2/s) higher than the natural stroma (McCarey and Schmidt, 1990) although adequate insulin and albumin transport (Schneider et al., 1999) were not measured. Furthermore, they were adequately robust for suturing during surgery. The collagen–TERP5 hydrogels have been used as corneal LKP replacements, sutured into one cornea of each of a series of Yucatan microswine (Figure 61.3b) with no adverse inflammatory or immune reaction was found after implantation. Contralateral untreated corneas and pig cornea allografts served as controls. Regrowth of corneal epithelial and stromal cells into the implanted hydrogel to reconstitute the cornea was reported (Figure 61.3c and d). In addition, regeneration of functional corneal nerves was observed in transplanted corneas but not in control allografts by 6 weeks post-operative, with concomitant recovery of touch sensitivity (Figures 61.3f and 61.4). Previous studies of restoration of touch sensitivity have indicated that only minimal function is detected even 10 years after partial-thickness lenticule transplantation from a human donor cornea (Kaminski et al., 2002). In addition, Goren et al. have reported that the quality of tear fluid in eyes

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Figure 61.3 Optical clarity of the (a) collagen–TERP co-polymer compared (b) to a translucent hydrogel containing only collagen. (c) It shows the implant grafted into host tissue. Clinical confocal microscope images of the implant show (d) corneal epithelialization, and the presence of (e) stromal cells, and (f) nerve fibers (arrows) within the implant. Scale bar is 25 μm.

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Figure 61.4 Recovery of touch sensitivity in implanted tissue engineered corneas. with even slightly diminished Cochet–Bonnet corneal sensitivity scores produce tears with a significantly decreased concentration of the antimicrobial protein lactoferrin (Goren, 2002). The collagen–TERP5 implanted corneas showed restoration of the tear film mucin. Recent Advances in Corneal Tissue Engineering One of the more serious and recurring problems with the corneal scaffolds used to date is related to the mechanical properties. Although the Li et al. hydrogels were adequately robust enough for suturing, they comprised 3.5% collagen and were therefore still fairly weak. Also synthesis of the TERP co-polymer was fairly complex for scale-up. As the mechanical properties of various prototypes can be improved by using higher collagen concentrations (Trinkaus-Randall, 2000), a variety of cross-linking techniques using more widely available cross-linkers are becoming available. Duan and Sheardown have used multifunctional dendrimers as collagen cross-linkers, demonstrating that transparent collagen gels with the mechanical properties superior to those of the gels of Li et al. with lower concentration collagen solutions (Duan and Sheardown, 2005a, b) could be prepared. The presence of additional functional groups also allowed these gels to be modified with large and tunable amounts of biologically relevant functional groups. The maximum achievable YIGSR concentration of 3.1  10–2 mg/mg collagen is significantly greater than that obtained previously using the NIPAAM-based cross-linking agent at 1.6  10–6 mg/mg collagen (Li et al., 2005). Liu et al. (2005), using recombinant Type I human collagen (from Fibrogen Inc., CA), have increased both the collagen concentrations, and total solids content to about 15%, which is closer to those of human corneas. This has greatly improved the tensile strength and elasticity of the resulting tissue engineered corneas, allowing for placements of sutures without any microshearing, while maintaining optical clarity. When incorporated into interpenetrating networks with PEO-based synthetic polymers, these collagen-based hydrogels now show tensile strengths of over 550 KPa, elastic modulus of over 2,000 MPa, and approaching 50% elongation at break.

CONCLUSIONS AND FUTURE PERSPECTIVE Human corneal substitutes for transplantation can be fabricated from synthetic polymers to completely natural polymers to composites of the two. While corneal substitutes reconstructed using tissue engineering techniques from extracted to recombinant collagen appear promising, it is clear that for clinical applications of these tissue engineered corneas, further investigation is required and several technical difficulties must yet be overcome. For example, more precise interaction between the scaffolds and surrounding host tissue will be needed to ensure long-term engraftment and proper regeneration (e.g. no hypo- or hyper-proliferation of cells, or malignant transformation). So, in addition to proper mechanical and optical properties, tissue engineered corneal substitutes will

Regenerative Medicine in the Cornea

need to include or be modified by various growth factors for all in-growing host cells to behave normally in this three-dimensional system. In addition, in high risk patients, there may not be a sufficient population of progenitor or stem cells that can engraft the implant and therefore incorporation of stem cells into these constructs will be a consideration. Nevertheless, it has been shown that tissue engineered corneas may in the near future be able to supplement the supply of post-mortem human corneas harvested for transplantation, thereby meeting the demand for donor corneas.

REFERENCES Amis, E.J., Carriere, C.J., Ferry, J.D. and Veis, A. (1985). Effect of pH on collagen flexibility determined from dilute solution viscoelastic measurements. Int. J. Biol. Macromol. 7: 130–134. Aucoin, L., Griffith, C.M., Pleizier, G., Deslandes, Y. and Sheardown, H. (2002). Interactions of corneal epithelial cells and surfaces modified with cell adhesion peptide combinations. J. Biomater. Sci. Polym. Ed. 13: 447–462. Auger, F.A., Remy-Zolghadri, M., Grenier, G. and Germain, L. (2002). A truly new approach for tissue engineering: the LOEX self-assembly technique. Ernst Schering Res. Found. Workshop, 35: 73–88. Beekhuis, W.H. (1995). Current clinicians’ opinions on risk factors in corneal grafting. Results of a survey among surgeons in the eurotransplant area. Cornea 14: 39–42. Benedek, G.B. (1971). Theory of transparency of the eye. Appl. Opt. 10: 459. Bruining, M.J., Pijpers, A.P., Kingshott, P. and Koole, L.H. (2002). Studies on new polymeric biomaterials with tunable hydrophilicity, and their possible utility in corneal repair surgery. Biomaterials 23: 1213–1219. Carlsson, D.J., Li, F., Shimmura, S. and Griffith, M. (2003). Bioengineered corneas: how close are we? Curr. Opin. Ophthalmol. 14: 192–197. Chapekar, M.S. (2000). Tissue engineering: challenges and opportunities. J. Biomed. Mater. Res. (Appl. Biomater.) 53: 617–620. Chirila, T.V. (2001). An overview of the development of artificial corneas with porous skirts and the use of PHEMA for such an application. Biomaterials 22: 3311–3317. Crawford, G.J., Hicks, C.R., Lou, X., Vijayasekaran, S., Tan, D., Mulholland, B., Chirila, T.V. and Constable, I.J. (2002). The chirila keratoprosthesis: phase I human clinical trial. Ophthalmology 109: 883–889. Crawford, G.J., Eguchi, H. and Hicks, C.R. (2005). Two cases of AlphaCor surgery performed using a small incision technique. Clin. Exp. Ophthalmol. 33: 10–15. Duan, X. and Sheardown, H. (2005a). Crosslinking of collagen with dendrimers. J. Biomed. Mater. Res. 75A: 510–518. Duan, X. and Sheardown, H. (2005b).Dendrimer crosslinked collagen as a corneal tissue engineering scaffold: mechanical properties and corneal epithelial cell interactions. Biomaterials (Submitted). Duffy, P., Wolf, J., Collins, G., DeVoe, A.G., Streeten, B. and Cowen, D. (1974). Possible person-to-person transmission of Creutzfeldt–Jakob disease. N. Engl. J. Med. 290: 692–693. Evans, M.D.M., Xie, R.Z., Fabbri, M., Madigan, M.C., Chaouk, H., Beumer, G.J., Meijs, G.F., Griesser, H.J., Steele, J.G. and Sweeney, D.F. (2000). Epithelialization of a synthetic polymer in the feline cornea: a preliminary study. Invest. Ophthalmol. Vis. Sci. 41: 1674–1680. Evans, M.D.M., Xie, R.Z. and Fabbri, M. (2002). Progress in the development of a synthetic corneal onlay. Invest. Ophthalmol. Vis. Sci. 43 3196–3201. Evans, M.D.M., Taylor, S., Dalton, B.A. and Lohmann, D. (2003). Polymer design for corneal epithelial tissue adhesion: pore density. J. Biomed. Mater. Res. 64A: 357–364. Freegard, T.J. (1997). The physical basis of transparency of the normal cornea. Eye 11: 465–471. Gagnon, N., Auger, F.A. and Germain, L. (2005). Porcine corneal endothelial cell culture improvement: effect of initial seeding density and presence of a feeder layer. Invest. Ophthalmol. Vis. Sci. 46(Suppl.): 5006. Gaudreault, M., Carrier, P., Larouche, K., Leclerc, S., Giasson, M., Germain, L. and Guerin, S.L. (2003). Influence of Sp1/Sp3 expression on corneal epithelial cells proliferation and differentiation properties in reconstructed tissues. Invest. Ophthalmol. Vis. Sci. 44: 1447–1457. George, A. and Pitt, W.G. (2002). Comparison of corneal epithelial cellular growth on synthetic cornea materials. Biomaterials 23: 1369–1373.

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Germain, L., Auger, F.A., Grandbois, E., Guignard, R., Giasson, M., Boisjoly, H. and Guérin, S.L. (1999). Reconstructed human cornea produced in vitro by tissue engineering. Pathobiology 67: 140–147. Goren, M.B. (2002). LASIK and dry eye. Ophthalmology 109: 1947–1948. Griffith, M., Hakim, M., Shimmura, S., Watsky, M.A., Li, F.F., Carlsson, D., Doillon, C.J., Nakamura, M., Shinozaki, N., et al. (2002). Artificial human corneas: scaffolds for transplantation and host regeneration. Cornea 21, S1–S8. Griffith, M., Li, F., Lohman, C., Sheardown, H., Shimmura, S. and Carlsson, D.J. (2005). Tissue engineering of the cornea. Scaffolding in Tissue Engineering. CRC Press. Boca Raton, FL, USA. Han, B., Schwab, I.R., Madsen, T.K. and Isseroff, R.R. (2002). A fibrin-based bioengineered ocular surface with human corneal epithelial stem cells. Cornea 21: 505–510. Hicks, C., Crawford, G., Chirila, T., Wiffen, S., Vijayasekaran, S., Lou, X., Fitton, J., Maley, M., Clayton, A., Dalton, P., Platten, S., et al. (2000). Development and clinical assessment of an artificial cornea. Prog. Retin. Eye Res. 19: 149–170. Hicks, C., Werner, L., Vijayasekaran, S., Mamalis, N. and Apple, D.J. (2005). Histology of AlphaCor skirts. Evaluation of biointegration. Cornea 24: 933–940. Hicks, C.R., Crawford, G.J., Tan, D.T., Snibson, G.R., Sutton, G.L., Gondhowiardjo, T.D., Lam, D.S. and Downie, N. (2002). Outcomes of implantation of an artificial cornea. AlphaCor: effects of prior ocular herpes simplex infection. Cornea 21: 685–690. Hicks, C.R., Crawford, G.J., Lou, X. Tan, D.T., Snibson, G.R., Sutton, G., Downie, N., Werner, L., Chirila, T.V. and Constable, I.J. (2003a). Corneal replacement using a synthetic hydrogel cornea, AlphaCorTM: device, preliminary outcomes and complications. Eye 17: 385–392. Hicks, C.R., Crawford, G.J., Tan, D.T., Snibson, G.R., Sutton, G.L., Downie, N., Gondhowiardjo, T.D., Lam, D.S.C., Werner, L., Apple, D. and Constable, I.J. (2003b). AlphaCor™ case: comparative outcomes. Cornea 22: 583–590. Hicks, C.R., Chirila, T.V., Werner, L., Crawford, G.J., Apple, D.J. and Constable, I.J. (2004). Deposits in artificial corneas: risk factors and prevention. Clin. Exp. Ophthalmol. 32: 185–191. Hicks, C.R., Crawford, G.J., Dart, J.K.G., Grabner, G., Holland, E.J., Stulting, R.D., Tan, D.T., Bulsara, M. (2006). AlphaCor – Clinical outcomes. Cornea 25: 1034–1042. Hoffman, A.S. (2002). Hydrogels for biomedical applications. Adv. Drug Deliv. Rev. 43: 3–12. Houff, S.A., Burton, R.C., Wilson, R.W., Henson, T.E., London, W.T., Baer, G.M., Anderson, L.J., Winkler, W.G., Madden, D.L. and Sever, J.L. (1979). Human-to-human transmission of rabies virus by corneal transplant. N. Engl. J. Med. 300: 603–604. Jacob, J.T., Rochefort, J.R., Bi, J. and Gebhardt, B.M. (2005). Corneal epithelial cell growth over tethered-protein/peptide surface-modified hydrogels. J. Biomed. Mater. Res. 72B: 198–205. Johnson, G., Jenkins, M., McLean, K.M., Griesser, H.J., Kwak, J., Goodman, M. and Steele, J.G. (2000). Peptoid-containing collagen mimetics with cell binding activity. J. Biomed. Mater. Res. 51: 612–624. Kaminski, S.L., Biowski, R., Lucas, J.R., Koyuncu, D. and Grabner, G. (2002). Corneal sensitivity 10 years after epikeratoplasty. J. Refract. Surg. 18: 731–736. Khan, B., Dudenhoefer, E.J. and Dohlman, C.H. (2001). Keratoprosthesis: an update. Curr. Opin. Ophthalmol. 12: 282–287. Kim, M.K., Park, I.S. and Park, H.D. (2001). Effect of poly(ethylene glycol) graft polymerization of poly(methyl methacrylate) on cell adhesion: in vitro and in vivo study. J. Cataract. Refract. Surg. 27: 766–774. Kim, M.K., Lee, J.L., Wee, W.R. and Lee, J.H. (2002). Comparative experiments for in vivo fibroplasias and biological stability of four porous polymers intended for use in the Seoul-type keratoprosthesis. Br. J. Ophthalmol. 86: 809–814. Klenkler, B.J., Griffith, M., Becerril, C., West-Mays, J.A. and Sheardown, H. (2005). EGF-grafted PDMS surfaces in artificial cornea applications. Biomaterials 26: 7286–7296. Kobayashi, H. and Ikada, Y. (1991). Corneal cell adhesion and proliferation on hydrogel sheets bound with cell-adhesive proteins. Curr. Eye Res. 10: 899–908. Lee, K.Y. and Mooney, D.J. (2001). Hydrogels for tissue engineering. Chem. Rev. 101: 1869–1879. Legeais, J.M. and Renard, G. (1998). A second generation of artificial cornea (BiokroII). Biomaterials 19: 1517–1522. Legeais, J.M., Drubaix, I., Briat, B., Renard, G. and Pouliquen, Y. (1997). Second generation bio-integrated keratoprosthesis. Implantation in animals. J. Fr. Ophthalmol. 20: 42–48. Li, F., Carlsson, D.J., Lohmann, C.P., Suuronen, E.J., Vascotto, S., Kobuch, K., Sheardown, H., Munger, R. and Griffith, M. (2003). Cellular and nerve regeneration within a biosynthetic extracellular matrix: corneal implantation. Proc. Natl. Acad. Sci. USA 100: 15346–15351.

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Li, F., Griffith, M., Li, Z., Tanodekaew, S., Sheardown, H., Hakim, M. and Carlsson, D.J. (2005). Recruitment of multiple cell lines by collagen-synthetic copolymer matrices in corneal regeneration. Biomaterials 26: 3093–3104. Liu, L. and Sheardown, H. (2005). Glucose permeable poly(dimethyl siloxane) poly(N-isopropyl acrylamide) interpenetrating networks as ophthalmic biomaterials. Biomaterials 26: 233–244. Liu, X., et al. (2005). Tissue substitutes for cornea transplantation from recombinant human collagen. Biomaterials (Submitted). Maurice, D.M. (1957). The structure and transparency of the cornea. J. Physiol. 136: 263–286. McCarey, B.E. and Schmidt, F.H. (1990). Modelling glucose distribution in the cornea. Curr. Eye Res. 9: 1025–1039. Merrett, K., Griffith, C.M., Deslandes, Y., Pleizier, G. and Sheardown, H. (2001). Adhesion of corneal epithelial cells to cell adhesion peptide modified pHEMA surfaces. J. Biomater. Sci. Polym. Ed. 12: 647–671. Merrett, K., Griffith, C.M., Deslandes, Y., Pleizier, G., Dubé, M.A. and Sheardown, H. (2003). Interactions of corneal cells with transforming growth factor β2-modified poly dimethyl siloxane surfaces. J. Biomed. Mater. Res. 67A: 981–993. Pierschbacher, M.D. and Ruoslahti, E. (1987). Influence of stereochemistry of the sequence Arg–Gly–Asp–Xaa on binding specificity in cell adhesion. J. Biol. Chem. 262: 17294–17298. Schneider, A.I., Maier-Reif, K. and Graeve, T. (1999). Constructing an in vitro cornea from cultures of the three specific corneal cell types. In Vitro Cell. Dev. Biol. Anim. 35: 515–526. Sit, M., Weisbrod, D.J., Naor, J. and Slomovic, A.R. (2001). Corneal graft outcome study. Cornea 20: 129–133. Smolin, G. and Goodman, D. (1988). Corneal graft rejection. Int. Ophthalmol. Clin. 28: 30–36. Stern, M.E., Beuerman, R.W., Fox, R.I., Gao, J., Mircheff, A.K. and Pflugfelder, S.C. (1988). A unified theory of the role of the ocular surface in dry eye. Adv. Exp. Med. Biol. 438: 643–651. Stoiber, J.S., Fernandez, V., Kaminski, S., Lamar, P.D., Dubovy, S., Alfonso, E. and Parel, J.M. (2004). Biological response to a supradescemetic synthetic cornea in rabbits. Arch. Ophthalmol. 122: 1850–1855. Sweeney, D.F., Xie, R.Z., Evans, M.D.M., Vannas, A., Tout, S.D., Griesser, H.J., Johnson, G. and Steele, J.G. (2003). A comparison of biological coatings for the promotion of corneal epithelialization of synthetic surface in vivo. Invest. Ophthalmol. Vis. Sci. 44: 3301–3309. Trinkaus-Randall, V. (2000). Cornea. In: Lanza, R.P., Langer, R. and Vacanti, J. (eds.), Principles of Tissue Engineering. San Diego: Academic Press, pp. 471–491. Vijayasekaran, S., Chirila, T.V., Robertson, T.A., Lou, X., Fitton, J.H., Hicks, C.R. and Constable, I.J. (2000). Calcification of poly(2-hydroxyethyl methacrylate) hydrogel sponges implanted in the rabbit cornea: a 3-month study. J. Biomater. Sci. Polym. Ed. 11: 599–615. Wallace, C., Jacob, J.T., Stoltz, A., Bi, J. and Bundy, K. (2005). Corneal epithelial adhesion strength to tetheredprotein/peptide modified hydrogel surfaces. J. Biomed. Mater. Res. 72A: 19–24. Whitcher, J.P., Srinivasan, M., Upadhyay, M.P. (2001). Corneal blindness: a global perspective Bulletin of the World Health Organization 79: 214–221.

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62 Alimentary Tract Mike K. Chen

INTRODUCTION Although the alimentary tract may appear to be a simple tubular structure that begins at the esophagus and ends at the rectum, it is a rather complex organ. Grossly, structural similarities are observed along the entire alimentary tract intrinsic to its tubular architecture. However, there are distinct cellular and functional differences among the various parts of the gut that allow food and byproducts to be propelled, digested, absorbed, and excreted in an organized and efficient fashion. Organizationally, functionally, and structurally, the gut is divided into the esophagus, stomach, small intestine, and colon. Ingested food is passed into the esophagus, which serves primarily as a conduit to transport food to the stomach. The stomach stores the food, initiates digestion, and controls the rate of emptying of food particles into the small intestine. The small intestine is the primary digestive organ where food is broken down into absorbable nutrient particles and where absorption actually occurs. Waste and water are passed into the colon where water is reabsorbed and waste is stored until a socially acceptable place is available for excretion. Although all parts of the gastrointestinal (GI) tract are useful and enhance the quality of life, the only absolutely essential portion of the alimentary tract is the small intestine. In addition to its essential function in food processing and absorption, the intestinal surface must provide a barrier against unwanted entry of toxins and organisms. The GI tract contains the largest surface area and is exposed to more than 400 bacterial species and a total of 1014 microbial cells (Smith and Gorbach, 1995). If invasion by pathogens occurs, the gut acts as an immune organ to minimize the incursion and protect the host. Not only is the alimentary tract vital for nutrient absorption, it is a crucial immune organ. This chapter will provide some insight into the complex issues involved in tissue engineering the neointestine. A broad understanding of the organization and function of the various parts of the alimentary tract can be used to formulate methods that enhance one’s ability to recreate this complex and multifunctional organ. Additionally, the embryology and the regulation of growth and repair of the gut will be briefly discussed. Even a rudimentary understanding of such an elaborate topic allows one to speculate on the application of peptides and other factors that may enhance the ability to regenerate neointestine. A dysfunctional alimentary tract may be deemed so for a variety of reasons because it has so many functions. Nevertheless, many of the dysfunctional parts may be removed completely with manageable morbidity or replaced using another portion of the GI tract. However, the small intestine is currently irreplaceable except through transplantation. This is where the essential nutrient digestion and absorption occurs. Because the absence of the esophagus, stomach and colon is not as critical as the lack of small intestine, most researchers have focused on tissue engineering small bowel. This chapter will discuss the current knowledge regarding tissue engineering of all parts of the alimentary tract but the primary focus will be on the current knowledge and capacity for tissue engineering small bowel. Successful strategies as well as failures will be explored to better understand the obstacles and challenges in this field.

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INTESTINAL DEVELOPMENT AND IMMUNE FUNCTION A comprehensive discussion of alimentary tract development and regeneration is beyond the scope of this chapter. But some insight into the embryology, growth, and maintenance of the alimentary tract can aid in devising strategies that may improve research models and enhance the ability to create functioning neointestine. Embryologically, the gut develops from the primitive endodermal tube. Formation of the intestine begins with the association of the visceral endoderm with the splanchnic mesoderm and subsequent organogenesis proceeds in a proximal to distal, or cephalad to caudal pattern (Kaufman and Bard, 1999). The endoderm gives rise to the epithelium that may be structurally similar throughout the entire GI tract, but is distinctly different on a cellular basis specific to each part of the gut. The mesoderm differentiates into connective tissues and muscle layers. Similar to the formation of other organs, part of the regulation of the development depends on reciprocal signaling between the epithelium and the mesenchyme and this epithelial–connective tissue cross-talk continues to control epithelial renewal and regeneration in mature intestine (Birchmeier and Bircheier, 1993; Mills and Gordon, 2001; Lees et al., 2005). Many regulatory systems are implicated including the Hedgehog, bone morphogenetic protein (BMP), Notch, and Wnt/-catenin signaling pathways; the Hox and Sox transcription factors, and the Eph receptor/ephrin ligand signaling system (De Santa et al., 2003; Lees et al., 2005). Numerous other growth and differentiation factors are involved and the discussion is beyond the scope of this chapter. The organization of the GI tract is similar throughout its length characterized by an inner layer of mucosa surrounded by submucosa, muscle, and serosa; except for the esophagus, which lacks a serosal covering. The mucosa is formed by a layer of columnar epithelium that folds upon itself and these invaginations are termed crypts. These crypts are thought to contain intestinal stem cells that give the epithelial lining a tremendous capacity for self-renewal and regeneration (Potten and Loeffler, 1990; Potten et al., 1997). As cells differentiate and mature, they migrate out of the crypts toward the lumen of the bowel where they become senescent over the course of a few days and are shed into the lumen of the bowel. Stem cells are described as having the capacity for self-renewal as well as the ability to generate the entire adult cell complement. Description and definition of stem cells continue to be modified as more discoveries are made. Traditionally, adult stem cells are thought to be tissue or organ-specific, but recent studies have shown that this may not be entirely correct. Adult stem cells may possess more plasticity than originally thought and have the capacity to transdifferentiate. An example is the ability for transplanted adult bone marrow stem cells to transdifferentiate to form cell lineages in the mouse and human GI tract (Krause et al., 2001). Markers for stem cells in most tissues have been described but are not well characterized for intestinal stem cells. Because of the difficulty in identifying the intestinal stem cells, there is a paucity of information about them. Recent studies have shown that Musashi-1 may be a marker of intestinal stem cells (Kayahara et al., 2003; Potten et al., 2003). As we learn more about the function of the intestinal stem cells, one can postulate that its regulation may be exploited to enhance the development of neointestine. In fact, Tait and his colleagues have implanted crypt cell aggregates that presumably contain intestinal epithelial stem cells onto the flank of adult recipients and have been able to demonstrate regeneration of neointestinal tissue with presence of digestive enzyme activities and glucose transport capacity similar to that of age matched controls (Tait et al., 1995). In addition to the complicated organization of the entire alimentary tract, the intestine is also a vital immune organ. The gut is exposed to approximately 1014 microbial cells and it must efficiently recognize benign commensal bacterial as well as identify and eliminate pathogens. The small intestine is exposed to approximately 103–9 bacterial per gram of intraluminal content whereas the colonic mucosa is home to 1011–12 bacteria per gram of feces (Hao and Lee, 2004). The intestine must function as a primary immune organ to prevent and minimize invasion and infection. As an immune organ, it is composed of an innate and an adaptive system. The adaptive system is comprised of the gut-associated lymphoid tissue, commonly referred to as 1073

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GALT. Components of GALT include Peyer’s patches, plasma cells, M cells, and T cells. The adaptive system recognizes pathogens and responds by upregulating formation of immunoglobulins and other effectors. The adaptive response takes time whereas the innate system maintains constant vigilance and can respond immediately. Innate immunity is provided by Paneth cells, which are specialized epithelial cells that reside in the crypts. Paneth cells produce several antimicrobial peptides, enzymes, and pro-inflammatory cytokines that are protective for the host (Ouellette, 1999; Porter et al., 2002). These cells normally reside in the small intestine but have been identified in the colon of patients with inflammatory bowel disease (Beil et al., 1995). Vacanti and his colleagues in Boston are pioneers and continue to be the most prolific group performing intestinal tissue engineering research. They demonstrated that neointestinal cysts can be formed that has an intact mucosal immune system containing an immunocyte population similar to that of native small intestine (Perez et al., 2002). They showed that exposure of the neomucosa to the luminal content was vital to the regeneration of the immune system by anastomosing the neointestinal cyst to the native bowel. Neointestinal cysts that were not anastomosed to the native bowel had no exposure to luminal contents and resulted in rudimentary formation of the immunocyte population. The researchers used a well-described polymer–organoid construct where neonatal Lewis rat intestinal organoids are harvested and seeded onto biodegradable polymer tubes. Tait et al. (1994) previously had demonstrated that the intestinal organoids may be used to regenerate the neomucosa because they contain all the elements of the intestine including the stem cells and mesenchyme. These polymer–organoid constructs were implanted into syngeneic adult animals and neointestinal cysts formed after a variable time. The polymer tubes used were 10 mm long and 5 mm in outer diameter with an internal diameter of 2 mm. The tubes were created from sheets of a nonwoven mesh of polyglycolic acid (PGA) fibers (Smith and Nephew, Heslington, York, UK) and sprayed on the outer surface with 5% polyl-lactic acid (Mooney et al., 1996). After that, the polymer was coated with 200 μl of 0.1% collagen solution (Vitrogen 100; Collagen Corp., Palo Alto, CA). A more detailed description of this model is in the small bowel section below.

THE ESOPHAGUS The esophagus functions primarily as a transport tube that directs the progression of food and fluids from the mouth to the stomach. It is not as complex when compared to other portions of the alimentary tract. The esophagus is lined by stratified squamous mucosa and submucosa, and it has a well-developed muscularis of striated muscle in the upper third and smooth muscle in the lower two-thirds. It has no serosa and its vascular supply is not as robust as the well-vascularized intra-abdominal portions of the gut. The paucity of vascular supply to the esophagus reduces its tolerance to injury and diminishes the quality of the healed tissue. The primary function of the esophagus requires that it maintain an ability to coordinate peristaltic contraction in response to swallowing, to propel the bolus of food into the stomach. Sphincters at the upper esophagus and gastroesophageal junction reduce reflux and regurgitation. The lower esophageal sphincter located at the gastroesophageal junction acts to curtail reflux of gastric contents into the esophagus because the acidic gastric secretion is injurious to the esophageal mucosa. The clinical need for esophageal replacement occurs as a result of congenital anomalies, injury, or malignancy. Due to its relatively poor vascular supply, esophageal injury often results in stricture formation and stenosis. Most strictures can be treated with dilation, injection with steroids to reduce recurrent stricture formation, or local resection (Baskin et al., 2004). However, there are occasional needs for an isolated segment of esophagus and a tissue-engineered structure would work well in this circumstance. When complete esophageal replacement is needed, other portions of the intestine may be used (Rodgers et al., 1981; Raffensperger et al., 1996). The operative goal is to create a tubular conduit for passage of food and the results from the use of stomach and colon to

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replace the esophagus have been reasonable. Reflux and poor peristalsis are some of the more common problems and can lead to recurrent aspiration resulting in pneumonitis and pneumonia. Current tissue engineering methodology does not allow for the production of a full length of esophagus. Fortunately, as noted above, this is not an essential need because there are surgical options that work well. Nevertheless, a discussion about the obstacles is useful because the same hurdles exist for creating full-length segments of other portions of the gut. One significant obstacle is the poor vascular supply of the native esophagus, which makes regenerating a complete length of tissue-engineered neoesophagus in the mediastinum highly unlikely to be successful. Placing an engineered construct and waiting for that to regenerate in a relatively avascular area such as the mediastinum is doomed to be unproductive. Forming a neoesophagus in a heterotopic locale would be challenging as well, since one would have to find a way to maintain the new vascular supply during the transfer of the newly formed tube to the mediastinum. The use of a prosthetic tube has been tried by researchers. Fukushima et al. (1983) attempted to use a Dacron tube as a replacement for the esophagus in a dog. The Dacron tubes measured 5–7 cm in length, 1.5–2 cm in diameter and 1.5 mm in thickness. The tubes were placed as an esophageal replacement in 16 dogs; 7 of the 16 dogs survived over a year with 4 alive at 6 years. In 6 of the 7 survivors, the tubes were extruded by 6 months, but it had provided an adequate base for the formation of a thin layer of squamous epithelium. The submucosa near the anastomotic sites was robust and appeared similar to native tissue but at the central portion of the prosthetic tube, the regenerated tissue was primarily fibrous scarring without muscle or mucous glands. The regenerated tissue had architectural disarray and the best result was a stenotic tubular structure with moderate function. Nonabsorbable prostheses appear to be poor choices for tissue ingrowth and are not reasonable alternatives to current surgical techniques. The capability for tissue engineering a small segment or a patch of esophageal tissue may be less onerous. Surgeons have transferred skin and other tissues on a vascular pedicle to patch a portion of the esophagus with good results (Jurkiewicz, 1984; Harii et al., 1985; Kakegawa et al., 1987). These techniques are widely used for reconstruction after resection for strictures and malignancies. Researchers have experimented with the addition of various configurations of mesh, collagen, and silicone with moderate success (Shinhar et al., 1998; Yamamoto et al., 1999; Badylak et al., 2000, 2005; Lynen, et al., 2004). The full length replacement remained problematic with stenosis and leakage, but patching with various absorbable materials was successful. We have demonstrated that a resorbable biomaterial can serve as a patch for repairing a defect in the esophagus (Badylak et al., 2000). We fashioned a sheet of extracellular matrix (ECM) to patch a defect created in the cervical esophagus in a dog model. The two types of ECMs were small intestinal submucosa (SIS) and urinary bladder submucosa (UBS). Eighty percent of the study was based on SIS because of its established efficacy. SIS is an ECM harvested from porcine small intestine and has been used extensively in tissue engineering experiments (Matsumoto et al., 1966; Badylak et al., 1989; Kropp et al., 1995; Dalla et al., 1999; De Ugarte et al., 2003). It was initially described by Matsumoto (1966) when he used inverted small intestine to replace large veins in dogs. It has since been demonstrated as an effective scaffold for the regeneration of numerous tissues and is now commercially available in variable thickness and sizes and human uses have included repair of hernias, diaphragms, tympanic membranes, and for large wound coverage (Puccio et al., 2005; Spiegel and Kessler, 2005; Grethel et al., 2006; Smith and Campbell, 2006). In our experiment, a defect was created in the cervical esophagus approximately 2–3 cm in width and 5–6 cm long. The defects were repaired using SIS and UBS. The results from the two ECMs were similar. The patched area healed without stricture formation and the dogs ate normally and survived for the duration of the study. Early deaths occurred from leakage. The resulting histology showed relatively normal architecture and good function. Another group of dogs had a 5–6 cm segment of the cervical esophagus removed and repaired with a tubular construct made from SIS or UBS. These animals had poor outcomes. Some were

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euthanized early due to leakage and the survivors had significant stricture formation. For a variety of mechanical and biological reasons, this did not work well when a segmental replacement was done. The scaffold that had been constructed as a tube did not have enough structural integrity to maintain its shape and collapsed on itself. This collapse resulted in either leakage and death or if the animal survived, strictures were formed. Producing a scaffold that holds its shape better may allow ingrowth and tissue regeneration to occur without scar formation. In addition to the shape and structural integrity, the ECMs employed in our study are easily resorbed and may be degraded before completion of tissue ingrowth and regeneration. Finally, there is probably a finite distance over which tissue ingrowth may occur. The existing vascular supply and regulation of angiogenesis may place biological limitation on the rate at which tissues will renew itself. Thus, some of the challenges in tissue engineering large organs include the need to devise biomaterials that will enhance angiogenesis; that will last long enough for the ingrowth to occur; and that will hold its shape and maintain support during the regenerative process. Similar experiments using AlloDerm as a patch led to reepithelialization and formation of neovasculature at 1 month (Isch et al., 2001). By 2 months, the AlloDerm had been resorbed and the mucosa was completely regenerated. AlloDerm is another useful scaffold that is commercially available. It is primarily used as a dermal substitute for wound care and hernia reconstruction. AlloDerm is human skin that has been processed to remove all epidermal and dermal cells while preserving the remaining biological dermal matrix.

THE SMALL BOWEL The small bowel is the one essential part of the gut that cannot be replaced by another part of the alimentary tract. Patients may need small bowel for a variety of reasons. Babies can be born without an adequate length of intestine due to in utero ischemia. Patients with poor bowel motility cannot propel the luminal contents in an organized fashion through its length; stasis of the contents in a dysmotile tube leads to bacterial overgrowth, poor absorption, cramping abdominal pain, emesis, and even potential sepsis from bacterial translocation. These patients cannot rely on their bowel for nutrition and generally require long-term intravenous nutrition or hyperalimentation with its attendant morbidities (Scolapio et al., 1999). Many of these patients will have infections from chronic indwelling intravenous catheters and they will eventually run out of access sites due to stenosed and thrombosed veins. Additionally, long-term hyperalimentation can lead to liver dysfunction and may result in a need for liver transplantation. Intestinal ischemia and bowel resection for tumors and inflammatory bowel disease can result in short bowel syndrome. Short bowel syndrome results from an inadequate amount of small bowel to support nutritional needs and is generally defined as losing more than 75% of the small bowel. The loss of this vital absorptive mass leads to malabsorption, malnutrition, and eventual death. Surgical options for increasing the absorptive surface or slowing the transit time to enhance absorption have had limited success (Bianchi, 1999; Thompson, 1999; Weber, 1999; Javid, et al., 2005). The only therapy is small bowel transplantation and this has its own set of problems including a finite number of donors, the need for long-term immunosuppression, graft versus host disease, and potential posttransplant lymphoproliferative disorder (Botha and Horslen, 2006). Fortunately, most people are born with more small bowel than needed so that even patients with short bowel syndrome have some residual small bowel. They generally need additional absorptive surface and not a complete replacement of their small bowel. The absolute amount of small bowel required is unclear and it is dependent on the patient’s age, the amount of small bowel present, the presence or absence of the ileocecal valve, and the amount of large bowel present. This suggests that for most patients, there is no need to create an entire length of small bowel. Instead, tissue engineers may be able to offer patients relief from short bowel syndrome simply by creating additional neointestine to add to the patient’s own available absorptive surface area.

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Creating new small bowel may be done through a variety of techniques. Early researchers found that patching bowel defects using the serosal surface of another piece of intestine resulted in that serosal surface being covered with regenerated mucosa (Kobold and Thal, 1963; Binnington et al., 1973). This has led other researchers to investigate the use of biomaterials as a scaffold for ingrowth of intestine. Initial studies on the use of prosthetic materials as a patch for repair of a defect in the bowel were thought to be unsuccessful because the materials used were not resorbable. In 1979, Harmon et al. (1979) used Dacron as a patch to repair defects in the ileum of rabbits. The material provided a bridge for ingrowth of new intestinal tissue and once that occurred, the prosthetic material was extruded into the lumen. Others experimented with the placement of a polytetrafluoroethylene (PTFE) tube in continuity with the small bowel and were able to demonstrate minimal ingrowth of mucosa (Watson et al., 1980). The first report on the use of an absorbable biomaterial as a patch was by Thompson et al. (1986). They compared the use of Dacron, PTFE, and PGA mesh. They found that Dacron and PTFE grafts were extruded and PGA meshes were resorbed. Dacron and PTFE grafts were shown to be poor biological scaffolds. We have also used a biomaterial to patch defects created in the small bowel and observed regeneration of neointestine (Chen and Badylak, 2001). The biomaterial was SIS which was resorbed and the resulting neointestine was quite similar to normal small bowel. Histological evaluation showed the presence of mucosa, varying amount of smooth muscle, sheets of collagen, and an outer serosal layer. Attempts at using a tubular configuration of SIS were uniformly unsuccessful. The tubes either leaked or obstructed and this occurred primarily because there is no intrinsic integrity in SIS. In fact, it is rather a thin material that becomes soft when exposed to the moist luminal contents. Producing a scaffold that can maintain its shape and structure for a few weeks while tissue regeneration occurs may be an important component to the creation of a complete tube of neointestine. Wang et al. (2005) in Japan created an interesting model to evaluate the feasibility of creating a tubular segment of small bowel using SIS as a scaffold. A 2 cm tubular SIS graft was fashioned from Sprague Dawley rat donors and interposed in the middle of a 6 cm Thiery–Vella loop of Lewis rats. The Thiery–Vella loop is a defunctionalized segment of ileum that is brought out as a double ileostomy. A silicone stent was left in place for 3 weeks to keep the SIS open during healing. After stent removal, the loop was washed with saline. By taking this loop out of continuity with the rest of the intestine, the tubular SIS graft was given time to heal without being exposed to potentially harmful normal intraluminal contents. By 4 weeks, a mucosal epithelial layer had begun to form and this was completely covered by 12 weeks. The neomucosa had a typical morphology containing goblet cells, Paneth cells, enterocytes, and enteroendocrine cells. The outer walls of the grafts were covered with bundles of smooth muscle-like cells. Vacanti et al. (1988) first reported on the generation of intestine from minced pieces of fetal intestine. Subsequently, they created a model for generating neointestine by forming a polymer–organoid construct (Choi et al., 1997). The organoids were harvested from neonatal Lewis rats and seeded onto a biomaterial produced from polyglycolic and polylactic acid and the constructs implanted onto the omentum of an adult Lewis rat. The presumption was that the organoids contain all the elements of the small bowel including stem cells, immunocytes, and mesenchyme. The resulting cysts were observed to be attached to a vascular pedicle and contained mucoid material in the center with a histology that mirrored that of normal small bowel. Using this model, they characterized the neointestine cysts by investigating the presence of brush border enzymes, basement membrane components, electrophysiological properties, immune ontogeny, and lymphangiogenesis (Choi et al., 1998; Perez et al., 2002; Duxbury et al., 2004). Additionally, they anastomosed the cysts to native small bowel after massive small bowel resection and demonstrated increased weight gain and proposed that this may be one method for rescuing patients with short bowel syndrome (Grikscheit et al., 2004).

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Our laboratory has also employed this model to investigate the feasibility of using other biomaterials for creating neointestinal cysts. In particular, we investigated the practicality of using commercially available biomaterials as scaffold for regenerating neointestine. We selected commercially produced SIS (SurgisisR)and fibrin glue (TisseelR) because they represent two distinct types of biomaterials and are readily available for clinical use (unpublished data). SIS, described before, is an acellular matrix derived from porcine intestine and fibrin glue is most commonly used for controlling hemorrhage. The components of fibrin glue are natural biological factors that include fibrinogen, thrombin, calcium chloride, and fibrinolysis inhibitor (aprotinin). Recently, some investigators have used it as a cell delivery vehicle for urothelial cells (Bach et al., 2001), and chondrocytes mixed with fibrin glue have resulted in formation of cartilage structures (Westreich et al., 2004). Fibrin glue may be an ideal biomaterial because it can be used as a vehicle to deliver cells and its inherent adhesive property ensures that the seeded cells are adhered to the desired surface. The components are easily resorbed and nontoxic. Internal matrices are formed after the glue sets into an elastic coagulum which provides scaffolding for the cells to reside and for influx of nutrients. The glue and cell construct may also be molded to the desired form prior to implantation. Finally, there may be inherent growth factors present in the fibrin glue mixture that are yet to be defined. For the experiments, intestinal organoids were harvested from neonatal rodents and processed as reported by Tait et al. (1994). Briefly, the intestine is stripped of its mesentery and washed in cold Hanks’ balanced salt solution (HBSS) until clean. It is then cut into 2–3 mm fragments and washed 8 times with HBSS. The fragments are then minced into 1 mm3 pieces. These fragments are digested with dispase (0.1 mg/ml) and collagenase (300 U/ml) at room temperature on an orbital shaking platform (80 rpm) for 25 min. The mixture is run through the pipet approximately 150 times. HBSS is added to the contents at 5 times of the volume and transferred to a sterile tube. Sediment is obtained under gravity for 60 sec., and then the excess fluid is removed. This is repeated 3 times. Ten milliliter of Dulbecco’s modified eagle’s medium (DMEM) with 2.5% fetal bovine serum (FBS) and 2% sorbitol is then added and the solution is centrifuged at 300 rpm for 2 min. This is repeated 5–6 times until the supernatant is clear. The final pellet is then resuspended in the growth medium ready for seeding onto the biomaterials. An average of 40,000 organoid units are isolated per small intestine harvested from each neonatal rat and approximately 30,000–50,000 organoid units are used per construct. Both rats and mice were used in our studies. Vacanti and his colleagues demonstrated that this model worked well in rats and we decided to determine the viability of this model in mice in preparation for studies that involved transgenic animals. Approximately 30,000–50,000 organoids were seeded onto each construct. The organoids were placed on a 1 cm two sheet of SIS, which was then folded upon itself. Another construct was made by mixing the organoids with the fibrin glue. The glue was then allowed to solidify into a coagulum, which has a gelatin-like consistency. A final construct was formed by placing the organoids on the SIS and then adding fibrin glue to hold the construct together. Unlike Vacanti’s group, the constructs were not shaped into a tube or sphere. The constructs were then implanted onto the omentum of an adult syngeneic rat or mouse and the neointestinal cysts were evaluated at 6–12 weeks. Our results showed that the animals tolerated the procedure well and all survived the implantation. Cysts were identified at 6 weeks and the gross appearance was that of a normal segment of bowel with an inner layer of mucosa covered by the serosa (Figure 62.1). The constructs made using fibrin glue with SIS formed cysts that had the best histology which were best identified at 10 weeks (Figure 62.2). Despite not creating a tube or a sphere when forming the constructs, the regenerated cysts had organized architecture similar to normal intestine. The organoids had the capacity to organize itself so that epithelial cells lined the inner layer forming a mucosa with crypts and villi. The submucosa was intact and surrounded by smooth muscle and covered

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Figure 62.1 Opened neointestinal cyst formed from a fibrin glue/SIS/intestinal organoid construct. The cyst measured 1.5–2 cm with mucoid materials in the center. Grossly, the cyst resemble normal intestine with an inner mucosal layer and an outer serosal lining.

by a serosal layer. The well-developed cysts measured 1.5–3 cm in diameter and had a vascular pedicle attached to the omentum (Figure 62.3). Another important aspect of intestinal tissue engineering is the ability for the intestine to repair and regenerate. Numerous enterocyte mitogens have been investigated including epidermal growth factor, hepatocyte growth factor, fibroblast growth factor, neurotensin, growth hormone, transforming growth factor, interleukin-11, glucagon-like peptide-2 (GLP-2), and glutamine (Walters, 2004). A comprehensive discussion about these factors is beyond the scope of this chapter but suffice it to say that manipulation of these factors may also enhance our ability to generate enough functional neointestine to make it a clinically relevant therapeutic option. Ramsanahie from Vacanti’s laboratory evaluated the effects of GLP-2 on the neointestinal cysts (Ramsanahie et al., 2003). GLP-2 is an endogenous regulatory peptide with a specific potent trophic effect on intestinal mucosal growth and increases the expression of Naglucose cotransporter 1 (SGLT1). Indeed, they found that the administration of GLP-2 enhanced the mucosal growth and increased the expression of SGLT1 which suggests that growth factors may be used to stimulate the regeneration of neointestine.

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Figure 62.2 Neointestinal cyst mucosa in a fibrin glue/SIS/intestinal organoid construct at 10 weeks. The H&E stain shows proliferation of crypts, villi, and submucosal tissue.

Additionally, the study shows that tissue-engineered tissues are responsive to regulatory factors and can be used as a model to study its response to exogenously applied components.

THE STOMACH AND COLON As previously noted, the stomach and colon are not vital to life. Nevertheless, they perform important functions that enhance the quality of life. The stomach is an important storage organ that allows one to eat a few large meals daily and reduces the need for constant grazing. It initiates the digestive process that involves acid production and enzymatic degradation of the food bolus. The stomach is important in enteroendocrine regulation by producing gastrin and somatostatin. Additionally, vitamin B12 absorption is dependent on the production of intrinsic factor by parietal cells. Partial replacement of the stomach has been done using a collagen sponge scaffold as a patch after partial gastrectomy (Hori et al., 2001). The sponge was produced by a using a combination of collagen extracted from pig skin that is cross-linked and reinforced by PGA felt. Beagle dogs were used for the experiment and they were kept without oral food for 14 days. Additionally, a silicone sheet was used as a patch on the luminal side to protect the scaffold and subsequently removed endoscopically 4 weeks after placement. They showed some regeneration at 4 weeks with complete coverage by regenerated tissue by 16 weeks. The histological finding confirmed the presence of mucosa and a thin muscular layer. They demonstrated that acid production capacity was present in the regenerated stomach wall but the contractile response to acetylcholine was poor (Hori et al., 2002).

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Figure 62.3 The regenerated cyst has a well-formed vascular pedicle when the construct is placed in the omentum. The vascular pedicle allows this well-vascularized cyst to absorb nutrients and mobility for moving it to the desired area for anastomosis to the native bowel.

Vacanti’s group also evaluated the feasibility of creating new stomach tissue by using the same technique that they described for regenerating small bowel (Grikscheit et al., 2003). They harvested organoids from the stomach of neonatal and adult Lewis rats. The organoids were seeded onto PGA scaffolds to form constructs that were implanted into the omentum of adult syngeneic rats. They showed that 98% of all animals generated new stomach tissue and that immunohistochemistry for α-actin smooth muscle and gastrin verified the presence of a smooth muscle layer and a well-developed gastric epithelium. All elements of the native rat stomach including gastric pits and squamous layers were identified. The colon is important for water and sodium resorption and as a storage pouch for waste products. As discussed earlier there is a paucity of studies regarding colon tissue engineering because having an intact colon is not essential to life. Patients who undergo total colectomy can have ileal reservoir created but they can still suffer from inflammation of the pouch (pouchitis), malabsorption, diarrhea, cramping abdominal pain, and fever (Meagher et al., 1998). Grikscheit and associates (2005) have shown that colonic tissue may be created by using the similar technique employed for small bowel tissue engineering. They harvested organoid units from the sigmoid colon of

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neonatal Lewis rats, adult rats, and tissue-engineered colon itself. The donor tissues were seeded onto a polymer scaffold previously described and implanted into the omentum of syngeneic adult Lewis rats. They found that tissue-engineered colon was generated by all tissue sources and that the neocolon architecture was similar to native tissue. The muscularis propria stained positively for actin and acetylcholinesterase was detected in the lamina propria in a linear distribution with presence of ganglion cells. In vitro Ussing chamber studies indicated appropriate transport parameters and barrier function. When anastomosed to the native bowel, there was gross visualization of fluid absorption and no obstruction was observed.

CONCLUSION The alimentary tract is a complicated organ that serves several vital functions. At the most basic level, it must propel and absorb nutrients while acting as a barrier against unwanted entry. Investigators have used the gut’s inherent capacity for healing and tissue regeneration to produce neointestine. The ability to form a small patch of the alimentary tract is feasible employing a variety of techniques, but our goal should be to create enough intestinal tissue to make it clinically useful and relevant for the patients. Hopefully, that day is not too distant as we learn more about biomaterials, regulatory peptides, and principles of intestinal growth and regeneration.

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Kakegawa, T., Machi, J., Yamana, H., Fujita, H. and Tai, Y. (1987). A new technique for esophageal reconstruction by combined skin and muscle flaps after failure in primary colonic interposition. Surg. Gynecol. Obstet. 164: 576–578. Kaufman, M.H. and Bard, J.B.L. (1999). The gut and its associated tissues. In: Kaufman, M.H. and Bard, J.B.L. (eds.), The Anatomical Basis of Mouse Development. London: Academic Press, pp. 129–152. Kayahara, T., Sawada, M., Takaishi, S., Fukui, H., Seno, H., Fukuzawa, H., Suzuki, K., Hiai, H., Kageyama, R., Okano, H. and Chiba, T. (2003). Candidate markers for stem and early progenitor cells, Musashi-1 and Hes1, are expressed in crypt base columnar cells of mouse small intestine. FEBS Lett. 535: 131–135. Kobold, E.E. and Thal, A.P. (1963). A simple method for management of experimental wounds of the duodenum. Surg. Gynecol. Obstet. 116: 340–344. Krause, D., Theise, N., Collector, M., Henegariu, O., Hwang, S., Gardner, R., Neutzel, S. and Sharkis, S. (2001). Multiorgan, multi-lineage engraftment by a single bone marrow-derived stem cell. Cell 105: 369–377. Kropp, B.P., Badylak, S. and Thor, K.B. (1995). Regenerative bladder augmentation: a review of the initial preclinical studies with porcine small intestinal submucosa. Adv. Exp. Med. Biol. 385: 229–235. Lees, C., Howie, S., Sartor, R.B. and Satsangi, J. (2005). The hedgehog signaling pathway in the gastrointestinal tract: implications for development, homeostasis, and disease. Gastroenterology 129: 1696–1710. Lynen Jansen, P., Klinge, U., Anurov, M., Titkova, S., Mertens, P.R. and Jansen, M. (2004). Surgical mesh as a scaffold for tissue regeneration in the esophagus. Eur. Surg. Res. 36: 104–111. Matsumoto, T., Holmes, R.H., Burdick, C.O., Heisterkamp, C.A. and O’Connell, T.J. (1966). The fate of inverted segment of small bowel used for the replacement of major veins. Surgery 60: 739–743. Meagher, A.P., Farouk, R., Dozois, R.R., Kelly, K.A. and Pemberton, J.H. (1998). 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63 Liver Cell-Based Therapy – Bioreactors as Enabling Technology Jörg C. Gerlach, Mariah Hout, Keneth Gage, and Katrin Zeilinger

NEED FOR INNOVATIVE THERAPIES IN LIVER DISEASE Liver disease presents a significant social and economic burden. The American Liver Foundation estimates that one in ten people have some form of liver disease and 26,000 people die each year from liver disease (American Liver Foundation, 2001). Patients with liver disease, fulminant, and chronic can progressively worsen until they require orthotopic liver transplantation (OLT). In 2002, 16,929 people were listed for liver transplantation (LTx) in the United States, only 4,778 cadaveric donor liver transplants were made, and 1,756 patients died while awaiting transplant (United Network for Organ Sharing (UNOS), 2003). Acute liver failure (ALF) has a poor prognosis with mortality rates between 50% and 90% under conservative management (Chapman et al., 1990). Over the last 50 years, advances have been made in the understanding of the pathophysiology encompassing the clinical features of ALF: encephalopathy, cerebral edema, hemorrhage, electrolyte and metabolic disturbances, renal failure, cardiovascular instability, and increased risk of infection. While cerebral edema is the most common cause of death, multi-organ failure and sepsis are also associated with significant mortality. Depending on etiology, the survival rate in ALF under conservative treatment ranges from 79% (amanita intoxication) to 10% (cryptogeneic genesis). Introduction of LTx as therapeutic option reduced mortality to 20–40% (Caraceni and Van Thiel, 1995). Cell-based therapies involve the transplantation of regenerative cells to replace damaged tissue. In addition to liver disease, aimed treatments include heart disease (Strauer et al., 2001), diabetes (Kaczorowski et al., 2002), Parkinson’s disease (Isacson, 2003), and muscular dystrophy (Deasy and Huard, 2002). Successful control of hepatic cell isolation and in vitro culture will enable the use of bioartificial organs to support patients until organ recovery or whole organ transplantation. In addition to developing treatments for liver diseases (Sauer et al., 2003a), extracorporeal organ support could be of interest for kidney diseases (Humes et al., 2002a).

METHODS OF LIVER SUPPORT OLT is the only accepted effective treatment of acute fulminant or endstage liver disease. However, the procedure is associated with some risks, and there are side effects of the required immunosuppressive follow-up therapy. Moreover, the number of donor organs is limited and organs may be available to late in the clinical

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course of a patient (Lucey et al., 1997). A medical need for a treatment modality that can “bridge” patients to transplant or slow or reverse the progression of acute liver disease is clearly demanded. With the continued, growing disparity between the numbers of organ donations and patients waiting for LTx, efforts have been made to design extracorporeal methods to support or replace the failing liver. The frequent lack of donor livers for urgent transplantation in ALF highlights the need for a liver support therapy until an organ becomes available. Moreover, patients with the capacity for liver recovery could be bridged to regeneration and would not require transplantation at all. One area in this field is the development of extracorporeal blood artificial detoxification methods, similar to kidney dialysis (Stange et al., 2002). These techniques based on detoxification cartridges, for example with adsorbents, and bioreactors are not required. Another area is the development of cell transplantation methods, where regenerative cells are injected, for example into the hepatic artery, to enable repair of injured liver tissue in the organ (Strom et al., 1999). These developments may require bioreactors for growing, or expanding, of cells to address the availability and the logistics of progenitor cells applications.

PROSPECTS OF THE USE OF HEPATIC CELLS FOR EXTRACORPOREAL LIVER SUPPORT Progress in hepatocyte tissue engineering and in vitro maintenance of differentiated function of primary liver cells has led to initial clinical pilot studies of extracorporeal support of patients in ALF using bioreactors incorporating hepatocyte cultures (Watanabe et al., 1997; Mazariegos et al., 2001, 2002; Gerlach et al., 2002; Morsiani et al., 2002; Sauer et al., 2003a). While encouraging, the non-significant results from the only pioneering Phase-III efficacy trial conducted to date (Stevens et al., 2001) indicates that much further progress in understanding and controlling tissue engineering of complex hepatocyte bioreactors is required. One of the challenges is the creation of neo-vascularized tissue constructs at high cell density in vitro that avoid central necrosis and exhibit recapitulation of the sinusoidal microvasculature of the liver. Bioartificial liver devices (BAL) have undergone a rapid evolution in the last decade. “First generation” BAL devices utilize commercial, off-the-shelf filtration, and plasmapheresis units for the bioreactor core. Some clinical work with clinical BAL systems was performed using such commercially available hollow-fiber cartridges for dialysis or plasmaseparation (two-compartment systems). “Second-generation” bioreactors were developed as specialized units intended to address the shortcomings of commercial available cartridges. Table 63.1 summarizes literature, describing specific bioreactor constructions for regenerative medicine applications, tested in animal experiments; Table 63.2 summarized constructions which were used clinically; Table 63.3 describes culture models which were not yet scaled up to a bioreactor; and Table 63.4 describes bioreactor constructions, tested in vitro. Bioreactors that are currently used for primary liver cell culture and clinical liver support are hollow fiber based and exhibit two functional compartments (Sussman and Kelly, 1993; Watanabe et al., 1997). Xu et al. (1999) pointed out, that four bioreactor compartments are necessary to enable integral oxygenation and distributed mass exchange with low gradients. We took this challenge and developed a bioreactor specific for clinical liver support (Figure 63.1) that accommodated 400–800 g of primary cells (Gerlach et al., 2001) and can be used for growing liver progenitors. We focused on a spontaneous reassembly of primary cells inoculated into a bioreactor and their establishment of a scaffold or biomatrix. We have shown that a homogeneous mix of adult liver cells from organ collagenase digestion containing parenchymal hepatocytes, non-parenchymal cells (such as sinusoidal endothelial cells, stellate cells, and liver progenitor cells) will restructure after injection into specific bioreactors to form well-defined liver structures, such as neo-sinusoidal structures and neo-spaces of Dissé, reminiscent of the native liver. The vasculature of organs can be regarded as complex structures, which supply plasma to the cells surrounding the network of capillaries. Oxygen supply is enhanced by the perfusion of hemoglobin-containing

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Table 63.1 Bioreactor constructions tested in animal models Technology

Cell type

Citation

Hollow fiber-based bioartificial liver with integral oxygenation (Gerlach CellModule) Spirally wound flat sheet and hollow fiber-based bioartificial liver with integral oxygenation Flat plate bioartificial liver with integral oxygenation Hollow fiber-based renal tubule assist device

Porcine liver cells

Gerlach et al. (1993b, 1994e, 2001), Janke et al. (1997)

Porcine hepatocytes

Flendrig et al. (1999)

Porcine hepatocytes

Shito et al. (2003)

Human renal tubule cells

Humes et al. (2002b)

Table 63.2 Bioreactors for hybrid kidney support and bioartificial liver support, tested in clinical trials Bioreactor technology

Hollow fiber-based renal tubule assist device Hollow fiber-based, two compartments (Vital Therapies ELAD®) Hollow fiber-based, two compartments, with hepatocytes attached to dextran microcarriers (Arbios Systems HepatAssist®) Hollow fiber-based, two compartments (Excorp Medical BLSS®) Hollow fiber-based, four compartments, with integral oxygenation (Gerlach MLS CellModule) Perfused scaffolt with oxygenation fibers (two-compartment system)

Additional therapeutic components

Perfusate

Cell type

Citation

Human renal tubule cells Human hepatoblastoma cell line (C3A) Porcine hepatocytes

Humes et al. (2002a)

Blood

Porcine hepatocytes

Mazariegos et al. (2002)

Plasma

Porcine and human liver cells

Plasma

Porcine cells

Mundt et al. (2002a), Irgang et al. (2003), Sauer et al. (2003a, b) Van de Kerkhove et al. (2003)

Plasma

Charcoal perfusion chamber

MLS concept may involve dialysis and artificial detoxification

Plasma

Sussman et al. (1992) Watanabe et al. (1997)

erythrocytes. Mass exchange is enhanced by pulsation in blood flow and alterations in the autonomous capillary resistance. Using various independently perfused interwoven networks of hollow fibers, our approach was to mimic the native larger vascular structure of an organ using an artificial capillary bed. We addressed the lack of erythrocytes by flowing oxygen through one compartment. Mass exchange was addressed by perfusing two independent capillary systems with patient plasma or culture media. Small artificial hollow fiber

Liver Cell-Based Therapy – Bioreactors as Enabling Technology

Table 63.3 Culture models for in vitro cell growth and maintenance Technology

Cell type

Citation

Cell entrapment within calcium alginate Collagen sandwich culture Collagen gel immobilization, perfusable culture Hollow fiber-based bioartificial liver with integral oxygenation (Gerlach CellModule) Micropatterned borosilicate wafers

Rat hepatocytes Rat hepatocytes Human hepatocytes

Miura et al. (1986, 1990) Dunn et al. (1989) Koebe et al. (1994a, b)

Porcine liver cells

Gerlach and Neuhaus (1994), Gerlach et al. (1994b, 1995)

Rat hepatocytes and 3T3 fibroblasts

Bhatia et al. (1997)

Table 63.4 Bioreactor constructions tested in vitro Technology

Cell type

Citation

Early perfusion chambers

Chick heart fibroblasts, human malignant epithelial cells, Chinese hamster cells, hybridomas Bone marrow-derived osteoblasts

Mouse fibroblasts, human choriocarcinoma cells, Reuber hepatoma cells, human hepatocytes Human leukemic cell lines

Christiansen et al. (1953), Rose (1954), Freed (1963), Katinger (1985) Minucells (Minucells and Minutissue Vertriebs GmbH) VectraCell gas-permeable bags (Diagnostic Chemicals Limited), Rotary Cell Culture System (Synthecon, Inc.), Wave Bioreactor (Wave Biotech LLC), CELLine (Integra Biosciences AG), miniPERM Bioreactor (Sartorius AG), CellMax (Spectrum Laboratories, Inc.), Tecnomouse (Integra Biosciences AG) AastromReplicell (Aastrom Biosciences, Inc.) Knazek et al. (1972), Wolf and Munkelt (1975), Hager et al. (1978, 1983) Gloeckner and Lemke (2001)

Rat hepatocytes

Macdonald et al. (2001)

Porcine and human liver cells

Gerlach et al. (1994c, 1996d, 2003a, b), Sauer et al. (2002b) Flendrig et al. (1997)

Commercially available perfusion chambers Commercially available systems for non-adherent cells

Commercially available system for bone marrow expansion Hollow fiber-based bioreactors

Hollow fiber-based bioreactor with integral oxygenation Coaxial hollow fiber-based bioreactor with integral oxygenation Hollow fiber-based bioartificial liver with integral oxygenation (Gerlach CellModule) Flat sheet and hollow fiber-based bioartificial liver with integral oxygenation

Hybridomas

Hematopoetic stem cells

Porcine hepatocytes

(Continued )

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Table 63.4 (Continued) Technology

Cell type

Citation

Hollow fiber-based renal tubule assist device (Humes RAD) Flat membrane bioreactor with integral oxygenation Flat plate bioartificial liver with integral oxygenation Micropatterned borosilicate wafers Biodegradable polymer bioreactor constructed via 3D printing Microfabricated bioreactor constructed via ion etching of silicon wafers Titanium mesh bioreactor

Porcine renal tubule cells

Humes et al. (1999)

Porcine hepatocytes

De Bartolo et al. (2000)

Porcine hepatocytes

Shito et al. (2001)

Rat hepatocytes and 3T3 fibroblasts Rat liver cells

Bhatia et al. (1997) Kim et al. (1998)

Bioreactor containing hydrated polyester fibers and porcine autologous biomatrix Bioreactor containing non-woven polyurethane matrix with integral oxygenation

Rat hepatocytes

Powers et al. (2002a, b)

Rat bone marrow stromal osteoblasts Porcine hepatocytes (1010)

Bancroft et al. (2003)

Rat or pig hepatocytes

Linti et al. (2002)

Ambrosino et al. (2002)

Figure 63.1 Clinical use of our 600 g cell-compartment bioreactor (Gerlach & Sauer, Charitié, Berlin, Germany).

Liver Cell-Based Therapy – Bioreactors as Enabling Technology

MEDIUM I O2/CO2

MEDIUM II (a)

(b)

(c)

Figure 63.2 Schemetic depiction of the four-compartment stem cell bioreactor. (a) Smallest repeating cell culture unit within the bioreactor. (b) Analytical scale stem cell bioreactor. (c) Laboratory scale stem cell bioreactor (Mckeel and Gerlach).

capillary subunits, in which interwoven membranes represent the organs secondary level of vascular supply, are simultaneously perfused. The design focused on transforming the previously published two-compartment bioreactor constructions into a technology with up to five compartments. A construction principle was developed to form independently perfused subunits from interwoven synthetic capillaries, each with a different function, and arrange these subunits in parallel (Gerlach et al., 1996c). To scale up the bioreactor, additional subunits are added resulting in a capillary system (Gerlach and Neuhaus, 1994; Gerlach et al., 1994c), similar to the natural organ vasculature (Figure 63.2). The extracapillary space forms a compartment where the inoculated liver cells and hepatocyte supporting cells (Gerlach et al., 1993a, d, 1994a, 1996a) reside (Gerlach et al., 1989). Cell migration processes reform tissue-like structures between the capillaries (Figure 63.3). Under conditions of capillary perfusion, parenchymal and non-parenchymal tissue formation (Gerlach et al., 1995) is possible by spontaneous organization to cellular aggregates and their attachment to the surface of the capillaries (Gerlach et al., 1993c, 1994d). Within these aggregates a spontaneous reorganization of the smallest vascular unit of the liver, sinusoid-like structures, is regularly seen.

THE SOURCE OF CELLS FOR LIVER SUPPORT The question of cell source for extracorporeal liver support has been the subject of numerous controversial discussions and remains a topic of major importance. The functionality of the cell source can be considered from the standpoint of pure hepatocyte function or from overall liver function. From a biotechnology standpoint, primary cells can be isolated, including primary stem cells; and spontaneously developed or genetechnologically modified cell lines could be available. Cells may derive from the patient (autologous cells, including adult stem cells), from human tissue (homologous cells, including from embryonic stem cells), or from an animal cell source (xenogeneic cells). The ideal cell source would possess the following characteristics: functional equivalence to human liver cells, unlimited (but controllable) expansion capability, and minimal patient risk from either an immunogenic response or transfer of infection. Safety of the cell source can be viewed in terms of potential risk to the patient, which is the primary concern, or of risk to the BAL device and the incorporated cells, the failure of which can also lead to severe consequences for the patient. The route from BAL to patient is the major concern when considering the possibility of infectious transfer, although the reverse could conceivably result in BAL dysfunction during long-term applications. One concern of safety of the cell source is immunogenic response. In the traditional organ transplantation practice, immunogenic responses are categorized as either graft versus host and host versus graft. These

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Cell Module Primary human liver cells

Dialysis fluid

High flux dialysis filter

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HA-Solution

Detox Module CSPAD Dialysis Module CWHDF

Waste

Modular extracorporeal liver support

Figure 63.3 MLS concept with the cellbioreactor module, a detoxmodule (albumin dialysis), and a dialysis module (renal support).

responses can be further subdivided into those that are primarily cell mediated and those based on humoral factors. The distinction is important as a number of current BAL designs incorporate membrane barriers that minimize the likelihood of cell-mediated damage in either direction (host to graft and vice versa). However, it is misleading to consider the risks inherent in a particular cell source without also taking the design of BAL device into account. One open question in technology development in the field is whether extracorporeal whole blood bioreactor perfusion systems can be developed for long-term patient operation or whether this has to continue to base on plasma perfusion bioreactor technologies. Both modes of operation have advantages, as summarized in Table 63.5, and the associated risks for the patient and the device differ significantly. The level of differentiation present in a cell source provides a rough idea of the overall performance of the cells in terms of specific metabolic functions. Differentiation can be considered along a spectrum, with the prototypical undifferentiated cells being early embryonic cells which retain the capacity to become any cell type in the body (totipotency) and possess remarkable proliferative potential. In general, these two characteristics tend to be lost as a cell progresses through differentiation, resulting in terminally differentiated cells that express a specific repertoire of functions and have little, if any, proliferative potential. Adult (primary) hepatocytes have the advantage of terminal differentiation and can express the full range of hepatocyte-specific functions in the proper environment. Initial clinical studies employing primary adult porcine liver cells for extracorporeal liver support were performed by others and by our group (Allen et al., 2001; Sauer et al., 2003a). However, the use of primary porcine cells, which are used in most of the current

Liver Cell-Based Therapy – Bioreactors as Enabling Technology

Table 63.5 Advantages and disadvantages of whole blood perfusion of BALs Advantages

Disadvantages

Reduction in circuit complexity Improved O2 and CO2 exchange Improved mass exchange between patient and BAL

Potential for detrimental blood interactions (thrombosis, hemolysis) Increased potential for bioreactor loss

technologies, is controversial for various reasons, including the possible transfer of porcine endothelial retroviruses and possible immunologic reactions by the xenogeneic proteins produced (Baquerizo et al., 1999).

HUMAN LIVER CELLS IN BIOREACTORS FOR LIVER SUPPORT Although there is a remarkable degree of similarity in liver function between animal species, differences do exist. A multivariate analysis of variance of patients treated with extracorporeal liver perfusion between 1964 and 2001 revealed that only the use of baboon and human livers provides an independent positive prognostic marker for improved survival (Pascher et al., 2002). This analysis supports the thesis that xenogenic hepatocytes are not an optimal substitute for the complex tasks of a human liver. It is necessary to consider that in the studies cited, human livers of impaired quality and therefore unsuitable for LTx are compared with porcine livers of ideal quality. In order to avoid the drawbacks involved in using porcine cells and human liver tumor cell lines, the use of primary human liver cells seems to present a promising cell source. Primary human liver cells, obtained from explanted organs found to be unsuitable for transplantation, are an interesting cell source as they are ethically acceptable and are capable of performing human metabolism and regulation (Tsiaoussis et al., 2001). According to Eurotransplant data, and the data of the European and American organ procurement organizations, approximately 20–25% of all explanted livers are unsuitable for transplantation and therefore discarded. This number corresponds to the number of patients with ALF requiring bridging to LTx (Annual Report/Eurotransplant International Foundation, 2000). However, the initial viability of those cells is impaired by preceding organ preservation and the isolation procedure. Our hypothesis is that, with appropriate logistics, after recovery from the preservation and isolation injury in a four-compartment bioreactor, cells from these organs could serve the demand for initial clinical studies on extracorporeal liver support. In one of our studies, cells were isolated from 54 human livers from discarded transplants. A cell mass of 400–600 g was obtained enabling the clinical application of a liver lobe equivalent to a hybrid organ. Freshly isolated cell preparations contained about 10% of non-parenchymal cells, as estimated microscopically using size-selective criteria; 18 cell isolation procedures (33.3%) failed because of unsatisfactory cell viability (40%) and/or unsuccessful separation during the washing procedure, due to a higher grade of initial organ impairment. In 36 cases (66.7%), the isolation was performed successfully. From these isolation procedures, 2.8 to 6.4  1010 hepatocytes were co-cultured in the bioreactors (n  36) with the non-parenchymal cells of the same liver. Trypan blue viability of the cells prior to charging the bioreactors was 55.0 15.9%. After inoculation into bioreactors, metabolic activities of the cells in vitro were maintained over at least 3 weeks of culturing. BIOREACTOR CULTURE OF LIVER CELLS AND CLINICAL APPLICATIONS OF OUR OWN DEVELOPMENT Maintenance of non-dividing primary liver cells and their differentiated functions was demonstrated in the above-described own bioreactor development over a period of 2 months and may continue longer (Gerlach

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et al., 1994b, 1996d; Gerlach, 1997). This bioreactor can accommodate a number of cells that would allow studies in extracorporeal liver support (Gerlach et al., 1996b; Sauer and Gerlach, 2002a, b). Animal studies showed encouraging results for the clinical use of the bioreactor (Gerlach et al., 1993b, 1994e; Bornemann et al., 1996; Gerlach, 1996; Gleisner et al., 1997; Janke et al., 1997). Utilizing porcine cells, the device was used as a bridge to LTx in a phase-I study with eight patients in ALF, coma stage III–IV (Mundt et al., 2002b). To date, the patient and organ survival rate is 100% (Irgang et al., 2003; Sauer et al., 2003a). A pilot study in Berlin, Germany, utilized the device with primary human liver cells harvested from organs explanted for transplantation but subsequently discarded due to steatosis, cirrhosis, or mechanical injury (Sauer and Gerlach, 2002b). The initial results on patients demonstrate the feasibility of employing primary human liver cells in clinical applications with specific bioreactor systems, and that the non-dividing cells recover from organ preservation and cell isolation injury (Gerlach et al., 2002; Sauer et al., 2003c).

THE CONCEPT OF MODULAR LIVER SUPPORT An extracorporeal liver support system has to support or substitute the main functions of the liver, providing detoxification, synthesis, and regulation. To date, developments concentrate on either artificial detoxification systems (such as albumin dialysis (Stange et al., 1999) or adsorber suspension with activated charcoal particles and ion exchange resin (Ash et al., 1998; Willinger et al., 1999; Kramer et al., 2001)) or on biologic systems (like whole organ perfusion and bioreactors with liver cells (Gerlach, 1996)). For detoxification (e.g. removal of bilirubin, bile acids, and toxins) simple, dialysis-like artificial detoxification systems have been shown to be efficient. The complex tasks of regulation (e.g. CNS transmitter precursors) and synthesis (e.g. coagulation factors) remain to be addressed by the use of human liver cells. The Modular Liver Support (MLS) concept combines different extracorporeal therapy units, tailored to suit the individual clinical needs of each patient (Figure 63.4). The CellbioreactorModule is a specific bioreactor charged with primary human liver cells, harvested from explanted livers found to be unsuitable for transplantation. The DetoxModule enables albumin dialysis for the removal of albumin-bound toxins. The DialysisModule for continuous veno-venuous hemofiltration can be added to the system if required. The mobile trolley-mounted liver support system consists of a blood circuit with a continuous plasma separation unit (e.g. CRRT, B. Braun, Melsungen, Germany), a high-flux dialysis filter (Fresenius Medical Care, Bad Homburg, Germany), and a second circuit for plasma perfusion of the bioreactor (800 ml-bioreactor, Hybrid Organ, Berlin, Germany). The simple set-up allows the integration of any other standard renal replacement therapy device. Venous access is gained by placing a double-lumen dialysis catheter into either the internal jugular or the femoral vein. Blood is pumped through a hollow fiber plasma filter (Plasmaselect 0.4, Braun) at a rate of 150–250 ml/min. If necessary, a continuous infusion of heparin was performed for anticoagulation in order to achieve an activated clotting time of 160–180 s (ACT, Fresenius, Oberursel, Germany). For continuous exchange with the CellbioreactorModule, the bioreactor is connected to the plasma circuit in counter-directional flow mode at 150–200 ml/min. The total extracorporeal volume is approximately 110 ml in the blood circuit and 900 ml of plasma in the bioreactor and associated circuitry. The CellbioreactorModule is the multi-compartment bioreactor for extracorporeal liver support therapy, as described above. In our actual concept, this module is loaded with human liver cells, harvested from organs explanted for LTx but subsequently discarded due to steatosis, cirrhosis or mechanical injury. Cells are obtained through a five-step collagenase liver perfusion (Gerlach et al., 1994a). Under conditions of capillary perfusion, parenchymal and non-parenchymal cells form tissue by spontaneous organization to cellular aggregates, immobilized to the surface of the capillaries. Within these aggregates, channels are formed, representing neosinusoidal structures with a reformation of a neo-space of Dissé. The cells produce their own biomatrix – the use

Liver Cell-Based Therapy – Bioreactors as Enabling Technology

SEM

Capillary membranes replace the liver lobuli vasculature and enable mass exchange at high density

LM

Perfusion channels, representing sinusoids, are rebuilt by the cells at high-density formations

Toluidine-blue

Non-parenchymal cells migrate on the surface of the neo-hepatocyte plates

A neo-space of Dissé is formed spontaneously TEM TEM

lto cells produce natural biomatrix in the neo-space of Dissé (serum free culture)

Figure 63.4 Microscopic “zoom” through the self-reassembled tissue structures formed by human primary liver cells in a four-compartment hollow fiber-based bioreactor.

of fetal calf serum and additional (animal-derived) biomatrix collagen can be avoided. A cell mass of 400–600 g enables the clinical application of a liver lobe equivalent to a hybrid organ. The bioreactor constructions were investigated in vitro (Gerlach et al., 1995) and in vivo (Gerlach et al., 1994e). The results are summarized above. During the stand-by phase of 21 days (mean) prior to therapeutic use, the bioreactors are characterized routinely concerning metabolic activity on a daily basis and contamination can be excluded. In liver failure the insufficient metabolism of endogenous toxins has been shown to be fatal. Most of these toxins are albumin bound. In previous reports, liver assist devices based on albumin dialysis were found to eliminate these toxins (Seige et al., 1999; Mitzner et al., 2001). The DetoxModule enables albumin dialysis via a standard high-flux dialysis cartridge. Single-pass albumin dialysis (SPAD) is a simple implementation of albumin dialysis using standard renal replacement therapy machines (Kreymann et al., 1999): The patient’s blood flows through a circuit with a high-flux hollow fiber hemodiafilter (Fresenius HdF 100S polysulfone high-flux

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hemodiafilter, Fresenius AG, Bad Homburg). The other side of this membrane is cleansed by an albumin solution in counter-directional flow and discarded after passing the filter; 1l of a 4.5 l bag with standard bicarbonate buffered dialysis solution is replaced by 1,000 ml of 20% human albumin solution, resulting in 4.44% albumin solution. During therapy, the blood pump speed is adjusted to 130–180 ml/min; the dialysis pump speed is 600 ml/h. 30–75% of all patients in ALF show renal failure with fluid overload, electrolyte derangement, and high levels of creatinin. Therefore, the MELS concept facilitates an integrated renal replacement therapy in terms of a continuous veno-venous hemodiafiltration via the high-flux hollow fiber hemodiafilter as part of the DetoxModule. A standard buffered aqueous solution is used (added after the filter, “postdilution”) with a flow rate of 1,000–3,000 ml/h. The proposed system for extracorporeal liver support is modular and combines cell-based therapy with different forms of dialysis and detoxification. One modular system for all relevant extracorporeal liver support therapies facilitates handling on the intensive care unit and reduces costs for disposables, service, and technical support. Modularity allows the clinician to adapt the extracorporeal liver support to the individual, actual needs of the patient suffering from liver failure. Reduction of albumin-bound toxins or drugs, reduction of the biochemical burden prior to LTx, or reduction of bilirubin levels in cases of pruritic hyperbilirubinemia are easily treated with albumin dialysis in the DetoxModule. In severe cases of liver failure with hepatic encephalopathy, the CellbioreactorModule is added. The additional DialysisModule facilitates the support of the failing kidneys in hepato-renal syndrome. Therapy can be started with the DetoxModule and the DialysisModule in every intensive care unit before the more sophisticated and logistically more demanding CellbioreactorModule is added. A liver support system has to provide detoxification (e.g. toxins, ammonia, bilirubin, endotoxins), synthesis (e.g. albumin, amino acids, coagulation factors), and regulation (e.g. acid–base status, electrolytes, amino acids, CNS energy supply, CNS transmitter precursors). In this modular approach we consider the principal task of the CellbioreactorModule in regulation and synthesis, secondary in detoxification. The DetoxModule enables detoxification in terms of removal of toxins from the patient’s blood, as well as the reduction of the biochemical burden of the cells inside the CellbioreactorModule. In addition, it replaces the bile excretion of the cells inside the bioreactor. The DialysisModule as renal replacement therapy permits regulation and detoxification. The novel bioreactor technology enabled spontaneous cell reorganization in the 600 g range as well as human tissue function over the period of several weeks. By replacing the secondary capillary structure of the liver lobuli with artificial hollow fibers for nutrition, mass exchange and oxygenation (Gerlach et al., 1990), the cells spontaneously reformed the primary capillary structure of liver sinusoids thus supporting a larger cell mass at a higher density (Sauer et al., 2002a). However, the limiting factor of liver support still remains the availability of primary human cells, as the present source is from discarded organs intended for transplantation. Thus, although an attractive technology with therapeutic potential is available, the availability of the cell source is crucial to enable further clinical studies. Stem cell research seeks to provide therapeutic solutions for this problem.

LIVER PROGENITOR CELLS AND BIOREACTOR CULTURES TO STUDY CELL SOURCES FOR EXTRACORPOREAL LIVER SUPPORT Stem cells – unspecialized cells that perpetuate themselves through self-renewal and generate specialized cells through differentiation (Reya et al., 2001) – may be the ideal cell source for cell-based therapies developed to treat debilitating and life-threatening diseases (Ringe et al., 2002). Stem cells are characterized by their potential to

Liver Cell-Based Therapy – Bioreactors as Enabling Technology

differentiate into mature cell types that may have therapeutic benefit in the support, repair, and regeneration of tissues damaged by disease or trauma (Odorico et al., 2001). To realize the full therapeutic potential of stem cells, many investigators are seeking in vitro methods for maintaining the unspecialized state of stem cells, promoting the proliferation of stem cells, and directing the differentiation of stem cells to the needed specialized cells. Objective of our work was to down scale the four-compartment 3D stem cell bioreactor technology platform to study embryonic, fetal, and adult stem cells in a tightly controlled environment (Figure 63.2). During development, hepatic induction in mice begins as early as E9 and a liver bud is recognizable at this stage (Jung et al., 1999; Zaret, 2000). In collaboration with the laboratories of Paul Monga, Steve Strom, and Eric Lagasse, Department of Pathology at the University of Pittsburgh, we have utilized microscopically dissected embryonic livers from different stages of mice liver development and performed collagenase cell dissociation. A progenitor pool was maintained over at least 3 weeks in bioreactors, exhibiting a growing cell mass, differenciation toward liver tissue while the early progenitors were also increasing in number (Monga and Gerlach, 2006). Initial bioreactor cultures of human fetal liver cell preparations exhibited comparable results to those of mice fetal liver cells.

CHALLENGES REGARDING STEM CELL RESEARCH FOR LIVER SUPPORT DEVELOPMENT Further advancement of stem-cell-related research may evolve new methods in transplantation medicine such as cell transplantation and extracorporeal bioreactor application for temporary support of a failing organ. In selected indications initial efforts are being made to replace the actual organ transplantation with adult liver cell transplantation. However, stem cell transplantation may be more suitable than transplantation of the fully differentiated adult derivatives since these cells still exhibit significant proliferation activities. The current challenge that prevents stem cells from use in these applications is directing the full differentiation of their progeny in vitro. In order to further develop new therapies, adult and embryonic stem cell research focuses on maintenance, proliferation and differentiation, and tissue formation in vitro. Numerous groups already work on derivation and characterization of specific stem cell lineages, but the underlying mechanisms are only partly understood. Several fundamental questions remain to be answered, for example:

• • • • • • • • • •

Can we control proliferation/differentiation of selected stem cells in vitro? Can we maintain the genotype/phenotype stability of selected stem cells in vitro? Can specific microenvironmental conditions be used to control in vitro maintenance of selected stem cells? Can, after inoculation of selected isolated stem cells, tissue restructuring be achieved by the cells themselves? Can tissue formation by selected stem cells be induced and controlled in vitro? Can we establish in vitro a tissue density of a larger number of selected stem cells without central necrosis? Can a reproducible proliferation/differentiation and utilization of selected stem cells be achieved in vitro? Can specific macroenvironmental conditions, such as 3D high-density co-culture with integrated oxygenation and decentralized mass exchange, be used to control in vitro maintenance of human embryonic stem cells in a larger cell mass? Can a phenotypic stabilization of selected stem cells and the maintenance of a progenitor cell pool be achieved by 3D high-density co-culture with integrated oxygenation and decentralized mass exchange? Does the integration of a more physiological tissue macroenvironment by the hollow fiber membranes into a growing cell mass result in a genotypic/phenotypic stabilization of proliferating human embryonic stem cells?

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• •

Can a compartmentalized co-culture of feeder cells and human embryonic stem cells better support human embryonic stem cells (hESC) in a growing cell mass? Can we maintain the stability of human embryonic stem cells in vitro, while they are proliferating as a larger mass?

For several topics, experimental animal source or biopsied human tissue, as well as conventional petridish in vitro culture methods seem to provide appropriate tools for investigation. However, investigations focusing on the impact of exogenous factors could benefit from the use of purpose-built bioreactors that enable 3D high-density perfusion co-culture. There is a considerable need for such in vitro stem cell systems, since the stem-cell-derived tissues must be capable of stable and long-term integration; into a bioreactor or, after transplantation, into existing physiological tissues, at least until they are replaced by the body’s own tissue repair process, or permanently if self-repair is not possible (Auchincloss and Bonventre, 2002). At least some of such studies require reproducible and controllable in vitro conditions. We believe that to guide tissue assembly by the cells themselves, maintenance, and proliferation/differentiation of stem cells, a switch from the conventional two- or three-compartment systems toward a fourcompartment technology is necessary. Our working hypothesis states that 3D cell–cell contact of various stem cell populations in a perfused macroenvironment, providing decentralized mass exchange at physiological tissue gradients and integral oxygenation, will allow a better approximation to the in vivo situation. There, embryonic stem cells form microvascular structures and adult stem cells reside and act in such microvascularized tissues. A fourcompartment bioreactor technology platform allows addressing the technical realization. Furthermore we hypothesize that stem cells themselves can create their own typical microenvironment in such in vitro culture models, and adult stem cells benefit from parenchymal/non-parenchymal cell co-culture in such systems for the creation of an organo-typical microenvironment.

BIOREACTOR TECHNOLOGY OVERVIEW FOR STEM CELL CULTURE In vitro control of adult liver stem cells differentiation by mediators is studied by many groups. Of the progenitors, the “small hepatocytes/oval cells” were described more in detail (Evarts et al., 1987; Mitaka et al., 1992, 1999). The effects of growth factors, cytokines or mediators and their role on growth/differentiation of fetal liver stem cells, as well as establishing of cell lines was also studied (Yoon et al., 1999). Hepatocyte growth factor (HGF), for example, is generally known to stimulate hepatocytes to replicate in vitro, or oncostatin M, a multifunctional cytokine of the interleukin-6 family, has been shown to both induce maturation of mice fetal hepatocytes in vitro (Kamiya et al., 1999; Kojima et al., 2000; Sakai et al., 2002), and to attenuate fetal liver hematopoiesis (Kinoshita et al., 1999). Aim of advanced bioreactor technology development for mamalian tissue culture is to allow cells to restructure functional tissue with organ functions ex vivo. Additionally, bioreactor technology developers aim to further the understanding of the behavior of cells and the mechanisms of tissue formation. Many different culture models for possible bioreactor constructions have been developed and tested in vitro, as summarized in Table 63.4. It was pointed out by several authors that 3D culture is important not only in accommodating an organlike mass but also in enhancing physical cell-to-cell contact, accumulation of extra cellular matrices, and local growth factor delivery, resulting in much higher biochemical effects on the maturation of fetal mouse hepatocytes then those in monolayer cultures (Miyoshi et al., 2000; Jiang et al., 2002). The general aim of bioreactor technology development for primary stem cells is the understanding of modeling of cells to allow recovery from isolation injury in vitro and to restructure functional tissue with

Liver Cell-Based Therapy – Bioreactors as Enabling Technology

organ functions ex vivo. Furthermore, bioreactor technology aims to contribute to further the understanding of the behavior of cells, and to control mechanisms of tissue formation. This implies to engineer technology to create functional tissue in vitro. Several technologies have been developed, as summarized in Table 63.4. The development of specific 3D tissue-density bioreactor technology for stem cells and co-cultures appears important in the context of the development of culture models and specific cells, as well as for subsequent translation of the findings to work on animal models, establish cell banks and possible clinical trials. Only few data are available about theoretical considerations in the development of such bioreactors. Current bioreactor designs span a wide range of geometries, flow arrangements, and exchange conditions. Selected bioreactor motifs (e.g. multi-axial hollow fiber (MacDonald et al., 2001) and flat-plate bioreactors (Shito et al., 2001)) have been modeled using simple engineering equations that relate operating parameters to conditions within the local cellular environment (e.g. plasma flow rates in cell chambers, oxygen tension profiles, etc.) and specific biological performance measures (e.g. detoxification, albumin production, etc.). These idealized geometries are excellent tools for investigating the impact of the microenvironment upon cell behavior and function; however, the extension of these designs toward clinical dimensions has been a challenge. In contrast, the four-compartment hollow fiber-based bioreactor technology presented here has been scaled up from small analytical devices (cell mass of 0.1 g) toward clinical bioartificial liver support units (cell mass of 600 g). It is only in the past decade that advanced numerical techniques such as computational fluid dynamics (CFD), systems analysis, and other approaches have been applied to the study of artificial organs. To date, much of the focus has been on artificial organs with less complexity than that exhibited by BAL, and yet questions remain about the validity of models for these systems. However, it can be expected that CFD methods will help to improve and further develop bioreactors for hepatic cell culture and clinical application.

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Sauer, I.M., Kardassis, D., Zeilinger, K., Pascher, A., Grunwald, A., Pless, G., Irgang, M., Kraemer, M., Puhl, G., Frank, J., Müller, A.R., Steinmüller, T., Denner, J., Neuhaus, P. and Gerlach, J.C. (2003a). Clinical extracorporeal hybrid liver support – phase I study with primary porcine liver cells. Xenotransplantation 10: 460–469. Sauer, I.M., Zeilinger, K., Pless, G., Kardassis, D., Theruvath, T., Pascher, A., Goetz, M., Neuhaus, P. and Gerlach, J.C. (2003b). Extracorporeal liver support based on primary human liver cells and albumin dialysis – treatment of a patient with primary graft non-function. J. Hepatol. 39: 649–653. Sauer, I.M., Zeilinger, K., Pless, G., Kardassis, D., Theruvath, T., Pascher, A., Mueller, A.R., Steinmueller, T., Neuhaus, P. and Gerlach, J.C. (2003c). Extracorporeal liver support based on human liver cells and albumin dialysis – treatment of a patient with primary graft non-function. J. Hepatol. 39(4): 649–653. Seige, M., Kreymann, B., Jeschke, B., Schweigart, U., Kopp, K.F. and Classen, M. (1999). Long-term treatment of patients with acute exacerbation of chronic liver failure by albumin dialysis. Transplant. Proc. 31: 1371–1375. Shito, M., Kim, N.H., Baskaran, H., Tilles, A.W., Tompkins, R.G., Yarmush, M.L. and Toner, M. (2001). In vitro and in vivo evaluation of albumin synthesis rate of porcine hepatocytes in a flat-plate bioreactor. Artif. Organs 25: 571–578. Shito, M., Tilles, A.W., Tompkins, R.G., Yarmush, M.L. and Toner, M. (2003). Efficacy of an extracorporeal flat-plate bioartificial liver in treating fulminant hepatic failure. J. Surg. Res. 111: 53–62. Spectrum Laboratories, Inc., Rancho Dominguez, California, USA. Stange, J., Mitzner, S.R., Risler, T., Erley, C.M., Lauchart, W., Goehl, H., Klammt, S., Peszynski, P., Freytag, J., Hickstein, H., Lohr, M., Liebe, S., Schareck, W., Hopt, U.T. and Schmidt, R. (1999). Molecular adsorbent recycling system (MARS): clinical results of a new membrane-based blood purification system for bioartificial liver support. Artif. Organs 23: 319–330. Stange, J., Hassanein, T.I., Metha, R., Mitzner, S.R. and Bartlett, R.H. (2002). The molecular absorbents recycling system as a liver support system based on albumin dialysis: a summary of preclinical investigations, prospective, randomized, controlled clinical trial, and clinical experience from 19 centers. Artif. Organs 26: 103–110. Stevens, A.C., Busuttil, R., Han, S., et al. (2001). An interim analysis of a phase II/III prospective randomized, multicenter, controlled trial of the HepatAssist® bioartificial liver support system for the treatment of fulminant hepatic failure. Hepatology 34(Suppl S, Pt 2): 509. Strauer, B.E., Brehm, M., Zeus, T., Gattermann, N., Hernandez, A., Sorg, R.V., Kogler, G. and Wernet, P. (2001). [Intracoronary, human autologous stem cell transplantation for myocardial regeneration following myocardial infarction]. Dtsch Med. Wochenschr. 126: 932–938. Strom, S.C., Chowdhury, J.R. and Fox, I.J. (1999). Hepatocyte transplantation for the treatment of human disease. Semin. Liver Dis. 19: 39–48. Sussman, N.L. and Kelly, J.H. (1993). Improved liver function following treatment with an extracorporeal liver assist device. Artif. Organs 17(1): 27–30. Sussman, N.L., Chong, M.G., Koussayer, T., He, D.E., Shang, T.A., Whisennand, H.H. and Kelly, J.H. (1992). Reversal of fulminant hepatic failure using an extracorporeal liver assist device. Hepatology 16: 60–65. Synthecon, Inc., Houston, Texas, USA. Tsiaoussis, J., Newsome, P.N., Nelson, L.J., Hayes, P.C. and Plevris, J.N. (2001). Which hepatocyte will it be? Hepatocyte choice for bioartificial liver support systems. Liver Transpl. 7: 2–10. United Network for Organ Sharing (UNOS). http://www.optn.org/latestData/ 2003. Van de Kerkhove, M.P., Di Florio, E., Scuderi, V., Mancini, A., Belli, A., Bracco, A., Scala, S., Zeuli, L., Di Nicuolo, G., Amoroso, P., Calise, F. and Chamuleau, R.A. (2003). Bridiging a patient with acute liver failure to liver transplantation by the AMC-bioartificial liver. Cell Transplant. 12(6): 563–568. Watanabe, F.D., Mullon, C.J., Hewitt, W.R., Arkadopoulos, N., Kahaku, E., Eguchi, S., Khalili, T., Arnaout, W., Shackleton, C.R., Rozga, J., Solomon, B. and Demetriou, A.A. (1997). Clinical experience with a bioartificial liver in the treatment of severe liver failure. A phase I clinical trial. Ann. Surg. 225: 484–491; discussion 491–484. Wave Biotech LLC, Bridgewater, New Jersey, USA. Willinger, M., Schima, H., Schmidt, C., Huber, L., Vogt, G., Falkenhagen, D. and Losert, U. (1999). Microspheres based detoxification system: in vitro study and mathematical estimation of filter performance. Int. J. Artif. Organs 22: 573–582.

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Wolf, C.F. and Munkelt, B.E. (1975). Bilirubin conjugation by an artificial liver composed of cultured cells and synthetic capillaries. Trans. Am. Soc. Artif. Intern. Organs 21: 16–27. Xu, A.S.L., Luntz, T.L., Macdonals, J.M., Kubota, H., Hsu, E., London, R.E. and Reid, L.M. (1999). Lineage and biology and liver. In: Lanza, R.P., Langer, R., Vacanti, J. (eds.), Principles of Tissue Engineering Chapter 41, 2nd edn, San Diego, CA: Academic Press, 559–598. Yoon, J.H., Lee, H.V.-S., Lee, J.S., Park, J.B. and Kim, C.Y. (1999). Development of non-transformed liver cell line with differentiated-hepatocyte and urea-synthetic functions: applicable for bioartificial liver. Int. J. Artif. Organs 18: 2127–2136. Zaret, K.S. (2000). Liver specification and early morphogenesis. Mech. Dev. 92: 83–88.

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64 Intracorporeal Kidney Support James J. Yoo, Akira Joraku, and Anthony Atala

INTRODUCTION End stage renal failure is a devastating disease which involves multiple organs in affected individuals. Although currently available treatment modalities, including dialysis and transplantation, can prolong survival for many patients, problems such as donor shortage, complications, and graft failure remain a continued concern (Chazan et al., 1991; Feldman et al., 1996; Amiel et al., 2000). Numerous investigative efforts have been commenced in order to improve, restore, or replace renal function. Cell-based approaches for kidney tissue regeneration have been proposed recently as an alternative method. In this chapter, we describe various cell-based approaches to achieve functional intracorporeal kidney support. The kidney is considered as one of the more challenging organs to reconstruct in the genitourinary system, due to its complex structure and function. Normal renal function includes synthesis of 1,25-vitamin D3, erythropoietin, glutathione, and free radical scavenging enzymes. The kidney also participates in the catabolism of low molecular weight proteins and in the production and regulation of cytokines (Frank et al., 1993; Stadnyk, 1994). Because these functions can not be replaced with dialysis therapy, long-term consequences, such as anemia and malnutrition, are prevalent in these patients. Limitation of current therapies for renal functional augmentation has led investigators to pursue alternative therapeutic modalities. The concept of cell transplantation using tissue engineering techniques has been proposed as a method to improve, restore, or replace renal function (Atala, 1997, 1999; Amiel and Atala, 1999; Humes et al., 1999). The emergence of tissue engineering and regenerative medicine strategies has presented alternative possibilities for the management of pathologic renal conditions (Figure 64.1). Augmentation of either isolated or total renal function with kidney cell expansion may be a feasible solution. We have followed an approach which involves the development of intracorporeal support systems for renal functional replacement.

BASIC PRINCIPLES OF KIDNEY TISSUE REGENERATION Components required to achieve partial or total kidney function are renal cells, three-dimensional scaffolds, and an in vivo environment. The challenge associated with renal cell culture is due to the unique structural and cellular heterogeneity present within the kidney. The system of nephrons and collecting ducts is composed of multiple functionally and phenotypically distinct segments. For this reason, appropriate conditions need to be provided for the long-term survival, differentiation, and growth of the cells. Extensive research has been performed in order to determine optimal growth conditions for renal cell enrichment (Milici et al., 1985; Carley et al., 1988; Horikoshi et al., 1988; Humes and Cieslinski, 1992; Schena, 1998). Isolation of particular cell types that produce specific factors, such as erythropoietin, may be a feasible approach for selective cell therapies. However, total renal function would not be achieved if specific cell types were separately isolated. To reconstitute kidney tissue that would deliver renal function, cells composing the functional

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Figure 64.1 A strategy for engineering of renal tissue. A patient with end stage renal failure undergoes a percutaneous biopsy. Renal cells are grown, expanded in culture, and seeded onto three-dimensional support system, which is then implanted into the same patient. nephron units may be preferable. Based on the literature and our experience, we were able to obtain optimal growth conditions for a stable cell culture system for kidney tissue reconstitution. Renal cells grown in culture are able to maintain their cellular characteristics (Lanza et al., 2002). When primary renal cells are placed in a collagen-based three-dimensional culture system, they are able to reconstitute into renal structures. Recent efforts in the area of kidney tissue regeneration were focused toward the development of a reliable cell source. Multipotent or progenitor cells have been proposed as a promising source due to their potential to differentiate into several cell lineages. Bone marrow-derived human mesenchymal stem cells have been shown to exhibit plasticity and differentiation potential into several different cell types (Prockop, 1997). These cells have been shown to participate in the kidney development when they are placed in a rat embryonic niche that allows for continued exposure to repertoire of nephrogenic signals (Yokoo et al., 2005). Another potential cell source is circulating stem cells, which are also known to participate at the site of kidney regeneration such as tubular and glomerular epithelial cells, podocytes, mesangial cells, and interstitial cells after renal injury (Ito et al., 2001; Poulsom et al., 2001; Gupta et al., 2002; Iwano et al., 2002; Kale et al., 2003; Lin et al., 2003; Rookmaaker

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et al., 2003). Although bone marrow cells were found to contribute to regeneration of damaged glomerular endothelial cells, the major cell source of kidney regeneration was found to originate from intrarenal cells in an ischemic renal injury model (Ikarashi et al., 2005; Lin et al., 2005). Although isolated renal cells are able to retain their phenotypic and functional characteristics in culture, transplantation of these cells in vivo may not result in structural remodeling. In addition, cell or tissue components may not be implanted in large volumes due to the limited diffusion (Folkman and Hochberg, 1973). Thus, a cell-support matrix is necessary to allow diffusion of nutrients across the entire implant. A variety of synthetic and naturally derived materials has been examined in order to determine the ideal support structures for the regeneration of urologic tissue (Tachibana et al., 1985; Atala et al., 1995, 2006; Oberpenning et al., 1999; El-Kassaby et al., 2003). Biodegradable synthetic materials, such as poly-lactic and glycolic acid polymers, have been used to provide structural support for cells. Synthetic materials can be easily fabricated and configured in a controlled manner. Naturally derived materials, such as collagen, laminin, and fibronectin, are biocompatible and provide a similar extracellular matrix environment as normal tissue. For this reason, collagen-based scaffolds have been preferred in many tissue applications (Hubbell et al., 1991; Wald et al., 1993; Freed et al., 1994; Mooney et al., 1996).

CREATION OF RENAL STRUCTURES IN VIVO The kidney is responsible not only for urine excretion but also for several other important metabolic functions. Our initial study involved in investigating the feasibility of achieving renal cell growth, expansion, and in vivo reconstitution using tissue engineering techniques (Atala et al., 1995). New Zealand white rabbits underwent nephrectomy and renal artery perfusion with a non-oxide solution which promoted iron particle entrapment in the glomeruli. Homogenization of the renal cortex and fractionation in 83 and 210 μm sieves with subsequent magnetic extraction yielded three separate purified suspensions of distal tubules, glomeruli, and proximal tubules. The cells were plated separately in vitro and seeded onto biodegradable polyglycolic acid scaffolds. Polymer scaffolds were implanted subcutaneously into host athymic mice. This included implants of proximal tubular cells, glomeruli, distal tubular cells, and a mixture of all three cell types. Polymers alone served as controls. Animals were sacrificed at 1 week, 2 weeks, and 1 month after implantation and the retrieved scaffolds were analyzed. An acute inflammatory phase and a chronic foreign body reaction were seen, accompanied by vascular ingrowth by 7 days after implantation. Histologic examination demonstrated progressive formation and organization of the nephron segments within the polymer fibers with time. Renal cell proliferation in the cell–polymer scaffolds was detected by in vivo labeling of replicating cells with the thymidine analog bromodeoxyuridine (BrdU). BrdU incorporation into renal cell DNA was identified immunocytochemically with monoclonal anti-BrdU antibodies. These results demonstrated that renal specific cells can be successfully harvested, survive in culture, and attach to artificial biodegradable polymers. The renal cell–polymer scaffolds can be implanted into host animals where the cells replicate and organize into nephron segments, as the polymer, which acts as a cell delivery vehicle, undergoes biodegradation. The initial experiments demonstrated that implanted cell–polymer scaffolds gave rise to renal tubular structures. However, it was unclear whether the tubular structures reconstituted de novo from dispersed renal elements, or they merely represented fragments of donor tubules, which survived intact. Further investigation was conducted in order to examine the tubular reconstitution process (Fung et al., 1996). Mouse renal cells were harvested, grown, and expanded in culture. Subsequently, single isolated cells were seeded on biodegradable polymers and implanted into syngeneic hosts. Renal epithelial cells were observed to reconstitute into tubular structures in vivo. Sequential analyses of the retrieved implants over time demonstrated that renal epithelial cells first organized into a cord-like structure with a solid center. Subsequent canalization into a hollow tube

Intracorporeal Kidney Support

could be seen by 2 weeks. Histologic examination with nephron segment specific lactins showed successful reconstitution of proximal tubules, distal tubules, loop of Henle, collecting tubules, and collecting ducts. These results showed that single suspended cells are capable of reconstituting into tubular structures, with homogeneous cell types within each tubule.

REGENERATION OF FUNCTIONAL RENAL TISSUE IN VIVO The kidneys are critical to body homeostasis because of their excretory, regulatory, and endocrinologic functions. The excretory function is initiated by filtration of blood at the glomerulus, and the regulatory function is provided by the tubular segments. Although our prior studies demonstrated that renal cells seeded on biodegradable polymer scaffolds are able to reconstitute into renal structures in vivo, complete renal function could not be achieved due to the type and structural configuration of polymers used. In our subsequent study we explored the feasibility of creating a functional artificial renal unit, wherein urine production could be achieved (Yoo et al., 1996). Mouse renal cells were harvested and expanded in culture. The cells were seeded onto a tubular device constructed from polycarbonate (4 μm pore size), connected at one end with a silastic catheter which terminated into a reservoir. The device was implanted in the subcutaneous space of athymic mice. Animals were sacrificed at 1, 2, 3, 4, and 8 weeks after implantation and the retrieved specimens were examined histologically and immunocytochemically. Fluid was collected from inside the implant, and uric acid and creatinine levels were determined. Histologic examination of the implanted device revealed extensive vascularization, formation of glomeruli, and highly organized tubule-like structures. Immunocytochemical staining with anti-osteopontin antibody, which is secreted by proximal and distal tubular cells and the cells of the thin ascending loop of Henle, stained the tubular sections. Immunohistochemical staining for alkaline phosphatase stained proximal tubule-like structures. Uniform staining for fibronectin in the extracellular matrix of newly formed tubes was observed. The fluid collected from the reservoir was yellow and contained 66 mg/dl uric acid (as compared to 2 mg/dl in plasma) suggesting that these tubules are capable of unidirectional secretion and concentration of uric acid. The creatinine assay performed on the collected fluid showed an 8.2-fold increase in concentration, as compared to serum. These results demonstrated that single cells form multicellular structures and become organized into functional renal units that are able to excrete high levels of solutes through a urine-like fluid. To determine whether renal tissue could be formed using an alternative cell source, nuclear transplantation was performed to generate histocompatible tissues. The feasibility of engineering syngeneic renal tissues in vivo using cloned cells was investigated (Lanza et al., 2002). In this study nuclear material from bovine dermal fibroblasts were transferred into unfertilized enucleated donor bovine eggs. Renal cells from the cloned embryos were harvested, expanded in vitro, and seeded onto three-dimensional renal devices (Figure 64.2a). The devices were implanted into the back of the same steer from which the cells were cloned, and were retrieved 12 weeks later. Functioning renal units were created from cells cloned from bovine fibroblasts. Urine production and viability were demonstrated after transplantation back into the nuclear donor animal despite expressing a different mtDNA haplotype (Figure 64.2b). Chemical analysis suggested unidirectional secretion and concentration of urea nitrogen and creatinine. The devices revealed formation of organized glomeruli and tubular structures (Figure 64.2c). Immunohistochemical and reverse transcription-polymerase chain reaction (RT-PCR) analysis confirmed the expression of renal mRNA and proteins, whereas delayed-type hypersensitivity testing and in vitro proliferative assays showed that there was no rejection response to the cloned cells. This study indicates that the cloned renal cells are able to form and organize into functional tissue structures, which are genetically same as the host. Generating immune-compatible cells using therapeutic cloning techniques is feasible and may be useful for the engineering of renal tissues for autologous applications.

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(a)

Polyethylene reservoir (3.5 × 3.5 cm)

(b)

(c) Glomerulus

Renal unit (3 cm) Tubule Silastic catheter Polycarbonate membrane Coated collagen

Polycarbonate membrane

Figure 64.2 Formation of functional renal tissue in vivo. (a) Renal device. (b) Tissue-engineered renal unit shows the accumulation of urine-like fluid. (c) There was a clear unidirectional continuity between the mature glomeruli, their tubules, and the polycarbonate membrane.

In our previous study, we showed that renal cells seeded on synthetic renal devices with a collecting system are able to form functional renal structures with urine-like fluid excretion. However, a naturally derived tissue matrix with existing three-dimensional kidney architecture would be preferable, in that it would allow for transplantation of a large number of cells for the creation of greater renal tissue volumes. We developed an acellular collagen-based kidney matrix, which is identical to the native renal architecture. In a subsequent study we investigated whether the collagen-based matrices could accommodate large volumes of renal cells which could proliferate and form kidney structures in vivo (Amiel et al., 2000). Acellular collagen matrices, derived from porcine kidneys, were obtained through a multiple step decellularization process. Serial evaluation of the matrix for cellular remnants was performed using histochemistry, scanning electron microscopy (SEM), and RT-PCR. Mouse renal cells were harvested, grown, and seeded on 80 collagen matrices at a concentration of 30  106 cells/ml. Forty cell–matrix constructs grown in vitro were analyzed 3 days, 1, 2, 4, and 6 weeks after seeding. The remaining 40 cell–matrices were implanted in the subcutaneous space of 20 athymic mice. The animals were sacrificed 3 days, 1, 2, 4, 8, and 24 weeks after implantation for analyses. Gross, SEM, histochemical, immunocytochemical, and biochemical analyses were performed. SEM and histologic examination confirmed the acellularity of the processed matrix. RT-PCR performed on the kidney matrices demonstrated the absence of any RNA residues. Renal cells seeded on the matrix adhered to the inner surface and proliferated to confluency 7 days after seeding, as demonstrated by SEM. Histochemical and immunocytochemical analyses performed using H & E, periodic acid Schiff, alkaline phosphatase, antiosteopontin and anti-CD-31 identified stromal, endothelial, and tubular epithelial cell phenotypes within the matrix. Renal tubular and glomeruli-like structures were observed 8 weeks after implantation. MTT proliferation and radioactive thymidine incorporation assays performed 6 weeks after cell seeding demonstrated a cell population increase of 116% and 92%, respectively, as compared to the 2-week time points. This study demonstrates that renal cells are able to adhere and proliferate on the collagen-based kidney matrices. The renal cells reconstitute renal tubular and glomeruli-like structures. The collagen-based kidney matrix system seeded with renal cells may be useful in the future for augmenting renal function. Our prior studies demonstrated that culture expanded primary renal cells seeded on artificial renal devices with a collecting system are able to form functional renal structures with urine-like fluid excretion. However, creation of renal structures without the use of an artificial device system would be preferable. In addition, implantation procedures are invasive and may result in unnecessary complications. In a subsequent study we investigated the feasibility of creating three-dimensional renal structures for in situ implantation within the native kidney tissue. Primary renal cells from 4-week-old mice were grown and expanded in

Intracorporeal Kidney Support

(a)

(b)

Figure 64.3 Reconstitution of kidney structures. The implanted renal cells self-assembled into (a) glomerular and (b) tubular structures within the kidney tissue.

culture. Culture expanded renal cells were labeled with fluorescent markers and injected into mouse kidneys in a collagen gel for in vivo formation of renal tissues. Collagen injection without cells and sham operated animals served as controls. In vitro reconstituted renal structures and in vivo implanted cells were retrieved and analyzed. The implanted renal cells formed tubular and glomerular structures within the kidney tissue, as confirmed by the fluorescent markers. There was no evidence of renal tissue formation in the control and the sham operated groups. These results demonstrate that single renal cells are able to reconstitute into organized kidney structures when placed in a collagen-based scaffolding system. The implanted renal cells are able to self-assemble into tubular and glomerular structures within the kidney tissue (Figure 64.3). These findings suggest that this system may be the preferred approach to engineer functional kidney tissues for the treatment of end stage renal disease.

SUMMARY Renal transplantation remains as the gold standard of treatment for end stage renal failure. The increasing demand and shortage of donor organs have ignited tremendous efforts to seek for alternative treatment modalities. Cell-based approaches have been proposed recently as an alternative method for kidney tissue regeneration. In this chapter, various tissue engineering and regenerative medicine approaches were presented in an effort to achieve functional intracorporeal kidney support. Research progress in the regeneration of kidney tissues has been somewhat successful toward augmenting tissue function. However, clinical application of this technology is still distant. Although it has been demonstrated that renal cells are able to reconstitute into functional kidney tissues in vivo, numerous challenges need to be worked out for clinical translation. Some of these include the generation of a large tissue mass that would augment systemic renal function, integration of engineered renal tissue into the host with adequate vascularization and excretory systems, and development of a reliable renal failure model system for testing cell-based technologies. Current work in our institute is aimed at addressing these challenges.

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Atala, A. (1997). Tissue Engineering in the Genitourinary System, Chap. 8. Boston: Birkhauser Press. Atala, A. (1999). Future perspectives in reconstructive surgery using tissue engineering. Urol. Clin. North Am. 26: 157–165, ix–x. Atala, A., Schlussel, R.N. and Retik, A.B. (1995). Renal cell growth in vivo after attachment to biodegradable polymer scaffolds. J. Urol. 153: 4. Atala, A., Bauer, S.B., Soker, S., Yoo, J.J. and Retik, A.B. (2006). Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet 367: 1241–1246. Carley, W.W., Milici, A.J. and Madri, J.A. (1988). Extracellular matrix specificity for the differentiation of capillary endothelial cells. Exp. Cell Res. 178: 426–434. Chazan, J.A., Libbey, N.P., London, M.R., Pono, L. and Abuelo, J.G. (1991). The clinical spectrum of renal osteodystrophy in 57 chronic hemodialysis patients: a correlation between biochemical parameters and bone pathology findings. Clin. Nephrol. 35: 78–85. El-Kassaby, A.W., Retik, A.B., Yoo, J.J. and Atala, A. (2003). Urethral stricture repair with an off-the-shelf collagen matrix. J. Urol. 169: 170–173; discussion 173. Feldman, H.I., Kobrin, S. and Wasserstein, A. (1996). Hemodialysis vascular access morbidity. J. Am. Soc. Nephrol. 7: 523–535. Folkman, J. and Hochberg, M. (1973). Self-regulation of growth in three dimensions. J. Exp. Med. 138: 745–753. Frank, J., Engler-Blum, G., Rodemann, H.P. and Muller, G.A. (1993). Human renal tubular cells as a cytokine source: PDGF-B, GM-CSF and IL-6 mRNA expression in vitro. Exp. Nephrol. 1: 26–35. Freed, L.E., Vunjak-Novakovic, G., Biron, R.J., Eagles, D.B., Lesnoy, D.C., Barlow, S.K. and Langer, R. (1994). Biodegradable polymer scaffolds for tissue engineering. Biotechnology (NY) 12: 689–693. Fung, L.C.T., Elenius, K., Freeman, M., Donovan, M.J. and Atala, A. (1996). Reconstitution of poor EGFr-poor renal epithelial cells into tubular structures on biodegradable polymer scaffold. Pediatrics 98(Suppl): S631. Gupta, S., Verfaillie, C., Chmielewski, D., Kim, Y. and Rosenberg, M.E. (2002). A role for extrarenal cells in the regeneration following acute renal failure. Kidney Int. 62: 1285–1290. Horikoshi, S., Koide, H. and Shirai, T. (1988). Monoclonal antibodies against laminin A chain and B chain in the human and mouse kidneys. Lab. Invest. 58: 532–538. Hubbell, J.A., Massia, S.P., Desai, N.P. and Drumheller, P.D. (1991). Endothelial cell-selective materials for tissue engineering in the vascular graft via a new receptor. Biotechnology (NY) 9: 568–572. Humes, H.D. and Cieslinski, D.A. (1992). Interaction between growth factors and retinoic acid in the induction of kidney tubulogenesis in tissue culture. Exp. Cell Res. 201: 8–15. Humes, H.D., Buffington, D.A., MacKay, S.M., Funke, A.J. and Weitzel, W.F. (1999). Replacement of renal function in uremic animals with a tissue-engineered kidney. Nat. Biotechnol. 17: 451–455. Ikarashi, K., Li, B., Suwa, M., Kawamura, K., Morioka, T., Yao, J., Khan, F., Uchiyama, M. and Oite, T. (2005). Bone marrow cells contribute to regeneration of damaged glomerular endothelial cells. Kidney Int. 67: 1925–1933. Ito, T., Suzuki, A., Imai, E., Okabe, M. and Hori, M. (2001). Bone marrow is a reservoir of repopulating mesangial cells during glomerular remodeling. J. Am. Soc. Nephrol. 12: 2625–2635. Iwano, M., Plieth, D., Danoff, T.M., Xue, C., Okada, H. and Neilson, E.G. (2002). Evidence that fibroblasts derive from epithelium during tissue fibrosis. J. Clin. Invest. 110: 341–350. Kale, S., Karihaloo, A., Clark, P.R., Kashgarian, M., Krause, D.S. and Cantley, L.G. (2003). Bone marrow stem cells contribute to repair of the ischemically injured renal tubule. J. Clin. Invest. 112: 42–49. Lanza, R.P., Chung, H.Y., Yoo, J.J., Wettstein, P.J., Blackwell, C., Borson, N., Hofmeister, E., Schuch, G., Soker, S., Moraes, C.T., et al. (2002). Generation of histocompatible tissues using nuclear transplantation. Nat. Biotechnol. 20: 689–696. Lin, F., Cordes, K., Li, L., Hood, L., Couser, W.G., Shankland, S.J. and Igarashi, P. (2003). Hematopoietic stem cells contribute to the regeneration of renal tubules after renal ischemia-reperfusion injury in mice. J. Am. Soc. Nephrol. 14: 1188–1199. Lin, F., Moran, A. and Igarashi, P. (2005). Intrarenal cells, not bone marrow-derived cells, are the major source for regeneration in postischemic kidney. J. Clin. Invest. 115: 1756–1764. Milici, A.J., Furie, M.B. and Carley, W.W. (1985). The formation of fenestrations and channels by capillary endothelium in vitro. Proc. Natl. Acad. Sci. USA 82: 6181–6185.

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Mooney, D.J., Mazzoni, C.L., Breuer, C., McNamara, K., Hern, D., Vacanti, J.P. and Langer, R. (1996). Stabilized polyglycolic acid fibre-based tubes for tissue engineering. Biomaterials 17: 115–124. Oberpenning, F., Meng, J., Yoo, J.J. and Atala, A. (1999). De novo reconstitution of a functional mammalian urinary bladder by tissue engineering. Nat. Biotechnol. 17: 149–155. Poulsom, R., Forbes, S.J., Hodivala-Dilke, K., Ryan, E., Wyles, S., Navaratnarasah, S., Jeffery, R., Hunt, T., Alison, M., Cook, T., et al. (2001). Bone marrow contributes to renal parenchymal turnover and regeneration. J. Pathol. 195: 229–235. Prockop, D.J. (1997). Marrow stromal cells as stem cells for nonhematopoietic tissues. Science 276: 71–74. Rookmaaker, M.B., Smits, A.M., Tolboom, H., Van’t Wout, K., Martens, A.C., Goldschmeding, R., Joles, J.A., Van Zonneveld, A.J., Grone, H.J., Rabelink, T.J. and Verhaar, M.C. (2003). Bone-marrow-derived cells contribute to glomerular endothelial repair in experimental glomerulonephritis. Am. J. Pathol. 163: 553–562. Schena, F.P. (1998). Role of growth factors in acute renal failure. Kidney Int. Suppl. 66: S11–S15. Stadnyk, A.W. (1994). Cytokine production by epithelial cells. Faseb J. 8: 1041–1047. Tachibana, M., Nagamatsu, G.R. and Addonizio, J.C. (1985). Ureteral replacement using collagen sponge tube grafts. J. Urol. 133: 866–869. Wald, H.L., Sarakinos, G., Lyman, M.D., Mikos, A.G., Vacanti, J.P. and Langer, R. (1993). Cell seeding in porous transplantation devices. Biomaterials 14: 270–278. Yokoo, T., Ohashi, T., Shen, J.S., Sakurai, K., Miyazaki, Y., Utsunomiya, Y., Takahashi, M., Terada, Y., Eto, Y., Kawamura, T., et al. (2005). Human mesenchymal stem cells in rodent whole-embryo culture are reprogrammed to contribute to kidney tissues. Proc. Natl. Acad. Sci. USA 102: 3296–3300. Yoo, J.J., Ashkar, S. and Atala, A. (1996). Creation of functional kidney structures with excretion of kidney-like fluid in vivo. Pediatrics 98S: 605.

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65 The Kidney William H. Fissell and H. David Humes

INTRODUCTION The kidney is unique in that it is the first organ for which long-term ex vivo substitutive therapy has been available and lifesaving. Renal failure prior to the era of hemodialysis and transplantation resulted in certain death, and this outcome of renal failure is still common outside the industrialized world. In the United States, 424,802 patients were listed as having end-stage renal disease (ESRD) by the 2004 US Renal Data System (USRDS) database, of whom 308,910 were receiving maintenance dialysis (US Renal Data System, 2002). The prevalence of ESRD in the United States is rising at approximately 8% per year (US Renal Data System, 2004; Neilson et al., 1997). The financial cost of dialysis is immense, estimated at $54,900 per hemodialysis patient per year and $46,121 per peritoneal dialysis patient per year. In contrast, transplant patients cost an average of $17,227 per patient per year (US Renal Data System, 2004). The higher cost of maintenance dialysis when compared with transplantation does not translate into better results; annual mortality for patients listed for transplant and awaiting a kidney is 6.3%, compared with only 3.8% for patients listed for transplant who did receive a kidney. These statistics compare favorably to the 16.7% annual mortality for ESRD patients not listed for transplant (Wolfe et al., 1999). Hemodialysis provides clearance of small molecules in blood by diffusive flow across a semipermeable membrane, and control of volume status by bulk flow of water and solutes through that membrane. These short-term effects are sufficient to abrogate the lethal acidosis, volume overload, and uremic syndromes which accompany renal failure but do not protect the patient from the increased mortality associated with dialysis-treated renal failure in either the acute or chronic form. Thus the metabolic, endocrine, and immune roles of the functioning kidney are candidate mechanisms for the difference in survival noted above. The dialytic clearance of glutathione, a key tripeptide in free radical scavenging and protection against oxidant stress, the negative nitrogen balance and energy loss in the clearance of peptides and amino acids in dialysate, loss of oxidative deamination and gluconeogenesis in the tubule cell, and loss of cytokine and hormone metabolic activity by the kidney each impose substantial stress upon the dialyzed patient and as such are appropriate targets for improved renal replacement therapy to the host that intermittent hemodialysis does not address. The kidney’s functional unit, the nephron, provides for the elimination of wastes and toxins without the need for specific enzymes and transporters for each toxin. All but the large proteins and cellular elements in the blood are filtered, and a system of cells reclaims filtered substances needed by the body, and allows all others to pass as urine. Teleologically, this allows each organism to cope with novel insults its genetic forebears may not have encountered. Filtration is accomplished by the glomerulus, a tuft of capillaries supported by a basement membrane and specialized epithelial cells called podocytes. The renal proximal tubule, a hollow tube of cells surrounded by capillaries, receives the filtrate from the glomerulus and accomplishes the bulk of reclamation of salt, water, glucose, small proteins, amino acids, glutathione, and other substances. The tubule also accomplishes metabolic functions, including excretion of acid as ammonia, hydroxylation of 25-hydroxy-vitamin-D3, and others. 1114

Intermittent hemodialysis is thought to replace the filtration function of the glomerulus. Our attention is drawn to duplicating the function of the proximal tubule. The transport of solutes and water is accomplished by ATP-driven electrolyte transporters in the luminal cell membrane. Reabsorption of small proteins and peptides in the filtrate stream is accomplished by membrane-bound proteases and specific amino acid transport proteins within the luminal membrane of the tubule cell. These amino acids are either used for protein and peptide synthesis in the tubule cell or transported into the capillaries for transport to and use by the body. The diversity and specificity of the functions of the proximal tubule cell argue against the development of an electromechanical or polymeric substitute, and so a number of years ago, our research group turned its attention on the isolation and culture of renal proximal tubule cells, which research has culminated in the hollow-fiber bioreactor discussed below.

ISOLATION AND CULTURE OF PROXIMAL TUBULE CELLS Our laboratory developed experience in the isolation and culture of porcine proximal tubule cells until we could reproducibly harvest cells and maintain them stably in culture (Humes and Cieslinski, 1992; Humes et al., 1996). In brief, Yorkshire breed pigs were sacrificed at 4–6 weeks of age and their kidneys harvested. The renal cortices were dissected, minced, digested with collagenase, and the resulting mixture was separated on a Percoll density gradient. Renal proximal tubule fragments were isolated and grown in a serum-free hormonally derived medium. After reaching confluence at third passage on 100-mm culture dishes, cells were mobilized with trypsin into a suspension and seeded into polysulfone single hollow-fiber bioreactors for in vitro assessment of cell viability and metabolic activity. Cellular attachment, stability, and confluence on the interior lumen of the bioreactor is of paramount importance. To promote attachment of the cells, the luminal surface of the polysulfone membrane was coated with ProNectin-L, a synthetic protein bearing the integrinbinding domains of laminin. Laminin and collagen type IV, key components of the tubular basement membrane, also provide an effective biomatrix for cell attachment and growth. After seeding of the hollow fiber with tubule cells, the hollow fibers were perfused with culture media. As newly seeded cells need time to attach, perfusion was initially performed via diffusion from the exterior through the polysulfone membrane, and after time for attachment, convective flow through the interior of the fiber was initiated. A graduated increase in flow (and thus shear forces) was used to condition the cells and minimize cellular detachment. Studies demonstrated confluence was reached in 7–10 days. After 14 days in culture the hollow-fiber bioreactors were assessed for cellular confluency and viability. Light microscopy of fixed sections showed evidence of a confluent monolayer formed on the inside of the hollow fibers, and intercellular tight junction formation was verified by measuring low inulin leak rates across the monolayer (Humes et al., 1999a).

TRANSPORT AND METABOLIC CHARACTERISTICS OF HOLLOW-FIBER BIOREACTORS As initial experiments using the single hollow-fiber model were promising, the design was scaled up to use commercially available polysulfone hollow-fiber dialysis cartridges from the manufacturers of the single hollow fibers. Single hollow-fiber measurements of transport and metabolic activity were repeated with 97 cm2 and 0.4 m2 surface area cartridges. We further explored the metabolic characteristics of the cultured proximal tubule cells. We examined the transport of glucose, bicarbonate, and glutathione and expressed the data in terms of fractional reabsorption accomplished by the bioreactor. For each of the molecules listed, fractional excretion was measured in the absence and presence of a known inhibitor of an enzyme essential for the reabsorption. In each case, there was evidence of active transport and specific inhibition (Humes et al., 1999a).

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The synthesis and secretion of ammonia into the tubule is essential for renal excretion of an acid load, as it buffers secreted protons. Proximal tubule cells are able to upregulate their ammoniagenesis in response to a decline in pH, and the proximal tubule cells in the bioreactor demonstrated a stepwise increase in ammonia production with changes in pH (Humes et al., 1999a). The experiments detailed above were performed with porcine tubule cells, and our laboratory has demonstrated similar results in culture, attachment, and activity with human proximal tubule cells from cadaveric organs. The final selection of cell type for use in a renal tubule device rests not only on supply and safety of cells, but also depends on the ability of xenotransplanted cells to participate in the homeostasis of the host. The above data suggest that our laboratory has successfully isolated and cultured renal proximal tubule cells, established stable confluent monolayers within hollow-fiber bioreactors, and scaled the initial construct to a level approximating the number of proximal tubule cells in a single kidney.

PRECLINICAL CHARACTERIZATION OF THE RENAL TUBULE ASSIST DEVICE (RAD) In keeping with its role as a metabolically active replacement for the renal proximal tubule, an extracorporeal circuit was devised that recapitulated nephron anatomy. A conventional hollow-fiber dialyser and a hollowfiber bioreactor were connected in series, so that a portion of the ultrafiltrate exiting the hemofilter was directed into the luminal spaces of the bioreactor, and thus presented to the apical aspect of the cultured proximal tubule cells. Concentrated blood exiting the hemofilter was directed to the extraluminal space (the dialysate compartment) of the bioreactor, just as post glomerular blood surrounds the renal tubules via the peritubular capillaries. To allow independent control of the subject’s volume status and clearance parameters during experiments, the balance of the ultrafiltrate was discarded and the subject infused with a balanced electrolyte solution, as in a conventional hemofiltration circuit. The bioartificial kidney setup consists of a filtration device (a conventional hemofilter) followed in series by the RAD unit. Specifically, blood is pumped out of a large animal using a peristaltic pump. The blood then enters the fibers of a hemofilter, where ultrafiltrate is formed and delivered into the fibers of the tubule lumens within the RAD downstream to the hemofilter. Processed ultrafiltrate exiting the RAD is collected and discarded as “urine.” The filtered blood exiting the hemofilter enters the RAD through the extracapillary space port and disperses among the fibers of the device. Upon exiting the RAD, the processed blood travels through a third pump and is delivered back to the animal; this additional pump is required to maintain appropriate hydraulic pressures within the RAD. Heparin is delivered continuously into the blood before entering the RAD to diminish clotting within the device. The RAD is oriented horizontally and placed into a temperature-controlled environment. The temperature of the cell compartment of the RAD must be maintained at 37°C throughout its operation to ensure optimal functionality of the cells. Maintenance of a physiological temperature is a critical factor in the functionality of the RAD. The tubule unit is able to maintain viability because metabolic substrates and lowmolecular weight growth factors are delivered to the tubule cells from the ultrafiltration unit and the blood in the extracapillary space. Furthermore, immunoprotection of the cells grown within the hollow fiber is achieved because of the impenetrability of immunoglobulins and immunologically competent cells through the hollow fibers. Rejection of the cells, therefore, does not occur. This arrangement thereby allows the filtrate to enter the internal compartments of the hollow-fiber network lined with confluent monolayers of renal tubule cells for regulated transport and metabolic function. Large animal studies have been completed with the use of this extracorporeal circuit (Humes et al., 1999b). Dogs were made uremic by performing bilateral nephrectomies. A double lumen catheter was placed into the internal jugular vein, extending into the heart. After 24 h of postoperative recovery, the dogs were

The Kidney

treated either with hemofiltration and the RAD or with hemofiltration and a sham control cartridge containing no cells. The blood flow rate to the hemofiltrator was maintained at 80 ml/min, with a controlled ultrafiltration rate of 5–7 ml/min. Dogs were treated daily for either 7 or 9 h or for 24 h continuously. The dogs in these experiments developed acute renal failure (ARF) with average blood urea nitrogen (BUN) and plasma creatinine levels of 68 and 6.6 mg/dl, respectively. The RADs maintained viability and functionality when connected in series to a hemofiltration cartridge within an extracorporeal perfusion circuit in an acutely uremic animal. During a 24-h perfusion period, fewer than 105 cells were lost from the RAD, which contained more than 109 cells. Treatment with the RAD and hemofiltration maintained BUN and plasma  creatinine levels similar to those of sham controls. In addition, plasma HCO 3 , Pi, and K levels were more readily maintained near normal values in RAD treatment than in sham treatment. The RADs were able to reabsorb 40–50% of ultrafiltrate volume presented to the devices. Furthermore, active transport of K, HCO 3 , and glucose was accomplished by the RAD in this ex vivo situation. Metabolic activity of the RAD was also shown in these experiments. Virtually no ammonia excretion occurred in the processed ultrafiltrate of the sham control group, in contrast with an ammonia excretion level as high as 100 μmol/h in the RAD-treated group. Glutathione processing by the RAD was also shown, with greater than 50% glutathione removal from the ultrafiltrate presented to the RAD. Finally, uremic animals treated with the RAD attained 1,25-(OH)2-D3 levels of 19.5 0.5 pg/ml, a value no different from the normal levels of the prenephrectomy condition. In contrast, sham treatment resulted in a further fall of 4.0 2.4 pg/ml from the already low plasma levels of 1,25-(OH)2-D3 in the acutely uremic animals. Thus, these experiments clearly showed that the combination of a synthetic hemofiltration cartridge and a RAD in an extracorporeal circuit successfully replaced filtration, transport, and metabolic and endocrinological functions of the kidney in acutely uremic dogs.

CELL THERAPY OF ACUTE RENAL FAILURE DUE TO SEPSIS After a series of experiments demonstrating bioactivity, longevity, and systemic activity of the proximal tubule cells in a large animal model, a series of experiments was designed to examine the impact of cell therapy on the course of sepsis complicated by renal failure (Fissell et al., 2002a, 2003). Septic shock with ARF was chosen as an experimental model for several reasons. First, there is no established animal model of chronic renal failure, largely due to the cost involved in maintaining a herd of animals on dialysis. Second, the time course of welldialyzed chronic renal failure is months to years of subject survival, which would be prohibitively expensive. Lastly, animal models of septic shock have been well established and can serve as a starting point for understanding the disease physiology. After two initial studies supported a systemic effect and hemodynamic benefit from cell therapy in large animal models of sepsis, we pursued further evidence that cell therapy with renal proximal tubule cells alters the physiological response to sepsis (Humes et al., 2003). A porcine model of septic shock was developed from the previous work (Natanson et al., 1989; Natanson, 1990; Dinarello, 1991). It had been noted in non-nephrectomized animals that in response to bacteria, the animals rapidly became oligoanuric. We hypothesized that septic shock rapidly induced tubule cell injury, and that replacement of the function of injured tubule cells would confer benefit upon the animals. Purpose-bred pigs were anesthetized and administered an intraperitoneal dose of bacteria, causing shock and renal failure. An hour later continuous venovenous hemofiltration (CVVH) was initiated with either cell or sham RAD. Urine output and mean arterial pressure declined within the first few hours after insult. Cell-treated animals survived 9.0 0.83 h versus 5.1 0.4 h (P  0.005) for sham-treated animals. Serum cytokines were similar between the two groups, with the striking exception of interleukin (IL)-6 and interferon (IFN)-γ. Treatment with the cell RAD resulted in

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significantly lower plasma levels of both IL-6 (P  0.04) and INF-γ (P  0.02) throughout the experimental time course compared to sham RAD exposure. This controlled trial of cell therapy of renal failure in a realistic animal model of sepsis has several findings not immediately expected from a priori assumptions regarding renal function. Heretofore, although renal failure has been strongly associated with poor outcome in hospitalized patients, and chronic renal failure is associated with specific defects in humoral and cellular immunity, a direct immunomodulatory effect of the kidney had not been accepted. In this trial clear differences in survival and clear differences in a serum cytokine associated with mortality in sepsis were found between animals treated with cells and with sham cartridges. This hearkens back to statements made earlier in the chapter: The increased mortality in renal failure has not been conclusively attributed to inadequate clearance but may arise from other bioactivity of the kidney.

CLINICAL EXPERIENCE WITH A HUMAN RENAL TUBULE ASSIST DEVICE With a series of preclinical experiments demonstrating the extracorporeal circuit containing the bioreactor, the US Food and Drug Administration (FDA) granted Investigational New Drug approval to the bioreactor, and a multicenter phase I/II clinical trial was begun to study the RAD containing human cells in patients with acute tubular necrosis (ATN) and multisystem organ failure who were receiving continuous renal replacement therapy. Human kidney cells were isolated from kidneys donated for cadaveric transplantation but found unsuitable for such purpose because of anatomic or fibrotic defects. The initial results in the first 10 treated patients demonstrated that this experimental treatment can be delivered safely under study protocol guidelines in this critically ill patient population for up to 24 h when used in conjunction with CVVH (Humes et al., 2004). These data also indicated that the RAD maintains and exhibits viability, durability, and functionality in this ex vivo clinical setting. Cardiovascular stability of the patients was maintained, and increased native renal function, as determined by elevated urine outputs, temporally correlated with RAD treatment, a finding that requires additional study. The isolated and expanded human cells also demonstrated differentiated metabolical and endocrinological activity in this ex vivo treatment. Glutathione degradation and endocrinological conversion of 25-(OH)-Vit D3 to 1,25-(OH)2-Vit D3 by the RAD tubule cells were demonstrated. All 10 patients were critically ill with ARF and multiple organ dysfunction syndrome, with predicted hospital mortality rates between 80% and 95%. One patient expired within 12 h after RAD treatment due to his family’s request to withdraw ventilatory life support. Another patient expired after a surgical catastrophe (toxic megacolon) required discontinuation of RAD treatment after only 12 h. Of the remaining eight patients, six survived past 28 days with renal function recovery. The other two patients died from non-recoverable complications unrelated to RAD therapy and ARF, including fungal pericarditis and vancomycin-resistant enterococcus septicemia in one patient and ischemic colitis with bowel perforations in the other patient. Plasma cytokine levels suggest that RAD therapy produces dynamic and individualized responses in patients depending on their unique pathophysiological conditions. For the subset of patients who had excessive proinflammatory levels, RAD treatment resulted in significant declines in granulocyte-colony stimulating factor (G-CSF), IL-6 (a pro-inflammatory cytokine), IL-10 (an anti-inflammatory cytokine), and especially IL-6/ IL-10 ratio, suggesting a greater decline in IL-6 relative to IL-10 level and the less pro-inflammatory state. The phase I/II results were encouraging and led to an FDA-approved, randomized, controlled, open-label phase II investigation at ten clinical sites to assess the safety and early efficacy of this cell therapy approach. This study involved 58 intensive care unit patients with dialysis-dependent ARF. Forty patients were randomized to receive both CVVH and RAD therapy, and 18 received only CVVH. Interim analysis found that the 28-day allcause mortality in patients receiving RAD therapy was 33%, significantly lower than the 61% rate in CVVH-only

The Kidney

patients, suggesting a treatment advantage with RAD therapy (Tumlin et al., 2005). A large, multicenter phase III clinical trial is being planned to confirm the efficacy of this therapy.

BIOARTIFICIAL KIDNEY IN ESRD A bioartificial kidney for long-term use in ESRD, similar to short-term use in ARF, would integrate tubular cell therapy and the filtration function of a hemofilter. As noted above, ESRD patients on conventional renal replacement therapy are at high risk for cardiovascular and infectious diseases. A recent clinical trial failed to show survival benefit from increased doses of hemodialysis above what is now standard care (Eknoyan et al., 2002), suggesting that there are important metabolic derangements not adequately treated with conventional dialytic treatment. Data from the survival of renal transplant recipients, which far exceed those from the survival of age-, sex-, and risk-matched controls awaiting transplant, also suggest that there is some metabolic function provided by the kidney that transcends this organ’s filtration function. Patients with ESRD display elevated levels of C-reactive protein (CRP), an emerging clinical marker, and pro-inflammatory cytokines, including IL-1, IL-6, and tumor necrosis factor (TNF) (Bologa et al., 1998; Kimmel et al., 1998; Zimmermann et al., 1999). All these parameters are associated with enhanced mortality in ESRD patients. Specifically, IL-6 has been identified as a single predictive factor closely correlated with mortality in hemodialysis patients (Bologa et al., 1998). Although all ESRD patients could conceivably benefit from a bioartificial kidney, patients in the inflammatory stage who display elevated levels of certain markers of chronic inflammation (most notably IL-6 and CRP) would likely benefit most and will be the target population for clinical study in the near future. For the ESRD patient population, however, there are obvious limitations in using an extracorporeal RAD connected to a hemofiltration circuit. Ideally, a bioartificial kidney suitable for long-term use in ESRD patients would be capable of performing continuously, like the native kidney, to reduce risks from fluctuations in volume status, electrolytes, and solute concentrations and to maintain acid–base and uremic toxin regulation. As in ARF, cell therapy currently can be administered only intermittently. These additional functions require the eventual design and manufacture of a compact implantable or wearable bioartificial kidney. The advent of silicon bulk and surface micromachining offers hope for the development of such a device (Fissell et al., 2002b). These technologies allow the fabrication of pores, beams, gears, and pressure sensors and the patterned deposition of cells in engineered micro- and nanoenvironments. Surface micromachining has been used to provide capillary-like conduits for blood flow in hepatocyte tissue cultures, and micropatterning has been used to control cell growth and differentiation (Chen et al., 1998; Kaihara et al., 2000). We look forward to the possible use of these emerging technologies in the manufacture of a bioartificial kidney for universal, optimal, renal replacement therapy for ESRD patients.

CELL SOURCING The application of a specific cell-based therapy requires several important methodological choices and the solution a number of technological hurdles. The most critical initial decision is cell sourcing. For cell-based therapies, cells need to be expanded in large quantities while maintaining uniformity in activity and being pathogen free. Current approaches to ensure robust cell expansion and uniformity requirements are dependent on either stem/progenitor cells or transformed cells. The use of human embryonic stem (ES) cells versus adult stem cells is under rigorous societal debate, with the current political environment strongly favoring adult stem cell processes (Brower, 1999; Wertz, 2002). The plasticity of adult stem cells to transdifferentiate from one lineage pathway to another is also under careful scientific scrutiny. The early support for stem cell

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plasticity appears to be questioned by recent reports demonstrating stem cell fusion with tissue-specific differentiated cells resulting in polyploidy rather than true stem cell transdifferentiation with normal diploid chromosomal numbers (Terada et al., 2002; Wang et al., 2003; Vassilopoulos et al., 2003). The ability of bone marrow stem cells to differentiate into a variety of cell types within the kidney, including glomerular, vascular, and tubular elements, has been demonstrated (Brodie and Humes, 2005). These reports, however, demonstrate highly variable engraftment rates and inconsistent phenotypic differentiation. The issue of cell fusion in these experiments has not been addressed. Current cell-based approaches, therefore, are directed toward utilizing adult tissue-specific stem cell expansion, but the potential use of ES cells is being aggressively pursued. The utilization of transformed cells, including applications to deliver a gene product with gene therapy, has recently come under intense scrutiny due to safety concerns. The autologous transplantation of genetically modified hematopoietic stem cells in children with adenosine deaminase deficiency, which leads to severe immunodeficiency, resulted in the development of acute leukemia in some of the patients due to genetic integration of the vector in the hematopoietic stem cells (Hacein-Bey-Abina et al., 2003). The need to retrieve or deactivate these transformed cells following cell implantation is required to mitigate this high risk. Even the use of non-transformed cells may have safety concerns. Implantation of nerve cells in patients with Parkinson’s disease leads to a high rate of severe and uncontrollable dyskinetic activity (Dunnett et al., 2001); implantation of myoblasts into the heart has resulted in high rates of cardiac arrhythmias (Menasché et al., 2003). A choice between autologous or non-autologous human cells is also critical in the formulation of a cellbased application. Non-autologous cells must overcome natural host immunological rejection processes. Since most indications preclude the use of immunosuppressant drugs to accommodate the discordant cell implant, immunoprotection of non-autologous cells has been approached with microencapsulation techniques using ultrathin synthetic membranes to prevent entry of antibodies and immunocompetent cells of the host. Implantation of cellular microcapsules has had limited success because of poor long-term functional performance secondary to progressive loss of cell viability (Orive et al., 2003). Success with short-term cell therapy utilizing hollow-fiber bioreactors in an extracorporeal blood perfusion circuit for organ replacement therapy in acute disorders, including ATN, has been more forthcoming (Humes et al., 1999b; Fissell et al., 2003). The use of autologous cells, although overcoming the immunological barrier, has its own set of problems. Autologous approaches require obtaining the patient’s own cells, expanding them in vitro in large quantities over several weeks, then reintroducing the cells in a site-specific manner. Thus, each treatment is an individualized and non-scalable process with substantial logistical and regulatory hurdles, including maintenance of the uniform quality of cells, avoidance of introduced pathogens during cell processing, and potential retrievability after implantation. A final technological hurdle for cell-based therapies is the maintenance of cell viability during long-term implantation. Maintenance of cell function is dependent upon adequate nutrient and oxygen delivery to the cellular implant (Avgoustiniatos and Colton, 1997). Creative approaches to induce and maintain formation of a neovascular capillary bed in and around the cell implant with various drug-delivery and cell scaffold formulations have been demonstrated experimentally but have not yet been successfully translated to the clinic (Richardson et al., 2001). An intravascular cell encapsulation implant approach looks promising but requires further experimentation to determine its successful implementation (Humes, 1998). Although cell-based therapy has substantial technological, regulatory, and ethical barriers, the potential to develop innovative treatments for a large number of clinical disorders, including acute and chronic renal diseases, is expanding rapidly. Progress in this field is highly dependent upon an interdisciplinary approach at the

The Kidney

interface of a number of scientific disciplines. This approach is at times empiric rather than reductive in nature, without full understanding of the manner in which cells may alter the complex pathophysiology of a systemic disorder. This limitation should not preclude continued efforts in this field. In fact, this empiric approach may result in unanticipated insights into basic biology, similar to those obtained in immunology as the field of human solid organ transplantation evolved. Cell therapy has the potential to creatively leverage nature’s ability to provide new and much needed treatments to patients with acute and chronic diseases.

ULTRAFILTRATION MEMBRANE DEVELOPMENT In parallel with progress in stem cell work, there has been considerable interest in applying novel technology to membrane engineering. The challenges in membrane development have been discussed at length (for a review, see Shettiggar, 1998). These efforts have included silicon nanopore membranes, directed etching of alumina, blood flow and dialysate flow enhancements, and computer modeling of theoretic artificial ion channels embedded in synthetic membranes. The filtration barrier in the kidney has been extensively studied (for overviews of anatomy and theoretical modeling, see Jennette et al., 1998; Deen et al., 1980). In brief, tufts of capillaries in the kidney called glomeruli are perfused with blood and allow leakage of water, electrolytes, and small molecules, while retaining cellular elements of blood and proteins larger than about 50,000 Daltons (50 kD) in the bloodstream. Depending on the ambient physiology, 10–20% by volume of the feed solution of blood is filtered by the glomerulus into the urinary space, where it is carried into the renal tubule and processed into urine. The filtration barrier in the kidney is widely considered to be a trilaminate structure, with an endothelial cell layer, a specialized basement membrane, and an elaborate epithelial layer bearing a specialized cell–cell junction called the glomerular slit diaphragm. The relative contributions of each layer of the filtration structure continue to be debated (Smithies, 2003), but there have long been established histological correlates between disruption of the cell–cell junction between glomerular epithelial cells (“podocytes”) and the appearance of protein in the urine. Recent advances in molecular biology have further characterized the glomerular slit diaphragm, and heritable mutations in the proteins that form the slit diaphragm, and genetic knockout experiments in animal models have demonstrated that the slit diaphragm is necessary for solute retention in the kidney’s filter. Unfortunately, the cell considered responsible for the permselectivity barrier in the kidney, the glomerular podocyte, is a terminally differentiated cell with limited regenerative capacity. It is remarkably neuron-like, with a central body and slender arborizing processes that interdigitate with like processes from surrounding cells. As has been observed clinically with neurons, podocytes that are damaged or shed in the urine are not replaced by expansion of neighboring podocytes. Similarly, primary cultures of podocytes do not assume a differentiated phenotype in the laboratory dish, nor do they easily divide and expand in number. Despite progress with a conditionally transformed cell line derived from mouse podocytes, there seems little immediate prospect of a cellbased bioartificial glomerulus. Existing polymer membranes used in dialysis and ultrafiltration have been extensively studied. The pores in such membranes are formed by extrusion or solvent casting techniques. The geometry and surface chemistry of the pores arise from the chemistry of the polymers and the fluid dynamics of the casting process. In general, the hollow-fiber membranes are fairly thick or employ a multilayer scaffold for mechanical support, and have a distribution of pore sizes rather than a regular array of uniform pores. Pores in conventional polymeric membranes tend to be either roughly cylindrical, have a round orifice terminating a larger channel, or have a structure resembling an open-cell sponge. An extensive description of porous structures

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used in commercial ultrafiltration and microfiltration is available (Dunleavy, 1996; Zeman and Zydney, 1996). It is not clear that any of these structures provide optimal geometries for membrane filtration, for two reasons. First, a wide dispersion in pore sizes within a membrane leads to imperfect retention of molecules larger than the mean pore size of the membrane. This is remedied in practice by engineering the mean pore size of the membrane to be sufficiently small that negligibly few pores are so large as to allow passage of a solute above the desired molecular weight cutoff of the membrane. This has the undesired effect of reducing the mean pore size in the membrane and thus reducing the hydraulic permeability of the membrane. Engineering narrower pore size distributions ameliorates this dilemma, allowing sharper transitions from passage to retention and maximizing the mean pore size of the membrane. Second, the round shape of conventional pores dictates a fourth-power dependence of hydraulic permeability on pore radius r: Q πr 4  ΔP 8μ L

(65.1)

where Q denotes volumetric flow, P is hydrostatic pressure, r is the radius of the pore, μ is viscosity, and L is the length of the pore, which may or may not be the same as the thickness of the membrane. A pore that is slit-shaped allows steric hindrance to solute passage dictated by the smallest critical dimension of the pore, while increasing hydraulic permeability by a factor of the long dimension of the pore w: Q wh3  ΔP 12μ L

(65.2)

where w is the long dimension of the slit, h is the thickness of the slit, and L is again the length of the pore. Consequently, it might be predicted that filtration structures with parallel slit-shaped pores might have superior performance when compared to structures with round pores. With that in mind, it is interesting although speculative to note that natural selection has produced filtration structures with elongated, slit-shaped geometries in the kidney, in the beaks of filter-feeding birds such as the flamingo, and in the baleen of filterfeeding whales.

Single pore

Figure 65.1 Scanning electron micrograph of an array of nanofabricated slit pores.

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Validation of the theoretic advantages of slit-shaped porous membranes has begun with testing of a novel silicon membrane. Silicon membranes with 10–100 nm  45 μm slit pores have been designed and manufactured (Fissell et al., 2002b). Silicon chips bearing 1-mm  1-mm arrays of approximately 104 slit pores were fabricated via sacrificial layer techniques (Figure 65.1). The pore structure is defined by deposition and patterning of a polysilicon film on the silicon wafer. The critical submicron pore dimension is defined by the thickness of a sacrificial SiO2 layer, which can be grown with unprecedented control to within 1 nm. The oxide layer is etched away in the final processing step to create the porous polysilicon membrane. The membrane pores are exceptionally smooth walled and monodisperse in pore size. The critical dimension of the pore size varies by 1 nm over the surface of a 100-mm silicon wafer. Hydraulic permeability of these membranes matches first-principles predictions exactly. Although the porosity of the protoype membrane is two to three orders of magnitude less than conventional polymer membranes, liquid hydraulic permeabilities for an 8-nm membrane have been measured and found to be similar to those published for modern commercial hemodialysis membranes. Although the work to date has been directed to treatment of ARF, a portable, continuous, dialysate-free artificial kidney remains the holy grail of renal tissue engineering. The enabling platform technologies discussed in this review advance this goal from a dream to the laboratory bench and the bedside. Future research in renal tissue engineering will need to focus on reproducing mechanisms of whole-body homeostasis. A high priority must be given to sensing and regulating extracellular fluid volume, even if only at the crude level of having the patient weigh him- or herself daily and adjust ultrafiltration and reabsorption by the bioartificial kidney. Chemical-field effect transistors (ChemFETs) offer the possibility of measuring electrolyte levels in a protein-free ultrafiltrate and reading out the potassium level to the patient, who could then alter diet or treat him- or herself with potassium-absorbing resins. The critical building blocks of an autonomous bioartificial kidney are advancing from the laboratory bench to the bedside in revolutionary clinical trials underway at multiple medical centers now. The technology with which to adapt these advances to a more autonomous, dialysate-free system is developing rapidly. The next decade, like the previous, will see quantum advances in renal tissue engineering.

REFERENCES Avgoustiniatos, E.S. and Colton, C.K. (1997). Effect of external oxygen mass transfer resistances on viability of immunoisolated tissue. Ann. NY Acad. Sci. 831: 145–167. Bologa, R.M., Levine, D.M., Parker, T.S., Cheigh, J.S., Serur, D., Stenzel, K.H. and Rubin, A.L. (1998). Interleukin-6 predicts hypoalbuminemia, hypocholesterolemia, and mortality in hemodialysis patients. Am. J. Kidney Dis. 32: 107–114. Brodie, J.C. and Humes, H.D. (2005). Stem cell approaches for the treatment of renal failure. Pharmacol. Rev. 57: 299–313. Brower, V. (1999). Human ES cells: can you build a business around them? Nat. Biotechnol. 17: 139–142. Chen, C.S., Mrksich, M., Huang, S., Whitesides, G.M. and Ingber, D.E. (1998). Micropatterned surfaces for control of cell shape, position, and function. Biotechnol. Prog. 14: 356–363. Deen, W.M., Satvat, B. and Jamieson, J.M. (1980). Theoretical model for glomerular filtration of charged solutes. Am. J. Physiol. 238: F126–F139. Dinarello, C.A. (1991).The proinflammatory cytokines interleukin-1 and tumor necrosis factor and the treatment of the septic shock syndrome. J. Infect. Dis. 163: 1177–1184. Dunleavy, M. (1996). Polymeric membranes: a review of applications. Med. Device Technol. 7: 18–21. Dunnett, S.B., Bjorklund, A., Lindvall, O. (2001). Cell therapy in Parkinson’s disease – Stop or go? Nat. Rev. Neurosci. 2: 365–369. Eknoyan, G., Beck, G.J., Cheung, A.K., Daugirdas, J.T., Greene, T., Kusek, J.W., Allon, M., Bailey, J., Delmez, J.A., Depner, et al. (2002). Effect of dialysis dose and membrane flux in maintenance hemodialysis. N. Engl. J. Med. 347: 2010–2019.

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Fissell, W.H., Dyke, D.B., Buffington, D.A., Weitzel, W.F., Westover, A.J., MacKay, S.M., Gutierrez, J.M. and Humes, H.D. (2002a). Bioartificial kidney alters cytokine response and hemodynamics in endotoxin-challenged uremic animals. Blood Purif. 20: 55–60. Fissell, W.H., Humes, H.D., Roy, S. and Fleischman, A. (2002b). Initial characterization of a nanoengineered ultrafiltration membrane. J. Am. Soc. Nephrol. 13: 602A. Fissell, W.H., Lou, L., Abrishami, S., Buffington, D.A. and Humes, H.D. (2003). Bioartificial kidney ameliorates gram-negative bacteria-induced septic shock in uremic animals. J. Am. Soc. Nephrol. 14: 454–461. Hacein-Bey-Abina, S., von Kalle, C., Schmidt, M., Le Deist, F., Wulffraat, N., McIntyre, E., Radford, I., Villeval, J.L., Fraser, C.C., Cavazzana-Calvo, M. and Fischer, A. (2003). A serious adverse event after successful gene therapy for X-linked severe combined immuno-deficiency. N. Engl. J. Med. 348: 255–256. Humes, H.D. (1998). Implantable device and use thereof. US Patent 5,704,910, January 6. Humes, H.D. and Cieslinski, D.A. (1992). Interaction between growth factors and retinoic acid in the induction of kidney tubulogenesis in tissue culture. Exp. Cell Res. 201: 8–15. Humes, H.D., Krauss, J.C., Cieslinski, D.A. and Funke, A.J. (1996). Tubulogenesis from isolated single cells of adult mammalian kidney: clonal analysis with a recombinant retrovirus. Am. J. Physiol. Renal Physiol. 271(40): F42–F49. Humes, H.D., MacKay, S.M., Funke, A.J. and Buffington, D.A. (1999a). Tissue engineering of a bioartificial renal tubule assist device: in vitro transport and metabolic characteristics. Kidney Int. 55: 2502–2514. Humes, H.D., Buffington, D.A., MacKay, S.M., Funke, A.J. and Weitzel, W.F. (1999b). Replacement of renal function in uremic animals with a tissue-engineered kidney. Nat. Biotechnol. 17: 451–455. Humes, H.D., Buffington, D.A., Lou, L., Abrishami, S., Wang, M., Xia, J. and Fissell, W.H. (2003). Cell therapy with a tissue-engineered reduces the multiple-organ consequences of septic shock. Crit. Care Med. 31: 2421–2428. Humes, H.D., Weitzel, W.F., Bartlett, R.H., Swaniker, F.C., Paganini, E.P., Luderer, J.R. and Sobota, J. (2004). Initial clinical results of the bioartificial kidney containing human cells in ICU patients with acute renal failure. Kidney Int. 66: 1578–1588. Jennette, J.C., Olson, J.L., Schwartz, M.M. and Silva, F.G. (1998). Heptinstall’s Pathology of the Kidney, 5th edn. Philadelphia: Lippincott-Raven, pp. 25–36. Kaihara, S., Borenstein, J., Koka, R., Lalan, S., Ochoa, E.R., Ravens, M., Pien, H., Cunningham, B. and Vacanti, J.P. (2000). Silicon micromachining to tissue engineer branched vascular channels for liver fabrication. Tissue Eng. 6: 105–117. Kimmel, P.L., Phillips, T.M., Simmens, S.J., Peterson, R.A., Weihs, K.L., Alleyne, S., Cruz, I., Yanovski, J.A. and Veis, J.H. (1998). Immunologic function and survival in hemodialysis patients. Kidney Int. 54: 236–244. Menasché, P., Hagège, A.A., Vilquin, J.T., Desnos, M., Abergel, E., Pouzet, B., Bel, A., Sarateanu, S., Scorsin, M., Schwartz, K., et al. (2003). Autologous skeletal myoblast transplantation for severe postinfarction left ventricular dysfunction. J. Am. Coll. Cardiol. 41: 1078–1083. Natanson, C. (1990). Studies using a canine model to investigate the cardiovascular abnormality of and potential therapies for septic shock. Clin. Res. 38: 206–214. Natanson, C., Danner, R.L., Elin, R.J., et al. (1989). Role of endotoxemia in cardiovascular dysfunction and mortality: Escherichia coli and Staphylococcus aureus challenges in a canine model of human septic shock. J. Clin. Invest. 83: 243–251. Neilson, E.G., Hull, A.R., Wish, J.B., Neylan, J.F., Sherman, D. and Suki, W.N. (1997). The Ad Hoc Committee Report on estimating the future workforce and training requirements for nephrology. J. Am. Soc. Nephrol. 8(5 Suppl 9): S1–S4. Orive, G., Hernandez, R.M., Gascon, A.R., Calafiore, R., Chang, T.M., De Vos, P., Hortelano, G., Hunkeler, D., Lacik, I., Shapiro, A.M., et al. (2003). Cell encapsulation: promise and progress. Nat. Med. 9: 104–107. Richardson, T.P., Peters, M.C., Ennett, A.B. and Mooney, D.J. (2001). Polymeric system for dual growth factor delivery. Nat. Biotechnol. 19: 1029–1034. Shettiggar, U.R. (1998). Innovative extracorporeal membrane systems. J. Membrane Sci. 44: 89–114. Smithies, O. (2003). Why the kidney glomerulus does not clog: a gel permeation/diffusion hypothesis of renal function. Proc. Nat. Acad. Sci. 100: 4108–4113. Terada, N., Hamazaki, T., Oka, M., Hoki, M., Mastalerz, D.M., Nakano, Y., Meyer, E.M., Morel, L., Petersen, B.E. and Scott, E.W. (2002). Bone marrow cells adopt the phenotype of other cells by spontaneous cell fusion. Nature 416: 542–545.

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Tumlin, J., Wali, R., Brennan, K. and Humes, H.D. (2005). Effect of the renal assist device (RAD) on mortality of dialysisdependent acute renal failure: a randomized, open-labeled, mulitcenter, Phase II trial. J. Am. Soc. Nephrol. 16(Abstracts): 46A. US Renal Data System (2004). USRDS 2002 Annual Data Report, Atlas of End-Stage Renal Disease in the United States. National Institutes of Health, National Institute of Diabetes and Digestive and Kidney Diseases, Bethesda, MD. Vassilopoulos, G., Wang, P.R. and Russell, D.W. (2003). Transplanted bone marrow regenerates liver by cell fusion. Nature 422: 901–904. Wang, X., Willenbring, H., Akkari, Y., Torimaru, Y., Foster, M., Al-Dhalimy, M., Lagasse, E., Finegold, M., Olson, S. and Grompe, M. (2003). Cell fusion is the principal source of bone-marrow-derived hepatocytes. Nature 422: 897–901. Wertz, D.C. (2002). Embryo and stem cell research in the United States: history and politics. Gene Ther. 9: 674–678. Wolfe, R.A., et al. (1999). Comparison of mortality in all patients on dialysis, patients on dialysis awaiting transplantation, and recipients of a first cadaveric transplant. N. Eng. J. Med. 341: 1725–1730. Zeman, L. and Zydney, A. (1996). Microfiltration and Ultrafiltration. New York: Marcel Dekker. Zimmermann, J., Herrlinger, S., Pruy, A., Metzger, T. and Wanner, C. (1999). Inflammation enhances cardiovascular risk and mortality in hemodialysis patients. Kidney Int. 55: 648–658.

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66 Genitourinary System Anthony Atala

INTRODUCTION From the time the fetus develops, the genitourinary system may be exposed to a variety of possible injuries. Individuals may suffer from infection, congenital disorders, trauma, cancer, inflammation, iatrogenic injuries, or other conditions that may lead to genitourinary organ damage or loss and necessitate eventual reconstruction. Whenever there is a lack of native urologic tissue, reconstruction may be performed with native non-urologic tissues (such as skin, gastrointestinal segments, or mucosa from multiple body sites), homologous tissues (cadaver fascia, cadaver, or donor kidney), heterologous tissues (bovine collagen), or artificial materials (silicone, polyurethane, Teflon). The tissues used for reconstruction may lead to complications because of their inherently different functional parameters. In most cases, the replacement of lost or deficient tissues with functionally equivalent tissues would improve the outcome for these patients. This goal may be attainable with the use of tissue-engineering techniques. RECONSTITUTION STRATEGIES The goal of tissue engineering is to develop biologic substitutes that can restore and maintain normal function. Tissue engineering may involve matrices alone, wherein the body’s natural ability to regenerate is used to orient or direct new tissue growth, or it may use matrices with cells. When cells are used for tissue engineering, donor tissue (heterologous, allogeneic, or autologous) is dissociated into individual cells, which are implanted directly into the host or expanded in culture, attached to a support matrix, and reimplanted after expansion. Ideally, this approach allows lost tissue function to be restored or replaced in toto and with limited complications (Atala et al., 1992, 1993a, b; Atala et al., 1994; Atala, 1996, 1997, 1998, 1999; Cilento et al., 1994; Yoo, 1997; Fauza et al., 1998a, b; Machluf, 1998a; Yoo et al., 1998; Amiel, 1999; Kershen, 1999; Oberpenning et al., 1999; Park et al., 1999). THE ROLE OF BIOMATERIALS Biomaterials in genitourinary tissue-engineering function as an artificial extracellular matrix (ECM) and elicit biologic and mechanical functions of native ECM found in body tissues. Biomaterials facilitate the localization and delivery of cells and/or bioactive factors (such as cell adhesion peptides and growth factors) to desired sites in the body; define a three-dimensional space for the formation of new tissues with appropriate structure; and guide the development of new tissues with appropriate function (Kim, 1998). While direct injection of cell suspensions without biomaterial matrices has been used (Ponder et al., 1991; Brittberg et al., 1994). it is difficult to control the localization of transplanted cells. The ideal biomaterial should be biocompatible, promote cellular interaction and tissue development, and possess proper mechanical and physical properties. Generally, three classes of biomaterials have been used for engineering of genitourinary tissues: naturally derived materials, such as collagen and alginate; acellular

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tissue matrices, such as bladder submucosa and small-intestinal submucosa (SIS) and synthetic polymers, such as polyglycolic acid (PGA), polylactic acid (PLA), and poly(lactic-co-glycolic acid) (PLGA). While naturally derived materials and acellular tissue matrices have the potential advantage of biologic recognition, synthetic polymers can be produced reproducibly on a large scale with controlled properties of strength, degradation rate, and microstructure.

VASCULARIZATION A restriction of tissue engineering is that cells cannot be implanted in volumes exceeding 3 mm3 because of the limitations of nutrition and gas exchange (Folkman, 1973). To achieve the goals of engineering large complex tissues, and possibly internal organs, vascularization of the regenerating cells is essential. Three approaches have been used for vascularization of bioengineered tissue: (1) incorporation of angiogeneic factors in the bioengineered tissue; (2) seeding embryonal carcinoma (EC) with other cell types in the bioengineered tissue; and (3) prevascularization of the matrix prior to cell seeding. Angiogeneic growth factors may be incorporated into the bioengineered tissue prior to implantation, in order to attract host capillaries and to enhance neovascularization of the implanted tissue. Many obstacles must be overcome before large entire tissue-engineered solid organs are produced. Recent developments in angiogenesis research may provide important knowledge and essential materials to accomplish this goal. PROGRESS IN TISSUE ENGINEERING OF UROLOGIC STRUCTURES Tissue-engineering techniques are currently being investigated for the replacement of lost or deficient genitourinary structures, including urethra, bladder, male and female genital tissues, ureter, and renal structures. Urethra Various strategies have been proposed to regenerate urethral tissue. Woven meshes of PGA, without cells, have been used to reconstruct urethras in dogs (Bazeed et al., 1983; Olsen et al., 1992). PGA has been used as a cell transplantation vehicle to engineer tubular urothelium in vivo (Atala et al., 1992). SIS without cells was used as an onlay patch graft for urethroplasty in rabbits (Kropp et al., 1998), and a homologous free graft of acellular urethral matrix was also used in a rabbit model (Sievert et al., 2000). Bladder-derived acellular collagen matrix has proven to be a suitable graft for repairing urethral defects in rabbits. The created neourethras demonstrated a normal urothelial luminal lining and organized muscle bundles (Chen, 1999). Results were confirmed clinically in a series of patients with a history of failed hypospadias reconstruction whose urethral defects were repaired with human bladder acellular collagen matrices (Figure 66.1) (Atala, 1999). An advantage of this material over non-genital tissue grafts for urethroplasty is that it is “off the shelf,” eliminating the need for additional surgical procedures for graft harvesting and decreasing operative time and potential morbidity from the harvest procedure. The above techniques, using non-seeded acellular matrices, were successfully applied experimentally and clinically for onlay urethral repairs. However, when tubularized repairs were attempted experimentally, adequate urethral tissue regeneration was not achieved, and complications, such as graft contracture and stricture formation ensued (le Roux, 2005). Seeded tubularized collagen matrices have performed better than their non-seeded counterparts in animal studies. In a rabbit model, entire urethral segments were resected, and urethroplasties were performed with tubularized collagen matrices, either seeded or non-seeded. The tubularized collagen matrices seeded with autologous cells formed new tissue which was histologically similar to native urethra. Those without cells lead to poor tissue development, fibrosis, and stricture formation. These findings were recently confirmed clinically (DeFilippo et al., 2002).

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Figure 66.1 Tissue-engineered urethra using a collagen matrix. (a) Representative case of a patient with a bulbar stricture. (b) Urethral Repair.Strictured tissue is excised, preserving the urethral plate on the left side, and matrix is anastomosed to the urethral plate in an onlay fashion on the right. (c) Urethrogram 6 months after repair. (d) Cystoscopic view of urethra before surgery on the left side and 4 months after repair on the right side.

Bladder Gastrointestinal segments are commonly used as tissues for bladder replacement or repair. However, these tissues are designed to absorb specific solutes, and when they come in contact with the urinary tract, multiple complications may ensue, including infection, metabolic disturbances, urolithiasis, perforation, increased mucus production, and malignancy (McDougal, 1992; Kaefer et al., 1997, 1998). Because of these problems, investigators have attempted alternative reconstructive procedures for bladder replacement or repair such as the use of tissue expansion, seromuscular grafts, matrices for tissue regeneration, and tissue engineering with cell transplantation. Tissue Expansion A system of progressive dilation for ureters and bladders has been proposed as a method of bladder augmentation but has not yet been attempted clinically (Lalias, 1996; Satar, 1999). Augmentation cystoplasty performed with the dilated ureteral segment in animals has resulted in increased bladder capacity ranging from 190% to 380% (Lalias, 1996). A system to progressively expand native bladder tissue has also been used for augmenting bladder volumes in animals. Within 30 days after progressive dilation, neoreservoir volume was expanded at least 10-fold. Urodynamic studies showed normal compliance in all animals, and microscopic examination of the expanded neoreservoir tissue confirmed a normal histology. A series of immunocytochemical studies demonstrated that the dilated bladder tissue maintained normal phenotypic characteristics (Satar, 1999). Seromuscular Grafts Seromuscular grafts and de-epithelialized bowel segments, either alone or over a native urothelium, have also been attempted (Blandy, 1961, 1964; Harada et al., 1965; Oesch, 1988; Salle et al., 1990; Cheng et al., 1994;

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Dewan, 1998).Keeping the urothelium intact avoids complications associated with the use of bowel in continuity with the urinary tract (Blandy, 1961; Harada, 1965).An example of this strategy is to combine the techniques of autoaugmentation and enterocystoplasty. An autoaugmentation is performed, and the diverticulum is covered with a demucosalized gastric or intestinal segment (Dewan, 1998). Matrices Non-seeded allogeneic acellular bladder matrices have served as scaffolds for the ingrowth of host bladder wall components. The matrices are prepared by mechanically and chemically removing all cellular components from bladder tissue (Sutherland et al., 1996; Probst et al., 1997; Piechota et al., 1998; Yoo et al., 1998). The matrices serve as vehicles for partial bladder regeneration, and relevant antigenicity is not evident. For example, SIS (a biodegradable, acellular, xenogeneic collagen-based tissue-matrix graft) was first used in the early 1980s as an acellular matrix for tissue replacement in the vascular field. It has been shown to promote regeneration of a variety of host tissues, including blood vessels and ligaments (Badylak et al., 1989). Animal studies have shown that the non-seeded SIS matrix used for bladder augmentation can regenerate in vivo (Kropp, 1996; Kropp et al., 2004). In multiple studies using various materials as non-seeded grafts for cystoplasty, the urothelial layer regenerated normally, but the muscle layer, although present, was not fully developed (Sutherland et al., 1996; Kropp et al., 1996; Probst et al., 1997; Yoo et al., 1998). Often the grafts contracted to 60–70% of their original sizes (Portis et al., 2000) with little increase in bladder capacity or compliance (Landman et al., 2004). Studies involving acellular matrices that may provide the necessary environment to promote cell migration, growth, and differentiation are being conducted. With continued research, these matrices may have a clinical role in bladder replacement in the future. Recently, bladder regeneration has been shown to be more reliable using SIS derived from the distal ileum (Kropp et al., 2004). Cell Transplantation Cell-seeded allogeneic acellular bladder matrices have been used for bladder augmentation in dogs. Trigonesparing cystectomy was performed in dogs randomly assigned to one of three groups. One group underwent closure of the trigone without a reconstructive procedure; another underwent reconstruction with a nonseeded bladder-shaped biodegradable scaffold; and the last underwent reconstruction using a bladder-shaped biodegradable scaffold that delivered seeded autologous urothelial cells and smooth muscle cells (Oberpenning et al., 1999). The cystectomy-only and non-seeded controls maintained average bladder capacities of 22% and 46% of preoperative values, respectively, compared with 95% in the cell-seeded tissue-engineered bladder replacements (Figure 66.2). The subtotal cystectomy reservoirs that were not reconstructed and the polymer-only reconstructed bladders showed a marked decrease in bladder compliance (10% and 42% total compliance). The compliance of the cell-seeded tissue-engineered bladders showed almost no difference from preoperative values that were measured when the native bladder was present (106%). Histologically, the non-seeded scaffold bladders presented a pattern of normal urothelial cells with a thickened fibrotic submucosa and a thin layer of muscle fibers (Figure 66.2b). The retrieved tissue-engineered bladders showed a normal cellular organization, consisting of a trilayer of urothelium, submucosa, and muscle (Figure 66.2c) (Oberpenning et al., 1999). Preliminary clinical trials for the application of this technology have been performed and are under evaluation. Genital Tissues Tissue-engineering techniques have been used to reconstruct male and female genital tissues.

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Figure 66.2 Comparison of tissue-engineered neo-bladders. Gross specimens and cystograms at 11 months of the cystectomy-only, non-seeded controls, and cell-seeded tissue-engineered bladder replacements. The cystectomy-only bladder had a capacity of 22% of the preoperative value and a decrease in bladder compliance to 10% of the preoperative value. The non-seeded controls showed significant scarring with a capacity of 46% of the preoperative value and a decrease in bladder compliance to 42% of the preoperative value. An average bladder capacity of 95% of the original precystectomy volume was achieved in the cell-seeded tissue-engineered bladder replacements and the compliance showed almost no difference from preoperative values that were measured when the native bladder was present (106%).

Corporal Smooth Muscle Because one of the major components of the phallus is corporal smooth muscle, the creation of autologous functional and structural corporal tissue de novo would be beneficial. To examine functional parameters of engineered corpora, acellular corporal collagen matrices were obtained from donor rabbit penis, and autologous corpus cavernosal smooth muscle, and endothelial cells were harvested, expanded, and seeded on the matrices. The entire rabbit corpora was removed and replaced with engineered scaffolds. The experimental corporal bodies demonstrated intact structural integrity by cavernosography and showed similar pressure by cavernosometry when compared with normal controls. The control rabbits (without cells) failed to show normal erectile function throughout the study. Mating activity in the animals with the engineered corpora appeared normal by 1 month after implantation. The presence of sperm was confirmed during mating and was present in all the rabbits with the engineered corpora. The female rabbits mated with the animals implanted with engineered corpora and also conceived and delivered healthy pups. Animals implanted with the matrix alone were unable to demonstrate normal mating activity and failed to ejaculate into the vagina (Chen, 2005). Engineered Penile Prostheses Although silicone is an accepted biomaterial for penile prostheses, biocompatibility remains a concern (Thomalla et al., 1987; Nukui et al., 1997). The use of a natural prosthesis composed of autologous cells may be advantageous.

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A recent study using an autologous system investigated the feasibility of applying the engineered cartilage rods in situ (Yoo et al., 1999). Autologous chondrocytes harvested from rabbit ear were grown and expanded in culture. The cells were seeded onto biodegradable poly-L-lactic acid-coated PGA polymer rods and implanted into the corporal spaces of rabbits. Examination at retrieval 1 month later showed the presence of well-formed, milky-white cartilage structures within the corpora. All polymers were fully degraded by 2 months. There was no evidence of erosion or infection in any of the implantation sites. Subsequent studies assessed the long-term functionality of the cartilage penile rods in vivo (Yoo et al., 1999). To date, the animals have done well and can copulate and impregnate their female partners without problems. Female Genital Tissues Congenital malformations of the uterus may have profound implications clinically. Patients with cloacal exstrophy or intersex disorders may not have sufficient uterine tissue for future reproduction. We investigated the possibility of engineering functional uterine tissue using autologous cells (Wang et al., 2003). Autologous rabbit uterine smooth muscle and epithelial cells were harvested, then grown, and expanded in culture. These cells were seeded onto preconfigured uterine-shaped biodegradable polymer scaffolds, which were then used for subtotal uterine tissue replacement in the corresponding autologous animals. Upon retrieval 6 months after implantation, histological, immunocytochemical, and Western blot analyses confirmed the presence of normal uterine tissue components. Biomechanical analyses and organ bath studies showed that the functional characteristics of these tissues were similar to those of normal uterine tissue. Breeding studies using these engineered uteri are currently being performed. Several pathologic conditions, including congenital malformations and malignancy, can adversely affect normal vaginal development or anatomy. Vaginal reconstruction has traditionally been challenging due to the paucity of available native tissue. Vaginal epithelial and smooth muscle cells of female rabbits were harvested, grown, and expanded in culture. The cells were seeded onto biodegradable polymer scaffolds, which were then implanted into nude mice for up to 6 weeks. Immunocytochemical, histological, and Western blot analyses confirmed the presence of vaginal tissue phenotypes. Electrical field stimulation studies in the tissue-engineered constructs showed similar functional properties to those of normal vaginal tissue. When these constructs were used for autologous total vaginal replacement, patent vaginal structures were noted in the tissue-engineered specimens, while the non-seeded structures were noted to be stenotic (De Filippo, 2003). Ureter Ureteral non-seeded matrices have been used as a scaffold for the ingrowth of ureteral tissue in rats. On implantation, the acellular matrices promoted the regeneration of the ureteral wall components (Dahms et al., 1997). Ureteral replacement with polytetrafluoroethylene (Teflon) grafts was also attempted in dogs, but with poor functional results (Baltaci, 1998). In a more recent study, non-seeded ureteral collagen acellular matrices were tabularized, but attempts to use them to replace 3 cm segments of canine ureters were unsuccessful (Osman, 2004). Cell-seeded biodegradable polymer scaffolds have been used with more success to reconstruct ureteral tissues. In one study, urothelial and smooth muscle cells isolated from bladders and expanded in vitro were seeded onto PGA scaffolds with tubular configurations and implanted subcutaneously into athymic mice. After implantation, the urothelial cells proliferated to form a multilayered luminal lining of tubular structures, while the smooth muscle cells organized into multilayered structures surrounding the urothelial cells. Abundant angiogenesis was evident. Polymer scaffold degradation resulted in the eventual formation of natural urothelial tissues. This approach has also been used to replace ureters in dogs (Yoo, 1995).

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Renal Structures Due to its complex structure and function, the kidney is possibly the most challenging organ in the genitourinary system to reconstruct using tissue-engineering techniques. However concepts for a bioartificial kidney are emerging. Some investigators are pursuing the replacement of isolated kidney function parameters with the use of extracorporeal units, while others are working toward the replacement of total renal function by tissue-engineered bioartificial structures. Ex vivo Renal Units Although dialysis is currently the most prevalent form of renal replacement therapy, the relatively high rates of morbidity and mortality have spurred investigators to seek alternative solutions involving ex vivo systems. To assess the viability and physiologic functionality of a cell-seeded device to replace the filtration, transport, metabolic, and endocrinologic functions of the kidney in acutely uremic dogs, researchers introduced a synthetic hemofiltration device combined with a renal tubular cell therapy device (containing porcine renal tubules in an extracorporeal perfusion circuit). Levels of potassium and blood urea nitrogen (BUN) were controlled during treatment with the device. The fractional reabsorption of sodium and water was possible, and active transport of potassium, bicarbonate, and glucose and a gradual ability to excrete ammonia was observed. These results demonstrated the technologic feasibility of an extracorporeal assist device that is reinforced by the use of proximal tubular cells (Humes et al., 1999). Using similar techniques, a tissue-engineered bioartificial kidney – consisting of a conventional hemofiltration cartridge in series with a renal tubule assist device containing human renal proximal tubule cells – was used in patients with acute renal failure in the intensive care unit. Initial clinical experience with the bioartificial kidney and the renal tubule assist device suggests that such therapy may provide a dynamic and individualized treatment program as assessed by acute physiologic and biochemical indices (Humes et al., 2003). In vivo Renal Structures Another method of improving renal function involves augmenting renal tissue with kidney cell expansion in vitro and subsequent autologous transplantation. The feasibility of achieving renal cell growth, expansion, and in vivo reconstitution with tissue-engineering techniques has been explored. Most recently, an attempt was made to harness the reconstitution of renal epithelial cells to generate functional nephron units. Renal cells harvested and expanded in culture were seeded onto a tubular device constructed from a polycarbonate membrane connected at one end to a Silastic catheter terminating into a reservoir. The device was implanted into athymic mice. Histologic examination of the implanted devices over time revealed extensive vascularization, with formation of glomeruli and highly organized tubule-like structures. Immunocytochemical staining confirmed the renal phenotype. Yellow fluid consistent with the makeup of dilute urine in its creatinine and uric acid concentrations was retrieved from inside the implant (Yoo, 1996). Further studies using nuclear transfer techniques have been performed showing the formation of renal structures in cows (Figure 66.3) (Lanza et al., 2002). Challenges facing this technology include the expansion to larger, three-dimensional structures.

ADDITIONAL APPLICATIONS Tissue engineering and cell therapy hold promise for a number of additional genitourinary applications. Fetal Tissue Engineering Improved prenatal diagnostic techniques have led to the use of intervention before birth to reverse potentially life-threatening processes. Several strategies may be pursued to facilitate the future prenatal management of

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Figure 66.3 Creation of kidney tissue from therapeutic cloning and tissue-engineering strategies. (a) Illustration of the tissue-engineered renal unit. (b) Renal unit seeded with cloned cells, 3 months after implantation, showing the accumulation of urine-like fluid. (c) There was a clear unidirectional continuity between the mature glomeruli, their tubules, and the polycarbonate membrane. (d) Elispot analyses of the frequencies of T cells that secrete IFN-gamma after primary and secondary stimulation with allogeneic renal cells, cloned renal cells, or nuclear donor fibroblasts. urologic disease. Having a ready supply of urologic-associated tissue for surgical reconstruction at birth may be advantageous. Theoretically, once the diagnosis of the pathologic condition is confirmed prenatally, a small tissue biopsy could be obtained under US guidance. These biopsy materials could then be processed and the various cell types expanded in vitro. Using tissue-engineering techniques, reconstituted structures in vitro could then be readily available at the time of birth for reconstruction (Fauza et al., 1998a). Injectable Therapies Both urinary incontinence and vesicoureteral reflux (VUR) are common conditions affecting the genitourinary system that can be treated with injectable bulking agents. The ideal substance for endoscopic treatment of VUR and incontinence should be injectable, non-antigenic, non-migratory, volume stable, and safe for human use. Animal studies have shown that chondrocytes can be easily harvested and combined with alginate in vitro; the suspension can be easily injected cystoscopically; and the elastic cartilage tissue formed is able to correct VUR without any evidence of obstruction (Atala, 1994). The first human application of cell-based tissue-engineering technology for urologic applications occurred with the injection of chondrocytes for the correction of VUR in children and for urinary incontinence in adults (Figure 66.4) (Diamond, 1999; Bent et al., 2001). Using cell therapy techniques, the use of autologous smooth muscle cells was explored for both urinary incontinence and VUR applications (Cilento, 1995). The potential use of injectable, cultured myoblasts for

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Figure 66.4 Tissue-engineered bulking agent. Chondrocytes are harvested and combined with alginate in vitro, and the suspension is injected cystoscopically as a bulking agent to treat urinary incontinence.

the treatment of stress urinary incontinence (SUI) has also been investigated (Yokoyama et al., 1999; Chancellor et al., 2000). Use of injectable muscle precursor cells has also been studied for treatment of urinary incontinence due to irreversible urethral sphincter injury or maldevelopment (Yiou, 2003). A clinical trial involving the use muscle-derived stem cells (MDSC) to treat SUI has also been performed with good results. Biopsies of skeletal muscle were obtained, and autologous myoblasts and fibroblasts were cultured. Under US guidance, myoblasts were injected into the rhabdosphincter, and fibroblasts mixed with collagen were injected into the submucosa. One year following injection, the thickness and function of the rhabdosphincter had significantly increased, and all patients were continent (Strasser et al., 2004). These are the first demonstrations of the replacement of both sphincter muscle tissue and its innervation by the injection of muscle precursor cells. In addition, injectable muscle-based gene therapy and tissue engineering were combined to improve detrusor function in a bladder injury model, and may have potential as a novel treatment option for urinary incontinence (Huard et al., 2002). Testicular Hormone Replacement Patients with testicular dysfunction require androgen replacement for somatic development. Conventional treatment consists of periodic intramuscular injections of chemically modified testosterone or, more recently, skin patch applications. However, long-term non-pulsatile testosterone therapy is not optimal and can cause multiple problems, including erythropoiesis and bone density changes. A system was designed wherein Leydig cells were microencapsulated for controlled testosterone replacement. Purified Leydig cells were isolated and encapsulated in an alginate-poly-L-lysine solution. The encapsulated Leydig cells were injected into castrated animals, and serum testosterone was measured serially; the animals were able to maintain testosterone levels in the long term (Machluf et al., 1998b). These studies suggest that microencapsulated Leydig cells may be able to replace or supplement testosterone in situations where anorchia or testicular failure is present.

CONCLUSION Tissue-engineering efforts are currently being undertaken for every type of tissue and organ within the urinary system. Most of the effort expended to engineer genitourinary tissues has occurred within the last decade. Tissue-engineering techniques require a cell culture facility designed for human application. Recent progress suggests that engineered urologic tissues and cell therapy may have clinical applicability. Personnel who have

Genitourinary System

mastered the techniques of cell harvest, culture, and expansion as well as polymer design are essential for the successful application of this technology. Before these engineering techniques can be applied to humans, further studies need to be performed in many of the tissues described.

ACKNOWLEDGMENTS The author would like to express his appreciation for editorial assistance provided by Diane Q. Mann, M.S. Dr. Atala is the W.H. Boyce Professor and Chair, Department of Urology, and Director, Wake Forest Institute for Regenerative Medicine, Wake Forest University School of Medicine, Winston-Salem, NC.

REFERENCES Amiel, G.E. and Atala, A. (1999). Current and future modalities for functional renal replacement. Urol. Clin. N. Am. 26: 235–246. Atala, A. (1994). Use of non-autologous substances in VUR and incontinence treatment. Dial. Pediatr. Urol. 17: 11–12. Atala, A. (1995). Commentary on the replacement of urologic associated mucosa. J. Urol. 156: 338–339. Atala, A. (1997). Tissue engineering in the genitourinary system. In: Atala, A., Mooney, D. (eds.), Tissue Engineering. Boston: Birkhauser Press, p. 149. Atala, A. (1998). Autologous cell transplantation for urologic reconstruction. J. Urol. 159: 2–3. Atala, A. (1999). Future perspectives in reconstructive surgery using tissue engineering. Urol. Clin. N. Am. 26: 157–165. Atala, A., Cima, L.G., Kim, W., Paige, K.T., Vacanti, J.P., Retik, A.B. and Vacanti, C.A. (1993a). Injectable alginate seeded with chondrocytes as a potential treatment for vesicoureteral reflux. J. Urol. 150: 745–747. Atala, A., Freeman, M.R., Vacanti, J.P., Shepard, J. and Retik, A.B. (1993b) Implantation in vivo and retrieval of artificial structures consisting of rabbit and human urothelium and human bladder muscle. J. Urol. 150: 608–612. Atala, A., Guzman, L. and Retik, A. (1999). A novel inert collagen matrix for hypospadias repair. J. Urol. 162: 1148–1151. Atala, A., Kim, W., Paige, K.T., Vacanti, C.A. and Retik, A.B. (1994). Endoscopic treatment of vesicoureteral reflux with chondrocyte-alginate suspension. J. Urol. 152: 641–643. Atala, A., Vacanti, J.P., Peters, C.A., Mandell J., Retick, J.B. and Freeman, M.R. (1992). Formation of urothelial structures in vivo from dissociated cells attached to biodegradable polymer scaffolds in vitro. J. Urol. 148: 658–662. Badylak, S.F., Lantz, G.C., Coffey, A. and Geddes, L.A. (1989). Small intestinal submucosa as a large diameter vascular graft in the dog. J. Surg. Res. 47: 74–80. Baltaci, S., Ozer, G., Ozer, E., Soygur, T., Besalti, O. and Anafarta, K. (1998). Failure of ureteral replacement with Gore-Tex tube grafts. Urology 51: 400–403. Bazeed, M.A., Thüroff, J.W., Schmidt, R.A. and Tanagho, E.A. (1983). New treatment for urethral strictures. Urology 21: 53–57. Bent, A., Tutrone, R., McLennan, M., Lloyd, L.K., Kennelly, M.J. and Badlani, G. (2001). Treatment of intrinsic sphincter deficiency using autologous ear chondrocytes as a bulking agent. Neurourol. Urodynam. 20: 157–165. Blandy, J.P. (1961). Neal pouch with transitional epithelium and anal sphincter as a continent urinary reservoir. J. Urol. 86: 749–767. Blandy, J.P. (1964). The feasibility of preparing an ideal substitute for the urinary bladder. Ann. R. Coll. Surg. 35: 287–311. Brittberg, M., Lindahl, A., Nilsson, A., Ohlsson, C., Isaksson, O. and Peterson, L. (1994). Treatment of deep cartilage defects in the knee with autologous chondrocyte transplantation. N. Engl. J. Med. 331: 889–895. Chancellor, M.B., Yokoyama, T., Tirney, S., Mattes, C.E., Ozawa, H., Yoshimura, N., de Groat, W.C. and Huard, J. (2000). Preliminary results of myoblast injection into the urethra and bladder wall: a possible method for the treatment of stress urinary incontinence and impaired detrusor contractility. Neurourol. Urodynam. 19: 279–287. Chen, F., Yoo, J.J. and Atala, A. (1999). Acellular collagen matrix as a possible off the shelf biomaterial for urethral repair. Urology 54: 407–410. Chen, K.L., Yoo, J.J. and Atala, A. (2005). Total penile corpora cavernosa replacement using tissue engineering techniques. Regenerate (abstract). Cheng, E., Rento, R., Grayhack, T.J., Oyasu, R. and McVary, K.T. (1994). Reversed seromuscular flaps in the urinary tract in dogs. J. Urol. 152: 2252–2257.

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Cilento, B.G. and Atala, A. (1995). Treatment of reflux and incontinence with autologous chondrocytes and bladder muscle cells. Dial. Pediatr. Urol. 18: 11–15. Cilento, B.G., Freeman, M.R., Schneck, F.X., Retik, A.B. and Atala, A. (1994). Phenotypic and cytogenetic characterization of human bladder urothelia expanded in vitro. J. Urol. 152: 655–670. Dahms, S.E., Piechota, H.J., Nunes, L., Dahiya, R., Lue, T.F. and Tanagho, E.A. (1997). Free ureteral replacement in rats: Regeneration of ureteral wall components in the acellular matrix graft. Urology 50: 818–825. DeFilippo, R.E., Pohl, H.G., Yoo, J.J. and Atala, A. (2002). Total penile urethra replacement with autologous cell-seeded collagen matrices. 167: 152–153. DeFilippo, R.E., Yoo, J.J. and Atala, A. (2003). Engineering of vaginal tissue in vivo. Tissue Eng. 9: 301–306. Dewan, P.A. (1998). Autoaugmentation demucosalized enterocystoplasty. World J. Urol. 16: 255–261. Diamond, D.A. and Caldamone, A.A. (1999). Endoscopic correction of vesicoureteral reflux in children using autologous chondrocytes: preliminary results. J. Urol. 162: 1185–1188. Fauza, D.O., Fishman, S., Mehegan, K. and Atala, A. (1998a). Videofetoscopically assisted fetal tissue engineering: bladder augmentation. J. Pediatr. Surg. 33: 7–12. Fauza, D.O., Fishman, S, Mehegan, K. and Atala, A. (1998b). Videofetoscopically assisted fetal tissue engineering: skin replacement. J. Pediatr. Surg. 33: 357–361. Folkman, J. and Hochberg, M. (1973). Self-regulation of growth in three dimensions. J. Exp. Med. 138: 745–753. Harada, N., Yano, H., Ohkawa, T., Misse, T., Kurita, T. and Nagahara, A. (1965). New surgical treatment of bladder tumors: mucosal denudation of the bladder. Br. J. Urol. 37: 545–547. Huard, J., Yokoyama, T., Pruchnic, R., Ou, A., Li, Y., Lee, J.Y., Somogyi, G.T., de Groat, W.C. and Chancellor, M.B. (2002). Muscle-derived cell-mediated ex vivo gene therapy for urological dysfunction. Gene Ther. 9: 1617–1626. Humes, H.D., Buffington, D.A., MacKay, S.M., Funke, A.J. and Weitzel, W.F. (1999). Replacement of renal function in uremic animals with a tissue engineered kidney. Nat. Biotech. 17: 451–455. Humes, H.D., Weitzel, W.F., Bartlett, R.H., Swaniker, F.C. and Paganini, E.P. (2003). Renal cell therapy is associated with dynamic and individualized responses in patients with acute renal failure. Blood Purif. 21: 64–71. Kaefer, M., Hendren, H., Bauer, S., Goldenblatt, P., Peters, C.A., Atala, A. and Retik, A.B. (1998). Reservoir calculi: a comparison of reservoirs constructed from stomach and other enteric segments. J .Urol. 160: 2187–2190. Kaefer, M., Tobin, M., Hendren, H., Bauer, S.B., Peters, C.A., Atala, A., Colodny, A.H., Mandell, J. and Retik, A.B. (1997). Continent urinary diversion: the Children’s Hospital experience. J. Urol. 157: 1394–1399. Kershen, R.T. and Atala, A. (1999). Advances in injectable therapies for the treatment of incontinence and vesicoureteral reflux. Urol. Clin. N. Am. 26: 81–94. Kim, B.S. and Mooney, D.J. (1998). Development of biocompatible synthetic extracellular matrices for tissue engineering. Trends Biotechnol. 16: 224–230. Kropp, B.P., Cheng, E.Y., Lin, H.K. and Zhang, Y. (2004). Reliable and reproducible bladder regeneration using unseeded distal small intestinal submucosa. J. Urol. 172:1710–1713. Kropp, B.P., Ludlow, J.K., Spicer, D., Rippy, M.K., Badylak, S.F., Adams, M.C., Keating, M.A., Rink, R.C., Birhle, R. and Thor, K.B. (1998). Rabbit urethral regeneration using small intestinal submucosa onlay grafts. Urology 52: 138–142. Kropp, B.P., Rippy, M.K., Badylak, S.F., Adams, M.C., Keating, M.A., Rink, R.C. and Thor, K.B. (1996). Small intestinal submucosa: urodynamic and histopathologic evaluation in long term canine bladder augmentations. J. Urol. 155: 2098–2104. Lailas, N.G., Cilento, B. and Atala, A. (1996). Progressive ureteral dilation for subsequent ureterocystoplasty. J. Urol. 156: 1151– 1153. Landman, J., Olweny, E., Sundaram, C.P., Andreoni, C., Collyer, W.C., Rehman, J., Jerde, T.J., Lin, H.K., Lee, D.I., Nunlist, E.H., et al. (2004). Laparoscopic mid sagittal hemicystectomy and bladder reconstruction with small intestinal submucosa and reimplantation of ureter into small intestinal submucosa: 1-year followup. J. Urol. 171: 2450–2455. Lanza, R.P., Chung, H.Y., Yoo, J.J., Wettstein, P.J., Blackwell, C., Borson, N., Hofmeister, E., Schuch, G., Soker, S., Moraes, C.T., et al. (2002). Generation of histocompatible tissues using nuclear transplantation. Nat. Biotechnol. 20: 689–696. le Roux, P.J. (2005). Endoscopic urethroplasty with unseeded small intestinal submucosa collagen matrix grafts: a pilot study. J. Urol. 173: 140–143. Machluf, M. and Atala, A. (1998a). Emerging concepts for tissue and organ transplantation. Graft 1: 31–37. Machluf, M., Boorjian, S., Caffaratti, J., Kershen, R., Orsola, A. and Atala, A. (1998b). Microencapsulation of Leydig cells: a new system for the therapeutic delivery of testosterone. Pediatrics 102(Suppl.): 32.

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McDougal, W.S. (1992). Metabolic complications of urinary intestinal diversion. J. Urol. 147: 1199–1208. Nukui, F., Okamoto, S., Nagata, M., Kurukawa, J. and Fukui, J. (1997). Complications and reimplantation of penile implants. Int. J. Urol. 4: 52–54. Oberpenning, F.O., Meng, J., Yoo, J. and Atala, A. (1999). De novo reconstitution of a functional urinary bladder by tissue engineering. Nat. Biotechnol. 17: 149–155. Oesch, I. (1988). Neourothelium in bladder augmentation: an experimental study in rats. Eur. Urol. 14: 328–329. Olsen, L., Bowald, S., Busch, C., Carlsten, J. and Eriksson, I. (1992). Urethral reconstruction with a new synthetic absorbable device. Scan. J. Urol. Nephrol. 26: 323–326. Osman, Y., Shokeir, A., Gabr, M., El-Tabey, N., Mohsen, T. and El-Baz, M. (2004). Canine ureteral replacement with long acellular matrix tube: is it clinically applicable? J. Urol. 172: 1151–1154. Park, H.J., Kershen, R., Yoo, J. and Atala, A. (1999). Reconstitution of human corporal smooth muscle and endothelial cells in vivo. J. Urol. 162: 1106–1109. Piechota, H.J., Dahms, S.E., Nunes, L.S., Dahiya, R., Lue, T.F. and Tanagho, E.A. (1998). In vitro functional properties of the rat bladder regenerated by the bladder acellular matrix graft. J. Urol. 159: 1717–1724. Ponder, K.P., Gupta, S., Leland, F., Darlington, G., Finegold, M., DeMayo, J., Ledley, F.D., Chowdbury, J.R. and Woo, S.L. (1991). Mouse hepatocytes migrate to liver parenchyma and function indefinitely after intrasplenic transplantation. Proc. Natl. Acad. Sci. USA 88: 1217–1221. Portis, A.J., Elbahnasy, A.M., Shalhav, A.L., Brewer, A., Humphrey, P., McDougall, E.M. and Clayman, R.V. (2000). Laparoscopic augmentation cystoplasty with different biodegradable grafts in an animal model. J. Urol. 164: 1405–1411. Probst, M., Dahiya, R., Carrier, S. and Tanagho, E.A. (1997). Reproduction of functional smooth muscle tissue and partial bladder replacement. Br. J. Urol. 79: 505–515. Salle, J., Fraga, C., Lucib, A., Lampertz, M., Jobim, G., Jobim, G. and Putten, A. (1990). Seromuscular enterocystoplasty in dogs. J. Urol. 144: 454–456. Satar, N., Yoo, J. and Atala, A. (1999). Progressive bladder dilation for subsequent augmentation cystoplasty. J. Urol. 162: 829–831. Sievert, K.D., Bakircioglu, M.E., Nunes, L., Tu, R., Dahiya, R. and Tanagho, E.A. (2000). Homologous acellular matrix graft for urethral reconstruction in the rabbit: histological and functional evaluation. J. Urol. 163: 1958–1965. Strasser, H., Berjukow, S., Markstiner, R., Margreiter, E., Hering, S., Bartsch, G. and Hering, S. (2004). Stem cell therapy for urinary stress incontinence. Exp. Gerontol. 39: 1259–1265. Sutherland, R.S., Baskin, L.S., Hayward, S.W. and Cunha, G.R. (1996). Regeneration of bladder urothelium, smooth muscle, blood vessels and nerves into an acellular tissue matrix. J. Urol. 156: 571–577. Thomalla, J.V., Thompson, S.T., Rowland, R.G. and Mulcahy, J.J. (1987). Infectious complications of penile prosthetic implants. J. Urol. 138: 65–67. Wang, T., Koh, C.J. and Yoo, J.J. (2003). Creation of an engineered uterus for surgical reconstruction Proceedings of the American Academy of Pediatrics Section on Urology, New Orleans, LA. Yiou, R., Yoo, J.Y. and Atala, A. (2003). Restoration of functional motor units in a rat model of sphincter injury by muscle precursor cell autografts. Transplantation 76: 1053–1060. Yokoyama, T., Chancellor, M.B. and Watanabe, T. (1999). Primary myoblasts injection into the urethra and bladder as a potential treatment of stress urinary incontinence and impaired detrusor contractility: long term survival without significant cytotoxicity. J. Urol. 161: 307. Yoo, J.J., Ashkar, S. and Atala, A. (1996). Creation of functional kidney structures with excretion of urine-like fluid in vivo. Pediatrics 98(Suppl.): 605. Yoo, J.J. and Atala, A. (1997). A novel gene delivery system using urothelial tissue engineered neo-organs. J. Urol. 158: 1066–1070. Yoo, J.J., Meng, J., Oberpenning, F. and Atala, A. (1998). Bladder augmentation using allogenic bladder submucosa seeded with cells. Urology 51: 221–225. Yoo, J., Park, H.J., Lee, I. and Atala, A. (1999). Autologous engineered cartilage rods for penile reconstruction. J. Urol. 162: 1119–1121. Yoo, J.J., Satar, N., Retik, A.B. and Atala, A. (1995). Ureteral replacement using biodegradable polymer scaffolds seeded with urothelial and smooth muscle cells. J. Urol. 153(Suppl.): 375A.

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67 Tissue Engineering of the Reproductive System Stefano Giuliani, Laura Perin, Sargis Sedrakyan, and Roger De Filippo

INTRODUCTION Human organ replacement is limited by a donor shortage, and problems of tissue compatibility, and rejection. Tissue engineering combines the principles and methods of the life sciences with those of engineering to elucidate a fundamental understanding of structure–function relationships of normal and diseased tissues, in order to facilitate the development of materials and methods to repair damaged or diseased tissues, and to create entire tissue replacements. A large number of materials, including naturally derived and synthetic polymers have been utilized to facilitate prostheses for genitourinary system. Porous, absorbable matrices made of natural or synthetic polymers are currently being investigated as scaffolds for genitourinary tissue transplantation. These biodegradable polymers include poly(glycolic acid) (PGA), polylactide (PLA), poly(glycolide-co-lactide) (PGLA), poly(caprolactone) (PCL), poly (glycolide-co-ε-caprolactone), collagen, alginate, hyaluronate, and laminin. These scaffolds require proper biocompatibility, degradability, mechanical stability, high surface area/volume ratio, and proper pore size. A highly porous scaffold is desirable to allow large number of cell seeding or migration throughout the material and the pore size affects both tissue ingrowth and the internal surface area available for cell attachment (Mikos et al., 1994). The success of using cell transplantation strategies for genitourinary reconstruction depends on our ability to use donor tissue efficiently and to provide the right conditions for long-term survival, differentiation, and growth (Patrick et al.,1998). Tissue engineering follows the principles of cell transplantation, materials science, and engineering toward the development of biological substitutes, which would restore and maintain normal function. Tissue engineering may involve matrices alone, wherein the body’s natural ability to regenerate is used to orientate or direct new tissue growth, or the use of matrices with cells. When cells are used for tissue engineering, donor tissue is dissociated into individual cells, which are either implanted directly into the host, or expanded in culture, attached to a support matrix, and reimplanted after expansion. The implanted tissue can be heterologous, allogeneic, or autologous. Ideally, this approach might allow lost tissue function to be restored or replaced in total and with limited complications (Atala, 1997). The use of autologous cells would avoid rejection, wherein a biopsy of tissue is obtained from the host, the cells are dissociated and expanded in vitro, reattached to a matrix, and implanted into the same host.

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MALE Urethra Congenital or acquired disorders of the urethra remain a challenge in the urology field. Various urethral conditions such as strictures, traumatic defects, congenital defects, and cancer often require additional tissue for reconstruction. Under circumstances in which there is a lack of urethral mucosa for adequate reconstruction, tissue from other sources has been used, such as genital and extragenital skin flaps or grafts (Xu et al., 2002). Complications such as hair growth, graft shrinkage, stricture, stone formation, and diverticuli have been associated with skin grafts (Hendran and Reda, 1986; Ozcan and Kahveci, 1987). Anatomy In men the urethra is divided into three parts, named after the location: the prostatic urethra crosses through the prostate gland; the membranous urethra is a small (1 or 2 cm) portion passing through the external urethral sphincter, this is the narrowest part of the urethra; the spongy (or penile) urethra runs along the length of the penis on its ventral (underneath) surface. It is about 15–16 cm in length, and travels through the corpus spongiosum (Figure 67.1). Bladder

Urethral crest Openings of prostatic utricle and ejaculatory ducts Prostatic part of urethra Membranous part of urethra

Small lacuna

Lacuna magna

Ext. urethral orifice

Figure 67.1 Anatomy of Urethra.

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The epithelium of the urethra starts off as transitional cells as it exits the bladder. Further along the urethra there are stratified columnar cells, then stratified squamous cells near the external meatus (exit hole). There are small mucus secreting urethral glands that help to protect the epithelium from the corrosive urine. Cell Growth One of the initial limitations of applying cell-based tissue engineering techniques to urologic organs had been the previously encountered inherent difficulty of growing genitourinary-associated cells in large quantities. Normal urothelial cells could be grown in the laboratory setting, but with limited expansion. Several protocols were developed over the last two decades, which improved urothelial growth and expansion (Cilento et al., 1994; Liebert et al., 1997; Scriven et al., 1997; Puthenveettil et al., 1999). A system of urothelial cell harvest was developed which does not use any enzymes or serum and has a large expansion potential. Using these methods of cell culture, it is possible to expand an urothelial strain from a single specimen which initially covers a surface area of 1 cm2 to one covering a surface area of 4202 m2 within 8 weeks. These studies indicated that it should be possible to collect autologous urothelial cells from human patients, expand them in culture, and return them to the human donor in sufficient quantities for reconstructive purposes. Normal human bladder epithelial and muscle cells can be efficiently harvested from surgical material, extensively expanded in culture, and their differentiation characteristics, growth requirements, and other biological properties can be studied (see Protocols I.A and B) (Liebert et al., 1991; Tobin et al., 1994). The cells were seeded onto non-woven meshes of PGA (Cilento et al., 1995). Partial urethrectomies were performed in rabbits and a segment of the polymer mesh of the appropriate diameter was interposed to form the neourethra in each animal. There was no evidence of voiding difficulties or any other complications. Retrograde urethrograms showed no evidence of stricture formation. Histologic examination of the neourethras demonstrated complete reepithelialization of the polymer mesh implanted sites by day 14, and reepithelialization continued for the entire duration of the study. Polymer fiber degradation was evident 14 days after implantation. Protocol I.A: Urothelial Cell Culture 1. Materials and medium

(a) Tissue source: bladder tissue. (b) Medium: keratinocyte and serum-free medium (Gibco/BRL), bovine pituitary extract (25 mg/500 ml medium), and recombinant epidermal growth factor (2.5 g/500 ml medium). 2. Tissue harvest (a) Obtain bladder specimen. (b) Gently rinse the specimen several times with medium in culture plates. (c) Mechanically scrape urothelial surface gently with a No. 10 scalpel blade. Be sure to use gentle short strokes and avoid cutting into the specimen. Urothelial cell clumps should be visible as tiny opaque material dispersing into the medium. (d) Aspirate urothelial cell-medium suspension and plate the cells in a 24-well cell culture plate with approximately 0.5 ml of the suspension in each well. Add an additional 0.5 ml to make a final volume of 1 ml. Incubate cells at 37°C with 5% CO2. (e) On the following day, aspirate the medium from the wells and replace with fresh medium. (f) Centrifuge the cells in the aspirate medium at 1000 rpm for 4 min. (g) Remove the supernatant and resuspend the cells in 3–4.5 ml of fresh medium. Replate the cells in new wells. 3. Maintenance of urothelial cells (a) Replace the medium with fresh warm (37°C) medium every 3 days, depending on the cell density. (b) Tripsinize cells when they are 80–90% confluent.

Tissue Engineering of the Reproductive System

4. Subculture of corporal smooth muscle cells

(a) Remove medium and add 1 ml of phosphate-buffered saline–ethylenediaminetetraacetic acid (PBS–EDTA) (0.5 M) to each well or 10 ml to each 10-cm culture plate. Observe the cells under a phase contrast microscope. (b) When cell–cell junctions are separated for the majority of the cells, remove PBS–EDTA and add 300 μl of trypsin–EDTA to each well or 5 ml to each 10-cm culture plate. (c) Periodically agitated the plates. When 80–90% of the cells are detached, add 30 μl of soybean trypsin inhibitor (Gibco/BRL, 294 mg of inibitor) to 20 ml of PBS to each well, or 700 μl to each 10-cm plate. Add 0.5 ml of medium to each well or 3 ml to each 10-cm plate. (d) Aspirate and centrifuge the cell suspension at 1000 rpm for 4 min, and remove the supernatant. (e) Resuspend cells and count the number of viable cells by means of trypan blue exclusion. (f) Aliquot the desired number of cells on the plate and place the cells in the incubator. Protocol I.B: Bladder Smooth Muscle Cell Culture 1. Materials and medium

(a) Tissue source: bladder tissue. (b) Medium: Dulbecco’s Modified Eagle’s Medium (DMEM), 10% fetal bovine serum (FBS), and antibiotic (penicillin (100 U/ml)-streptomycin (100 l/ml), and amphotericin B (0.25 l/ml). 2. Tissue harvest (a) Obtain fresh bladder tissue specimen. (b) Use sharp tenotomy scissors to cut muscle tissue into small fragments (2–3 mm). (c) Space muscle fragments evenly onto a cell culture plate (100 mm). (d) Allow muscle fragments to dry and adhere to the plate (5–10 min). (e) Add 15 ml of DMEM and incubate for 5 days. (f) Change medium on the sixth day and remove non-adherent tissue fragments. (g) When small islands of cell colony are formed, remove the tissue fragments and the change medium. (h) When sufficient cells are grown, trypsinize, count, and plate them onto new plates. 3. Maintenance of bladder smooth muscle cells (a) Feed cells every 3 days, depending on the cell density. (b) Trypsinize cells when they are 80–90% confluent. 4. Subculture of bladder smooth muscle cells (a) Remove medium and add 10 ml of PBS–EDTA (0.5 M) for 4 min. Confirm the separation of cell junction under a phase microscope. (b) Remove PBS–EDTA and add 5 ml of trypsin–EDTA. (c) Add 5 ml of medium when 80–90% of the cells lift under microscope. (d) Aspirate the cell suspension into a 15-ml test tube. (e) Aliquot the desired number of cells on the plate and make the volume of medium to a total of 10 ml. (f) Place the cells in the incubator. A variety of synthetic grafts composed of silicone, Teflon, or polyvinyl have been proposed for urethral reconstruction and erosion, dislodgment, fistula, stenosis, extravasation, and calcification have been associated with synthetic grafts (Guzman, 1999; Vozzi et al., 2002). In the past years, in case of limited urethral mucosa for adequate reconstruction, tissues from other sources have been used, such as genital and extragenital skin flaps or grafts, mucosal grafts from the bladder or buccal regions, tunica vaginalis, and peritoneal grafts (Humby, 1941; Ehrlich et al., 1989; Dessanti et al., 1992). Various biomaterials without cells have been used experimentally

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(in animal models) for the regeneration of urethral tissue, including PGA, and acellular collagen-based matrices from small intestine and bladder (Bazeed et al., 1983; Atala et al., 1992; Olsen et al., 1992; Kropp et al., 1998; Chen et al., 1999; Sievert et al., 2000). Some of these biomaterials, like acellular collagen matrices derived from bladder submucosa, have also been seeded with autologous cells for urethral reconstruction. Atala’s laboratory has been able to replace tubularized urethral segments with cell-seeded collagen matrices. Acellular collagen matrices derived from bladder submucosa have been used experimentally and clinically. In animal studies, segments of the urethra were resected and replaced with acellular matrix grafts in an onlay fashion. The animals were able to void through the neourethras (Chen et al., 1999). These results were confirmed clinically in a series of patients with hypospadias and urethral stricture disease (Atala, 1999; El kassaby et al., 2003). Cadaveric bladders were microdissected and the submucosal layers isolated. The submucosa was washed and decellularized. The matrix was used for urethral repair in patients with stricture disease (n  33; 28 adults, 5 children) and hypospadias (n  7 children). The matrices were trimmed to size and the neourethras were created by anastomosing the matrix in an onlay fashion to the urethral plate. The size of the neourethras ranged from 2 to 16 cm. Voiding histories, physical examination, retrograde urethrography, uroflowmetry, and cystoscopies were performed serially, pre- and postoperatively, with up to a 7-year follow-up. After a 4- to 7-year follow-up, 34 of the 40 patients had a successful outcome. Six patients with a urethral stricture had a recurrence, and one patient with hypospadias developed a fistula. The mean maximum urine flow rate significantly increased postoperatively. Cystoscopic studies showed adequate caliber conduits. Histologic examination of the biopsies showed the typical urethral epithelium. The use of an off the shelf matrix appears to be beneficial for patients with abnormal urethral conditions, and obviates the need for obtaining autologous grafts, thus decreasing operative time and eliminating donor site morbidity. Unfortunately, the above techniques are not applicable for tubularized urethral repairs. The collagen matrices are able to replace urethral segments when used in an onlay fashion. However, if a tubularized repair is needed, the collagen matrices need to be seeded with autologous cells (De Fillipo et al., 2002a, b). Autologous bladder epithelial and smooth muscle cells from male rabbits were grown and seeded on to preconfigured tubular matrices. The entire anterior urethra was resected and urethroplasties were performed with tubularized collagen matrices seeded with cells in nine animals, and without cells in six animals. Serial urethrograms showed a wide urethral caliber without strictures in the animals implanted with the cell-seeded matrices, and collapsed urethral segments with strictures within the unseeded scaffolds. Gross examination of the urethral implants seeded with cells showed normal appearing tissue without any evidence of fibrosis. Histologically, a transitional cell layer surrounded by muscle cell fiber bundles with increasing cellular organization over time were observed on the cell-seeded constructs. The epithelial and muscle phenotypes were confirmed with pAE1/AE3 and smooth muscle specific a-actin antibodies. A transitional cell layer with scant unorganized muscle fiber bundles and large areas of fibrosis were present at the anastomotic sites on the unseeded constructs. Therefore, tubularized collagen matrices seeded with autologous cells can be used successfully for total penile urethra replacement; whereas, tubularized collagen matrices without cells lead to poor tissue development and stricture formation. The cell-seeded collagen matrices form new tissue, which is histologically similar to native urethra. This technology may be applicable to patients requiring tubularized urethral repair. A variety of synthetic grafts composed of silicone, Teflon, and Dacron have been proposed for urethral reconstruction and then left behind because of their complications. These materials have been associated with erosion, dislodgment, fistula, stenosis, extravasation, or calcification (Hakky, 1976, 1977; Anwar et al., 1984). Biodegradable substitutes like a polyglactin fiber mesh tube coated with poly(hydroxybutyric acid) and hyaluronan benzyl ester has been used experimentally. Complete regeneration of the urethral epithelium and the adjacent connective tissue was achieved as a consequence of the fact that the scaffolds guided urothelial and connective tissue regeneration (Olsen et al., 1992; Italiano et al., 1997). Although several innovative tissues

Tissue Engineering of the Reproductive System

have been proposed as possible free grafts for urethral repair, it is evident that all have specific advantages and disadvantages. Free grafts of tubularized peritoneum were used as urethral tissue substitutes experimentally in rabbit. Organized multilayered graft epithelialization occurred; however, fistula formed in two of the animals (Shaul et al., 1996). Later, porcine small intestine submucosa (SIS) was used for urethral repair in a rabbit model to determine whether this material can evoke urethral regeneration. The SIS onlay grafts were shown to promote regeneration of the normal rabbit epithelium supported by a vascularized collagen and smooth muscle backing (Kropp et al., 1998). More recently, Nuininga et al. partially resected a 0.5–1 cm segment of the native urethra in 24 rabbits and a novel molecularly defined collagen-based biocompatible and biodegradable matrix graft was sewn into place and compared with SIS. They did not notice any differences between the two biomatrics and the major advantage is that the new biometrics proposed can be modulated in different ways such as variation in the porous matrix structures, incorporation of growth factors and binding of glycosaminoglycans (Nuininga et al., 2003). A naturally derived acellular collagen-based tissue substitute was developed from donor porcine bladder (see Protocol II). The acellular collagen matrix had been initially developed in our laboratory as a biomaterial for bladder augmentation. The results from this study demonstrated that the acellular matrix was biocompatible and was able, upon in vivo implantation, to form bladder tissue similar to the native bladder (Yoo et al., 1998). Protocol II: Acellular Collagen Matrix Preparation 1. Obtain donor bladder tissue. 2. Isolate the submucosa from the muscular and serosal layers means of microdissection techniques. 3. Treat tissue with distilled water in a magnetic stirring flask set at moderate speed for 24–48 h at 4°C. 4. Remove distilled water and treat with Triton X-100 (0.5%) and ammonium hydroxide (0.05%) in fresh 5.

6.

7. 8. 9. 10.

distilled water for 72 h in a stirring flask at 4°C. Wash with distilled water in a stirring flask for 24–48 h at 4°C. After this washing step, take a small piece of tissue for histological analysis to confirm any cellular remnants. Tissue matrix is usually decellularized at this time. After confirmation of decellularization, wash with distilled water in a stirring flask for 24–48 h at 4°C. Tissue retaining cellular components should undergo an additional cycle of treatment. Repeat Steps 4 and 5, and perform another histological analysis. After the washing cycle with distilled water, rinse with 1  PBS overnight. Freeze-dry the tissue sample overnight. Pack the samples and sterilize in ethylene oxide. Store until used. When ready to use equilibrate the tissue in 1  PBS or normal saline.

Penis The indication for extended phalloplastic procedures results from the severe congenital malformation, penile tissue loss from malignancies, trauma or other diseases, and gender dysphoria. Owing to the shortage of autologous penile tissue, multiple staged surgeries using non-genital tissues and silicone prostheses have been the mainstay in phallic reconstruction. However, graft failure and prothesis-related complications remain a problem. Creation of penile structures composed of autologous tissue would be preferable treatment approach for these patients. Replacement of penile tissue with alternative materials has been a challenge due to the unique anatomical architecture of the corporal bodies. One of the major limitations of penile tissue reconstruction is the availability of sufficient autologous tissue. Non-genital tissue sources have been used over the years; however, complications such as infection, graft failure,

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and donor site morbidity have posed continuing problems (Goodwin and Scott, 1952; Puckett and Montie, 1978; Chang and Hwang, 1984; Gilbert et al., 1988; Horton and Dean, 1990; Sharaby et al., 1995). The ability to engineer penile tissue composed of autologous cells would be beneficial. Anatomy The anatomy of the penis is complex and is comprised primarily of three separate cylinders. The two paired cylinders called the corpora cavernosa make up the majority of the bulk and the erectile functioning of the penis. Each of these cylinders is encased in a very tough thick sheath called tunica albuginea. The third cylinder of the penis is called corpus spongiosum, and it contains the urethra. The tissue around this erectile body is much thinner, and the cylinder actually sits in a groove created by the other two cylinders. As this structure approaches the end of the penis, it becomes swollen and is known as the glans, or the head of the penis. As this layer gets closer to the body, it expands to form the bulb. Covering all three of these cylinders is a thick tough membrane called Buck’s fascia. Finally, a final layer covers this area called Colles fascia, or the superficial layer. This is actually continuous with the abdominal wall and makes this whole supporting structure of the penis very tough, allowing it to take quite a bit of force and trauma. The shaft is covered by nearly hairless skin. Under skin lies the dense connective tissue of penile fascia. Tunica albuginea encircles all three corpora; divides corpora proximally but is incomplete distally. Corpora cavernosa are paired columns of erectile tissue located dorsally. Each column consists of a network of large venous sinuses separated by dense connective tissue septae, the trabeculae. Blood empties from central artery into helicine arteries into sinuses and is then drained by veins emptying into dorsal vein. Corpus spongiosum has similar arrangement as corpora cavernosa except contains penile uretra. Same arterial and venous relationship described for corpora cavernosa. Corpus Cavernosum Reconstruction Although consisting only of two important functional cell types (i.e. smooth muscle and endothelial cells) the tissue engineering of autologous penile tissue remain a challenge. Our initial effort was focused on the formation of corporal tissue, since corpus cavernosum is one of the major tissue components of the phallus. Human corporal smooth muscle cells were isolated, grown, and expanded in culture (see Protocol I.A). The cells were seeded on biodegradable PGA polymers for implantation. Multilayers of corporal smooth muscle cells were identified grossly and histologically. This study provided the evidence that cultured human corporal smooth muscle cells could be used in conjunction with biodegradable polymers to create cavernosal smooth muscle tissue in vivo. Protocol I.A: Corpus Cavernosal Smooth Muscle Cell Culture 1. Materials and medium

(a) Tissue source: human corpus cavernosum. (b) Medium: DMEM, 10% FBS, and antibiotic (penicillin (100 U/ml-streptomycin (100 μg/ml), amphotericin B (0.25 μg/ml)). 2. Tissue harvest (a) Obtain fresh cavernosal tissue specimen. (b) Use sharp tenotomy scissors to cut muscle tissue into small fragments (2–3 mm). (c) Space muscle fragments evenly onto a cell culture plate (100 mm). (d) Allow muscle fragments to dry and adhere to the plate (5–10 min). (e) Add 15 ml of DMEM and incubate for 5 days. (f) Change medium on the sixth day and remove non-adherent tissue fragments.

Tissue Engineering of the Reproductive System

(g) When small islands of cell colony are formed, remove the tissue fragments and change the medium. (h) When sufficient cells are grown, trypsinize, count, and plate the cells onto new plates. 3. Maintenance of corporal smooth muscle cells (a) Feed cells every 3 days, depending on the cell density. (b) Trypsinize cells when they are 80–90% confluent. 4. Subculture of corporal smooth muscle cells (a) Remove medium and add 10 ml of PBS–EDTA (0.5 M) over 4 min. Confirm the separation of cell junction under phase contrast microscope. (b) Remove PBS–EDTA and add 5 ml of trypsin–EDTA (c) Add 5 ml of medium when 80–90% of the cells lift under the microscope. (d) Aspirate the cell suspension into a 15-ml test tube. (e) Centrifuge the cells at 1000 rpm for 4 min and remove the supernatant. (f) Resuspend cells and use trypan blue exclusion to count the number of viable cells. (g) Aliquot the desired number of cells in the plate and make the volume of medium to a total of 10 ml. (h) Place the cells in the incubator. The main cellular components of corporal tissue consist of cavernosal smooth muscle and endothelial cells. In a subsequent study, we investigated the possibility of developing corporal tissue by combining smooth muscle and endothelial cells. Normal human cavernosal smooth muscle cells and ECV 304 human endothelial cells were seeded on biodegradable polymer for implantation (Park et al., 1999). ECV 304 endothelial cells were used in the study, to allow the investigator to distinguish the implanted cells from the host endothelial cells. The retrieved structures showed formation of distinct tissue structures, consisting of organized smooth muscle tissue adjacent to endothelial cells. Presence of vascular structures was evident. Each cell type was confirmed by means of various assessment methods. This study showed that human corporal muscle and endothelial cells seeded on biodegradable polymer scaffolds are able to form vascularized cavernosal tissue when implanted in vivo. Endothelial cells can act in concert with the native vasculature. These results suggest that the creation of well-vascularized autologous corpus-like tissue consisting of smooth muscle and endothelial cells may be possible. We developed a naturally derived collagen matrix, which is structurally similar to the native corporal architecture (Faike et al., 2003). Acellular collagen matrices, derived from rabbit corpora, were obtained by means of cell lysis technique (see Protocol I.B). Human corpus cavernosal muscle and endothelial cells were grown and expanded in culture (Protocol I.C). We have used human capillary cells, isolated from newborn foreskin via Ulex europaeus I (UEA-I)-coated Dynabeads (Jackson et al., 1990; Kraling and Bischoff, 1998). Primary human cavernosal smooth muscle and endothelial cells were seeded in a stepwise fashion. Cavernosal smooth muscle cells were initially seeded on the collagen matrices at a concentration of 30  106 cells/ml. The cells were allowed to attach and grow for 3 days in culture. Endothelial cells were then seeded at a concentration of 3  106 cells/ml. Cell matrices seeded with corporal cells were implanted in vivo. The implanted cell matrices showed neovascularity into the sinusoidal spaces by 1 week after implantation. Increased organization of smooth muscle and endothelial cells lining the sinusoidal walls was observed at 2 weeks and continued with time. The matrices were covered with the appropriate cell architecture 4 weeks after implantation (Atala, 1999). This study demonstrates that human cavernosal smooth muscle and endothelial cells seeded on three-dimensional acellular collagen matrices derived from donor corpora are able to form a well-vascularized corporal architecture in vivo. Protocol I.B: Acellular Collagen Matrix Preparation 1. Obtain corpus cavernosum from donor rabbits. 2. Take cross-sectional corporal fragments 0.5 cm in thickness.

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3. Treat tissue with distilled water in a magnetic stirring flask (moderate speed) for 24–48 h at 4°C. 4. Remove distilled water and treat with TritonX-100 (0.5%) and ammonium hydroxide (0.05%) in fresh 5. 6.

7. 8. 9. 10.

distilled water for 72 h in a stirring flask at 4°C. Wash with distilled water in a stirring flask for 24–48 h at 4°C. After this washing step, take a small piece of tissue for histology to confirm any cellular remnants. A small tissue mass is usually decellularized at this time. After confirmation of decellularization, repeat Step 5. Dense tissue may require another cycle of treatment. Repeat Steps 4 and 5, and perform another histological analysis. After washing the tissue with distilled water, rinse with 1  PBS overnight. Freeze-dry the tissue samples overnight. Pack and sterilize in ethylene oxide. Store until used. When ready to use, equilibrate the tissue in culture medium overnight prior to cell seeding.

Protocol I.C: Human Endothelial Cell Culture from Foreskin 1. Materials and media

(a) Medim A (for primary culture and first passage after UEA-I bead selection): 38.5 ml of endothelial basal medium 131 (Clonetics Corp., cat. No. CC 3121), 10 ml of 20% FBS, 0.5 ml (2 mM) L-glutamine, 0.5 ml of PFS (antibiotic–antimycotic) (Gibco, cat. No. 600-5240AG), 0.5 ml (0.5 mM) dibutyryl cyclic adenosine-3,5-cyclic monophosphate (AMP) (Sigma, cat. No. D-0627), and 50 μl (1 μg/ml) hydrocortisone (Sigma, cat.No. H-0888). (b) Medium B (for passage 2 and all following passages): endothelial basal medium 131, 1  GPS, 10% FBS, and 2 μg/ml basic fibroblast growth factor (25 μg/ml stock solution) (Scios Nova). (c) Gelatin coating (1% Difco Bacto Gelatin in PBS): dissolve gelatin in PBS; autoclave to sterilize, and filter to remove particles. 2. Processing foreskin (a) Prepare foreskin collecting medium: Four hundred and fifty milliliter of DMEM, 25 ml of PBS (5%), 20 ml of antibiotic–antimycotic (400 U/ml penicillin, 400 μg/ml streptomycin, 1 μg/ml fungizone), 5 ml of L-glutamine (2 mM), and 1 ml of gentamicin sulfate (100 μg/ml). (b) Place the collecting medium with the foreskin in a culture plate (100 mm) in a tissue culture hood. (c) Rinse two or three times with the collecting medium. (d) Add 30 ml of collecting medium to a new 50-ml Falcon tube. Add an additional 2 ml of antibiotic–antimycotic. (e) Separate the skin and subcutaneous tissue with a sterile scalpel blade and transfer the segments into the collecting medium in a 50-ml tube. (f) Agitate the segments in the collecting medium at room temperature for at least 4–5 h to kill bacteria and spores that reside on the skin. 3. Isolation of endothelial cells (a) Prepare digestion solution: 7.5 ml of 1:250 trypsin, 2.7 ml of 0.5 M EDTA, pH 8.0, and 40 ml of Hanks’ Balanced Salt Solution (HBSS). (b) Prepare 10  HBSS without Ca2+ and Mg2+: 40 g of NaCl, 2 g of KCl, 240 mg of Na2HPO4, 300 mg of KH2PO4, 1750 mg of NaHCO3, 5 g of glucose, and 100 mg of phenol red. (c) Prepare wash solution (HBBS with 1  Ca2+ and Mg2+): 50 ml of 10  HBSS, 92.7 mg of CaCl2 2H2O (1.26 mM final), 100 mg of MgSO4 7H2O (0.8 mM final), 25 ml of FBS (5% final), and 5 ml of PSF (antibiotic–antimycotic). (d) Coat a petri dish (100 mm) for each one or two foreskins with 8.0 ml of 1% gelatin–PBS. Remove excessive gelatin before plating.

Tissue Engineering of the Reproductive System

(e) (f) (g) (h) (i)

Autoclave a Teflon homogenizer (2.5 cm diameter) and gauze. Remove the collecting medium from the foreskin segments. Transfer the tissue segments into a sterile culture plate (100 mm). Cut the foreskin segments into 4-mm2 fragments with a sterile scalpel blade. Transfer the tissue fragments to a sterile 50 ml Falcon tube and add 6.0 ml of digestion solution for 1–2 foreskins. Agitate vigorously at 37°C for 10 min. (j) Allow the skin fragments to sediment by gravitational force and aspirate the digestion medium. Wash once with 20 ml of wash solution, swirl vigorously, and remove the wash solution. (k) Add 10 ml of fresh wash solution and squeeze the fragments with the homogenizer. (l) Filter through 8–10 layers of sterile gauze into a 50 ml Falcon tube (mesh filter). (m) Repeat Steps k and l, and collect the expelled cells into the same Falcon tube. (n) Centrifugate cells at 1000 rpm for 10 min at room temperature. (o) Aspirate the supernatant and plate the cells with 10 ml of endothelial cell basal medium (EBM) 131 (culture medium A) in a gelatin-coated culture dish (100 mm). Place the cells in an incubator overnight with 5% CO2. (p) Wash the cells vigorously three or four times with PBS. Feed the cells with 10 ml of culture medium A. (q) Change the medium every 2 days. The primary culture will be subconfluent after 7–8 days. They will be ready for the UEA-I isolation procedure at this point. 4. UEA-I selection of endothelial cells (a) Coating of Dynabeads with UEA: mix together 250 μl of Dynabeads (4  108 beads/ml) (Dynal, cat. No. 140.03), M-450, tosylactivated, 50 μg of unconjugated UEA-I (Vector, cat. No. L-1060), and 225 μl of 0.5 M boric acid, pH 9.5. The bead/lectin ratio should be 2.0  106 beads per microgram of lectin. The volume ratio of Dynabeads to boric acid with lectin should be 1:1. (b) Reconstitute the UEA-I with 1 ml of sterile PBS-0.1 mM CaCl2 to 2 mg/ml and store at 4°C (UEA-I is quite stable); 50 μg  25 μl. (c) Mix Dynabeads, lectin, and boric acid in a sterile 2.0-ml screw-cap tube and agitate on a rotor at room temperature overnight. (d) Pipette the bead–lectin mixture (in 10 ml of HBSS) into a 15-ml Falcon tube. Wash with 10 ml of HBSS (plus Ca2+/Mg2+, 1% BSA) on the rotator for 15 min at room temperature. (e) Place the tube in a magnetic particle concentrator (MPC) (MPC-1, Dynal, cat. No. 12001) and wait 1 min for the beads to be collected onto the magnet. Aspirate the supernatant with a Pasteur pipette. Take the tube out of the MCP, rinse three times at room temperature for 15 min, and once overnight at 4°C. (f) Resuspend the beads in 250 μl of HBSS (plus Ca2+/Mg2+, 5% FBS, 1  PBS) and store at 4°C in a sterile 2.0-ml screw-cap tube. The beads will be stable for several months. 5. Purification of endothelial cells from primary cultures (a) Trypsinize subconfluent cell cultures (7–8 days) with 1  trypsin–EDTA. (b) Centrifuge the trypsinize cells at 208 g (1000 rpm) for 10 min. (c) Resuspend the cell pellet from one 100-mm petri dish in 190 μl of HBSS buffer. Pipette up and down several times with a 200 μl pipetman to break up the cell clusters. Transfer the cell suspension into a sterile 2-ml screw-cap tube and add 5 μl UEA-I-coated Dynabeads. (d) Incubate cells and the beads for 3–5 min. Hold the tube in your hand and roll it between your palms gently to keep the beads in suspension. Endothelial cells and beads will form visible tiny clusters. (e) Transfer the cell–bead mixture to a 15-ml Falcon tube. Add 5 ml of HBSS buffer and pipette the cells several times up and down with the buffer. Place the Falcon tube into the MPC and collect the beads onto the magnet for about 1 min. Aspirate the wash solution with a Pasteur pipette while the tube is in the MCP. Take the tube out of the MCP. Repeat this wash four times with 5 ml HBSS wash buffer.

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(f) Resuspend the cells in 6 ml of EBM 131 growth medium A and place 3 ml onto each gelatin-coated 60-mm petri dish. This passage is designated as passage 1. Let the cells grow to confluence at 37°C and 5% CO2. Change the medium every 3–4 days or twice a week. (g) When endothelial cells become confluent, trypsinize, and split the cells 1:3–1:4. From now on (passage 2 and all the following passages), endothelial call are cultured in growth medium B. (h) The endothelial cells should be fed every 2–3 days and split every 5–7 days (at least once a week). Penile Prothesis for Reconstruction Early attempts at penile reconstruction involved use of rib cartilage as a stiffener but this method was discouraged due to the unsatisfactory functional and cosmetic results (Frumpkin, 1944; Goodwin and Scott, 1952; Small, 1976; Bretan, 1989). However, biocompatibility has been a problem in some patients (Kardar and Pettersson, 1995; Nukui et al., 1997). Of the tissue existing in the human body, cartilage would serve as an ideal prothesis for penile reconstruction, owing to its biomechanical properties (Yoo et al., 1998, 2000). Initial studies performed in our laboratory showed that chondrocytes suspended in biocompatible polymers form cartilage structures when implanted in vivo (Atala et al., 1993). A feasibility study of engineering natural penile prothesis made with cartilage was attempted. Chondrocytes, harvested from bovine articular cartilage tissue, were grown and seeded onto preformed cylindrical PGA polymer rods for implantation in vivo (Protocol II.A) (Yoo et al., 1998). Chondrocytes were seeded onto preformed cylindrical PGA polymer rods at a concentration of 50  106 chondrocytes/cm3. The cell-polymer were implanted in vivo. The retrieved implants formed milky white rod-shaped cartilaginous structures, maintaining their preimplantation size and shape. Biomechanical properties of the engineered cartilage rods, including compression, tension, and bending, showed that the cartilage tissues were readily elastic and could withstand high degrees of pressure. These results indicate that the engineered cartilage rods possessed the mechanical properties required to maintain penile rigidity. Histomorphological analyzes confirmed the presence of mature and well-formed cartilage in all the cell-seeded implants. Protocol II.A: Cartilage Tissue Harvest 1. Materials and medium

2. 3. 4. 5. 6. 7. 8. 9. 10.

(a) Medium: Ham’s F-12 nutrient medium, 10% FBS, vitamin C (5 μg/ml), and antibiotic (penicillin (100 U/ml)-streptomycin (100 μg/ml), amphotericin B (0.25 μg/ml)). (b) Digestion solution: 3% collagenase type II (Worthington Biochemical, Lakewood, NJ). Obtain cartilage tissue in a sterile manner. Use either a surgical blade or sharp tenotomy scissors to cut the tissue into small fragments (2–3 mm). Digest cartilage tissue fragments in 3% collagenase type II solution. Place the tube containing the digesting solution and cartilage fragments in a 37°C agitating incubator for 6–8 h. Be sure to check the tissue fragments periodically for overtreatment. When digestive step is complete, filter through a nylon mesh to remove undigested cartilage tissue. Wash the filtered cells twice with 1  PBS. Centrifuge the cells at 1200 rpm for 10 min. Use a hemocytometer to count the viable cells by means of trypan blue exclusion. Plate the cells in culture dishes at a desiderable density. Incubate the cells at 37°C in the presence of 5% CO2 and maintain the cells in a routine manner.

In a subsequent study using an autologous system, the feasibility of applying the engineered cartilage rods in situ was investigated (Yoo et al., 1999). Autologous cartilages harvested from rabbit ear were dissected into small

Tissue Engineering of the Reproductive System

fragments (2  2 mm2). The technique describe in Protocol II.A was used to harvest chondrocytes under sterile conditions (Atala et al., 1993, 1994). The chondrocytes were expanded until sufficient cell quantities were available. The cells were trypsinized, collected, washed, and counted for seeding. Chondrocytes were seeded onto performed poly(L-lactic acid) coated PGA polymer rods at a concentration of 50  106 chondrocytes/cm3. The chondrocyte–polymer scaffolds were implanted in the corporal spaces of rabbit. Bilateral intracorporal implantation of the cell–polymer scaffolds were performed. The implants were retrieved and analyzed grossly and histologically 1, 2, 3 and 6 months after surgery. Gross examination at retrieval showed the presence of well-formed milky white cartilage structures within the corpora at 1 month. There was no evidence of erosion or infection in any of the implant sites. Histological analysis demonstrated the presence of mature and well-formed chondrocytes in the retrieved impants. Autologous chondrocytes seeded on preformed biodegradable polymer structures are able to form cartilage structures within the rabbit corpus cavernosum. The technology appears to be useful for creation of autologous penile protheses. Testes In males, androgens, in particularly testosterone, are known to have many important physiological actions, including effects on muscle, bone, central nervous system, prostate, bone marrow, and sexual function. Testicular dysfunction and hypogonadal disorders evolve from different pathophysiological conditions such as Klinifelter’s syndrome, bilateral mump orchitis, toxic damage from alcohol or chemotherapy, and orchiectomy (Griffen and Willson, 1998). Patients with such conditions require lifelong androgen replacement therapy to maintain physiological levels of serum testosterone. Such therapy may increase muscle strength, stabilize bone density, improve osteoporosis, and restore secondary sexual characteristics, including libido and erectile function (Bhasin and Bremner, 1997). Anatomy The testes are two glandular organs, which secrete the semen; they are suspended in the scrotum by the spermatic cords. In mammals, the testes are located outside the body due to the fact that spermatogenesis in mammals is more efficient at a temperature some what less than the core body temperature (37°C for humans). When the temperature needs to be lowered, the cremasteric muscle relaxes and the testicles is lowered away form the warm body and are able to cool. Under a tough fibrous shell, the tunica albuginea, the testis contains very fine coiled tubes called seminiferous tubules. The tubes are lined with a layer of cells that form puberty into old age, produce sperm cells. The sperm travel form the seminiferous tubules to the rete testis, the efferent ducts, and then to the epididymis where newly created sperm cells mature (spermatogenesis). The sperm move into the vas deferens (also called the ductus deferens), which opens into the urethra. Upon any sufficient sexual arousal, the sperm cells move through the ejaculatory duct and into the prostatic urethra, where the prostate, through muscular contractions, ejaculates the sperm, mixed with other fluids, out through the penis. From the cellular point of view the human testis is a complex organ comprising germ cells and a variety of somatic cells such as Sertoli, Leydig, endothelial, fibroblast, macrophage, and peritubulat myoid cells. Testicles are component of both the reproductive system (being gonads) and the endocrine system (being endocrine glands). The testis has two functions: spermatogenesis, which occurs in the seminiferous tubules, and secretion of steroid hormones (androgens) by Leydig cells in the interstitial tissue. Transplantation of Testes The first authenticated record of gonadal transplantation is attributed to an eighteenth century Scottish anatomist and surgeon, John Hunter, who grafted chicken testes to the body cavity of birth male and female

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hosts. Full details of this work have not survived, and is difficult to evaluate its outcome. Berhold was the first to report on a successful testicular transplant, since he used autografts and avoided the risk of rejection. When he replaced the testes of capons in their own body cavity, he found that the growth of comb. and plumage, and courting behavior, all of which are androgen dependent, were maintained (Berhold, 1849). A century later, interest in testicular transplant increased as a result of the misapprehension that somatic aging is caused by withdrawal of sex hormones. Lydston had published a series of testicular transplantation experiments performed in his patients (Lydston, 1916). He and others following him believed that transplanted of midlife and later. sex glands produced a hormone that was a “cell stimulant, nutrient and regenerator” capable of prolonging life and restoring waning sexual functions, arteriosclerosis and other infirmities Voronoff in 1923 was the first to use chimpanzee and baboon organs for treating patients (Brinster and Zimmermann, 1994). This approach was taken by other surgeons, but none of them used microsurgery to join blood vessels of the graft to the host’s circulation, resulting in ischemic necrosis preceded by organ rejection. Later, successful testicular transplantation could be achieved when ischemia time was reduced to less than an hour by using vascular anastomosis in dogs (Attaran et al., 1966; Lee et al., 1971; Gittes et al., 1972). The first convincing human testicular transplant was published by Silber (1978), who grafted a patient with a testis from the patient’s genetically identical twin brother (Silber, 1978). However, with time the stringent requirements for success have precluded a surge in demand for this operation. Moreover, carefully conducted grafting trials failed to confirm former claim; the new synthetic sex steroids were shown not to affect the life span of experimental animals (Parkes, 1966). Nevertheless, testicular transplantation may still be regarded as having clinical potential for example, who in carriers of genetic disease, who can receive normal germ cells from donors. Transplantation of Testicular Tissue The problems arising from the size of the testis and its fibrous capsule led some transplanters to use sliced or minced organs. Kearns (1941), who reimplanted testicular tissue subcutaneously in a victim of accidental castration, reported the most plausible case (Kearns, 1941). According to this report, testosterone was being produced by the autograft, but without the normal architecture of the seminiferous epithelium, it hard to understand how germ cell transfer could have restored spermatogenesis. Furthermore, injecting spermatogonial stem cells from donor testes into atrophic tubules is daunting the testes must produce millions of spermatozoa per day to be fertile. Therefore, efforts to develop tissue grafting for the purpose of improving testosterone levels in hypogonadal men are more likely to succeed that are attempts at restoring fertility. The former goal appears to be simple, requiring the transfer of interstitial cells (Leydig cells), which are readily isolated from the donor testes by means of collagenase. Interstitial cells grafted in castrated rodents resulted in partial restoration of body weight, and testosterone levels above those controls (Fox et al., 1973; Boyle et al., 1975; Tai et al., 1992). A number of vehicles and several implantation sites for interstitial cells have been tried, but none fully replaced testicular androgen production. Testosterone Delivery Systems The main goal of androgen replacement therapy is to maintain physiological levels of serum testosterone and also its metabolites, dihydrotestosterone, and estradiol. Hypogonadal states secondary to hypothalamic–pituitary disorders, gonadal abnormalities, and defects in androgen action or secretion may benefit from androgen replacement. Androgen replacement modalities include oral administration of testosterone tablets, or capsules (Franchimont et al., 1978; Snyder and Lawrence, 1980; Sokol et al., 1982; Canteril, 1984; Fujioka et al., 1986; Stuenkel et al., 1991; Ferrini and Barret-Conner, 1998; Wilson et al., 1998; McClella and Goa, 1998; Bennet, 1998). When taken orally, testosterone preparations are largely rendered metabolically inactive during the “first

Tissue Engineering of the Reproductive System

pass” through the liver. This metabolic inactivation requires large oral doses of testosterone (200 ng/day) to reach normal serum levels. These large doses of testosterone may be toxic to the liver and may lead to hepatitis, hepatoma, or hepatocarcinoma (Snyder and Lawrence, 1980; Gooren, 1994; Bagatelle and Bremner, 1996). Parenteral depot preparation include testosterone enanthante (delatestryl) and testosterone cypionate (depot testosterone cypionate). These preparations are based in 17B-hydroxyl esters, which are given intramuscularly (IM), with slow-release, oil-based injection vehicles every 10–21 days. Testosterone levels with these preparations rise to supernormal levels for 1 or 2 days, after which they gradually fall within the normal range for 10–12 days, reaching baseline at approximately 21 days. This fluctuation in testosterone levels may produce significant swings in mood, libido, and sexual function (Sokol et al., 1982; Bhasin and Bremner, 1997). Transdermal testosterone therapy includes both scrotal and non-scrotal patches. Testoderm and androderm are multilayered skin patches that deliver measured doses of testosterone across the scrotal skin acting thanks to the 5α-reductase activity present within this site. When used in non-scrotal skin, the patch has to be applied twice daily reducing frequency of administration. However, despite these advantages, the transdermal systems have been associated with adverse effects, such as transient erythema, pruritis, induration, burning, rash, and skin necrosis (Hogan and Maibach, 1990; McClellan and Goa, 1998; Bennet, 1998). Long-term exogenous testosterone therapy has been associated with several complications, such as fluid and nitrogen retention, erythropoiesis, hypertension, and bone-density changes. In addition, fluctuating serum testosterone levels may occur, and frequent treatments may be required. Due to these problems, alternate treatment modalities, involving more physiological and longer-acting systems for androgen delivery, have been pursued. Cell Encapsulation for Testosterone Therapy Cell transplantation has long been proposed as a treatment for several diseases involving hormone or protein deficiencies. Cell rejection by the host immune system, however, has limited the use of this strategy. Encapsulation of living cells in a protective, biocompatible, and semipermeable polymeric membrane has been proven to be an effective method of immunoprotection of the desired cells, regardless of the type of recipient (allograft, xenograft) (Chang, 1993). A majority of the implantation work using microencapsulated cells as delivery vehicles employs two polymers: sodium alginate and poly(L-lysine) (PLL) (Lim and Sun, 1980). Alginate microcapsule have been used for various applications (Chang, 1998; Joki et al., 2001) particularly for the encapsulation of the pancreatic islet cells/or insulin delivery (Lim and Sun, 1980; Wang et al., 1997) and recombinant cells have served for the delivery of therapeutic gene products (Tai and Sun, 1993). The Leydig cells of the testes are the major source of testosterone in men (95%). Implantation of heterologous Leydig cells has been proposed as a method for chronic testosterone replacement. However, these approaches were limited by tissue and cell failure to produce long-term testosterone and dissemination of the implanted cells. Therefore, encapsulation of Leydig cells might be useful for testosterone replacement therapy. Such a system might be able to stimulate the normal diurnal pattern of testosterone release by the testes, therapy avoiding side effects such as those associated with chemically modified testosterone administration. Leydig cell transplantation may be also beneficial not only for testosterone replacement but also for the secretion of other associated hormones and growth factors such as melanocytes, β-andorpilin, prostaglandins, insulin-like growth factor 1 (IGF-1), and interleukins (Verhoeven, 1992). Studies in our laboratory have been focused on the encapsulation and implantation of isolated Leydig cells for long-term testosterone delivery. Leydig cells were isolated from male Sprague–Dawley rats, 56–70 days old, by means of collagenase and Percoll gradient separation. The isolated Leydig cells were encapsulated within microspheres composed of calcium-alginate, coated with the positively charged polyelectrolyte PLL, and recoated with alginate. Based on the molecular weight of testosterone (300 Da), PLL having a molecular weight of 21 kDa and 1.2% sodium alginate with a high glucuronic acid content (65%) were chosen. PLL with

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a molecular weight ranging between 16 and 22 kDa produced a semipermeable membrane with a molecular weight cutoff of 70 kDa (i.e. preventing the diffusion of cells and metabolites). Methyl(this) tetrazole assays performed daily on the microencapsulated cells showed that the cells remained viable during the experiments. Testosterone secretion from cultured encapsulated Leydig cells in response to human chorionic gonadotropin (hCG) was the highest 24 h after hCG stimulation (0.06 IU/ml) (Figure 92.2). There was no significant difference in testosterone secretion when cells were cultured at either 32 or 37°C (Figure 92.2). The encapsulated and nonencapsulated Leydig cells were found to be resistant to temperature changes. This finding broadens the list of possible in vivo sites for Leydig cell transplantation. The intraperitoneal cavity was the first site chosen for cell implantation because of its generous vascular and nutritional capabilities. In vivo studies performed in castrated rats showed that the total testosterone levels measured in the serum of castrated rats that were injected intraperitonally with 5  106 encapsulated Leydig cells were between 0.23 and 0.51 ng/ml for more than 3 weeks (Figure 92.3). These animals did not receive any exogenous hCG stimulation. Similar testosterone levels (0.24–0.48 ng/ml) were obtained when encapsulated Leydig cells were injected subcutaneously (Figure 92.3b). However, testosterone was detected for a longer time period in the subcutaneous group (43 days) than in the intraperitoneal group (35 days). These testosterone levels were lower than the ones detected in the control rats (1.7 ng/ml), which were not castrated. However, only 5  106 microencapsulated cells were implanted in each animal, representing only 10% of the normal adult rat Leydig cell population (Machluf et al., 2003). Methods for Encapsulation Microencapsulation is currently the optimal immuno-isolation technique. Different approaches and polymers are being used for encapsulating cells and tissue for therapeutic applications. The technique of microencapsulation used by our laboratory utilizes two polymers: highly purified calcium-alginate (Pronova, Norway) and low molecular weight (23.6 kDa, Sigma) PLL. This procedure is described as follows. Protocol I: Cell Encapsulation 1. Isolated cells are suspended in sodium alginate (1.2%) (60% glucuronic acid content) in 0.9% saline for 2. 3. 4. 5. 6. 7.

5 min. The cell–alginate suspension is extruded through a 22-gauge airjet-needle into a calcium chloride–4-(2hydroxyethyl)-1-piperazineethanesulfonic acid (CaCl2–HEPES) solution (1.5%). The beads are stirred for 20 min in the CaCl2–HEPES solution. Gelled droplets are transferred to ecno-colums (Bio-Rad) and decanted. The columns are filled with 15 ml of PLL solution in 0.9% saline, sealed, and rotated gently for 12 min. The PLL solution is decanted from the columns and washed three times with HEPES solution. A 0.125% alginate solution is added, and the mixture is rotated for 10 min. Then the alginate solution is decanted and the supernatant is washed three times with HEPES prior to culturing.

FEMALE Vagina A variety of pathological and congenital disorders affect the vagina and require extensive surgical intervention (Machluf et al., 2003). The choice of operation and outcome depend critically on correct identification of the underlying disorder. Three basic categories of anomalies must be distinguished, namely, vaginal agenesis and its variants, ambiguous genitalia, and imperforate anus and urogenital sinus variants. Vaginal reconstruction is an

Tissue Engineering of the Reproductive System

uncommon and a challenging procedure that varies considerably by specialty, with plastic surgeons and gynecologists generally recommending skin graft/dilation procedures and pediatric urologists recommending bowel vaginoplasty (Rajimwale et al., 2004). Various procedures have been used in the past for vaginal reconstruction and different tissue sources have been employed for reconstructive surgery. Traditionally, the reconstructions have been performed with non-urologic tissues or synthetic prostheses. The non-urologic tissues include gastrointestinal segments (Leong and Ong, 1972; Hendren and Atala, 1994), skin (Draper and Stark, 1956), peritoneum (Hutschenreiter et al., 1978), fascia (Neuhof, 1917), omentum (Goldstein et al., 1967), pericardium (Kambic et al., 1992), and dura (Kelami, 1971). The majority of surgical options require the use of non-genital tissues for vaginal replacement too. However, the use of non-vaginal tissue for surgical reconstruction is not ideal in terms of normal vaginal function (Machluf et al., 2003). Tissue engineering may offer a solution for challenging cases when shortage of local tissue exists. While tissue engineering has been applied to many tissue–organ reconstructions, there is a paucity of information regarding the engineering of female reproductive and genital tissues. This chapter summarizes the known and recently developed tissue engineering applications for total vaginal reconstruction. Anatomy The vagina (Kelami, 1971) is a muscular, highly expandable, tubular cavity that connects the vulva at the outside to the cervix of the uterus on the inside. The vagina consists of an internal mucous lining and a muscular coat separated by a layer of erectile tissue. It does not have any glands and is kept moist by the lubrication provided by the cervical and uterine glands. The mucous membrane (tunica mucosa) is continuous above with that lining the uterus. Its inner surface presents two longitudinal ridges, one on its anterior and one on its posterior wall. These ridges are called the columns of the vagina and from them numerous transverse ridges or rugae extend outward on either side. These rugæ are divided by furrows of variable depth, giving to the mucous membrane the appearance of being studded over with conical projections or papillae; they are most numerous near the orifice of the vagina, especially before parturition. The epithelium covering the mucous membrane is of the stratified squamous variety. The submucous tissue is very loose, and contains numerous large veins, which by their anastomoses form a plexus, together with smooth muscular fibers derived from the muscular coat; it is regarded by Gussenbauer as an erectile tissue. It contains a number of mucous crypts, but no true glands. The muscular coat (tunica muscularis) consists of two layers: an external longitudinal, which is by far the stronger, and an internal circular layer. The longitudinal fibers are continuous with the superficial muscular fibers of the uterus. The strongest fasciculi are those attached to the rectovesical fascia on either side. The two layers are not distinctly separable from each other, but are connected by oblique decussating fasciculi, which pass from the one layer to the other. In addition to this, the vagina at its lower end is surrounded by a band of striped muscular fibers, the Bulbocavernosus. External to the muscular coat is a layer of connective tissue, containing a large plexus of blood vessels. The erectile tissue consists of a layer of loose connective tissue, situated between the mucous membrane and the muscular coat; imbedded in it is a plexus of large veins, and numerous bundles of unstriped muscular fibers, derived from the circular muscular layer. The arrangement of the veins is similar to that found in other erectile tissues. Vaginal Tissue Engineering Clinically related studies have already demonstrated encouraging results with regard to the applicability of tissue engineering in genitourinary reconstruction (Atala, 1999). Atala et al. have also demonstrated that in vitro expansion of vaginal epithelial and smooth muscle cells followed by seeding them onto synthetic matrices and

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placing them in vivo can reconstitute to de novo tissues (Mikos, et al., 1994). In this study expanded cells of muscle and epithelial cells seeded onto PGA scaffolds at a concentration of 10  106 and 20  106 cells/cm3 were co-cultured at 37°C with 5% CO2 for 24 to 48 h and implanted subcutaneously into athymic mice. The cells were able to survive and replicate in vivo for prolonged periods and could self-organize toward seemingly normal structural orientation. By the sixthth week of implantation the constructs have shown to organize into a distinguishable layer of both the vaginal epithelial and smooth muscle cell types. Penetrating native vasculature was also observed. Further analysis of the tissue engineered vaginal constructs has shown to produce contractile forces similar to those seen with native vaginal tissue when simulated with a series of electrical impulses. Protocol: Methods of Cell Culture 1. Materials and medium

(a) Tissue source: vaginal tissue from New Zealand White rabbits (b) Medium: (i) Smooth muscle cells DMEM supplemented with 10% FBS (ii) Epithelial cells – keratinocytes Keratinocyte serum-free medium (K-SFM) supplemented with bovine pituitary extract and epidermal growth factor. 2. Tissue harvest and cell culture (a) Obtain vaginal tissue. (b) Wash the specimen several times with PBS containing EDTA. 3. Smooth muscle (a) Mechanically microdissect the muscle from the seromuscular layer with sterile instruments under loop magnification. (b) Individually place small portions of the dissected samples onto culture dishes, and allow them to dry and adhere to the surface. (c) Incubate the pieces with medium at 37°C in air and 5% CO2 undisturbed until a sufficient colony of progenitor cells grows from the tissue islets. (d) Remove the tissue explants by gentle suction when sufficient amount of cells are established. 4. Epithelial cells (a) Digest the vaginal specimen with collagenase type IV by immersing them into the enzymatic solution and shake vigorously for 30 min at 37°C. (b) Centrifuge the cell-fluid suspension at low revolutions for 5 min. (c) Resuspend the supernatant in K-SFM and distribute onto culture dishes. C. Cell expansion (a) Remove the culture medium and wash the cells with PBS–EDTA (b) Incubate the cells with a 0.05% trypsin–EDTA solution, (0.5 g trypsin and 0.2 g EDTA per 1.0 L of stock solution) and monitor under the microscope until cell separation is observed. (c) With a pipette gently transfer the cell–trypsin solution in to a 50-ml Falcon tube with serum containing medium to inactivate the trypsin. (d) Centrifuge the cells at 1500-rpm for 5 min. (e) Resuspend the cell pellet into a predetermined volume of fresh medium and partition equally among several more culture dishes for expansion. D. Cell maintenance (a) Replace the medium with fresh warm (37°C) medium every 24–48 h.

Tissue Engineering of the Reproductive System

Uterus Tissue engineering is a relatively new and rapidly expanding field of biological research. It is also a clinically applicable discipline that aims to provide a repository of alternative tissue sources when reconstructive surgery is necessary(Skalak and Fox, 1998). Congenital malformations of the uterus may have profound implications clinically. Patients with cloacal exstrophy or intersex conditions may not have sufficient uterine tissue for future reproduction. With developing aspects of tissue engineering it may be possible to solve this kind of problems in the future. (Figure 67.2) Anatomy The uterus (Gray, 1918) is a hollow, thick-walled, muscular organ situated deeply in the pelvic cavity between the bladder and rectum. Into its upper part the uterine tubes open, one on either side, while below, its cavity communicates with that of the vagina. The uterus measures about 7.5 cm in length, 5 cm in breadth, at its upper part, and nearly 2.5 cm. in thickness; it weighs from 30 to 40 g. It is divisible into two portions. On the surface, about midway between the apex and base, is a slight constriction, known as the isthmus, and corresponding to this in the interior is a narrowing of the uterine cavity, the internal orifice of the uterus. The portion above the isthmus is termed the body, and that below, the cervix. The part of the body, which lies above a plane passing through the points of entrance of the uterine tubes, is known as the fundus. The cavity of the uterus is small in comparison with the size of the organ. It is a mere slit, flattened anteroposteriorly. It is triangular in shape, the base being formed by the internal surface of the fundus between the orifices of the uterine tubes, the apex by the internal orifice of the uterus through which the cavity of the body communicates with the canal of the cervix. The canal of the cervix (canalis cervicis uteri) is somewhat fusiform, flattened from before backward, and broader at the middle than at either extremity. It communicates above through the internal orifice with the cavity of the body, and below through the external orifice with the vaginal cavity. The wall of the canal presents an anterior and a posterior longitudinal ridge, from each of which proceed a number of small oblique columns, the palmate folds, giving the appearance of branches from the stem of a tree; to this arrangement the name

Uterine tube Uterine tube

l wal

Anal canal

rine

Rectum

Round ligament of uterus Bladder

Ute

Cavity of uterus Sigmoid colon

Cavity of body

Internal orifice

Symphysis pubis Urethra Vagina

External orifice Vagina

Figure 67.2 Uterus.

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arbor vitae uterina is applied. The folds on the two walls are not exactly opposed, but fit between one another so as to close the cervical canal. The uterus is composed of three coats: an external or serous, a middle or muscular, and an internal or mucous. The serous coat (tunica serosa) is derived from the peritoneum; it invests the fundus and the whole of the intestinal surface of the uterus; but covers the vesical surface only as far as the junction of the body and cervix. The muscular coat (tunica muscularis) forms the chief bulk of the substance of the uterus. It consists of bundles of unstriped muscular fibers, disposed in layers, intermixed with areolar tissue, bloodvessels, lymphatic vessels, and nerves. The layers are three in number: external, middle, and internal. The external and middle layers constitute the muscular coat proper, while the inner layer is a greatly hypertrophied muscularis mucosæ. The arteries of the uterus are the uterine, from the hypogastric; and the ovarian, from the abdominal aorta. They are remarkable for their tortuous course in the substance of the organ, and for their frequent anastomoses. The veins are of large size, and correspond with the arteries. They end in the uterine plexuses. In the impregnated uterus the arteries carry the blood to, and the veins convey it away from, the intervillous space of the placenta. Uterine Tissue Reconstruction The first report of tissue engineering of human uterine smooth muscle cells was reported in 2003(Atala, 2004). In this study, primary cell lines were initiated from human myomerium obtained at the time of term cesarean delivery. After several passages the cells were seeded onto a polyglactin-910 (Vicryl) mesh and maintained in culture. This system provides a three-dimensional myocyte culture where cells are attached to each other instead of to a culture dish. The resulting “tissue” contains cells in an environment that approximates whole tissue, but grown under controlled conditions. Similar experiments have been reported for urinary bladder(Vozzi, et al., 2002) and vascular smooth muscle cells(Dessanti, et al., 1992). In addition to this, double-mesh experiments were performed to build thicker sections of tissue. The mechanical strength of the bridging myocytes was determined by hanging the two-mesh complexes in the muscle bath, with one mesh fixed and the other attached to the force transducer. The meshes were subjected to increasing force until separation. The constructs were able to maintain a maximum force of 5 g/cm2. The bridging myocytes were also tested for contractile activity by hanging a two-mesh complex in the muscle bath and applying a 2  g of force. Addition of oxytocin (100 nM) to the bath produced small, irregular contractions, which remained stable for 25 min. Addition of 140 mM KCl to a final concentration of about 50 mM resulted in loss of contractile behavior. Although no repetitive pattern reminiscent of human labor was observed, these observations represent the first example of a group of cultured human uterine myocytes exhibiting coordinated contraction. Protocol: Uterine Cell Culture 1. Materials and medium

(a) Tissue source: human myometrium. (b) Medium: Dulbecco’s modified essential medium supplemented with 10% FBS 2. Tissue harvest (a) Obtain human myocytes from the upper margin of the uterine incision (b) Mince the collected tissue (c) Perform double digestion at 37°C for 45 min. each (i) Prepare and perform the first digestion containing collagenase-dispase (1.5 mg/ml), trypsin inhibitor (1 mg/ml), and bovine serum albumin (2 mg/ml) in calcium-free Hanks’ solution. (ii) Prepare and perform the second digestion containing collagenase (1 mg/ml), trypsin inhibitor (0.3 mg/ml), and bovine serum albumin (2 mg/ml) in the same Hanks’ solution.

Tissue Engineering of the Reproductive System

(d) Centrifuge the cell-digestion solution mix at low revolutions for 5 min wash with PBS, and resuspend in culture medium (e) Culture the cells onto culture flasks in an atmosphere of 95% O2 and 5% CO2 at 37°C. 3. Cell expansion (a) Follow the protocol for vaginal cell culture expansion. 4. Cell maintenance (a) Replace the medium with fresh warm (37°C) medium every 2–3 days. In the subsequent study the possibility of engineering functional uterine tissue using autologous cells was investigated (Kim et al., 1999). Autologous rabbit uterine smooth muscle and epithelial cells were harvested, then grown and expanded in culture. These cells were seeded onto preconfigured, uterine-shaped, biodegradable polymer scaffolds, which were then used for subtotal uterine tissue replacement in the corresponding autologous animals. Upon retrieval 6 months after implantation, histologic, immunocytochemical, and Western blot analyses confirmed the presence of normal uterine tissue components. Biomechanical analyses and organ bath studies showed that the functional characteristics of these tissues were similar to those of normal uterine tissue. Breeding studies using these engineered uteri are currently being performed. Ovary Anatomy The ovaries (Gray, 1918) are homologous with the testes in the male. They are two nodular bodies, situated one on either side of the uterus in relation to the lateral wall of the pelvis, and attached to the back of the broad ligament of the uterus. The surface of the ovary is covered by a layer of columnar cells, which constitutes the germinal epithelium of Waldeyer. This epithelium gives to the ovary a dull gray color as compared with the shining smoothness of the peritoneum; and the transition between the squamous epithelium of the peritoneum and the columnar cells, which cover the ovary, is usually marked by a line around the anterior border of the ovary. The ovary consists of a number of vesicular ovarian follicles imbedded in the meshes of a stroma or frame-work. The development and maturation of the follicles and ova continue uninterruptedly from puberty to the end of the fruitful period of woman’s life, while their formation commences before birth. Before puberty the ovaries are small and the follicles contained in them are disposed in a comparatively thick layer in the cortical substance; here they present the appearance of a large number of minute closed vesicles, constituting the early condition of the follicles; many, however, never attain full development, but shrink and disappear. At puberty the ovaries enlarge and become more vascular, the follicles are developed in greater abundance, and their ova are capable of fecundation (Figure 67.3). The follicles, after attaining a certain stage of development, gradually approach the surface of the ovary and burst; the ovum and fluid contents of the follicle are liberated on the exterior of the ovary, and carried into the uterine tube by currents set up by the movements of the cilia covering the mucous membrane of the fimbriæ. After the discharge of the ovum the lining of the follicle is thrown into folds, and vascular processes grow inward from the surrounding tissue. In this way the space is filled up and the corpus luteum formed. The arteries of the ovaries, each anastomoses freely in the mesosalpinx, which traverse the mesovarium and enter the hilum of the ovary. In Vitro Culture of Ovarian Follicles The fundamental role of the ovary is to produce oocytes capable of fertilization and subsequent development into viable offspring (Wang et al., 2003). Number of pathological conditions such as polycystic ovarian syndrome (PCOS), premature ovarian failure, or definitive sterility (postoncotherapy) may affect ovarian function

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Fibro-vascular coat Membrane granulos

Zona striata Germinal vesicle

Discus proligerus

Figure 67.3 The Ovary. and severely compromise their reproductive potential. Recently, new in vitro culture methods involving tissue engineered matrices have been developed to study the maturation of ovarian follicles (Pangas et al., 2003). Unlike the two-dimensional culture systems supporting the production of immature mouse follicles or granulose cell-oocyte complexes where the granulose cells attach to the culture substrate, and migrate away from the oocyte (Spears et al., 1994; Cortvrindt et al., 1996; Rowley et al., 1999; Smitz and Cortvrind, 2002; O’brien et al., 2003; Kreeger et al., 2006). This research study has developed a three-dimensional culture system for mouse granulose-oocyte complexes, which maintains cell–cell connections and provides an environment that supports follicle development (Wang et al., 2003). Protocol: Follicle Isolation And Culture 1. Materials and medium

(a) Tissue source: C57BL/6  CBA F1mouse (b) Medium: αMEM supplemented with 3 mg/ml BSA, 5 μg/ml insulin, 5 μg/ml transferring, and 5 μg/ml selenium. 2. Tissue harvest and culture (a) Obtain two-layered (100–130 μm) and multilayered secondary follicles (150–180 μm) using insulin gauge needles in L-15 media, while maintaining them at 37°C and pH 7. (b) Encapsulate the follicles into alginate or alginate-ECM matrices (i) Suspend droplets (2–3 μL) of alginate or alginate-ECM solution on a polypropylene mesh (0.1 mm opening) (ii) Pipette a single follicle into each droplet in a minimal amount of media. (c) After all the droplets are filled, immerse the mesh in sterile 50 mM CaCl2 for 2 min. (d) Rinse the mesh in L-15 media. (e) Plate individual beads in 96 well plates in 100 μl of culture media. (f) Culture the follicles at 37°C in 5% CO2 for 8 days. (g) Change half of the media volume every 2 days. For the preservation of fertility for women or young girls, cryopreservation of ovaries has been proposed; however, there is a critical limitation in obtaining a sufficient supply of meiotically competent oocytes (Cortvrindt et al., 1996). By merging principles from tissue engineering with those from follicle biology, this

Tissue Engineering of the Reproductive System

research team has developed synthetic matrices that promote follicle maturation to produce meiotically competent oocytes, which may provide mechanisms to preserve fertility. It was concluded that, this alginate culture system serves as a tool for fundamental studies that correlate the composition of the cellular microenvironment to the properties of the developing tissue, which may ultimately provide design principles for scaffold-based approaches to tissue engineering. In the subsequent study, similarly, in vitro cultures of immature ovarian follicles were used to examine the factors that regulate the follicle development (Wang et al., 2003). In this system, individual granulose celloocyte complexes were incorporated into a three-dimensional culture system based on an alginate hydrogel. Briefly, ovaries from 12 day old mice were dissected and dissociated in μ-MEM containing 0.3% bovine serum albumin, 0.1% type I collagenase, and 0.02% deoxyribonuclease I at 37°C and 5% O2, 5% CO2, 90% N2 for approximately 1 hour. GOCs were collected and washed 3 times in α-MEM, manually counted and encapsulated into alginate. The growth medium consisted of α-MEM supplemented with 0.3% BSA, bovine pancreatic insulin (5 μg/ml), human transferring (5 μg/ml), sodium selenite (ITS, 5 ng/ml), penicillin (5 U/ml), and streptomycin (5 μg/ml). Alginate beads were then cultured, and fed every other day by replacement of one-third the volume of growth medium. After 10 days the morphology of GOCs was assessed by TEM. They showed no signs of degeneration, had cortical granules around the periphery, and contained an intact zona pellucida. No evidence of cellular apoptosis was detected in intact GOCs containing an oocyte. At this time point the GOCs were also analyzed for their ability to undergo in vitro maturation. On average, 40% of oocytes retrieved from in vitro growth in alginate beads underwent germinal vesicle breakdown and proceeded to meiosis II. In these studies, alginate exhibits minimal cellular interactions with mammalian cells, and thus likely provides only mechanical support. However, cell adhesion ligands can be incorporated onto the alginate backbone (Huet et al., 2001), which are known to influence granulose cell morphology, differentiation, and signaling (Huet et al., 2001). In conclusion, this three-dimensional culture system allows immature GOCs to be maintained in culture.

REFERENCES Atala, A. (2004). Tissue engineering and regenerative medicine: concepts for clinical application. Rejuvenation Res. 7(1): 15–31. Anwar, H., Dave, B. and Seebode, J.J. (1984). Replacement of partially resected canine urethra by polytetrafluoroethylene. Urology 24: 583. Atala, A. (1997). Tissue engineering in the genitourinary system. In: Atala, A. and Mooney, D. (eds.), Tissue Engineering. p. 149. Atala, A. (1999). Engineering tissues and organs. Curr. Opin. Urol. 9(6): 517–526. Atala, A. (1999). Future perspectives in reconstructive surgery using tissue engineering. Urol. Clin. N. Am. 26: 157. Atala, A. (1999). Tissue engineering applications for erectile dysfunction. Int. J. Impot. Res. 11(Suppl 1): S41. Atala, A., Cima, L.G., Kim, W., et al. (1993). Injectable alginate seeded with chondrocytes as a potential treatment for vesicoureteral reflux. J. Urol. 150: 745. Atala, A., Kim, W., Paige, K.T., et al. (1994). Endoscopic treatment of vesicoureteral reflux with a chondrocyte-alginate suspension. J. Urol. 152: 641. Atala, A., Vacanti, J.P., Peters, C.A., et al. (1992).: Formation of urothelial structures in vivo from dissociated cells attached to biodegradable polymer scaffolds in vitro. J. Urol. 148: 658. Attaran, S.E., Hodges, C.V., Crary Jr., L.S., et al. (1966). Homotransplants of the testis. J. Urol. 95: 387. Bagatelle, C. and Bremner, W. (1996). Drug therapy: androgen in men, use and abuses. N. Eng. J. Med. 334: 707. Bazeed, M.A., Thuroff, J.W., Schmidt, R.A., et al. (1983). New treatment for urethral strictures. Urology 21: 53. Bennett, N.J. (1998). A burn-like lesion caused by a testosterone transdermal system. Burns 24: 478. Berhold, A. (1849). Transplantation der hoden. Arch. Anat. Physiol. Wiss. Med. 16: 42.

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Bhasin, S. and Bremner, W.J. (1997). Clinical review 85: Emerging issues in androgen replacement therapy. J. Clin. Endocrinol. Metab. 82: 3. Boyle, P.F., Fox, M. and Slater, D. (1975). Transplantation of interstitial cells of the testis: effect of implant site, graft mass and ischaemia. Br. J. Urol. 47: 891. Bretan Jr., P.N. (1989). History of the prosthetic treatment of impotence. Urol. Clin. N. Am. 16: 1. Brinster, R.L. and Zimmermann, J.W. (1994). Spermatogenesis following male germ-cell transplantation. Proc. Natl. Acad. Sci. USA 91: 11298. Canteril, J., al, e. (1984). Which testosterone therapy? Clin. Endocrinol. 21: 97. Chang, T.M. (1993). Bioencapsulation in biotechnology. Biomater. Artif. Cell Immobil. Biotechnol. 21: 291. Chang, T.M. (1998). Pharmaceutical and therapeutic applications of artificial cells including microencapsulation. Eur. J. Pharm. Biopharm. 45: 3. Chang, T.S. and Hwang, W.Y. (1984). Forearm flap in one-stage reconstruction of the penis. Plast. Reconstr. Surg. 74: 251. Chen, F., Yoo, J.J. and Atala, A. (1999). Acellular collagen matrix as a possible off the shelf biomaterial for urethral repair. Urology 54: 407. Cilento, B.G., Freeman, M.R., Schneck, F.X., et al. (1994). Phenotypic and cytogenetic characterization of human bladder urothelia expanded in vitro. J. Urol. 152: 665. Cilento, B., Retik, A. and Atala, A. (1995). Uretheral reconstruction using a polymer mesh. J. Urol. 153: 371A. Cortvrindt, R., Smitz, J. and Van Steirteghem, A.C. (1996). In-vitro maturation , fertilization and embryo development of immature oocytes from early preantral follicles from prepuberal mice in a simplified culture system. Hum. Reprod. 11(12): 2656–2666. De Filippo, R.E., Yoo, J.J. and Atala, A. (2002a). Urethral replacement using cell seeded tubularized collagen matrices. J. Urol. 168: 1789. De Filippo, R., Pohl, H.G., Yoo, J., et al. (2002b). Total penile urethral replacement with autologous cell-seeded collagen matrices. J. Urol. 168: 1789 (abstract). De Filippo, R.E., Yoo, J.J. and Atala, A. (2003). Engineering of Vaginal Tissue in Vivo. Tissue Eng. 9: 301–306. Dessanti, A., Rigamonti, W., Merulla, V., et al. (1992). Autologous buccal mucosa graft for hypospadias repair: an initial report. J. Urol. 147: 1081. Draper, J.W. and Stark, R.B. (1956) End results in the replacement of mucous membrane of theurinary bladder with thick-split grafts of skin. Surgery 39(3): 434–440. Ehrlich, R., Reda, E. and Kyle, M. (1989). Complications of bladder mucosal graft. J. Urol. 142: 626. El kassaby, A., Retik, A., Yoo, J., et al. (2003). Urethral stricture repair with an “off the shelf” collagen matrix. J. Urol. 169: 170. Falke, G., Yoo, J.J., Kwon, T.G., et al. (2003). Formation of corporal tissue architecture in vivo using human cavernosal muscle and endothelial cells seeded on collagen matrices. Tissue Eng. 9: 871. Ferrini, R.L. and Barrett-Connor, E. (1998). Sex hormones and age: a cross-sectional study of testosterone and estradiol and their bioavailable fractions in community-dwelling men. Am. J. Epidemiol. 147: 750. Fox, M., Boyle, P.F. and Hammonds, J.C. (1973). Transplantation of interstitial cells of the testis. Br. J. Urol. 45: 696. Franchimont, P., Kicovic, P.M., Mattei, A., et al. (1978). Effects of oral testosterone undecanoate in hypogonadal male patients. Clin. Endocrinol. (Oxf.) 9: 313. Frumpkin, A. (1944). Reconstruction of male genitalia. Am. Rev. Sov. Med. 2: 14. Fujioka, M., Shinohara, Y., Baba, S., et al. (1986). Pharmacokinetic properties of testosterone propionate in normal men. J. Clin. Endocrinol. Metab. 63: 1361. Gilbert, D.A., Williams, M.W., Horton, C.E., et al. (1988). Phallic reinnervation via the pudendal nerve. J. Urol. 140: 295. Gittes, R.F., Altwein, J.E., Yen, S.S., et al. (1972). Testicular transplantation in the rat: long-term gonadotropin and testosterone radioimmunoassays. Surgery 72: 187. Goldstein, M.B., Dearden, L.C. and Gualtieri, V. (1967). Regeneration of subtotally cystectomized bladder patched with omentum: an experimental study in rabbits. J. Urol. 97(4): 664–668. Goodwin, W.E., Scott, W.W. (1952). Phalloplasty. J. Urol. 68: 903. Gooren, L.J. (1994). A ten-year safety study of the oral androgen testosterone undecanoate. J. Androl. 15: 212. Gray, H. (1918). Anatomy of the Human Body. Philadelphia: Lea & Febiger,

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68 Therapeutic Opportunities for Bone Grafting Jeffrey O. Hollinger, John P. Schmitz, Gary E. Friedlaender, Chris R. Brown, Scott D. Boden, and Samuel Lynch

INTRODUCTION Bone will regenerate. Regeneration may be defined as the restoration of form and function with tissue indistinguishable from that of the pre-injured state. Regeneration will occur in bone fractures and bone “gaps.” Healing progresses through a predictable set of discrete molecular and cellular stages that can lead to regeneration. However, fracture non-unions and bone gaps beyond a certain size (i.e. critical-sized defects (CSD)) need to be augmented with a “bone graft” to promote regeneration rather than scar formation. Regional location of either the fracture or gap, regardless if the gap is or is not a CSD, may also influence bone regeneration. Bone gaps and fractures can occur in different regional locations that will have distinctive vascular and biomechanical profiles. Consequently, the authors of this chapter wants to introduce the concept of regional anatomical domains (RADs). RADs are defined by the axial, appendicular, and craniofacial skeletons. Within RADs are regional mechanical cues coupled with molecular and cellular information that provide dual guidance for the regenerative process. Furthermore, RADs may have distinctive mechano-anatomical units (MAUs) based on regionally specific mechanical and anatomical input. RADs have both unique MAUs as well as common features, such as cell phenotypes (e.g. osteoblasts) and molecular signals (e.g. bone morphogenetic proteins (BMPs)). MAUs can be further defined by physiological parameters of force vectors, movement envelopes, tissue attachments, sensory organs, and gender distinctions. It is therefore not certain whether RADs may respond either predictably the same or differently to the same therapeutic intervention or may need RAD-specific bone graft therapies. The provocative notion of designer therapeutics may be related to developmental skeletal biology (Ferguson et al., 1999; Schneider et al., 1999; Thompson et al., 2002; Eames et al., 2003; Helms and Schneider, 2003; Franceschi, 2005) and the opinions offered by two bone icons, M.R. Urist (1980) and A.H. Reddi (1975). Reddi, in 1975, underscored the anatomical field from which bone matrix was derived and stated: “The transforming potency (of bone matrix) varies widely in matrices of different bones … .” Urist, in 1980, wrote: “Every bone and every part of the human skeleton responds to injury in its individual way and incorporates a bone graft at its own rate of repair. The factors intrinsic to the repair process are age, anatomical pattern of vascularity, immobolization, contact compression, and pathological condition.” It is nevertheless well-known and accepted, that bone regeneration across RADs will respond most favorably to the autograft. Therefore, the authors are mindful that, dividing the skeletal system into three components based on RADs and MAUs does not limit the opportunities for the same beneficial clinical outcome from a particular bone grafting therapy.

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This review chapter will emphasize both proven bone graft therapeutics as well as therapeutic opportunities for skeletal regeneration that could be effective across RADs. Further, we will note possible future opportunities for advancing the regenerative process.

APPENDICULAR SKELETON Fracture Repair: An Instructive Guide for Bone Regenerative Therapeutics Most available information on how bone heals is based on the appendicular skeleton, from both laboratory animals as well as clinical trials. Although there have been some recent efforts to develop strategies for the systemic enhancement of skeletal healing, most studies thus far have focused on local treatment. These local options for the repair and regeneration of bone include osteogenic materials (e.g. autogenous bone, marrow and blood concentrates, allogeneic bone), osteoconductive materials (such as calcium phosphates, calcium sulfate, calcium phosphate/collagen composites), tissue repair factors (e.g. recombinant proteins including fibroblast growth factor (FGF), platelet-derived growth factor (PDGF), vascular endothelial growth factor (VEGF), and osteoinductive factors (such as recombinant human bone morphogenetic proteins-2 and -7 (rhBMP-2, -7)). The two major processes that govern bone repair and regeneration are osteoconduction and osteoinduction. Osteoconduction supports the ingrowth of sprouting capillaries, perivascular tissues, and osteoprogenitor cells into the three-dimensional structure of an implant or a graft. Osteoconduction does not encompass the ability to directly regenerate bone. Osteoinduction includes recruitment of undifferentiated mesenchymal cells and their differentiation to osteoprogenitor cells followed by bone formation (Urist, 1988, 1997).

Autogenous Bone Autogenous bone is the most frequently used graft material for osseous clinical indications, and is the standard against which all other options are compared. The autogenous bone graft has an enviable profile of assets, including an unsurpassed biological potential; freedom from disease transmission from donor to recipient; the lack of immunological disparities and consequences; and desired biomechanical properties. Unless transferred on its vascular pedicle with immediate reanastomosis, autogenous bone grafts must be revascularized and repopulated with osteogenic cells. It is unlikely that a satisfactory threshold quantity of viable cells can be transferred and maintained in an autograft, with the notable exception of an immediately revascularized graft. Neovascularity and most cells crucial to graft incorporation and remodeling are provided by the host bed. Bone grafts are initially encompassed in an injury response, a hematoma, followed by inflammation and the development of fibrovascular tissue. Cortical grafts first undergo significant osteoclastic resorption of pre-existing matrix to accommodate the ingrowth of new blood vessels (i.e. neovascularization). The more porous nature of cancellous bone enables the process of graft incorporation. Depending on graft volume and location, resorption and replacement with host bone may take months to years (Goldberg and Akhavan, 2005). The disadvantages of autografts include limitations in the amount and shape of tissue available. Also, the process of recovering autogenous bone is subject to donor site morbidity, including significant and prolonged pain in more than 30% of cases, donor bone weakness or fracture, additional operative time and blood loss, and possible infection of an additional operative site (Younger and Chapman, 1989).

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Autologous Bone Marrow Autologous bone marrow can be an effective graft material for musculoskeletal applications and to restore and regenerate skeletal defects. Although most animal studies have suggested potential efficacy, there have been conflicting clinical reports regarding the success of freshly harvested bone marrow. A recent study described 60 patients with non-infected non-unions who underwent bone marrow aspiration from both iliac crests, with samples concentrated on a cell-separator and then reinjected into the non-union site. The outcome was union in 53 of 60 patients (Hernigou et al., 2005). The study suggested that autologous bone marrow can be effective, but efficacy will depend on the number and concentration of osteoprogenitor cells (Hernigou et al., 2005). An interesting and related report showed that individuals who are undergoing rapid growth or who have sustained a fracture of a major long bone, show significant numbers of osteoblast-lineage cells in their peripheral circulation (Eghbali-Fatourechi et al., 2005). Therefore, a potential direction for future research would be to develop a method to harvest osteoprogenitor cells from peripheral blood, expand them in culture, and subsequently implant them into impaired bone healing sites. Osteoconductive Materials Osteoconductive materials for musculoskeletal applications have been available for over 25 years. There is extensive literature on their use in all three RADs, which will be briefly reviewed. Most osteoconductive materials used in applications in the three RADs have been composed of calcium phosphate, calcium hydroxyapatite, or calcium sulfate. These materials do not regenerate bone, but can serve as scaffolds that in combination with a growth factor such as rhBMP, for example, will stimulate osteoprogenitor cells to attach, proliferate, and differentiate to osteoblasts. Ceramics Ceramic calcium phosphate is formed by heating and pressurizing stoichiometric ratios of calcium and phosphate. Hydroxyapatitic calcium phosphate (either laboratory synthesized, bovine derived, or as a coralline by product) and tricalcium phosphate are the most widely used. They can serve as either bone graft extenders or by themselves for cranioplasties, ridge augmentation, sinus lifts, and to deliver the recently Food and Drug Administration (FDA)-approved recombinant human platelet-derived growth factor (rhPDGF) known as GEM21S (tricalcium phosphate plus rhPDGF) (vide infra). Ceramics have the advantage of inducing little inflammatory response and pose little or no risk of disease transmission. They are available in unlimited quantities and require no donor site. The disadvantages of ceramics include low fracture resistance and tensile strength. Several animal studies using ceramics as adjuncts to posterolateral fusions have produced conflicting results. There are clinical data showing comparable results for posterolateral fusions in idiopathic scoliosis using ceramics and local bone versus local bone and autograft (Delecrin et al., 2000). The best clinical experience with the ceramics has been in the setting of anterior interbody fusion in the cervical spine. Most series report a near 100% fusion rate with the use of this graft material in the presence of rigid internal fixation (McConnell et al., 2003). Application to the appendicular skeleton has been limited. Current and future directions for the advancement of this field focus on the development of injectable calcium phosphate-based pastes that could biodegrade in synchrony with bone formation and deliver growth factors at a predictable, calibrated dose and rate, as well as aid in fracture fixation (Lobenhoffer et al., 2002; Cassidy et al., 2003; Seeherman et al., 2005). Collagen The structure of type I collagen will support mineral deposition, vascular ingrowth, and growth factor binding, but provides no structural support nor precisely calibrates growth factor release. Contemporary processes

Therapeutic Opportunities for Bone Grafting

that cross-link collagens and add calcium phosphate have yielded pre-clinical products with sufficient resistance to compression that osteoconduction through pores can be obtained. Collagen has not been effective as a stand-alone bone graft substitute. Its future role will most likely be as an ingredient in bone graft substitute composites. Beta-tricalcium Phosphate (β-TCP) A porous β-TCP bone void filler has been developed to mimic the trabecular structure of cancellous bone (Vitos, Orthovita, Malvern, PA) which enables vascularization and bone ingrowth. This material will most likely play a role as a bone graft extender in spinal arthrodeses. In the craniofacial skeleton, especially for bone and periodontal regeneration, β-TCP appears to be an effective osteoconductive matrix when used in conjunction with rhPDGF. An advantage of β-TCP compared to the apatitic calcium phosphates, is its ability to biodegrade. Apatitic calcium phosphates may take years to biodegrade and therefore inhibit bone regeneration. Tissue Repair Factors Several biological signaling molecules have been investigated for their potential efficacy for fracture repair and bone regeneration. Those that have received the most attention include FGF, PDGF, VEGF, BMP, and insulin-like growth factor (IGF). Data from one study appear to indicate that FGF may enhance skeletal defect healing in laboratory animals (Kawaguchi et al., 2001). rhPDGF has been found to accelerate fracture repair and increase bone density across RADs (Nash et al., 1994; Mitlak et al., 1996; Goldberg and Akhavan, 2005). Howes et al. (Goldberg and Akhavan, 2005) studied subcutaneous implantation of demineralized bone matrix (DBM), augmented with PDGF in young and mature rats and demonstrated enhanced cellular events, induced bone formation, and up-regulated biochemical bone markers. PDGF–DBM increased bone formation and associated bone markers in older animals by two-fold compared to DBM alone. In the younger animals, the impact of PDGF amended DBM was not significant. Nash et al. (1994) delivered PDGF in a collagen gel to treat tibial osteotomies in rabbits. The authors reported a distinct increase in callus density and volume around the PDGF-treated osteotomies compared to non-treated controls receiving only collagen. Indeed, PDGF–collagen-treated tibiae were not significantly different than unoperated, contralateral tibiae. Histologically, PDGF–collagen produced a more robust and advanced osteogenesis, both endosteally and periosteally, than collagen alone. In another PDGF study, Mitlak et al. (1996) reported on the substantial increase in bone density following treatment with rhPDGF or rhPDGF in combination with alendronate. Quantitative computerized tomography of axial and appendicular bones indicated significant enhancement in bone mass. Histologically, the PDGF recipients had a substantial increase in osteoblast number and lining osteoblasts, without a change in osteoclast number when compared to the untreated group. Biomechanically, rats treated with PDGF had significantly enhanced vertebral body compressive strength and femoral shaft torsional stiffness. The combination of alendronate with PDGF further increased these indices. VEGF has been shown to promote angiogenesis and bone turnover in experimental animals (Street et al., 2002). Because this growth factor promotes the ingrowth of capillaries which may bring with them perivascular mesenchymal cells capable of responding to osteoinductive substances in the environment, the ability to combine VEGF with other growth factors represents an important research direction for developing therapies for bone regeneration. One interesting concept that has recently come forward is the use of synthetic small peptides for the enhancement of skeletal repair and regeneration. It is well known that prostaglandins in the E-series are strong stimulants of bone formation in trabecular sites (Raisz, 1999). Indeed, a report of prostaglandin E infusion into neonates demonstrated substantial periosteal bone formation at remote skeletal locations (Ueda et al., 1980). Because of the objectionable side effects of prostaglandin E, direct administration in patients may not be feasible. However, the ability to target the prostaglandin E receptor may allow for a means to achieve

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prostaglandin-type stimulation without exposure of other organs and tissues to this molecule. A recently developed prostaglandin EP-2 receptor agonist has been demonstrated to enhance fracture healing in animals and is currently undergoing investigation for use in patients with tibial fractures (Li et al., 2003). Osteoinductive Materials The concept of osteoinduction was first introduced by Marshall R. Urist in 1965 (Urist, 1965). Urist recognized the capacity for bone to regenerate and hypothesized that a naturally occurring protein in bone was responsible for this effect; he named the protein bone morphogenetic protein (Urist and Strates, 1971). Over the course of the next 25 years, investigators will identify a family of proteins with bone-inducing properties. There are now at least 15 BMPs, most of which have some bone-inductive activity. Currently, BMP-2 and -7 (also known as osteogenic protein-1 (OP-1)) have received FDA approval for clinical use. BMP-7 (OP-1) was tested in a randomized controlled trial of 122 patients with 124 tibial non-unions. All nonunions were at least 9 months old with no progress toward healing for the 3 months prior to the study. Each patient was treated with an intramedullary nail plus recombinant human OP-1 (rhOP-1) in a type I collagen carrier or with autogenous bone graft alone. After 9 months’ follow-up, 81% of the patients treated with OP-1 and 85% of the patients treated with autogenous bone graft were judged to be healed by clinical criteria. The investigators concluded that rhOP-1 implanted with a type I collagen carrier was safe and effective in the treatment of tibial nonunions (Friedlaender et al., 2001). In another investigation using rhBMP-2, a prospective randomized controlled trial in 450 patients with open tibial fractures was conducted. Patients received irrigation and debridement and treatment with a statically locked intramedullary nail or supplementation with one or two doses of rhBMP-2. After 12 months follow-up, the patients treated with the higher dose of BMP-2 showed a 44% reduction in the risk of secondary interventions and showed fewer hardware failures and fewer infections (Govender et al., 2002). Future Directions Currently approved clinical applications of BMPs require supraphysiological doses, and this escalates the cost of therapy. Consequently, there need to be strategies to enhance osteoinduction with substantially lower dosing and more effective carriers. Gene therapy may provide a more effective opportunity for bone regeneration than protein forms of BMP (Lieberman et al., 1999). A combination approach to augment BMP is yet another strategy. One attractive combination is that use of an angiogenic gene for a factor such as VEGF and an osteoinductive gene coding for a molecule such as BMP-4. Indeed, a recent report demonstrated the synergistic enhancement of bone formation and healing by stem cell expressed VEGF and BMP-4 (Peng et al., 2002). The challenge for the future is to design and develop an effective and efficient biologically programmed delivery system that will provide to the healing bone wound, the proper dose of growth factor at the appropriate rate and at the precisely required point in time during the regenerative process. This will require a better understanding of growth factor pharmacokinetics.

AXIAL SKELETON Currently there are over 250,000 spine fusions carried out in the United States each year. Since the 1990s, spinal arthrodesis has become the most common reason for autogenous bone grafting. The most common types of fusion can broadly be classified as interbody fusions, facet fusions, and posterolateral fusions. Despite instrumentation and used with autogenous grafts, non-union rates range from 5% to 44% with single level posterolateral fusions (Dodd et al., 1988). Several strategies have been developed to augment healing or to replace autogenous bone grafting, which is considered the gold standard in spinal arthrodesis. A successful spinal arthrodesis requires the incorporation

Therapeutic Opportunities for Bone Grafting

of the bone graft into the recipient site. Graft incorporation is influenced by the graft material and the biological and biomechanical environments of the healing process, which are neither fully understood nor adequately defined across RADs. The importance of the biological environment has been described by Boden and Schimandle (1995) where they stressed the need for bone marrow-derived osteoprogenitor cells, rigid internal fixation to avoid shearing of blood vessels, and adequate endplate strength to prevent subsidence (Boden and Schimandle, 1995). Boden and associates, using a rabbit intertransverse fusion model with iliac crest bone graft, revealed three distinct and temporal phases of spine fusion healing: inflammatory, reparative, and remodeling. These phases occur throughout the fusion mass but at different times. Maturation of the fusion mass was most advanced in the regions near the decorticated transverse processes, the site of major blood and progenitor cell supply to the fusion mass. In these transverse process zones, intramembranous formation dominated. In the central zone, a similar process occurred but was delayed in time and included a period of endochondral bone formation, where cartilage was replaced by bone. This lag effect may explain why non-unions most often occur at this central zone (Boden et al., 1995). Advances in molecular biology have shown a predicable cascade of gene expression during the maturation of a fusion bed. Various growth factors, including the (BMPs), have a temporal and spatial pattern of expression. Consistent with the histological analysis, peak expression of these genes was seen 1–2 weeks earlier in the outer/transverse process zones compared to the central zone (Morone et al., 1998). Graft Options for the Axial Skeleton Bone Grafts Autografts and allografts were mentioned earlier in this chapter. DBM DBM is produced by the decalcification of cortical bone (Urist, 1965, 1967). The ideal demineralization process removes the calcium and phosphate but leaves the extracellular matrix which consists predominantly of type I collagen and non-structural proteins, including growth factors, such as BMPs. DBM provides little biomechanical strength but functions as an osteoconductive and, to a lesser degree, an osteoinductive material. The qualitative and quantitative analyses to determine the types of BMPs and their concentrations in various DBM products have not been comprehensively ascertained. Moreover, DBM processing, donor selection, terminal sterilization, and additives (e.g. hyaluronate, plurionic acid, glycerine, gelatin) yield products with variable biological activities, and, consequently, unpredictable biological osteoinductivity and clinical performance. A number of animal studies have validated DBM as a bone graft extender or enhancer for posterolateral fusions. DBM has also been shown to have variable efficacy based on details of the preparation, composite form, terminal sterilization, and the environment being tested (Wang et al., 2000). There are little prospective clinical outcome data related to the use of DBM because it has been regulated as a minimally manipulated tissue rather than a device. Its current use is mostly as a bone graft extender in spinal fusions, trauma, and bone defects. BMP-2 and -7 Currently rhBMP-2 is available for anterior interbody fusions. Boden et al. have reported on the first pilot study examining the osteoinductive capacity of rhBMP-2 for a human spinal fusion application (Boden et al., 2000). In a limited randomized multicenter study involving 14 patients, threaded interbody fusion cages were filled with either rhBMP-2/collagen sponge or autogenous iliac crest bone graft and implanted for anterior lumbar interbody fusion. The rhBMP-2 patients, not requiring iliac crest harvest, had a shorter hospital stay compared to the autograft control patients (2 days versus 3.3 days). Of the rhBMP-2 patients, 10 of 11 were judged fused by 3 months after surgery, and all 11 were fused by 6 months. Of the three control patients, one

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was deemed a non-union after 1 year. Among the rhBMP-2-induced fusions, computed tomography (CT) scan reconstructed images consistently showed new bone growth through and anterior to the cages 6 and 12 months after surgery (Boden et al., 2000). The results were confirmed in a large prospective, randomized trial which resulted in FDA approval in 2002 (Burkus et al., 2002). The efficacy of rhBMP-7 in human spinal fusion procedures has also been demonstrated. Jeppsson et al. evaluated BMP-7 (also known asOP-1) in cervical spine posterior fusion in four patients with rheumatoid disease (Jeppsson et al., 1999). No bone formation occurred in three patients. Vaccaro et al. reported their 2 year experience with OP-1 for posterolateral spine fusion in humans (Vaccaro et al., 2005). In this study, 12 patients underwent decompression and fusion for spinal stenosis and degenerative spondylolisthesis. The patients received combined autograft and OP-1. No adverse effects were seen, and the radiographic fusion rate was slightly more that the 50% fusion rates using autograft alone. Any observed differences in clinical outcomes between the two BMPs (BMP-2 versus BMP-7) may underscore the different fundamental biological mechanisms between BMPs during bone healing, even within the same RAD, or reflect biological or biomechanical differences in the sites and condition under which these molecules were studied. Future Directions A strategy to achieve spinal fusion could involve gene therapy. In vivo gene therapy transfers a gene to cells on site, while ex vivo gene therapy transfers a gene to a cell outside the body that is then implanted. If the gene is for a secreted protein (e.g. BMP) the transfer cells do not need to be bone forming cells. They simply serve as a factory for the gene. Despite gene therapy being a compelling option, daunting regulatory hurdles may likely challenge clinical realization. Recently, a novel protein has been discovered and its gene elucidated. The protein appears to function early in the cascade of events that lead to bone formation. It is an intracellular signaling protein named LIM mineralization protein-1 (LMP-1). This gene has been transfected into bone marrow cells of rats and reimplanted into posterolateral fusion beds. It proved effective for promoting a solid fusion (Boden et al., 1998). This study validated the possibility of gene therapy for inducing spinal fusions.

CRANIOFACIAL SKELETON The previous sections provided the reader an overview on the grafting opportunities that may find utility across appendicular and axial RADs. Therefore, this section will focus on growth factor opportunities, emphasizing the two growth factors that have generated the most clinical enthusiasm: rhBMP and rhPDGF. These two growth factors also have been selected for emphasis because they have received FDA approval although only PDGF has received FDA approval for use in the craniofacial skeleton. We further choose to emphasize the oral and maxillofacial (OMF) segment of the craniofacial skeleton. This particular anatomical entity may prove to be one of the most stringent reconstructive challenges of the RADs, in that the surgeon and patient must come to terms with bacterial contaminants, the ginglymo–arthroidal temporomandibular joint, the dento–alveolar gomphotic joints (i.e. tooth–cementum–periodontal ligament–alveolar bone), as well as the mucosal, glandular, neuronal, vascular, and muscular anatomies and focal emphasis on the facial “personality.” BMP-2 and -7 BMPs-2 and -7 are potent osteoblast differentiating factors with widespread therapeutic skeletal potential (Urist, 1971, 1997). There have been a number of pre-clinical and human clinical studies employing rhBMP-2 in the craniofacial skeleton (Toriumi et al., 1991; Mayer et al., 1996; Boyne et al., 1997, 1998; Wikesjo et al., 1998; Cochran et al., 2000; Boyne, 2001; Boyne and Shabahang, 2001; Wikesjo et al., 2001; Bianchi et al., 2004; Fiorellini et al., 2005). To date, however, neither rhBMP-2 nor rhBMP-7 (rhOP-1) has received FDA approval

Therapeutic Opportunities for Bone Grafting

for the craniofacial skeleton. This delay may be the result of variable results in rigorous clinical trails in this RAD (i.e. the craniofacial skeleton). In the craniofacial skeleton, human trials with the BMPs have been conducted in periodontal defects, alveolar defects, and for maxillary sinus floor augmentation, a procedure to increase bone in the maxilla to facilitate placement and retention of dental implants (Azari et al., 2002), with the most promising published data in the latter indication. A collagen sponge (ACS, absorbable collagen hemostatic agent) vehicle has been used to deliver the rhBMP-2 (dose range 1.77–3.4 mg). This is the same “vehicle” employed in the rhBMP-2 InFuse spinal fusion product and the InDuctos tibial fracture repair product. In the maxillary sinus, approximately 6 months post-operatively, biopsy specimens from five patients exhibited a wide variation in woven bone quantity, undetectable or small populations of active bone cells and capillaries. No residual collagen matrix was noted in any of the biopsy specimens taken after 19 weeks postoperatively. The adverse reactions were consistent with usual morbidity for the sinus augmentation procedure. Following therapy with rhBMP-2, three patients did not meet the criteria for implant placement. The authors suggested that the method of access to the sinus may influence efficacy of rhBMP-2-induced bone formation (Goldberg and Akhavan, 2005). Another clinical target for rhBMP-2 is the alveolar ridge, which undergoes resorption following loss of the teeth. Edentulous ridges characteristically have decreased bone height and width – a complication for the placement of implants. A two center clinical trial evaluated rhBMP-2/ACS enhancement of alveolar ridge following tooth extraction in six patients. The mean dose for rhBMP-2 was 0.27 mg in the extraction sites and 0.83 mg in the augmentation sites. Two aspects for each procedure were evaluated: safety and osteointegration by tomographic analysis over a 4-month period. Both the Boyne (Goldberg and Akhavan, 2005) and Howell (Morone et al., 1998) studies lacked controls. Biopsies were not analyzed in the extraction graft and ridge augmentation study. Both studies reported relative ease in the use of the graft vehicle (Morone et al., 1998). Probing and palpation of the graft sites provided an additional clinical impression of bone integrity. However, ridge augmentation with the graft was not observed. Progress in the area of BMP research has resulted in the knowledge of the mechanisms of BMP signaling from the extracellular environment to changes in gene expression in the nucleus (Schmitt et al., 1999; Celil et al., 2005). Future investigations will need to determine the appropriate dose-response relationships with clinical targets and focus on the development of a safe and effective carrier system to predictably deliver BMP. PDGF PDGF is a powerful mitogen, chemoattractant, and co-factor with VEGF for angiogenesis (Heldin and Westermark, 1999). These biological properties are in contrast to BMP, which is a differentiating factor. The mitogenic property of PDGF for bone wound healing make this molecule an ideal candidate for conditions where compromised osseous regeneration may be encountered, such as is in aging (Doll, 2002). Moreover, augmenting DBM with PDGF maybe a clinical opportunity, since evidence suggests PDGF enhances DBM-induced bone formation (Howes et al., 1988). It is noteworthy to mention a recent study that contradicted the report by Howes et.al. (1988) that rhPDGF diminished the biological potential of DBM (Ranly et al., 2005). The report on decreased DBM activity may be flawed in that an immunocompromized mouse model was used, excess PDGF dosing may have actually down-regulated the biological signal of PDGF (Heldin and Westermark, 1999), and there is concern that human PDGF may lack sufficient homology in the mouse to be biologically effective. PDGF is accepted as a pivotal molecular cue for osseous wound healing. This has resulted in the therapeutic application of PDGF to the dental–alveolar complex. Compelling pre-clinical and clinical studies have led to the recent FDA approval of a product known as GEM21S (Biomimetic Therapeutics, Franklin, TN), which contains rhPDGF-BB delivered in tricalcium phosphate granules.

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In dental applications requiring regeneration of the alveolar bone, cementum, and periodontal ligament, PDGF has been profoundly successful (Camelo et al., 2003; Nevins et al., 2003). Nevins and colleagues reported demineralized freeze–dried bone allograft (DFDBA) amended with one of three concentrations of rhPDGF-BB (0.5 mg/ml, 1.0 mg/ml, 5 mg/ml) were used to treat nine patients with advanced periodontitis and whose teeth were destined for extractions (Nevins et al., 2003). At 9 months postoperatively, treated teeth and surrounding bone removed en bloc and histological and clinical evaluations revealed no unfavorable tissue reactions or safety concerns, either at the time of extraction or throughout the study duration. Further, there was complete regeneration of the periodontal attachment apparatus, including new cementum, periodontal ligament, and bone coronal to a fiduciary tooth notch in four of the six interproximal defects (Nevins et al., 2003). The Nevins report is the first one to show periodontal regeneration by histological sections in human class II furcation defects. Subsequently, these results were confirmed in another human histological study (Camelo et al., 2003). Nevins et al. in another study reported the results of a pivotal large-scale, prospective, randomized, double-blind, controlled, parallel-arm, human clinical trial evaluating the safety and efficacy of rhPDGF in combination with β-TCP, a biocompatible, biodegradable synthetic matrix to treat intrabony periodontal defects (Nevins et al., 2005). The Nevins study is the first to provide data on efficacy and safety of a growth factor (rhPDGF-BB) for periodontal regeneration that lead to the pre-market approval by the FDA. The outcome for the study was determined following surgeries to 180 patients across 11 clinical centers. The outcome revealed at doses of 0.2 mg/ml and 1.0 mg/ml rhPDGF was safe and efficacious over a 6-month period in treating periodontal osseous defects. The study demonstrated a significant increase in the clinical attachment levels and reduced gingival recession at 3 months post-surgery and improved osseous regeneration as compared to a β-TCP bone substitute at 6 months (Nevins et al., 2005). The periodontal wound model is considered to be an extraordinarily daunting regenerative challenge due to continuous bacterial and food debris challenges and the fact is that, the periodontal wound is an “open wound.” Periodontal disease (i.e. periodontitis) is a challenge to treat and is a widespread disease. The majority of adults throughout the world have periodontitis and according to the American Academy of Periodontology, about one in three people over 30 years of age have periodontitis and 50% of 55–64 year olds in the United States have evidence of severe periodontal disease (Reviewed (DeCarlo and Whitelock, 2006)). Consequently, an effective rhPDGF-based bone graft therapy for periodontal indications will provide compelling and profound therapeutic benefit to patients as well as offer opportunities for bone applications at other RADs. This emphasis is consistent with the universal participation of PDGF in wound healing. Future Directions The success of a therapeutic, regardless of the RAD and particular bone to be regenerated, will be determined by a suitable match between the biological profile of both the patient and the recipient RAD site and the biological and biomechanical properties of the therapy. Contemporary wound biology has not satisfactorily provided a comprehensive bone healing road map for the craniofacial, appendicular, and axial skeletons. Consequently, there is significant opportunity to design and develop biologicals and delivery systems based on a three-dimensional blue print for healing that can be tuned to the clinical target. The blue print is three-dimensional and includes the elements of time, dose, and position. Future opportunities therefore will include a temporal release of a growth factor or combination suite of growth factors that will be sequenced and calibrated precisely to the needs of the RAD, the patient gender and health status. For example, if the patient is an elderly osteoporotic with qualitative and quantitative osteoblast deficit, PDGF would be beneficial for osseous wound healing, regardless of the RAD. This fact is underscored as a consequence of the fundamental biological properties of PDGF: it is a chemoattractant and mitogen for osteoprogenitor cells and is a universal and pivotal wound healing growth factor.

Therapeutic Opportunities for Bone Grafting

We expect future directions in bone therapeutic design and development will be guided precisely by unique attributes of the RADs and MAUs and defined by comprehensive three-dimensional healing maps. Moreover, individualized and common thematic biological and cellular elements across and within RADs will enable effective and broad utilization of bone graft therapeutics.

ACKNOWLEDGMENT Partial support for this work was provided by the NIDCR R01-DE15392 (JOH)

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Nevins, M., Giannobile, W.V., McGuire, M.H., Kao, R., Mellonig, J., Hinrichs, J., McAllister, B., Murphy, K.G., McClain, P., Nevins, M.L., Paquette, D.W., Han, T., Reddy, M., Lavin, P.T., Genco, R. and Lynch, S. (2005). Platelet-derived growth factor stimulates bone fill and rate of attachment level gain: results of a large multicenter randomized controlled trail. J. Periodontol. 76: 2205–2215. Peng, H., Wright, V., Usas, A., Gearhart, B., Shen, H.C., Cummins, J. and Huard, J. (2002). Synergistic enhancement of bone formation and healing by stem cell-expressed VEGF and bone morphogenetic protein-4. J. Clin. Invest. 110: 751–759. Raisz, L.G. (1999). Prostaglandins and bone: physiology and pathophysiology. Osteoarthritis Cartilage 7(4): 419–421. Ranly, D., McMillan, J., Keller, T., Lohmann, C.H., Meunch, T., Cochran, D., Scwartz, Z. and Boyan, B.D. (2005). Plateletderived growth factor inhibits demineralized bone matrix-induced intramuscular cartilage and bone formation. A study of immunocompromised mice. J. Bone Joint Sur. Am. 87: 2052–2064. Reddi, A.H. (1975). The matrix of rat calvarium as transformant of fibroblasts (39028). Proc. Soc. Exp. Biol. Med. 150: 324–326. Schmitt, J.M., Hwang, K., Winn, S.R. and Hollinger, J.O. (1999). Bone morphogenetic proteins: an update on basic biology and clinical relevance. J. Orthop. Res. 17(2): 269–278. Schneider, D.J., Hu, D. and Helms, J. (1999). From head to toe: conservation of molecular signals regulating limb and craniofacial morphogenesis. Cell Tissue Res. 296: 103–109. Seeherman, H. and Wozney, J.M. (2005). Delivery of bone morphogenetic proteins for orthopaedic tissue regeneration. Cytokine Growth Factor Rev. 16(3): 329–345. Street, J., Bao, M., deGuzman, L., Bunting, S., Pearle, F.V., Ferrara, N., Steinmetz, N., Hoeffel, J., Cleland, J.L., Daugherty, A., van Bruggen, N., Redmond, H.P., Carano, R.A. and Filvato, E.H. (2002). Vascular endothelial growth factor stimulates bone repair by promoting angiogenesis and bone turnover. Proc. Natl Acad. Sci. USA 99(15): 9656–9661. Thompson, Z., Miclau, T., Hu, D. and Helms, J. (2002). A model for intramembranous ossification during fracture healing. J. Orthop. Res. 20: 1091–1098. Toriumi, D.M., Kotler, H.S., Luxunberg, D.P., Holtrop, M.E. and Wang, E.A. (1991). Mandibular reconstruction with a recombinant bone-inducing factor. Functional, histologic, and biomechanical evaluation. Arch. Otolaryngol. Head Neck Surg. 117(10): 1101–1112. Urist, M.R. (1965). Bone: formation by autoinduction. Science 150: 893–899. Urist, M.R. (1980). Fundamental and Clinical Bone Physiology, Philadelphia: J.B, Lippencott. Urist, M.R. (1988). Bone morphogenetic protein induced bone formation in experimental animals and patients with large bone defects, in cell and molecular biology of vertebrate hard tissue. In: Evered,D. and Harnett,S. (eds.), Symposium 136, London Ciba Foundation. pp. 281–284. Urist, M.R. (1997). Bone morphogenetic protein: The molecularization of the skeletal system. J. Bone Miner. Res. 12(3): 343–346. Urist, M.R. and Strates, B.S. (1971). Bone morphogenetic protein. J. Dent. Res. 50: 1392–1406. Urist, M.R., Silverman, M.F., Buring, K., Dubuc, F.L. and Rosenburg, J.M. (1967). The bone induction principle. Clin. Orthop. Relat. Res. 53: 243. Ueda, K., Sito, A., Nakano, H., Aoshima, M., Yokota, M., Muraoka, R. and Iway, T. (1980). Cortical hyperostosis following long-term administration of prostaglandin E1 in infants with cyanotic congenital heart disease. J. Pediatr. 97: 834–836. Vaccaro, A.R., Patel, T., Firschgrund, J., Anderson, D.G., Trumees, E., Herkowitz, H., Phillips, F., Hilibrand, A.S. and Albert, T.J. (2005). A 2 year follow up pilot study evaluation of the safety and efficacy of OP-1 putty as an adjunct to iliac crest autograft in posteriorlateral lumbar fusions. Eur. Spine J. September 14(7): 623–629. Wang, J.C., Davies, M., Kanin, L.E. and al, e. Prospective comparison of commercially available demineralized bone matrix for spinal fusion. in North American Spine Society. 2000. New Orleans, LA. Wikesjo, U.M., Razi, S.S., Sigurdsson, T.J., Tatakis, D.N., Lee, M.B., Ongpipattanakul, B., Nguyen, T. and Hardwick, R. (1998). Periodontal repair in dogs: effect of recombinant human transforming growth factor-beta1 on guided tissue regeneration. J. Clin. Periodontol. 25(6): 475–481. Wikesjo, U.M., Sorensen, R.G. and Wozney, J.M. (2001). Augmentation of alveolar bone and dental implant osseointegration: Clinical implications of studies with rhBMP-2. J. Bone Joint Surg. Am. 83A (Suppl 1(Pt 2)): S136–S145. Younger, E.M. and Chapman, M.W. (1989). Morbidity of bone graft donor sites. J. Orthop. Trauma 3: 192–195.

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69 Cartilage Tissue Engineering Paulesh Shah, Alexander Hillel, Ronald Silverman, and Jennifer Elisseeff

CLINICAL NEED FOR CARTILAGE REPAIR Cartilage is a mesenchymal tissue that functions to provide form, strength, and support. Cartilage is found in joints, throughout the head and neck, and adjacent to intervertebral discs. In joints, cartilage lines the surface of bones cushioning the impact caused by physical load during movement. In the head and neck it is unique in its structural support of the ear, nose, and throat. Cartilage is composed of cells, extracellular matrix (ECM) molecules, and water. Chondrocytes represent the cellular component and comprise 1–2% of adult cartilage tissue volume (Lohmander, 2003). The cartilage ECM is comprised of collagen fibers supporting glycoproteins and proteoglycans – which have a protein core associated with glycosaminoglycan molecules such as hyaluronic acid (HA) and chondroitin sulfate. The tissue fluid, mainly water, contributes to the mechanical properties of cartilage and provides nutrition and waste exchange with synovial fluid in the joint and with extracellular fluid of other adjacent tissues. There are three types of cartilage that are distinguished by their molecular components in the ECM, their anatomic location, and their function. Hyaline cartilage has a white glassy appearance and is found primarily in the joint. Its ECM is composed of water, proteoglycans, and type II collagen. Hyaline cartilage functions to provide stable movement with minimal friction. It demonstrates an excellent ability to provide resistance to compression and distribute loads – adapting to the load-bearing force as it changes over time (Temenoff and Mikos, 2000). Elastic cartilage is distinguished by the presence of elastin in the ECM. Elastic cartilage provides a structural function, represented by the support it provides the airway and the external ear. Lastly, fibrocartilage has a higher proportion of type I collagen in the matrix. Fibrocartilage is found at end of tendons and ligaments in apposition to bone providing tensile strength and countering compression and shear forces (Benjamin and Rauphs, 2004). In contrast to fibrocartilage, hyaline and elastic cartilage have a perichondrium, which channels a limited blood supply and contains a source of progenitor cells. All types of cartilage are distinguished by a limited ability for intrinsic repair after sustaining damage. This is due, in part, to a lack of vascularity, the sparse cellularity of native chondrocytes, and the inflammatory response. Furthermore, the perichondrial blood channels and progenitor cells are limited in their ability to regenerate tissue. In full-thickness articular defects the inflammatory response creates a fibrin clot and directs progenitor cells to migrate to the wound. However, the spontaneous repair by these cells results in inferior scar tissue with more fibrous characteristics that do not restore full function (Temenoff and Mikos, 2000; Solchaga et al., 2001). In 2001, the CDC estimated 69.9% of American adults suffered from arthritis (Prevention, 2002). Osteoarthritis frequently arises from cartilage damage due to sports injury, other trauma, and/or simply overuse. The resultant degeneration of cartilage combined with its inability to self-repair leads to further degradation of the joint (Temenoff and Mikos, 2000). Each year in the United States surgery is performed on greater than 3 million knees, hips, and other joints (Owings and Kozak, 1998). The majority of these procedures are indicated due to cartilage injury. In the head and neck, cartilage replacement or repair is needed for degenerative disease, traumatic injury, or agenesis. Over 500,000 surgical procedures are performed each year involving cartilage replacement in the

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head and neck (Owings and Kozak, 1998). Septal reconstruction represents the most common procedure. However, auricular reconstruction is needed in cases of traumatic amputation, injury, or cases of microtia, which affects 1 in 7,000 to 8,000 infants (Aguilar, 2001). Laryngotracheal reconstruction, though infrequently performed, requires cartilage to repair the airway in patients with subglottic stenosis. Currently, surgeons harvest cartilage from the auricle, thyroid ala or costal area to repair these defects. Current therapies for cartilage repair are inadequate at restoring form and function. Cartilage autografts suffer from many problems including limited donor tissue availability, donor site injury, scarring, and pain. Similarly, allogeneic and alloplastic implants have a high risk of infection, graft resorption, and structural failure. The implant materials do not integrate into the host tissue and have a limited lifetime. In addition, implantation generally requires invasive surgery. In many cases the procedure leaves a poor esthetic result. Therefore, there is a large opportunity for new therapies such as cartilage tissue engineering (Lalan et al., 2001; Solchaga et al., 2001; Stock and Vacanti, 2001).

TISSUE ENGINEERING OVERVIEW Tissue engineering is a multidisciplinary field that applies the principles of engineering, life sciences, cell and molecular biology toward the development of biological substitutes that restore, maintain, and improve tissue function (Mooney and Mikos, 1999). Three general components are involved in tissue engineering: (1) reparative cells that can form a functional matrix; (2) an appropriate scaffold for transplantation and support; and (3) bioreactive molecules, such as cytokines and growth factors that will support and choreograph formation of the desired tissue (Sharma and Elisseeff, 2004). These three components may be used individually or in combination to regenerate organs or tissues. Different cell sources are available to provide reparative tissue including differentiated cells, mesenchymal stem cells (MSCs), and embryonic progenitor cells. Differentiated cells may be autologous and procured from biopsy, grown ex vivo, and transplanted elsewhere in the body to heal a defect. A direct biopsy is taken from tissue that is the same as the engineered tissue type. An indirect biopsy may be taken from a similar tissue when a direct biopsy would cause considerable injury. For example, a peripheral vein biopsy provides the necessary tissue to reconstruct a heart valve (Stock and Vacanti, 2001). Currently the only FDA approved cellular-based therapy for cartilage defects uses differentiated chondrocytes (Brittberg et al., 1994). Autologous cells may be harvested from a small biopsy and expanded ex vivo to create a relatively large supply of cells for transplantation. While autologous expanded chondrocytes have a low risk of immune rejection, they have a tendency to dedifferentiate (lose their phenotype) in vitro. Preventing dedifferentiation must be achieved for these therapies to restore function. Furthermore, while indirect biopsies are feasible for some tissue types, there are many anatomical locations, such as nerves, where a biopsy would cause extensive damage and cannot be performed. A second cellular paradigm for engineering tissue is the use of undifferentiated progenitor cells that have the ability to proliferate and differentiate into the target tissue. Embryonic progenitor cells include (1) embryonic stem (ES) cells derived from the inner cell mass of the blastocyst (Barberi et al., 2005); (2) embryonic germ (EG) cells which are harvested from the gonadal ridge of fetal tissue (Kim et al., 2005); and (3) embryonic carcinoma cells which are isolated from teratocarcinomas (Taranger et al., 2005). These cells are pluripotent and theoretically have an unlimited life span and unlimited potential to become any differentiated cell in the body. However, there is some controversy in the United States about using these cells due to their origin and procurement. Furthermore, there are concerns about the potential immunogenicity and oncological transformation of these cells. Adult stem cells are a second form of undifferentiated progenitor cell. These multipotent cells are found in adult tissue, such as bone marrow, adipose, lung, and inner ear. Their physiological role is theorized to be

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directing the regenerative capabilities of injured tissue. MSCs derived from bone marrow have received significant attention in regenerative medicine for their ability to differentiate into numerous cell lineages and form tissues including cartilage and bone (Pittenger et al., 1999; Caplan and Bruder, 2001). MSCs have been transplanted for the repair of bone, cartilage, and myocardium among other tissues (Wakitani et al., 2002; Liu et al., 2004). MSC transplantation is attractive because cells are able to multiply without losing their capacity to differentiate, and may potentially be used without immunosuppressive therapy with autologous or allogeneic transfer (Wakitani et al., 2002; Liu et al., 2004). While these cells have traditionally been recognized as having limited differentiation capacity that view is now being debated. Recently, adult stem cells plasticity – their ability to differentiate into endodermal or ectodermal lineages, has been proposed as MSCs have been shown to differentiate into liver, kidney, skin, and lung cells (Herzog et al., 2003). To create distinct tissue types, specific control over the induction and maintenance of stem cell differentiation is imperative. Currently, we are establishing control of stem cell behavior to develop therapies, particularly in the area of the musculoskeletal system. The mechanisms of induction, stability, and permanence of differentiation and proliferation remain to be elucidated (Lalan et al., 2001). The use of scaffolds and biological signaling molecules will play an important role in the signaling mechanisms of stem cells. Tissue engineering scaffolds are designed to provide a three-dimensional (3D) environment to support and direct cellular processes in their migration, proliferation, and differentiation toward functional tissue. The scaffold provides mechanical stability and surface chemistry to orchestrate biological signals and influence cellular adhesion and function. Scaffolds must be biocompatible with host tissue, porous enough to allow nutrients to reach cells, and ultimately degrade without toxic effects as tissue develops. Scaffolds may be processed into solid structures, requiring ex vivo cell seeding and surgical implantation; while gel scaffolds, such as hydrogels, physically entrap the reparative cells and may be injected to take the form of a defect in situ (Sharma and Elisseeff, 2004). Scaffolds may be biological, synthetic, inorganic composites, or hybrid combinations. Biological scaffolds include natural collagen, alginate, HA, and acellular dermis. Biological scaffolds potentially allow for better regulation of cell adhesion and matrix production of the resident cells (Hubbell, 2003). However, biological scaffolds have a greater risk of contamination or immune reaction than synthetic scaffolds. On the other hand, synthetic materials are created de novo from molecules such as glycolic acid polymers, lactic acid polymers, and polypropylene fumarate. Synthetic scaffolds provide precise control over the structural properties, mechanical properties, and rates of resorption (Drury and Mooney, 2003). Their advantage over biological materials is that they may be engineered to better regulate cellular interactions with the scaffold (Seliktar, 2005). Biosynthetic hybrids have the potential to improve the structure and function of scaffolds (Seliktar, 2005). For example, a synthetic scaffold incorporated with proteins will enhance cell adhesion and proliferation while retaining mechanical and structural control (Lalan et al., 2001; Seliktar, 2005). Inorganic composite scaffolds such as metals and ceramics are commonly used for bone. Biological factors that guide cellular differentiation include soluble biochemical signals, transfection of gene vectors, insoluble ECM molecules, environmental factors like mechanical compression, and cell–cell interactions (Bottaro et al., 2002; Sekiya et al., 2002). A combination of these factors creates the physiological and mechanical factors that promote cellular differentiation and proliferation. It should be emphasized at this time that many signals, signaling pathways, and the rationale behind physiological design remain to be elucidated. Soluble signaling molecules include growth factors and cytokines. Identifying the correct factors as well as the timing and amount of their release plays a large role in the efficacy of tissue differentiation. For example, bone morphogenetic protein-2 (BMP-2) causes chondrogeneic differentiation in early embryonic distal digit formation, but causes cell death in a later phase (Zou and Niswander, 1996; Caplan, 2003). Furthermore, the effect of the signaling molecule may depend on the effector cell’s location within tissue, that is, the effect on a cell at the tissue edge may be different that that on a cell within the tissue center.

Cartilage Tissue Engineering

It is possible to produce signaling molecules by genetic manipulation (Trippel et al., 2004). Gene therapy holds promise as a cell programming method to induce differentiation. In contrast to measuring and monitoring growth factor administration, gene transfer provides a local and sustained supply of bioactive proteins. Gene therapy has encountered obstacles with delivery methods; however, upon development of a reliable delivery technique, genetic engineering will likely interface with tissue engineering (Nussenbaum et al., 2004). ECM molecules and growth substrates are a third form of signaling control. The integrins represent molecules that adhere to cell surface receptors and influence cell morphology, migration, and signal transmission (Bottaro et al., 2002). These molecules may be integrated with scaffolds and theoretically combine with growth factors to have a synergistic effect on intercellular signaling to regulate cellular migration, proliferation, and differentiation (Bottaro et al., 2002). Theoretically scaffolds can be designed to supply soluble signaling molecules and direct local and migrating signaling factors’ interaction with the cellular component of tissue engineering. Environmental factors, including mechanical stimulation and shear forces, create biological cues that exert an effect on cells. The influence of mechanical forces on cell function in vitro has been demonstrated on engineered cartilage, bone, smooth muscle, and vascular endothelium amongst other tissues (Kim et al., 1999; Bottaro et al., 2002; Riha et al., 2005). Tissues form in response to distinct mechanical stimulation. For example, pulsatile flow influences vascular endothelium and smooth muscle differentiation and proliferation. In cartilage production, dynamic mechanical stresses on chondrocyte and MSCs promote differentiation and increased matrix production (Darling and Athanasiou, 2003; Waldman et al., 2003). Lastly, cadherins and cellular adhesion molecules mediate cell–cell interactions that can initiate signaling cascades that, in turn, influence cell adhesion and migration. One theory models tissue engineering on embryonic tissue development (Solchaga et al., 2001; Caplan, 2003). Embryonic mesenchymal tissues are highly cellular with smaller amounts of ECM compared to adult tissues. As animals grow, the cell–matrix ratio declines with adult ECM increasing its molecular components to become more dense and specialized (Caplan, 2003). During development it is evident that the matrix has an active role in cellular communication to engineer cell differentiation and tissue formation in addition to its structural role. ECM molecules bind growth factors and enzymes such as metalloproteases, collagenases as well as progenitor and differentiated cells. Perhaps the use of biological scaffolds such as chondroitin sulfate will better harness scaffolds natural signaling abilities (Solchaga et al., 2001). Mechanical stimulation may also influence signaling in embryonic tissue development. Mikic et al. demonstrated mechanical loading influenced composition of embryonic cartilage development (Mikic et al., 2004). At this point it is unlikely that tissue engineering will be able to mimic the complex and ever shifting signaling between cells within tissue and different tissue types that occur during embryonic tissue development (Caplan, 2003). However, the greater understanding we have over the intricate signaling pathways during embryonic development, the greater insight and efficacy we may have engineering replacement tissue and organs.

CELL-BASED CARTILAGE ENGINEERING IN VITRO In vitro culture systems provide many benefits to studying cellular-based tissue engineering. First, it provides an isolated system with which to define the proliferation and differentiation of cells and allows for focused study into the cellular and molecular events of chondrogenesis. While results from monolayer culture may not be representative of tissue conditions, 3D culture systems better mimic in vivo cellular conditions. Second, in vitro results provide a springboard for in vivo concepts and ultimately clinical applications. Finally, in vitro culture will probably have a role in clinical therapy as cells are expanded ex vivo, similar to the model of autologous chondrocyte transplantation. This section will review cellular-based cartilage tissue engineering in vitro by both differentiated chondrocytes and undifferentiated progenitor cells (MSCs, ES cells, and EG cells).

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In addressing the current state of in vitro research for each cell type, we will describe effective biological signals and scaffolds, the limitations of each cell type, and potential directions of the field. Chondrocytes are readily available as they can be isolated from human cartilage and cultured ex vivo. For close to a decade, chondrocytes have been expanded ex vivo for clinical applications as an FDA approved therapy (Brittberg et al., 1994). There appears to be age-dependent differences in chondrocyte tissue forming capabilities and proliferation. Much of this data is in animals cells, as Tran-Khanh et al. demonstrated aged bovine chondrocytes to have decreased proliferation and ECM production when compared to juvenile and fetal bovine chondrocytes (Tran-Khanh et al., 2005). Furthermore, one of the major limitations of chondrocytes is a tendency to rapidly dedifferentiate in monolayer culture (Darling et al., 2004; Darling and Athanasiou, 2005a). Studies have demonstrated that plating chondrocytes at higher cell densities, medium components, using 3D scaffold materials, or adding growth factors and mechanical stimulation helps retain phenotype (Freed et al., 1993; Arevalo-Silva et al., 2001; Pei et al., 2002a; Vunjak-Novakovic et al., 2002; Darling and Athanasiou, 2003; Mauck et al., 2003; Shieh and Athanasiou, 2003; Kamil et al., 2004; Darling and Athanasiou, 2005b). Tissue culture material and scaffold type can also influence chondrocyte phenotype. In vitro monolayer expansion of chondrocytes demonstrated greater chondrocytic gene expression in aggrecan-coated polystyrene than collagen II-coated polystyrene and basic tissue culture dishes (Darling and Athanasiou, 2005a). 3D culture scaffolds promote chondrogenesis by producing increased collagen II and decreased collagen I compared to chondrocytes in monolayer culture (Benya and Shaffer, 1982; Freed et al., 1993). Darling and Athanasiou showed improved results when growing chondrocytes in a 3D agarose scaffold (Darling and Athanasiou, 2005b). They suggested aggrecan and agarose help maintain the round shape that associated with cell synthesis mode while a flat shape is associated with expansion in fibrocartilage phenotype (Darling and Athanasiou, 2005b). Altering medium conditions, especially through the addition of growth factors, has a substantial influence on preserving chondrocyte phenotype. Chondrocytic medium traditionally includes animal or human serum as a standard component with concentration ranging between 2% and 20% (Nerum, 1991). Sera includes hormones and growth factors that enhance cell survival and proliferation. However, lot to lot variation between serum batches has the potential to alter experimental results. There are reports of growth factors in serum resulting in chondrocyte dedifferentiation and even having a toxic effect on cells (Nerum, 1991; Malpeli et al., 2004). Optimizing serum free medium for in vitro chondrocyte expansion may be accomplished by including individual serum components like selenium, insulin, and transferrin as well as a balance of growth factors (Chua et al., 2005). Lastly, reusing growth medium may increase expansion speed. Kamil et al. speculated that the retention of smaller and more immature cells within the otherwise discarded medium accounts for the more rapid population doublings (Kamil et al., 2004a). Tissue engineers have experimented with many growth factor effects on chondrocytes. The transforming growth factor-beta (TGF-β) superfamily consistently promote chondrocyte phenotype and proliferation in monolayer and 3D culture (Arevalo-Silva et al., 2001; Elisseeff et al., 2001; Koch and Gorti, 2002; Pei et al., 2002a; Richmon et al., 2005). Other common factors that show promising results include insulin-like growth factor 1 (IGF-1) and the fibroblast growth factor (FGF) family. Arevalo-Silva et al. demonstrated that basic FGF increased proliferation of elastic chondrocytes (Arevalo-Silva et al., 2001). In a separate study, Richmon et al. concluded that BMP-2 and IGF-1 did not influence chondrocyte proliferation in monolayer culture, however were effective in retention of chondrocytic phenotype of nasal septal chondrocytes (Richmon et al., 2005). Sequential administration of growth factors may provide better regulation. Pei et al. showed early supplementation with TGF-β1 and FGF-2 doubled cell proliferation rates with subsequent addition of IGF-1 increasing ECM fractions (Pei et al., 2002a). As scientists discover more about the physiological signaling of chondrocytes, it is likely that timing and sequence of administration will enhance the retention of chondrocyte phenotype and expansion in vitro.

Cartilage Tissue Engineering

Biological signals may be directly incorporated into a chondrocyte by genetic transduction (Trippel et al., 2004). In one study of chondrocyte transduction, Grande et al. demonstrated chondrocyte transfection with BMP-7 or sonic hedgehog genes enhanced quality of cartilage produced (Grande et al., 2003). Research into gene transfer of TGF-β, bone morphogenic proteins, insulin-like growth factor-1, other growth factors, Sox family transcription factors, and SMAD signal transduction molecules is promising due to their role in chondrogenesis and maintenance of chondrocyte phenotype (Trippel et al., 2004). In vitro transduction could potentially improve cartilage repair with autologous chondrocyte transplant. To prove genetic transfer improves long-term cartilage repair future studies will need to compare the outcome of gene transfer to that of direct administration of growth factors. Mechanical stimulation influences chondrocyte morphology, biochemistry, biomechanical, and electrochemical properties when compared to chondrocytes grown in a static environment (Vunjak-Novakovic et al., 1999, 2002). 3D culture of chondrocytes in bioreactors creates an isolated in vitro system to measure the effect of dynamic mechanical loading on cartilage formation (Demarteau et al., 2003). In 1999, Vunjak-Novakovic et al. demonstrated that dynamic laminar flow patterns on chondrocytes grown on a polyglycolic acid (PGA) scaffold resulted in higher fractions of collagen and glycosaminoglycan as well as improved mechanical and electromechanical properties when compared to chondrocytes grown in static culture or turbulent flow conditions (VunjakNovakovic et al., 1999). Furthermore, bioreactors may be programmed to create both shear and compression forces. The principle effect of moderate mechanical loads are thicker cartilage due to increased synthesis of collagen and proteoglycans (Darling and Athanasiou, 2003; Shieh and Athanasiou, 2003; Waldman et al., 2003). In contrast, high shear conditions promote apoptosis in chondrocytes resulting in matrix degradation (Healy et al., 2005). In one study, Kisiday et al. used an alternating day mechanostimulation on chondrocytes to increase proteoglycan accumulation (Kisiday et al., 2004). Waldman et al. compared shear with compression stimulation and demonstrated a greater effect on ECM molecule synthesis with shear forces (Waldman et al., 2003). Increased ECM translates to an increased load-bearing capacity and stiffness by the cartilage construct (Mauck, 2003; Waldman, 2003). Mechanical forces affect chondrocytes via a variety of potential mechanisms. Benya theorized that dynamic stimulation maintains chondrocyte round cell shape which promotes protein synthesis (Benya and Shaffer, 1982). Shieh and Athanasiou propose a number of cellular methods for mechanotransduction: (1) transduction of biochemical signals in chondrocytes by changing physicochemistry including the osmotic pressure, pH, fluid flow, and electric potential of the matrix environment; (2) conformational change of ion pumps and channels; (3) via integrins which mediate cell–matrix interaction; (4) nuclear deformation changes the nuclear pore complex influencing DNA available for transcription; and (5) deformation of cytoskeleton (Shieh and Athanasiou, 2003). Further clarification of the signaling pathways involve in mechanotransduction should yield better control of cartilage repair. Future investigation into the types and amounts of forces is needed to better define the limitations of mechanical stimulation (Shieh and Athanasiou, 2003). Combining mechanical stimulation of chondrocytes with growth factor administration enhances matrix synthesis to a greater degree than either variable alone (Darling and Athanasiou, 2003; Mauck et al., 2003). Mauck demonstrated that dynamic deformational loading combined with TGF-β1 or IGF-1 increased ECM production in a 3D scaffold (Mauck et al., 2003). The synergy may be a result of mechanical compression causing increased transport and accessibility of growth factors (Shieh and Athanasiou, 2003). Just as sequential administration of growth factors increases chondrocyte proliferation, the investigation of mechanical stimulation and growth factor administration timing could yield future advancements in cartilage engineering (Darling and Athanasiou, 2003; Shieh and Athanasiou, 2003). MSCs represent a viable alternative to chondrocytes as a cell source for cartilage tissue engineering (Pittenger et al., 1999). MSCs have the advantage of being able to be expanded in vitro in an undifferentiated state, while retaining the ability to differentiate after exposure to suitable stimuli (Song et al., 2004). In contrast to the differentiated phenotype of chondrocytes, stem cells require the complexity of additional steps to initiate chondrogeneic differentiation and then maintain chondrogeneic phenotype without hypertrophy or

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dedifferentiation (Trippel et al., 2004). Monolayer and 3D seeding requires a higher density of stem cells, which is similar to the density of cells needed for embryonic limb bud development (Tuan et al., 2002; Song et al., 2004). With regard to medium conditions, most MSC culture uses serum free medium in contrast to chondrocyte culture (Tuan et al., 2002). As described earlier serum free medium provides a secure and reproducible in vitro environment though it may slightly decrease proliferation. The type of MSC may also influence proliferation. In a 2003 study, Winter et al. compared chondrogeneic gene expression and morphology from MSCs derived from bone marrow and adipose tissue (Winter et al., 2003). The study demonstrated similar partial differentiation in monolayer culture, however bone marrow-derived MSCs improved chondrogenesis in 3D culture (Winter et al., 2003). This study’s results combined with 3D culture results and established MSC isolation techniques resulted in the majority of research using bone marrow-derived MSCs. Lastly, one limitation of MSC proliferation in vitro is that these cells do not remain in the undifferentiated state forever. Most MSC populations lose their ability to differentiate and form tissues by 40–50 population doublings (Stenderup et al., 2003). A 3D culture environment for cartilage engineering with MSCs is superior to monolayer culture. This parallels 3D chondrocyte culture, as cell-to-cell contact promotes MSC acquisition of chondrogeneic phenotype. Culture pellets, natural and synthetic scaffolds have all proved successful, though synthetic scaffolds provide more uniform properties and purity (Sekiya et al., 2002; Tuan et al., 2002; Raghunath et al., 2005; Wang, et al., 2005).Wang et al. developed a highly porous 3D silk scaffold for MSC chondrogenesis that yielded cellular spacing and collagen type II distribution similar to that of native articular cartilage tissue (Wang et al., 2005). Composite scaffolds include pellet culture in polylactic acid (PLA) polymer and alginate–PLA – where the alginate improves cell loading and retention and the PLA provides mechanical support and stability (Caterson et al., 2001; Tuan et al., 2002). Li et al. manufactured an electrospun nanofibrous polylactic-co-glycolic acid (PLGA)/poly-e-caprolactone amalgam to better mimic the natural extracellular architecture of cartilage (Li et al., 2002). These composite scaffolds have the potential to be modified with natural or synthetic materials to improve growth factors delivery (Li et al., 2002). Studies consistently demonstrate that TGF-β promotes chondrogeneic differentiation in MSCs. Barry et al. compared TGF-β1, -β2, and -β3 and demonstrated that TGF-β2 and -β3 produced significantly more proteoglycans and collagen II than TGF-β1 (Barry et al., 2001). It should be stressed that all TGF-βs have a differential and proliferative effect on MSCs and their differences are negligible in vitro. Given the consistent effect of TGF-β on MSCs, studies of other growth factors are measured by the growth factor’s combined effect with TGF-β. BMPs have had mixed results, though Sekiya et al. demonstrated increased chondrogeneic differentiation by immunohistochemistry and weight of cartilage constructs when combining different BMPs with TGF-β3 (Sekiya et al., 2002, 2005; Indrawattana et al., 2004). Indrawattana et al. concluded that combining and cycling TGF-β3 with BMP-6 or IGF-1 improved chondrogenesis at induction as determined by gene expression of collagen II and aggrecan (Indrawattana et al., 2004). Evidence with FGF ligands and receptors are mixed (Raghunath et al., 2005). In one of the few MSC transfection studies, Palmer et al. showed that adenoviral-mediated expression of TGF-β1 or BMP-2 improved chondrogenesis (Palmer et al., 2005). Further investigation into the timing and synergistic effect of growth factors should yield a better understanding of MSC chondrogenesis. Additionally, identifying the signaling cascades initiated by growth factors represents an important area of investigation. Signaling pathways in MSC chondrogenesis may be elucidated through the study of limb development (Sekiya et al., 2002). The SMAD and mitogen-activated protein kinase (MAPK) cascades mediate TGF-β effects in MSC chondrogenesis (Tuan et al., 2002). A progressive increase in the WNT11 signaling molecule through 14 day culture in chondrogenic media led Sekiya et al. to identify it as a late factor MSC chondrogenesis. Though WNT signaling has undergone extensive study, further research into signaling cascades is needed to better elucidate MSC chondrogenesis and increase cartilage production (Sekiya et al., 2002; Tuan et al., 2002; Raghunath et al., 2005). Other genes that require more study include sonic hedgehog and Sox transcription factors (Tuan et al., 2002; Raghunath, et al., 2005).

Cartilage Tissue Engineering

ES chondrogenesis is a relatively new area of research. Kramer et al. first studied ES chondrogenesis in 2000 (Kramer et al., 2000). ES have demonstrated an ability to retain their undifferentiated state for up to 147 passages in vitro (Cowan et al., 2004). As with chondrocyte and MSC culture, 3D culture appears to enhance chondrogenesis of ES cells. Tanaka et al. demonstrated 3D micromass pellet culture enhanced chondrogeneic gene expression when compared to monolayer culture (Tanaka et al., 2004). Hwang et al. showed an upregulation of chondrogeneic transcription factors, Sox9 and cbfa1, when cultured in a poly(ethylene glycol) PEG hydrogel (Hwang et al., 2005). TGF-β superfamily signaling factors are effective at promoting ES chondrogenesis in 3D culture. Numerous studies have showed TGF-β1 to induce a strong chondrogeneic response in 3D culture (Levenberg et al., 2003; Nakayama et al., 2003). On the other hand, researchers found that BMP-2 enhances chondrogenesis in monolayer culture (Hwang et al., 2005; zur Nieden et al., 2005). The few studies investigating ES chondrogenesis demonstrate a field in its infancy that is currently establishing standard experimental methods. EG cell chondrogenesis represents a new area of investigation that developed as an alternative to ES cells (Kim et al., 2005). Kim et al. demonstrated musculoskeletal differentiation of EG in 3D culture micromass pellet culture by osteochondral gene markers (Kim et al., 2005). The addition of TGF-β3 and TGF-β1 increases proteoglycan accumulation in the pellets (Kim et al., 2005; Ferran and Elisseeff, unpublished data). Growth factors that influence EG in monolayer culture include TGF-β1 and BMP, both of which enhance cartilage-specific genes (Ferran and Elisseeff, unpublished data). Recent unpublished data from our laboratory demonstrated chondrogeneic differentiation by histology when EG cells were clustered in hydrogels and cocultured adjacent to chondrocytes (Ferran and Elisseeff, unpublished data). While EG cells are not immortal they retain marker expression for up to 70 population doublings in vitro (Shamblott, 2001). Research into embryonic progenitor cells may have an advantage based on previous results studying chondrocytes and MSCs, at least with respect to expansion capacity. However, extensive research into ES and EG chondrogenesis is necessary to define scaffold materials, growth factor effects, and signaling pathways.

BIOMATERIALS IN CARTILAGE REPAIR Overview As outlined above, a great deal of effort has been expended to identify the optimum cellular components with which to engineer functional cartilage, and much remains to be done. However, cells in vivo do not exist in isolation, but rather are intimately associated with an ECM. This matrix supports and regulates cellular proliferation, differentiation, and ultimately function (Benya and Shaffer, 1982; Aulthouse et al., 1989). The ECM is particularly important in the case of tissues such as bone or cartilage, as the structural properties of those tissues is highly dependent on the precise structure and organization of the ECM. In the case of engineered tissues, the scaffold functions as the ECM until it is replaced by native tissue. It may serve as the vehicle by which cells are delivered to heal the defect, or it may serve to recruit and support the ingrowth of tissue into the scaffold. It provides the primary strength of the engineered tissue, as well as the means by which the tissue is retained in the repair site. Furthermore, the scaffold may actively support healing and tissue formation by incorporating growth factors and adhesion molecules (Chenite et al., 2000; Fortier et al., 2002). Therefore, research into cartilage engineering must focus as much on the scaffold materials as it does on the cells that reside within those scaffolds. There are a number of characteristics that make up the “ideal” scaffold for cartilage engineering. It must be able to support (or at least not hinder) cellular functions. More specifically, the scaffold and any breakdown products must not be cytotoxic. They should not incite a marked immunological reaction or interfere with wound healing. It should be in a format that is easily handled and implanted. The scaffold should allow for

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adequate waste and nutrient transport. It should have sufficient structural integrity for its intended purpose (e.g. joint reconstruction), and ideally should degrade at the same rate at which it is being replaced by new tissue so as not to compromise the quality of the repair. Finally, it should be able to be produced with sufficient purity in commercial quantities with reproducible and consistent properties. Ultimately, the goal is to design an engineered tissue that can be used to repair defects, while stimulating and supporting the patient’s ability to heal defects with tissue that should be indistinguishable from native healthy tissue. To date, no scaffold material fulfills all these criteria, and as such there is considerable ongoing work in biomaterials research. It would be nearly impossible to summarize the vast body of this research in a single chapter, so what follows is a necessarily brief summary of selected topics. Autologous Scaffolds Early attempts at supporting the healing of injured cartilage focused on using autologous tissue, and some of these techniques are in clinical use at this time. The range of techniques include periosteal and perichondral grafts, autologous chondrocyte transplantation, and osteochondral autograft (also known as mosaicplasty). Each of these tissue sources has its advantages and disadvantages. Periosteal and perichondral lining tissues are ready-made sources of osteoprogenitor and chondroprogenitor cells in a natural scaffold (Homminga et al., 1990; Moran et al., 1992; O’Driscoll et al., 1994). These tissues are easily harvested and handled (Homminga et al., 1990; Moran et al., 1992). When implanted into a defect, they form a tissue which has some characteristics of cartilage, but does not usually heal the defect completely (Frenkel and Di Cesare, 2004). The quality of the tissue formed was variable, and it was noted that the grafts had a tendency to detach or ossify. Most importantly, age seemed to be a significant predictor of graft failure, with no patient over 40 experiencing a good clinical result (Seradge et al., 1984). Attempts to modify the procedure by recruiting bone marrow cells via subchondral drilling, or by adding other scaffold materials such as demineralized bone have shown similarly variable results (van Osch et al., 1999; Hunziker, 2002). Currently, only autologous chondrocyte transplantation and osteochondral autografting are widely accepted and approved techniques for articular cartilage regeneration (Lynn et al., 2004). Both techniques are associated with good outcomes. Autologous chondrocyte transplantation is a technique in which chondrocytes are harvested from an uninjured, non-weight-bearing region of the joint, expanded ex vivo, and reimplanted into a chondral defect in the weight-bearing area of the joint. The chondrocytes, which are not contained within a scaffold, are retained in place by a periosteal flap that is sutured to the surrounding cartilage (Brittberg et al., 1994). This technique is associated with hyaline-like repair tissue that is maintained up to 10 years post-operatively, with a reported good to excellent outcome rate of 60–70% (Minas, 2001). Disadvantages of the technique include donor site morbidity, limited supply of healthy autologous tissue, the difficulty of suturing the periosteal flap to the cartilage surrounding the implant, and the potential failure of dedifferentiated cultured chondrocytes (von der Mark et al., 1977) to regain a chondrocytic phenotype after implantation. Osteochondral autografting or mosaicplasty is a technique of autotransplantation in which osteochondral plugs are harvested from non-weight-bearing or low-weight-bearing regions of the joint and implanted into defects that have been prepared and sized. Survival of hyaline cartilage has been reported in 85% of patients, with a 91% good to excellent clinical outcome reported by patients followed for 3–6 years (Hangody et al., 1998, 2001). However, this technique suffers from an even greater limitation on tissue supply than autologous chondrocyte transplantation. There is evidence for potential long-term donor site morbidity (Hunziker and Quinn, 2003). Animal models suggest that the grafts may not survive in the long term (Hurtig et al., 2001). This may be due to direct trauma sustained by the grafts during harvest or implantation, or by long-term exposure

Cartilage Tissue Engineering

of the grafts to supraphysiological loads (i.e. transfer of tissue from an area of low-weight-bearing to an area of high-weight-bearing) (Lynn et al., 2004). Natural Scaffolds Collagen Collagen is the primary structural protein found in both bone and cartilage. As such, collagen-based scaffolds are theoretically capable of supporting chondrocyte attachment and function. They are also biocompatible and biodegradable. Collagen scaffolds have been used in a wide variety of forms such as gels, membranes, and sponges into which cells and/or bioactive factors may be introduced (Frenkel and Di Cesare, 2004). For example, Yokoyama et al. recently cultured MSCs in a collagen gel matrix in a chondrogeneic medium supplemented with BMP-2, TGF-β3, and dexamethasone (Yokoyama et al., 2005). Grossly, the collagen gels transformed from clear and fragile at day 0 to firm and opaque at day 21. Biochemically, the constructs were characterized by progressive increases in the amounts of type II collagen, chondroitin 4-sulfate, and chondroitin 6-sulfate secreted by cells. Polymerase chain reaction (PCR) analysis demonstrated downregulation of type I collagen, as well as upregulation of type II collagen and the cartilage-related proteoglycans aggrecan, biglycan, and decorin. The maximum size of cartilaginous tissue produced was 7 mm diameter  0.5 mm thick. A novel use of a collagen gel matrix was recently reported by Nawata et al. (Nawata et al., 2005). In this experiment, muscle-derived mesenchymal cells were harvested from rat embryos, cultured in monolayer, mixed with human BMP-2 and type I collagen, and packed into diffusion chambers, which were then implanted in adult rats under the abdominal fascia. The gels were then harvested and subjected to histological, biochemical, and molecular analysis. This analysis showed evidence of chondrogeneic differentiation and production of what appeared to be mature cartilage after 5 weeks of implantation. The 5-week constructs were then implanted into a full-thickness patellar groove defect, with reported full healing of those defects at 6 months post implantation. Collagen has also been used as a vehicle for bioactive factors alone, without cell transplantation. Sellers et al. describe the implantation of a type I collagen sponge impregnated with BMP-2 into a full-thickness osteochondral defect created in rabbits. At 1 year, the defects treated with BMP-2 appeared significantly better histologically, although the tissue was still not identical to normal cartilage (Sellers et al., 2000). The cell-based studies also highlight some of the disadvantages of collagen scaffolds. Collagen gels allow for uniform mixing of cells and matrix, and for extensive molding and shaping of tissue, but tend to be fragile until new matrix is laid down. Solid collagen scaffolds such as membranes or sponges exhibit greater initial mechanical strength, but at the cost of less flexibility in shaping and a greater risk of non-uniform cell seeding. Collagen remains a useful scaffold with which to study 3D cell culture, but the disadvantages noted above weigh against its use in clinical applications. Alginates Alginates are polysaccharides derived from seaweed. They comprise a family of linear mannuronate/guluronate copolymers that differ in their specific sequences and overall compositions (Rowley et al., 1999). When exposed to a divalent cation (usually calcium for sake of biocompatibility), the linear alginate polymers ionically crosslink to form a porous hydrogel. This allows the uniform seeding of chondrocytes and bioactive factors within the alginate hydrogel, as well as their release, if desired, by exposure to a cation chelating agent such as EDTA. Alginates may also be covalently modified in order to enhance properties such as cell adhesion (Sultzbaugh and Speaker, 1996; Rowley et al., 1999; Alsberg et al., 2001) or to fix bioactive factors in place (Suzuki et al., 2000; Gerard et al., 2005). Interestingly, alginates are being studied not only as an in vivo scaffold, but also as an in vitro culture medium (Diduch et al., 2000). Most cartilage engineering efforts using mature chondrocytes rely on monolayer expansion to obtain an adequate population for study and reimplantation. Alginate matrices may provide a means to preserve or reestablish characteristic chondrocyte phenotype and matrix production even during in vitro expansion (Homicz et al., 2003; Chia et al., 2005).

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Chitosan Chitosan is a polysaccharide, this time derived from chitin (found in arthropod exoskeletons) that has been partially or fully deacetylated (Chenite et al., 2000). It is composed of linear chains of β-linked D-glucosamine residues. Chitosan has been studied as both a scaffold and as a controlled delivery system for bioactive factors (Lee et al., 2004; Hoemann et al., 2005). There is interest in chitosan as a cell delivery vehicle as it demonstrates good biocompatibility, and some formulations exhibit the property of temperature-dependent gelation, in that they are liquid at room temperature but gel when exposed to physiological temperatures (Chenite et al., 2000). Thermosetting chitosan constructs injected subcutaneously into nude mice supported chondrocyte growth and matrix production, although the constructs were mechanically inferior to native cartilage (Hoemann et al., 2005). Interestingly, the degree of deacetylation of chitosan directly influences the degradation rate of the constructs as well as the inflammatory response. A lower degree of deacetylation was associated with an increased degradation rate and host inflammatory response (Chenite et al., 2000). Chitosan constructs were also injected into osteochondral defects created in rabbit knees. Retention of the constructs in the defects was observed at 1 week despite full mobility and weight-bearing (Hoemann et al., 2005). Hyaluronic Acid HA is a polysaccharide that is naturally found in both the ECM of articular cartilage and in synovial fluid. It is composed of alternating residues of N-acetyl-D-glucosamine and D-glucoronic acid (Nettles et al., 2004). As with collagen, interest focused on HA as a potential scaffold for cartilage engineering due to its intimate association with chondrocytes in vivo. Intra-articular HA injection has been used to treat symptoms of osteoarthritis, and HA has been shown to have a stimulatory effect on chondrocyte production of type II collagen and proteoglycan (Akmal et al., 2005). A novel use of HA was reported in which HA was modified by methacrylation to form a photocrosslinkable polymer, which was then used to encapsulate chondrocytes for in vitro and in vivo culture. Chondrocytes encapsulated within this matrix retained their phenotype and generated type II collagen (Nettles et al., 2004). Another study evaluated a chitosan–alginate–hyaluronate scaffold modified with a protein containing an RGD adhesion peptide motif. Chondrocytes were grown on chitosan–alginate scaffolds with and without HA and the RGD-containing protein. It was noted that glycosaminoglycan and collagen synthesis was greater in the chitosan–alginate–HA–RGD scaffold as compared to chitosan–alginate and chitosan–alginate–HA scaffolds (Hsu et al., 2004). Synthetic Scaffolds Biological scaffolds are an attractive option for tissue engineering applications due to their inherent biocompatibility. However, biological materials are notoriously difficult to generate in large quantities with acceptable consistency, and as noted above often exhibit poor mechanical characteristics (Frenkel and Di Cesare, 2004). Also, there is always the possibility of immune responses or disease transmission with materials derived from biological sources. Synthetic materials generally avoid these problems. These scaffolds may be precisely engineered and tailored with respect to mechanical, chemical, and biological properties, and may be produced with a great deal of batch-to-batch consistency. The most common synthetic polymers in use at this time are PGA, PLA, polyethylene oxide (PEO), and various derivatives and copolymers based on these entities (Frenkel and Di Cesare, 2004). These biodegradable polymers have a long history of medical usage, and are able to be fabricated and processed in a variety of ways (Sharma and Elisseeff, 2004). These materials provide for the adherence, growth, differentiation, and matrix production by chondrocytes or MSCs (Lu et al., 2001; Riley et al., 2001; Frenkel and Di Cesare, 2004; Lynn et al., 2004; Klein et al., 2005). In general, these materials exhibit many properties ideal for the production of engineered tissue: a high surface area to volume ratio if processed correctly, sufficient porosity to allow for nutrient and waste diffusion, the potential for surface modification, and the ability to control their degradation

Cartilage Tissue Engineering

rate via selection and modification of their chemical composition (Muschler et al., 2004). In particular, the ability to specifically control the rate of degradation is important. First, the scaffold must provide sufficient mechanical strength when first implanted, but should ultimately degrade to allow for replacement by growing tissue. If degradation is too rapid, then there is a risk of cell loss, scaffold failure, and inflammation of surrounding tissue due to rapid release of acidic breakdown products (Lu et al., 2001). Conversely, an overly slow rate of scaffold degradation would likely impede tissue incorporation. The synthetic polymers mentioned above, despite their proven biocompatibility, still form foreign bodies in vivo. Furthermore, the internal microstructure of these polymers does not closely resemble that of the ECM (Kisiday et al., 2002). Novel polymers based on self-assembling synthetic peptides have been studied as a potential scaffold for cartilage tissue engineering. These peptides spontaneously form hydrogels in response to changes in their environment, such as alterations in pH or ionic strength (Kisiday et al., 2002). The microstructure of these hydrogels is approximately three orders of magnitude smaller than that of conventional polymers, and more closely approximates the structure of native ECM (Kisiday et al., 2002). These materials have the potential for extensive modification by incorporation of peptide domains that influence cell adhesion, differentiation, and proliferation (Holmes, 2002). 3D culture of chondrocytes in peptide hydrogels results in maintenance of chondrocyte phenotype and secretion of cartilage-specific matrix, with increased proliferation and improved mechanical characteristics as compared to chondrocytes cultured in agarose (Kisiday et al., 2002, 2004). Synthetic scaffolds have been synthesized in a variety of configurations, from preformed fibers, meshes, and membranes, to photopolymerized injectable gels. Preformed solid scaffolds are seeded in vitro by incubation in a cell suspension. These scaffolds may be applied to large, shallow, or open defects. However, obtaining suitable cell densities and uniform cell seeding continues to be a challenge (Lu et al., 2001). A considerable amount of recent work, by our group and others, has focused on liquid polymer solutions that are polymerized or crosslinked in situ after incorporation of cells and bioactive factors. Such solutions allow for uniform incorporation of cells throughout the scaffold. Their liquid form allows for development of minimally invasive application techniques (Sims et al., 1996; Elisseeff et al., 1999a, b; Xu et al., 2004). Finally, these in situ polymerizable solutions offer the possibility of precise control of the final shape and composition of the scaffold. For example, there has been considerable work by our group and others studying bilayered constructs in which one layer contains MSCs and the other contains chondrocytes, in order to approximate the cell–cell interactions that would occur in native tissue. Our group has also investigated recapitulating the zonal architecture of native cartilage using sequentially photopolymerized hydrogel layers (Elisseeff et al., 1999a, b; Mercier et al., 2004; Nettles et al., 2004; Alhadlaq and Mao, 2005), to generate an engineered tissue that more closely approximates normal cartilage.

SUCCESSFUL IN VITRO CARTILAGE ENGINEERING There have been a number of challenges that have been identified with engineering cartilage replacements in the laboratory. One is that mature chondrocytes have a limited potential for expansion in vitro. Furthermore, articular chondrocytes have a tendency to dedifferentiate in monolayer tissue culture, and fail to synthesize cartilaginous matrix. The observed dedifferentiation may be overcome by the addition of various growth factors and drugs to the culture medium, but the degree of redifferentiation is variable (Homicz et al., 2003; Chia et al., 2005). Use of a multipotential cell type such as MSCs overcomes the issue of limited expansion, but requires specialized culture conditions to allow for differentiation into chondrogeneic cells capable of elaborating cartilaginous matrix. Successful in vitro cartilage engineering, then, depends on maintenance of a proper state of differentiation of the cultured cells as well as production of cartilage-specific matrix. A number of groups have reported chondrocyte proliferation, phenotype maintenance, and cartilage matrix production using a variety of scaffold materials, including alginate (Almqvist et al., 2001), collagen (Paige et al., 1996; Yokoyama et al., 2005), PGA/PLA and copolymers (Baek et al., 2002), and PEO (Riley et al., 2001; Alhadlaq et al., 2004; Sharma and

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Elisseeff, 2004). In most of these studies, the cell–scaffold constructs were subjected to biochemical, histological, and/or mechanical testing to support the hypothesis of in vitro cartilage generation (i.e. increased DNA content, presence of type II collagen and aggrecan, absence of type I collagen, etc.). Ultrastructural studies are somewhat less common, but seem to indicate that engineered tissue at least approximates native cartilage (Adkisson et al., 2001; Iwasaki et al., 2004; Yokoyama et al., 2005). At this point, cartilage-like tissue is able to be reliably generated in vitro in a variety of scaffold materials, using chondrocytes and/or MSCs. However, the mechanical properties of these engineered tissues indicate they are inferior to native cartilage, and would be unlikely to survive long term in a load-bearing application (Lynn et al., 2004). Current research is focusing on optimizing the materials and culture conditions to generate engineered tissues with improved mechanical characteristics. Mechanical Stimulation One area of active research is the effect of mechanical stimulation on chondrocyte growth and function. A fundamental characteristic of many musculoskeletal tissues is their responsiveness to mechanical stimuli (Hung et al., 2004). Articular cartilage is subject to complex forces through its range of motion, including shear, compression, and hydrostatic pressure, and it is reasonable to expect that those forces affect the growth and functioning of chondrocytes. This observation has been borne out by a number of in vitro studies. Hung et al. demonstrated that mechanical stimulation affected gene expression and biochemical and mechanical properties of bovine articular chondrocytes cultured in agarose. It was shown that the effect was proportional to the frequency of stimulation (which was varied between 0.005 and 1 Hz) and was synergistic with TGF-β1 and IGF-1 (Hung et al., 2004). Elder et al. demonstrated enhanced chondrogenesis of chick embryo mesenchymal cells in agarose culture when subjected to cyclic loading at a frequency between 0.15 and 0.33 Hz (Elder et al., 2001). Prior groups had noted a similar stimulatory effect of moderate dynamic stress applied to cultured chondrocytes, within a range of 0.1–1 Hz. Static compression tended to have an inhibitory effect, and dynamic compression frequencies outside the range of 0.1–1 Hz tended to be inhibitory or have no effect (Buschmann et al., 1995; Lee and Bader, 1997; Mauck et al., 2000; Mizuno et al., 2002). There are multiple theories to explain this observed effect, including direct cell deformation, alteration of cellular microenvironment, alteration of cell–matrix interactions, and enhanced mass transport within the matrix (Darling and Athanasiou, 2003). Although still unclear, it appears that enhanced mass transport plays a large role in the observed effects of dynamic compression (Mauck et al., 2003). Bioreactors One of the key observations regarding in vitro cartilage growth is the need for high cell densities within the scaffold, on the order of 20 to 100  106 cells/ml (Lu et al., 2001). The high cell density coupled with the need to maintain the cells in a 3D construct lead to potential problems with nutrient and waste transport, particularly as the constructs get larger and more matrix is deposited. Static culture which relies on passive diffusion of nutrients and waste may be inadequate to serve the needs of metabolically active tissue. There has been considerable interest in bioreactor culture of these constructs, which may help alleviate some of the mass transport issues that arise with engineered tissues (Lu et al., 2001). Furthermore, there may be a direct stimulatory effect by flowing media on the cultured cells, mediated by hydrostatic pressure and shear forces. Pei et al. cultured bovine chondrocytes in a variety of preformed scaffolds in static conditions and in a rotating bioreactor system (Pei et al., 2002b). Uniformity of cell seeding, cell numbers, mechanical properties, and molecular and biochemical markers of chondrogenesis were analyzed. Constructs cultivated in the bioreactor system demonstrated more uniform cell seeding, greater cell numbers, and enhanced chondrogenesis when compared with their static counterparts. Raimondi et al. utilized a novel forced-perfusion bioreactor system in order to expose the inner portions of their chondrocyte constructs to bulk fluid flow and hydrodynamic stresses, as opposed to the rotating bioreactors which expose the surface only to fluid stresses and convective mass transport (Raimondi et al., 2002). Analysis of

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the constructs, which included scanning electron microscopy, demonstrated cell loss as well as loss of structural integrity of the constructs incubated in static conditions. In comparison, constructs cultivated in the bioreactor demonstrated greater cell proliferation and better structural integrity.

TRANSLATION OF CARTILAGE TISSUE ENGINEERING The current state of in vitro research into cartilage engineering is promising. Improvements in culture conditions and scaffold materials, incorporation of bioactive factors, and recognition of the role of mechanical stimulation in cartilage formation have yielded results that look much like cartilage (Elisseeff et al., 1999b; Alhadlaq and Mao, 2003; Alhadlaq et al., 2004; Kamil et al., 2004; Alhadlaq and Mao, 2005). However, in vivo testing provides necessary information as to the ultimate clinical applicability of a given engineered tissue. In addition, novel techniques are being developed to construct complex tissues suitable for clinical use. Initial studies focused on the ability of chondrocyte-seeded scaffolds to generate cartilaginous tissue in vivo after subcutaneous implantation. These studies served to show that, in general, the in vivo environment was not hostile to the development of cartilage and osteochondral tissues. Elisseeff et al. utilized a photopolymerizable PEO-based hydrogel that exists in liquid form prior to polymerization with ultraviolet light (Elisseeff et al., 1999a, b). Isolated bovine articular chondrocytes were suspended in this polymer solution and injected subcutaneously into athymic mice. The mice were then exposed to ultraviolet light which resulted in the formation of a solid, chondrocyte-seeded hydrogel. After a 6 week in vivo incubation the constructs were harvested and subjected to histological and biochemical testing. The implanted constructs demonstrated evidence of fibrocartilaginous tissue production, with partial chondrocyte dedifferentiation, mixed type I and II collagen deposition, and chondroitin sulfate deposition. There was evidence of cellular proliferation, without any sign of necrosis. Control hydrogels without cells did not demonstrate any matrix deposition, suggesting that the chondrocytes in the hydrogel were responsible for the new tissue formation. A similar experiment was performed by Westreich et al., in which rabbit ear chondrocytes were isolated, suspended in fibrin glue, and injected subcutaneously back into the rabbit. After incubation, the constructs were removed and analyzed. There was cartilage formation in 85% of the samples, although the quality of the cartilage varied markedly (Westreich et al., 2004). Haisch et al. describe their efforts to produce tissue-engineered auricular cartilage for ear reconstruction (Haisch et al., 2002). Isolated nasal septal chondrocytes were expanded in monolayer culture, then suspended in fibrinogen in order to provide for uniform cell distribution and applied to a PLA–PGA fleece scaffold molded in the shape of a human auricle. The fibrinogen was then polymerized by addition of thrombin, and the resulting 3D constructs were cultured in vitro and in vivo. As compared to in vitro culture, the constructs incubated in vivo demonstrated increased matrix production, with a macroscopic and microscopic appearance consistent with hyaline cartilage, once again demonstrating the feasibility of generating cartilage-like tissues in vivo. A similar experiment was performed by Kamil et al., in which sheep auricular chondrocytes were harvested and seeded onto three different scaffold materials (alginate, Pluronic, and PGA) (Kamil et al., 2004). The scaffolds were then placed into an auricle-shaped hollow gold mold and implanted subcutaneously into the sheep. They were then harvested and analyzed for gross appearance and mechanical properties as well as histology. The cartilage formed was of variable quality, depending on the scaffold used. The Pluronic scaffold generated cartilage with the best histological appearance, while the alginate scaffold retained the most anatomic detail from the mold. The PGA scaffold did not produce useful cartilage. The challenges of generating cartilage for craniofacial reconstruction lie primarily in the generation of complex shapes and retention of anatomic detail. Structural cartilage such as laryngotracheal or articular cartilage presents a new set of challenges. In these cases, the implant must have similar mechanical properties as normal tissue so it may serve as a suitable replacement in a structural or load-bearing role. These properties must remain relatively constant as the scaffold degrades and is replaced by synthesized (and hopefully normal)

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matrix. Additionally, the complexity of these tissues, which may include multiple cell types in a specific spatial organization, may require the construction of a composite engineered tissue for optimal results. Glatz et al. describe a tissue engineering technique for laryngotracheal reconstruction. In their study, rabbit auricular cartilage was harvested and fashioned into strips (Glatz et al., 2003). These strips were then wrapped around a cylindrical silicone stent in order to fashion a structure similar to tracheal rings. The stent was then implanted in the sternocleidomastoid muscle in order to encourage a vascularized fibrous capsule to form around the implant. The silicone stent was removed after 2 weeks, leaving behind the cartilage and the surrounding fibrous capsule. An anterior cricoid split was then performed and reconstructed with a patch graft consisting of a vascularized segment of this neotrachea. The control was standard laryngotracheal reconstruction with autologous auricular cartilage. Initial results were positive, with no significant complications in the experimental group, and evidence of good survival of the implanted cartilage. In contrast, there was evidence of cartilage necrosis in the control group. This technique could be used to study the feasibility of using engineered cartilage, as opposed to autografts, in the construction of the neotrachea. This study also highlights the need to consider vascular supply to engineered tissue, especially as the size of the construct increases. The generation of composite tissues is also of significant interest. Alhadlaq describe the generation of an osteochondral construct in the shape of a mandibular condyle (Alhadlaq and Mao, 2003). Rat MSCs were harvested and subjected to culture in either chondrogeneic or osteogeneic media. Following predifferentiation, the MSCs were photoencapsulated in a bilayered construct containing discrete populations of osteogeneic and chondrogeneic cells, using a polyurethane mold to control shape during photopolymerization. These constructs were subjected to in vitro and in vivo culture. Following incubation, the constructs were subjected to histological analysis, which demonstrated discrete cartilage-like and bone-like layers with distinct histological staining patterns. The cells in the constructs maintained their state of differentiation even after removal from chondrogeneic or osteogeneic media and in vivo implantation. A more recent study by Alhadlaq et al. increased the cell seeding density four-fold and increased the in vivo incubation period to 12 weeks (Alhadlaq and Mao, 2005). They noted an increase in the maturational stage of the construct that was not solely explained by the longer incubation time. Again, the constructs had the appearance of integrated osteochondral tissues. Mechanical properties were not investigated, but prior reports from the same group indicated that the constructs had mechanical properties similar to neonatal articular cartilage (Alhadlaq et al., 2004). These experiments show the possibility of engineering complex, spatially organized tissues. At this time, the only primary surgical options for treatment of cartilage and osteochondral defects, short of joint replacement, are autologous chondrocyte implantation (ACI) and mosaicplasty. While the tissue engineering techniques above are not in clinical use, aspects of current research are being applied to current practice. Marcacci et al. report the results of a cohort of patients treated with ACI in which the in vitro chondrocyte expansion was performed using a 3D scaffold made from modified HA (Marcacci et al., 2005). The scaffold was then implanted into the cartilage defect via a mini-arthrotomy or an arthroscopic approach. Their cohort includes 141 patients followed for 2–5 years. Their results appear impressive, with improvement in subjective symptoms reported in over 90% of patients. Second-look arthroscopy was performed in 55 patients, and the cartilage repair was graded as normal or near-normal in over 95% of these patients. Biopsies were taken in 22 of these 55 patients, which revealed a hyaline appearance in 12 out of 22, with the remainder having a mixed or fibrocartilaginous appearance. Again, this only underscores the potential of tissue engineering to affect clinical outcomes.

CURRENT AND FUTURE TRENDS IN CARTILAGE ENGINEERING The range of studies outlined above is necessarily only a sampling of the past and current literature on cartilage engineering. Nevertheless, it is apparent that great progress has been made toward bringing engineered tissues

Cartilage Tissue Engineering

closer to clinical applicability. There are still an enormous number of questions that remain to be answered, and a great deal of potential for innovative solutions to clinical and research problems. It has been well established that cartilaginous and even osteochondral tissue may be grown in vitro and in vivo. The recurring theme throughout much of the current literature is that this tissue has the histological appearance and biochemical makeup of cartilage of varying stages of maturation. However, it has been reported that mechanically most of these constructs are inferior to native cartilage. As the basic techniques of chondrocyte, osteoblast, and MSC culture are elucidated, the focus shifts toward improving the mechanical properties of engineered tissues. There will need to be an increased focus on correlation between the biochemical markers of matrix production and the biomechanical and ultrastructural properties of native and engineered tissues (Lee and Bader, 1997; Elder et al., 2001; Lu et al., 2001; Pei et al., 2002b; Raimondi et al., 2002). In addition, work continues on the effects of scaffold materials on cellular function, as well as modifications of the scaffold such as alteration of surface chemistry and incorporation of adhesive peptides and/or growth factors (Suzuki et al., 2000; Cui et al., 2003; Grunder et al., 2004). One goal of this avenue of research is the development of a controlled delivery system which would have the ability to release bioactive factors within the scaffold over time, thus providing an element of temporal control of scaffold composition to match the current potential of spatial control (Saito et al., 2001; Lee et al., 2004). To develop successful clinical applications for cartilage tissue engineering, in vitro methods will need to become standardized and provide clear results (Song et al., 2004). There are several important basic questions that remain to be answered. What are the optimal types, amounts, and timing of the growth factor milieu? What are the cellular and genetic patterns needed to retain the chondrogeneic phenotype and lead to chondrogenesis? Is transdifferentiation possible – does stem cell plasticity exist? In vitro models provide the isolated environment necessary to clearly define genetic programming and signaling pathways that are involved in chondrogenesis. As basic research determines these steps it will allow cartilage tissue engineering to translate to the clinical realm (Caplan and Bruder, 2001; Tuan et al., 2002). Other areas of intense scrutiny include exploring ways to implement complex tissue elements such as spatial organization and vasculature in engineered tissues. The current work on osteochondral tissue generation has been discussed above. However, as the effects of 3D organization of tissues become more fully appreciated, there will be a need to regenerate those structures for the purposes of controlled laboratory study and clinical application. With the use of novel polymers that gel under controlled conditions there is the potential for fine control of the shape of engineered tissue, as well as of 3D spatial arrangement of heterogeneous cell populations within the scaffold. This has already been demonstrated by Liu et al. using photolithographic methods to control hydrogel configuration and cellular organization (Liu and Bhatia, 2002). A clinical application was demonstrated by Naumann et al., in which computer-aided design techniques were used to fashion a HA scaffold seeded with chondrocytes into the shape of a human ear. The construct demonstrated an acceptable shape, as well as evidence of cartilage production in vitro, but mechanical properties were not tested (Naumann et al., 2003). Another clinical application of coronary artery disease (CAD) was shown by Sidler et al., in which a bony defect was made in the talus of a human cadaver ankle. The defect was then analyzed by computed tomography, and an implant was fashioned using a computer-aided manufacturing device (Sidler et al., 2005). Of course, there are many more questions and challenges that remain before the promise of tissue engineering is fully realized. The contributions of scientists in fields as diverse as cell/molecular biology, materials science, chemistry, and mathematics will be required in order to answer these questions.

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70 Phalanges and Small Joints Makoto Komura, Daniel Eberli, James J. Yoo, and Anthony Atala

INTRODUCTION Digital amputation is a common injury with estimated 61,000 patients seen in emergency departments in the United States every year (Dubernard et al., 1999). These injuries often result in extensive functional disability and a substantial social and economic cost to the society. More importantly, the outcome of digital dysfunction is detrimental to patient’s daily activities, such as buttoning a shirt or unlocking a door. Therefore, the overall goal for these patients is to rebuild a finger with restoration of normal function, stability, length, and sensation. For sharp amputations where the removed digit is in a clean and fresh state, immediate re-implantation is considered the treatment of choice, which involves microsurgical anastomosis of nerves, blood vessels, bone, and tendon. Functional recovery is achievable if the blood supply to the distal portion is restored within hours and sensory/motor nerves are connected to the stump. The success rate for immediate digital re-implantation is reported to be up to 90% at specialized centers (Patradul et al., 1998). Unfortunately, patients who are eligible for immediate digit re-implantation consist of only a small subset of the total cases. A more complex approach to digit reconstruction is needed for patients with congenital malformation, such as adactyly (missing digit), brachydactyly (short digit), ectrodactylia (reduced number of one or more digital rays), and cleft hands (extended division between the fingers) (Sandzen, 1985). Moreover, surgical re-implantation procedures are not feasible if the distal phalanx is unavailable or heavily damaged after a crush injury. In patients with missing tissues, complex multiple stage surgeries are usually performed with the anticipation of limited functional outcome. The overall goal of functional restoration is to achieve a hand with opposable digits that permits a grip function and prehension. To achieve this goal, common practice is to amputate an entire finger from the other hand and re-implant onto the index position (autologous finger transfer) of the injured hand. Recently, toe-to-hand transfers have become a popular method to reconstruct a missing digit. However, this treatment is associated with possible impairment in foot stability, partial functional recovery of the hand, and poor cosmetic results. Due to the complexity of these procedures and poor outcome, many patients face amputation of their injured digit as the final treatment. This is especially true for soldiers in modern warfare where high energy weapons and improvised explosives are becoming more common. The injured digits are usually lacerated and not amenable to re-implantation. A recent report indicates that approximately 29% of all digit injuries in solders result in amputation (Jovanovic et al., 1999). In cases where an entire hand has been amputated, whole hand transplants were performed successfully in a few patients (Dubernard et al., 1999). Administration of immunosuppressive drugs has extended the viability of the tissues for up to 7 years. However, most of these patients suffered from severe side effects including neuro- and nephro-toxicity. Further, the availability of donor limbs is extremely rare, and this approach has become controversial. Replacement of a single diseased joint with a non-vascularized autogenous joint transplant was studied previously (Campbell, 1963, 1972). While the transplanted articular cartilage remained in place initially, the delayed vascularization led to subchondral collapse and disintegration of cartilage tissue.

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Prosthetic treatments are more common (Pereira et al., 1996). Unlike large limb prosthetics which are becoming increasingly advanced and functional, digit prostheses are primarily aimed toward improving cosmetic appearance. These prosthetic devices are often associated with complications, such as erosion, infection, and inflammation. Moreover, prosthetic devices are not suitable for pediatric or adolescent patients, as their tissues grow with age. Therefore, most patients decide not to wear the prosthesis, due to the lack of gain in function and its associated complications. Despite the rapid technological advances in medicine, treatment options for digit reconstruction are severely limited. Currently available modalities are involved with a prolonged treatment period, multiple stage surgeries, a lengthy rehabilitation process for a questionable restoration of normal function, and prostheses and surgery-associated complications. Limitation of current treatment modalities for digit reconstruction and functional restoration has led investigators to pursue alternative therapeutic approaches. The concept of cell transplantation using tissue engineering techniques has been proposed as a method to improve, restore or replace tissue or organ function (Oberpenning et al., 1999; Yoo et al., 1999; Amiel et al., 2001; Kim et al., 2002, 2004; Kwon et al., 2002; Lanza et al., 2002; De Filippo et al., 2003; El-Kassaby et al., 2003; Falke et al., 2003; Yiou et al., 2003; Atala et al., 2006). The emergence of tissue engineering and regenerative medicine strategies has presented alternative possibilities for the restoration of digit tissues. The ability to engineer physiological units comprised of bone, muscle, and connective tissue, with supporting vasculature and innervation, into functional appendages such as fingers and toes would be an evolutionary step toward the regeneration of missing limbs.

BASIC PRINCIPLES OF ENGINEERING PHALANGES AND SMALL JOINTS The digit is a complex structure that is composed of highly specialized tissues, including skin, bone, cartilage, tendon, muscle, fat, blood vessels, and nerves. To engineer a digit tissue that restores normal function, a composite tissue, consisting of all the necessary components, is necessary. One of the challenges in building such a complex tissue is the ability to pack all tissue components into a small compartment, while each tissue type delivers individual function in a coordinated manner. Although individual or a group of tissue components has been engineered separately for various applications, combination of multiple tissue types has not been used clinically (Campbell, 1972; Isogai et al., 1999, 2004; Yoo et al., 1999; Kim et al., 2002, 2004; Yiou et al., 2003). Recent efforts have been focused in engineering composite tissues that would allow for multiple functions (Isogai et al., 1999; Oberpenning et al., 1999; Kwon et al., 2002; Lanza et al., 2002; De Filippo et al., 2003). To achieve this goal, special considerations are mandated, some of which include the use of multiple cell types, designing of a scaffolding system that would allow for coordinated motion and achieving adequate vascularization for the survival of the implanted tissue. Although there are many cell types present in the digital compartment, these cells are abundantly found in other tissues. Cells and tissues composing the digit have been harvested from periosteum, skeletal muscles, skin, articular, and fibrous cartilage. Cells obtained from remote regions of the body are grown and expanded for reconstruction of digits (Isogai et al., 1999; Vacanti et al., 2001; Kim et al., 2002, 2004; Yiou et al., 2003; Landis et al., 2005). Recently, stem and progenitor cells have been proposed as an alternative cell source for the engineering of the digit tissues. While autologous somatic cells are preferable, these cells are known to have a limited expansion capability and require multiple tissue biopsies in order to obtain all cell types. For this reason, stem cells have been proposed as an attractive cell source for various tissue regeneration applications. Human embryonic stem cells are able to proliferate in the undifferentiated state (self-renewal) and have the ability to differentiate into many specialized cell types (Brivanlou et al., 2003). However, controversies surrounding their use have hindered research progress. Generation of stem cells through therapeutic cloning technology has been proposed as an alternative method to bypass ethical challenges. This method employs

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nuclear transfer techniques to generate stem cells that can be used to engineer tissues, such as skeletal muscle, cardiac muscle, and kidney tissues (Lanza et al., 2002). While the use of embryonic stem cells is ethically controversial, adult stem cells are being proposed as a new source. Although their potential and limitations are not fully evaluated, stem cells from fat, skin, muscle, placenta, umbilical cord, amniotic, and bone marrow have been used for various cell and tissue therapies (Bartsch et al., 2005; Deasy et al., 2005; Moise, 2005; Tholpady et al., 2006). A variety of synthetic and naturally derived materials have been examined in order to determine the ideal support systems for the regeneration of tissues (Kim and Mooney, 1998; Isogai et al., 1999; Amiel et al., 2001; Vacanti et al., 2001; Pariente et al., 2002; El-Kassaby et al., 2003; Falke et al., 2003; Derwin et al., 2004; Kim et al., 2004; Atala et al., 2006; Murray et al., 2006). Biodegradable synthetic materials, such as poly-lactic and glycolic acid polymers have been used to provide structural support for cells. Synthetic materials can be easily fabricated and configured in a reproducible manner. Naturally derived materials, such as collagen, laminin, and fibronectin are biocompatible and provide a similar extracellular matrix (ECM) environment as normal tissue. Biomaterials used for regenerative medicine are designed to replicate the biologic and mechanical function of native tissue structures and their ECM. They provide three-dimensional architecture for the cells to reconstitute into new tissues with appropriate structure and function, and allow for the delivery of cells and appropriate bioactive factors (e.g. cell-adhesion peptides, growth factors) to desired sites in the body (Kim and Mooney, 1998). As the majority of mammalian cell types are anchorage-dependent, biomaterials provide a cell-adhesion substrate that can deliver cells to specific regions of the body with high loading efficiency. Biomaterials also provide mechanical support against in vivo forces such that the predefined three-dimensional structure is maintained during tissue development. Furthermore, bioactive signals such as cell-adhesion peptides and growth factors can be loaded along with cells to help regulate cellular function. The ideal biomaterial for digit reconstruction should be biodegradable and bioabsorbable without eliciting inflammatory responses that interfere with tissue formation. Incompatible materials are destined for an inflammatory or foreign-body response that eventually leads to rejection and/or necrosis. Since biomaterials provide temporary mechanical support while the cells undergo spatial tissue organization, a properly chosen biomaterial should allow the engineered tissue to maintain adequate mechanical integrity to support itself in early development. Both the synthetic polymers (e.g. polyglycolic acid (PGA)) and naturally derived materials (e.g. bladder submucosa and small intestinal submucosa) have been used as biomaterials for digit reconstruction. These materials have been shown to be biocompatible and suitable for tissue engineering applications when appropriately configured (Pariente et al., 2002). Synthetic biomaterials have been used successfully for phalangeal reconstruction, which include PGA, PCL (poly ε-caprolactone) and PLLA (poly-L-lactate) (Rosenberg, 1971; Isogai et al., 2004). Non-woven PGA has porosity of greater than 95%, thus is able to accommodate a large quantity of cells and offers adequate environment for bone and cartilage formation. Further, PGA can easily be configured to structures similar to native tissue. Naturally derived materials, such as hydroxyapatite and small intestine submucosa (SIS) have been used to engineer phalangeal tissues (Vacanti et al., 2001; Derwin et al., 2004). Engineering of a viable tissue is the prime goal of cell-based technology, thus, obtaining adequate vascular supply is critical to cell viability and the development of tissues. Although the body has the ability to form new vessels over time, it is critical to provide nutrients and oxygen to the cell-constructs within the limited space, as in distal phalanges. In addition, it is a general conception that cell or tissue components may not be implanted in large volumes due to the limited diffusion (Folkman and Hochberg, 1973). Numerous efforts have been made to overcome this limitation and attempts to enhance vascularization within the host tissue have been pursued using several approaches. These include the use of angiogeneic factors, such as use of vascular endothelial growth factors (VEGF) and endothelial cells (EC), and cell-support matrices that permit

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enhanced easy diffusion of nutrients across the entire implant (Nomi et al., 2002; De Coppi et al., 2005; Kaigler et al., 2006). In one study, VEGF secreting myoblasts were implanted with vascular EC to enhance neovascularization of engineered tissues. In this study skeletal myoblasts from were cultured and transfected with an adenovirus encoding VEGF and combined with human vascular EC (De Coppi et al., 2005). The cell mixtures were injected subcutaneously in athymic mice, and the engineered tissues were retrieved up to 8 weeks after implantation. The transfected cells formed vascularized muscle tissue mass with evidence of adequate neovascularization by histology and immunohistochemical probing. The engineered muscle tissue, composed of non-transfected cells had a significantly smaller tissue mass with decreased muscle volume and less neovascularization. These results indicate that a combination of VEGF and EC may be useful for inducing neovascularization and volume preservation in the engineered tissues. In addition to the biological approach, incorporation of bioactive factors that enhance vascularization into the support matrices has been tried (Schuch et al., 2002; Kim et al., 2004; Ennett et al., 2006; Kaigler et al., 2006). VEGF was incorporated into poly(lactide-co-glycolide) (PLG) scaffolds or pre-encapsulated in PLG microspheres for therapeutic angiogenesis. In this study VEGF was positioned predominantly adjacent to scaffold pores when incorporated directly and was released rapidly. Pre-encapsulation led to the VEGF being more deeply embedded and resulted in a delayed release. In vivo, the released VEGF generated local protein concentrations for the 21 days of the experiment, with negligible release into the systemic circulation, and significantly enhanced local angiogenesis. These data indicate that VEGF can be administered in a sustained and localized fashion in vivo (Ennett et al., 2006). This angiogenic factor delivery system was applied to bone regeneration application and was demonstrated that VEGF scaffolds have the ability to enhance neovascularization and bone regeneration in irradiated osseous defects, which outlines a novel approach for engineering tissues in hypovascular environments (Kaigler et al., 2006).

CREATION OF COMPOSITE TISSUE STRUCTURES The goal of engineering of the digit is to achieve esthetic and functional tissues that would allow for adequate dexterity for daily activities. To permit this capability, composite tissues with all cellular and tissue components are necessary. Toward this goal, initial efforts were focused on engineering of individual tissues, such as the skin, bone, cartilage, tendon, muscle, fat, and nerve (Amiel et al., 2001; Lanza et al., 2002; Kim et al., 2002, 2004; Yiou et al., 2003; Derwin et al., 2004; Lee et al., 2006; Wood, 2006). In a case report, a tissue engineered distal phalanx was used to replace a distal bone in a patient who had a partial avulsion of the thumb. The procedure resulted in the functional restoration of a stable and biomechanically sound thumb of normal length, without the pain and complications that are usually associated with harvesting a bone graft (Vacanti et al., 2001). Although the formation of individual tissues has been successfully demonstrated experimentally and clinically, engineering of a more complex tissue, consisting of multiple cell types, remains a challenge. The first attempt to engineer a phalanx was demonstrated through the use of cartilage, bone, and tendon on a single construct (Isogai et al., 1999). In an ex situ model, bovine chondrocytes and tenocytes were seeded on biodegradable polymer scaffolds that were configured to the shape of human phalanx bones and wrapped with the bovine periosteum to serve as a joint. Subsequently, these constructs were implanted into the subcutaneous space of athymic mice and followed for up to 60 weeks. The retrieved phalangeal constructs showed the formation of bone, cartilage and tendon tissues, which were confirmed using polymerase chain reaction (PCR), histology, and immunohistochemistry. More importantly, the implanted phalanges maintained the shape of human phalanges (Landis et al., 2005). This experimentation demonstrated that phalangeal tissues can be engineered ex situ using a composite cell system and provide the possibility of engineering morphologically adequate phalanx. However, considerations for engineering of functional phalangeal units have not been suggested.

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(a)

(b)

(c)

Figure 70.1 Engineered composite tissue in vivo. (a) Gross examination of the explanted muscle and cartilage tissue. (b) Histology of cartilage and muscle tissue composite (H&E). (c) Formation of bone (B) and organized muscle bundles (M) was observed in the bone/muscle composite tissue.

In an effort to engineer functional digits for reconstruction, we have designed scaffolds that would permit the delivery of different cell types to target sites and would allow for joint movement. The PGA polymer was chosen to generate the rigid portion of the digit, made from cartilage or bone, and acellular collagen matrix derived from bladder submucosa to allow movement by muscle. The PGA polymer was composed of two tubular pieces connected with non-absorbable sutures, which served as a hinge between the two bony scaffolds. Multiple thin strips of acellular collagen matrix were attached to the distal ends of the bony scaffolds. This configuration was designed to allow for contraction, relaxation, and movement of the composite engineered digit-like structure. In the initial experiments, the PGA polymer scaffolds were seeded with bovine chondrocytes/osteoblasts, and bovine skeletal muscle cells were seeded onto the collagen matrix. The cellseeded composites were incubated in vitro for 2 weeks. Histologically, abundant muscle cells were present throughout the collagen matrix and the PGA polymers seeded with chondrocytes showed the presence of whitish ECM. Western blot analysis confirmed the tissue identity and showed expression of muscle specific genes such as actin, desmin, and tropomyosin in the muscle component and expression of collagen-1 in the chondrocyte-seeded PGA. Subsequently, the feasibility of engineering skeletal muscle and cartilage/bone composite structure was studied in vivo. Composite scaffolds consisting of synthetic PGA polymers (1.0  0.5  0.5 cm) and naturally derived collagen matrices obtained from the bladder submucosa (1.4  0.4  0.3 cm) were constructed. Bovine chondrocytes or osteoblasts were seeded onto the PGA polymer matrices and skeletal muscle cells onto the collagen matrices. The scaffolds containing both cell types were analyzed in vitro for cell viability and tissue formation. The engineered digits were implanted subcutaneously in athymic mice (n  36) and followed for up to 6 months. The cells seeded on the composite digit constructs readily attached to their designated region of the scaffold and remained viable. Grossly, the implanted scaffolds formed muscle and cartilage or bone tissues adjacent to each other (Figure 70.1). Each tissue type was confirmed histo- and immunohistochemically using cell specific antibodies. Biomechanical studies showed that the cartilage tissue was elastic and could withstand high degrees of pressure, which demonstrate the ability to preserve its structural integrity. Physiologic organ bath studies of the retrieved muscle tissues showed adequate contractility in response to electric field stimulation. These findings show that different tissue types can be engineered simultaneously using a composite scaffold system. The tissues retained their respective phenotypic and functional characteristics independent of the other. This study demonstrates that the engineering of functional digit tissues may be feasible. To determine the feasibility of replacing a missing digit segment, in situ implantation of digit segments were performed in a rabbit model. Autologous muscle cells and chondrocytes were grown, expanded, and seeded on the composite scaffolds, consisting of interconnecting bony segments, muscle, and tendon. The digit

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(a)

(d)

(b)

C

(c)

T

M Figure 70.2 Engineering of phalangeal tissues in vivo. (a) Bone scaffold, (b) muscle scaffolds composed of multiple thin collagen fiber strips, (c) tendon scaffold, (d) engineered phalangeal tissue at retrieval shows the formation of cartilage, muscle and tendon in a rabbit model.

segments with an interconnected joint were excised and the engineered digit segments of the same length and caliber were replaced in rabbits. The forelimb with the engineered digit was placed in a cast for approximately 4 weeks in order to protect the wound and enhance tissue maturation in situ. The engineered digit segments were able to form cartilage, muscle, and tendon at retrieval. Scaffolds without cells failed to form tissue structures (Figure 70.2).

SUMMARY AND FUTURE PERSPECTIVES IN CLINICAL TRANSLATION Efforts in engineering of the digit are focused toward the eventual clinical application. Engineering of the phalangeal tissues presents a unique challenge in tissue reconstruction due to the complexity of the functions required. Although individual or partial replacement of phalanges has been demonstrated experimentally and clinically, achieving a fully functional engineered digit requires continued investigations. One of the unsolved tasks that would accelerate the progress of this research is finding methods to integrate all tissue components, including the vascular and neural network, into a compact compartment of the distal limb. In addition, development of an intelligent composite bioscaffold system that would allow for the enhanced formation of individual tissue types in a controlled manner is critical. Recent progress made towards the engineering of phalanx suggests that achieving functional phalangeal tissues may have an expanded role in clinical medicine.

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REFERENCES Amiel, G.E., Yoo, J.J., Kim, B.S. and Atala, A. (2001). Tissue engineered stents created from chondrocytes. J. Urol. 165: 2091–2095. Atala, A., Bauer, S.B., Soker, S., Yoo, J.J. and Retik, A.B. (2006). Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet 367: 1241–1246. Bartsch, G., Yoo, J.J., De Coppi, P., Siddiqui, M.M., Schuch, G., Pohl, H.G., Fuhr, J., Perin, L., Soker, S. and Atala, A. (2005). Propagation, expansion, and multilineage differentiation of human somatic stem cells from dermal progenitors. Stem Cell. Dev. 14: 337–348. Brivanlou, A.H., Gage, F.H., Jaenisch, R., Jessell, T., Melton, D. and Rossant, J. (2003). Stem cells. Setting standards for human embryonic stem cells. Science 300: 913–916. Campbell, C.J., Ishida, H., Takahashi, H. and Kelly, F. (1963). The transplantation of articular cartilage. An experimental study in dogs. J. Bone Joint Surg. 45: 1579–1592. Campbell, C.J., Ishida, H., Takahashi, H. and Kelly, F. (1972). Homotransplantation of a half or whole joint. Clin. Orthop. Relat. Res. 87: 146–155. De Coppi, P., Delo, D., Farrugia, L., Udompanyanan, K., Yoo, J.J., Nomi, M., Atala, A. and Soker, S. (2005). Angiogenic gene-modified muscle cells for enhancement of tissue formation. Tissue Eng. 11: 1034–1044. De Filippo, R.E., Yoo, J.J. and Atala, A. (2003). Engineering of vaginal tissue in vivo. Tissue Eng. 9: 301–306. Deasy, B.M., Gharaibeh, B.M., Pollett, J.B., Jones, M.M., Lucas, M.A., Kanda, Y. and Huard, J. (2005). Long-term selfrenewal of postnatal muscle-derived stem cells. Mol. Biol. Cell 16: 3323–3333. Derwin, K., Androjna, C., Spencer, E., Safran, O., Bauer, T.W., Hunt, T., Caplan, A. and Iannotti, J. (2004). Porcine small intestine submucosa as a flexor tendon graft. Clin. Orthop. Relat. Res. 245–252. Dubernard, J.M., Owen, E., Herzberg, G., Lanzetta, M., Martin, X., Kapila, H., Dawahra, M. and Hakim, N.S. (1999). Human hand allograft: report on first 6 months. Lancet 353: 1315–1320. El-Kassaby, A.W., Retik, A.B., Yoo, J.J. and Atala, A. (2003). Urethral stricture repair with an off-the-shelf collagen matrix. J. Urol. 169: 170–173; discussion 173. Ennett, A.B., Kaigler, D. and Mooney, D.J. (2006). Temporally regulated delivery of VEGF in vitro and in vivo. J. Biomed. Mater. Res. A. Falke, G., Yoo, J.J., Kwon, T.G., Moreland, R. and Atala, A. (2003). Formation of corporal tissue architecture in vivo using human cavernosal muscle and endothelial cells seeded on collagen matrices. Tissue Eng. 9: 871–879. Folkman, J. and Hochberg, M. (1973). Self-regulation of growth in three dimensions. J. Exp. Med. 138: 745–753. Isogai, N., Asamura, S., Higashi, T., Ikada, Y., Morita, S., Hillyer, J., Jacquet, R. and Landis, W.J. (2004). Tissue engineering of an auricular cartilage model utilizing cultured chondrocyte-poly(L-lactide-epsilon-caprolactone) scaffolds. Tissue Eng. 10: 673–687. Isogai, N., Landis, W., Kim, T.H., Gerstenfeld, L.C., Upton, J. and Vacanti, J.P. (1999). Formation of phalanges and small joints by tissue-engineering. J. Bone Joint Surg. 81: 306–316. Jovanovic, S., Wertheimer, B., Zelic, Z. and Getos, Z. (1999). Wartime amputations. Mil. Med. 164: 44–47. Kaigler, D., Wang, Z., Horger, K., Mooney, D.J. and Krebsbach, P.H. (2006). VEGF scaffolds enhance angiogenesis and bone regeneration in irradiated osseous defects. J. Bone Miner. Res. 21: 735–744. Kim, B.S. and Mooney, D.J. (1998). Development of biocompatible synthetic extracellular matrices for tissue engineering. Trends Biotechnol. 16: 224–230. Kim, B.S., Yoo, J.J. and Atala, A. (2002). Engineering of human cartilage rods: potential application for penile prostheses. J. Urol. 168: 1794–1797. Kim, B.S., Yoo, J.J. and Atala, A. (2004). Peripheral nerve regeneration using acellular nerve grafts. J. Biomed. Mater. Res. A 68: 201–209. Kwon, T.G., Yoo, J.J. and Atala, A. (2002). Autologous penile corpora cavernosa replacement using tissue engineering techniques. J. Urol. 168: 1754–1758. Landis, W.J., Jacquet, R., Hillyer, J., Lowder, E., Yanke, A., Siperko, L., Asamura, S., Kusuhara, H., Enjo, M., Chubinskaya, S., et al. (2005). Design and assessment of a tissue-engineered model of human phalanges and a small joint. Orthod. Craniofac. Res. 8: 303–312.

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Lanza, R.P., Chung, H.Y., Yoo, J.J., Wettstein, P.J., Blackwell, C., Borson, N., Hofmeister, E., Schuch, G., Soker, S., Moraes, C.T., et al. (2002). Generation of histocompatible tissues using nuclear transplantation. Nat. Biotechnol. 20: 689–696. Lee, S.J., Lim, G.J., Lee, J.W., Atala, A. and Yoo, J.J. (2006). In vitro evaluation of a poly(lactide-co-glycolide)-collagen composite scaffold for bone regeneration. Biomaterials 27: 3466–3472. Moise Jr., K.J. (2005). Umbilical cord stem cells. Obstet. Gynecol. 106: 1393–1407. Murray, M.M., Forsythe, B., Chen, F., Lee, S.J., Yoo, J.J., Atala, A. and Steinert, A. (2006). The effect of thrombin on ACL fibroblast interactions with collagen hydrogels. J. Orthop. Res. 24: 508–515. Nomi, M., Atala, A., Coppi, P.D. and Soker, S. (2002). Principals of neovascularization for tissue engineering. Mol. Aspect. Med. 23: 463–483. Oberpenning, F., Meng, J., Yoo, J.J. and Atala, A. (1999). De novo reconstitution of a functional mammalian urinary bladder by tissue engineering. Nat. Biotechnol. 17: 149–155. Pariente, J.L., Kim, B.S. and Atala, A. (2002). In vitro biocompatibility evaluation of naturally derived and synthetic biomaterials using normal human bladder smooth muscle cells. J. Urol. 167: 1867–1871. Patradul, A., Ngarmukos, C. and Parkpian, V. (1998). Distal digital replantations and revascularizations. 237 digits in 192 patients. J. Hand Surg. (Edinburgh, Lothian) 23: 578–582. Pereira, B.P., Kour, A.K., Leow, E.L. and Pho, R.W. (1996). Benefits and use of digital prostheses. J. Hand Surg. 21: 222–228. Rosenberg, L. (1971). Chemical basis for the histological use of safranin O in the study of articular cartilage. J. Bone Joint Surg. 53: 69–82. Sandzen Jr., S.C. (1985). Classification and functional management of congenital central defect of the hand. Hand Clin. 1: 483–498. Schuch, G., Machluf, M., Bartsch Jr., G., Nomi, M., Richard, H., Atala, A. and Soker, S. (2002). In vivo administration of vascular endothelial growth factor (VEGF) and its antagonist, soluble neuropilin-1, predicts a role of VEGF in the progression of acute myeloid leukemia in vivo. Blood 100: 4622–4628. Tholpady, S.S., Llull, R., Ogle, R.C., Rubin, J.P., Futrell, J.W. and Katz, A.J. (2006). A dipose tissue: stem cells and beyond. Clin. Plast. Surg. 33: 55–62, vi. Vacanti, C.A., Bonassar, L.J., Vacanti, M.P. and Shufflebarger, J. (2001). Replacement of an avulsed phalanx with tissueengineered bone. N. Engl. J. Med. 344: 1511–1514. Wood, F.M., Kolybaba, M. L., Allen, P. (2006). The use of cultured epithelial autograft in the treatment of major burn wounds: Eleven years of clinical experience. Burns 32: 538–544. Yiou, R., Yoo, J.J. and Atala, A. (2003). Restoration of functional motor units in a rat model of sphincter injury by muscle precursor cell autografts. Transplantation 76: 1053––1060. Yoo, J.J., Park, H.J., Lee, I. and Atala, A. (1999). Autologous engineered cartilage rods for penile reconstruction. J. Urol. 162: 1119–1121.

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71 Functional Tissue Engineering of Ligament and Tendon Injuries Savio L.-Y. Woo, Alejandro J. Almarza, Sinan Karaoglu, and Steven D. Abramowitch INTRODUCTION Tendons and ligaments are soft connective tissues composed of closely packed, parallel collagen fiber bundles which connect bone to muscle and bone to bone, respectively. These unique tissues serve essential roles in the musculoskeletal system by transferring tensile loads to guide motion and stabilize diarthrodial joints. Injuries to tendons, such as the patellar tendon (PT) of the knee, or ligaments, such as the collateral and cruciate ligaments of the knee, upset the balance between mobility and stability of this joint. These injuries are often manifested in abnormal knee kinematics and damage to other tissues in and around the joint such as meniscus and articular cartilage, which may lead to morbidity, pain, and osteoarthritis. With the high incidence of ligament and tendon injuries in sports and work related activities, improvements on healing and repair of these tissues are of great interest (Beaty 1999). Interestingly, there is a dramatic variability in the propensity for healing of ligaments within the same knee joint, namely the medial collateral ligament (MCL) and anterior cruciate ligament (ACL). Clinical and laboratory studies have shown that injuries to the MCL generally heal sufficiently well such that non-surgical management has become the treatment of choice (Frank et al., 1983; Indelicato, 1983; Jokl et al., 1984; Woo et al., 1987b; Kannus, 1988; Scheffler et al., 2001). While most structural properties of the femur–MCL-tibia complex (FMTC) are restored within weeks, the mechanical properties of the healed MCL (i.e. the stress–strain curve) remain much different from those of the normal MCL, as are the altered histomorphological appearance (e.g. uniform distribution of small collagen fibrils) and biochemical composition (e.g. elevated type III and V collagens) (Adachi and Hayashi, 1986; Birk et al., 1990; Weiss et al., 1991; Frank et al., 1992; Hart et al., 1992; Marchant et al., 1996; Frank et al., 1997; Hart et al., 2000; Nakamura et al., 2000; Niyibizi et al., 2000; Birk, 2001). For the ACL, it is well known that a midsubstance tear would not heal and the success of non-surgical management is limited. Thus, surgical reconstruction of the ACL using autografts harvested from the PT or hamstring tendons is recommended. Issues affecting patient outcome from the use of bone–PT–bone (BPTP) autografts include a persistent palpable defect in the tendon, anterior knee pain, arthrofibrosis, changes to the remaining PT, and PT adhesion to adjacent tissues (i.e. the fat pad) (Coupens et al., 1992; Svensson et al., 2005). The problems associated with hamstrings tendon autografts include slower healing due to a development of a soft tissue to bone interphase, less long-term stability of the knee (Freedman et al., 2003), significant hamstring muscle weakness (Marder 1991; Aune et al., 2001), as well as the increased prevalence of bone tunnel enlargement after reconstruction (Nebelung et al., 1998; Clatworthy et al., 1999; Jansson et al., 1999;

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Feller and Webster, 2003). Hence, functional tissue engineering (FTE) efforts are aiming to improve the suboptimal properties of the healing MCL, as well as the issues related to ACL graft harvest and healing following reconstruction. With the knowledge gained, it is hoped that the same principles could be applied to aid the repair of other ligaments and tendons (Huang et al., 1993; Badylak et al., 1995; Hildebrand and Frank, 1998; Woo et al., 1999; Nakamura et al., 2000; Spindler et al., 2002; Shimomura et al., 2003). Thus, FTE offers many attractive approaches to enhance ligament and tendon healing. The goal is not only to restore the normal histomorphological appearance, biochemistry, and mechanical properties of the healing ligament, but most importantly restore its normal joint function. In this chapter, we will review the properties of normal and healing ligaments and tendons, and discuss the current FTE methods, which include the use of growth factors, gene delivery, stem cell therapy, and the use of scaffolding as well as external mechanical stimuli, aimed at enhancing tendon and ligament healing. To conclude, new technologies and research avenues that have the potential to enhance treatment strategies for ligament and tendon injuries are suggested.

NORMAL LIGAMENTS AND TENDONS Biology Ligaments and tendons consist of collagen, proteoglycans, elastin, glycolipids, water (65% to 70% of the total weight), and cells. Both tissues are hypocellular with less than 5% of the total volume occupied by cells (Woo et al., 2000; Lo et al., 2002; Hildebrand et al., 2004) and are relatively hypovascular (Manske, 1988; Bray et al., 1996; Lo et al., 2002). The cells in these tissues, fibroblasts and tenocytes, are arranged in rows along the fibers of the extracellular matrix (ECM). Both cell types produce both the fibrillar and the non-fibrillar components of the ECM and may also reabsorb collagen fibers (Birk and Trelstad, 1984; Maffulli, 1999). Roughly 70% to 80% of the dry weight of normal tendon or ligament is composed of type I collagen. Histologically, collagen displays a crimp pattern, which refers to a regular, wavy pattern of the matrix when viewed in unloaded conditions (Amiel et al., 1984; Woo et al., 2000; Thornton et al., 2002; Hildebrand et al., 2004). Both in ligaments and tendons, there is a bimodal distribution of collagen fibril diameters. One group of fibrils measures between 40 and 75 nm in diameter, the other is between 100 and 150 nm (Dyer and Enna, 1976; Eyden and Tzaphlidou, 2001; Goh et al., 2003). It has been proposed that a bimodal diameter distribution endows tendons and ligaments with better functional properties. The incorporation of a high fraction of small diameter fibrils would ensure a better interfibrillar binding by virtue of their higher surface/volume ratio, whereas the strength requirements would be satisfied by the inclusion of large diameter fibrils. A bimodal distribution would also improve fibril packing, the smaller fibrils wedging themselves in the spaces left among the larger ones (Ottani et al., 2001). There are many other collagen types, including III, V, X, XI, and XII, which exist only in minor amounts in ligaments and tendons. The significance of some of these minor collagen types has recently been elucidated. For example, type V collagen is believed to exist in association with type I collagen and serves as a regulator of collagen fibril diameter (Linsenmayer et al., 1993; Birk and Mayne 1997), whereas type III collagen is needed for wound healing (Liu et al., 1995). Our research has further identified that type XII collagen provides lubrication between collagen fibers (Niyibizi et al., 1995). Lastly, collagen types IX, X, and XI have been identified to exist with type II collagen at the fibrocartilaginous zone of the ligament–bone and tendon–bone interface (Niyibizi et al., 1996; Sagarriga Visconti et al., 1996; Fukuta et al., 1998). It is hypothesized that these collagens exist in this zone to minimize the stress concentrations when loads are transmitted from soft tissue to bone (Cooper and Misol, 1970; Matyas et al., 1995). The ground substance constituents of tendons or ligaments make up only a small percentage of the total dry tissue weight but are nevertheless quite significant because of their ability to imbibe water. The water and

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proteoglycans provide lubrication and spacing which are crucial to the gliding function of fibers in the tissue matrix. Elastin, which is present in ligaments and tendons in a few percent by weight, allows the tissue to return to its prestretched length following physiological loading, but the detailed significance has yet to be elucidated. Collectively, these constituents serve to maintain fiber orientation and separation for optimal load distributions. Although ligaments and tendons are morphologically similar to each other, there are some biochemical differences. When compared with tendons, ligaments are more metabolically active. They have more cellular nuclei, a higher DNA content, and greater amounts of reducible cross-links between collagen fibers (Amiel et al., 1984). Ligaments are composed primarily of water (65–70% of wet weight), cells and collagen (70–80% of dry weight). The most abundant collagen is type I collagen (90% dry weight). Type III collagen (8% dry weight) and type V collagen (12% dry weight) are other major components (Linsenmayer et al., 1993; Birk and Mayne, 1997). Collagen types II, IX, X, XI, and XII have also been found to be present (Niyibizi et al., 1996; Sagarriga Visconti et al., 1996; Fukuta et al., 1998; Woo et al., 2006). Tendons, on the other hand, generally contain less water (55% of wet weight), and slightly more type I collagen (85% of dry weight), along with much smaller amounts of other collagens, such as collagens type III, V, XII, and XIV (Goh et al., 2003). Biomechanics The major function of ligaments and tendons include maintaining the proper anatomic alignment of the skeleton and guiding joint motions. They accomplish this by transmitting forces along their longitudinal axis; hence their biomechanical properties are measured in uniaxial tension. They demonstrate non-linear behavior, which is governed by the recruitment of collagen. This allows ligaments to maintain normal joint laxity in response to low loads, and also stiffen dramatically in response to high loads, preventing excessive joint displacements. Ligaments and tendons also exhibit time- and history- dependent viscoelastic behavior which could be attributed to the complex interactions of tissue constituents such as collagen, proteoglycans, water, and ground substance (Fung et al., 1972; Woo et al., 1981b). Viscoelastic properties of ligaments and tendons are important and clinically relevant. For instance, during regular activities such as walking and jogging, tissues have the ability to soften over time. This phenomenon reduces the susceptibility to damage related to fatigue (Frank and Jackson, 1997). However, following injury, ligaments and tendons generally fail to recover their normal mechanical and viscoelastic behaviors. Thus, abnormal joint kinematics result which can directly lead to excessive forces in surrounding tissues (e.g. articular cartilage). This can lead to further injury to other structures either through traumatic mechanisms or degeneration (i.e. osteoarthritis). As the ultimate goal of FTE is to restore the function of ligaments and tendons, and thereby the function of the injured joint, it is necessary to understand their normal mechanical behavior and contributions to joint function. Testing methodologies for this purpose include (1) functional testing, which involves determining the contribution of the ligament or tendon to joint kinematics (i.e. the in situ forces in response to external loading conditions), and (2) tensile testing, which provides an assessment of the structural properties of the bone–ligament–bone complex and mechanical properties of the tissue substance. Tensile Testing Tendons are generally long and can be tested in their isolated state using sinusoidal-shape or frozen grips to limit slippage. Isolated ligaments, on the other hand, are short in length, making it difficult to clamp them independently. Hence, a tensile test is generally conducted on the entire bone–ligament–bone complex (e.g. FMTC) with tissue insertion sites left anatomically intact. With cross-sectional area (CSA) measurements and the utilization of tissue markers to measure tissue strain, the structural properties of the bone–ligament–bone

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complex as well as mechanical properties of the ligament substance can be measured from a load to failure test (Woo et al., 1983; Lee and Woo, 1988). Structural properties (Figure 71.1) of the bone–ligament–bone complex (i.e. a load-elongation curve) are generally described by four parameters including stiffness (slope of the linear portion of the load-elongation curve), ultimate load (maximum load at which the complex fails), ultimate elongation (elongation corresponding to the maximum load), and energy absorbed at failure (area under the curve to the maximum load). These data reflect behavior of the entire bone–ligament–bone complex which includes tissue size, orientation of collagen fibers to applied loads as well as the contribution of the bony insertions (Woo et al., 1983). Measuring the mechanical properties (Figure 71.2) of the ligament substance (i.e. a stress–strain curve), on the other hand, requires knowledge of the CSA of the ligament, commonly measured using a laser micrometer system (Lee and Woo, 1988), and tissue strain, commonly measured using video techniques to track two or more reflective markers placed on the tissue midsubstance (Scheffler et al., 2001). Stress in the tissue is obtained by dividing load by the CSA and strain is obtained by computing change in marker distance during the test relative to their original distance. Parameters describing the mechanical properties of the ligaments and tendons (Figure 71.2) include tangent modulus (slope of the linear portion of the stress–strain curve), tensile strength (stress at failure), ultimate strain (strain corresponding with the tensile strength), and strain energy density (area under the stress–strain curve until failure). These data represent the quality of the tissue, irrespective of tissue size.

Ultimate load

Failure

750 Load (N)

600

Linear stiffness

450 300

Ultimate elongation

Energy absorbed

150 0 0

1

2

3

4

5

6

Elongation (mm)

Figure 71.1 A typical load-elongation curve representing the structural properties of the femur– anterior–medial bundle-tibia complex of the human anterior cruciate ligament.

Tensile strength

Failure

Stress (MPa)

25 20

Tangent modulus

15

Strain energy density

10 5

Ultimate strain

0 0

5

10

15

Strain (%)

Figure 71.2 A typical stress–strain curve representing the mechanical properties of the anterior–medial bundle of the human anterior cruciate ligament.

THERAPEUTIC APPLICATIONS: TISSUE THERAPY

The viscoelastic properties of ligaments and tendons include stress relaxation (decrease in stress over time in response to a constant elongation), and creep (increase in elongation over time in response to a constant load). In addition, they also display a phenomenon called “hysteresis” in response to cyclic loading (Figure 71.3). This results from a loss of internal energy causing the loading and unloading paths to be different. The area of hysteresis reduces as the tissue undergoes several cycles of loading and unloading and the tissue is said to be “preconditioned,” a state desired for a tissue prior to mechanical testing. Non-linear viscoelastic models such as the quasi-linear viscoelastic theory, t

σ(t )  ∫ G(t  τ ) ∞

∂σ e (ε) ∂ε ∂τ ∂ε ∂τ

(71.1)

and single integral finite strain theory, T  pI  C0 {[1  μ I (t )]B(t)  μB2 (t )}  C0 (1  γ ) t

∫ G (t  s){[1  μI(s)]B(t )  μF(t )C(s)FT (s)}ds

(71.2)

0

have been utilized to model these behaviors in ligaments and tendons (Fung et al., 1972; Woo et al., 1981b; Johnson et al., 1994). These basic testing methodologies described in this section have been utilized for decades to examine soft tissues and much work has been done to define the appropriate testing procedures such as specimen orientation (Woo et al., 1991), handling, storage, and hydration (Woo et al., 1986a). They have lead to important findings regarding the physiological changes associated with growth and development (Woo et al., 1986b, 1991), the adaptation of ligaments and tendons to mobility (Woo et al., 1981c, 1982, 1987a), as well as the effects of injury and treatment (Woo et al., 1987b; Weiss et al., 1991). Contribution to Joint Function Joint motion is governed by the direction and magnitude of externally applied loads, ligament forces, contact between joint surfaces, and muscle activity. For the knee, motions include a combination of translations: 16 14 1st Cycle

12

10th Cycle

10 8 6

Loa

4

di

U nlo

2

ng

ng

Load (N)

1210

ad

i

0 0

0.2

0.4

0.6

0.8

1

Elongation (mm)

Figure 71.3 Hysteresis loops (1st and 10th cycle) obtained from cyclic loading of the femur–anterior– medial bundle-tibia complex of the human ACL in uniaxial tension. The area of hysteresis (i.e. the area between the loading and unloading curves) decreases with repetitive cycling demonstrating the phenomenon of “preconditioning.”

Functional Tissue Engineering of Ligament and Tendon Injuries 1211

proximal–distal, medial–lateral, and anterior–posterior, and rotations: internal–external, flexion–extension and varus–valgus. In total, these translations and rotations describe motion in six degrees of freedom (DOF). While evaluating joint function, it is important to note that constraining DOF of the knee can have significant impact on the results obtained (Inoue et al., 1987; Livesay et al., 1997). When knee motion was allowed in all directions, sectioning the MCL only resulted in small increases in valgus laxity (21%) suggesting that the ACL plays a significant role as a joint restraint to this knee motion. However, when anterior–posterior translation and internal–external rotation were constrained, valgus laxity increased significantly (171%) following sectioning of the MCL. For this reason, it is important to have a testing device which allows for unconstrained knee motion. For more than a decade, a robotic/universal force-moment sensor (UFS) testing system, which was developed by our research center, has been used to study knee kinematics as well as to directly measure the in situ forces in the knee ligaments in response to external loading conditions (Figure 71.4) (Fujie et al., 1993, 1995; Rudy et al., 1996). This methodology has been utilized to study many variables of ACL reconstruction. Most recently the limitations of single bundle reconstructions to restore rotatory stability along with the potential advantages of an anatomical reconstruction were demonstrated (Woo et al., 2002; Yagi et al., 2002). In addition, knee function following an isolated MCL injury in a goat model has also been studied (Scheffler et al., 2001; Abramowitch et al., 2003a).

HEALING OF LIGAMENTS AND TENDONS The events of healing of ligaments and tendons can be roughly divided into four overlapping phases:that is, hemorrhage, inflammation, repair (proliferation), and remodeling. Following injury, the hemorrhagic and inflammatory phases occur over the first several days. Minutes after the ligament injury, blood collects and forms a platelet-rich fibrin clot at the injury site. The hemorrhage phase of the injury forms a lattice for many

Figure 71.4 Schematic drawing illustrating the six degrees of freedom of motion of the human knee joint (permission pending from The Journal of Biomechanics).

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following cellular events. Triggered by cytokines released within the clot, polymononuclear leukocytes and lymphocytes appear within several hours. These cells respond to autocrine and paracrine signals to expand the inflammatory response and recruit other types of cells to the wound (Frank et al., 1994). The reparative phase takes place over the first couple of weeks to months following the injury. During this phase fibroblasts recruited to the injury site start forming healing tissue. Growth factors, including TGF-β and platelet-derived growth factor (PDGF) isoforms, are likely to be involved in modulating the healing environment in favor of effectively repairing the damaged ligament substance (Murphy et al., 1994). Meanwhile, increased neovascularization brings in circulating cells and nutrients to further enhance the healing process. The blood clot quickly turns into newly formed healing tissue which is composed of an aggregation of cells surrounded by a matrix whose histomorphological appearance and biochemical composition is different from that of the uninjured ligament. It is notably characterized by a homogenous distribution of smaller diameter collagen fibrils which is in stark contrast to the bimodal distribution of the normal ligament (Frank et al., 1992, 1997; Hart et al., 1992, 2000; Nakamura et al., 2000). Biochemically, it contains increased amount of proteoglycans, a higher ratio of type V to type I collagen, and a decrease in the number of mature collagen cross-links. The proliferative phase gives way to the remodeling phase. It occurs from months to years after the injury and is characterized by decreasing cellularity, decreasing levels of collagen type III, and realigning of the matrix to respond better to the forces applied to the tissues. On the other hand, the diameter of collagen fibrils have been found to remain small and levels of collagen type V have been found to remain elevated for years after injury (Adachi and Hayashi, 1986; Birk et al., 1990; Frank et al., 1992, 1997; Hart et al., 1992, 2000; Marchant et al., 1996; Nakamura et al., 2000; Niyibizi et al., 2000; Birk, 2001). Interestingly, the type V collagen has been shown to play a central role in the regulation of the lateral growth of collagen fibrils. The elevated type V collagen could be involved in the lack of large collagen fibrils which in turn are associated with the inferior mechanical properties of healing tissue (Parry et al., 1978; Doillon et al., 1985).

MCL OF THE KNEE The healing process of the MCL follows this general wound healing pathway and has been well studied. Thus, it serves as a good model for the histological, biochemical, and biomechanical events. The process of ligament healing is also greatly impacted by the selection of treatment (Clayton et al., 1968; Tipton et al., 1970; O’Donoghue et al., 1971; Woo et al., 1987b, 1990). Laboratory and clinical studies have shown mobilization is superior to immobilization (Woo et al., 1987a; Inoue et al., 1990; Weiss et al., 1991). Interestingly, non-operative repairs have an equivalent healing outcome to surgical repairs. A severe “mop-end” injury model in the rabbit, developed in our research center, that causes a midsubstance tear and damage at the insertion sites (Weiss et al., 1991), along with non-operative treatment with mobilization was used to compare to surgical repair with mobilization. After 52 weeks of healing, there were no significant differences in varus–valgus rotation of the knee, in situ force of the MCL, or tensile properties between repaired and non-repaired MCL (Weiss et al., 1991). Based on these and other studies clinical management has shifted from surgical repair with immobilization to non-operative management with early controlled range-of-motion exercises as soon as pain subsides (Reider et al., 1994; Indelicato, 1995). While the MCL heals with non-operative treatment and the stiffness of the healing FMTC begins to approach normal levels by 52 weeks after injury, the CSA of the healed tissue increases with time, measuring as much as 2.5 times its normal size (Inoue et al., 1990). Thus, the mechanical properties of the healing MCL midsubstance remain consistently inferior to those of the normal ligament and do not change with time. Thus, the healing process involves making a larger quantity of lesser quality ligamentous tissue. Moreover, the rate of healing between the ligament midsubstance and the insertion sites is asynchronous with the insertion sites demonstrating a lower stiffness and strength resulting from injury as well as a lack of stress during the healing process.

Functional Tissue Engineering of Ligament and Tendon Injuries 1213

There has been evidence that indicates that activity level influences the rate of healing (Abramowitch et al., 2003a). The goat model is a more clinically relevant model to study ligament healing than the rabbit due to its more robust activity level. In addition, its large size and the previously published success of ACL reconstructions using this animal (Ng et al., 1996), make it attractive to study more complex multiple-ligament injuries. Generally, these models display similar trends. However, it was noted that the stiffness and ultimate load of the healing goat FMTC are closer to control values at earlier time periods when compared to data from the rabbit model, suggesting that activity level may influence the healing response. ACL Reconstruction With the ultimate goal of ACL reconstruction being to restore knee function, the success of these procedures is dependent on a number of surgical, biomechanical, and biological factors. The most popular choice being autografts from the PT (i.e. bone-central third of the patellar tendon-bone (BPTB)), or hamstring tendons (i.e. semitendinosous plus gracilis tendons). Allografts, including the Achilles tendon, BPTB, and hamstring tendons, have seen limited use except in revision surgery or for multiple ligamentous injuries. BPTB grafts are generally considered the “gold standard” for ACL reconstruction because it facilitates better fixation and bone-to-bone healing inside the bone tunnels (Jones, 1970; Lambert, 1983; Noyes et al., 1984; Kurosaka et al., 1987; Aglietti et al., 1992; Cooper et al., 1993). However, the major drawback is that the open defect of the donor site remains visible and is not completely healed for months (Coupens et al., 1992; Rubinstein et al., 1994; Cerullo et al., 1995; Nixon et al., 1995). This contributes to a higher incidence of complications including donor site morbidity, patella baja, arthrofibrosis, adhesion to the fat pad, and patellofemoral pain (Paulos et al., 1987; Tibone and Antich, 1988; Sachs et al., 1989; Shelbourne et al., 1991; Breitfuss et al., 1996; Kartus et al., 1999). Efforts have been made to examine the healing PT after harvest, using animal models, after removal of the central third. Studies have found a deterioration of PT structural properties with a concomitant increase in the CSA of the PT tissue (Kamps et al., 1994; Linder et al., 1994; Beynnon et al., 1995; Awad et al., 2003; Tohyama et al., 2003). Specifically in the rabbit model, the ultimate load of the entire BPTB complex decreased by 38% (Beynnon et al., 1995), while a CSA increase of 83–108% was observed at 12 weeks post-harvesting (Awad et al., 2003; Tohyama et al., 2003). For the central healing tissue, its tangent modulus and ultimate tensile strength were only measured to be 15% and 18% of controls after 26 weeks, respectively (Awad et al., 2003). The mechanical properties of the remaining PT tissues also deteriorated compared to sham controls after 24 weeks (Tohyama et al., 2003). Following implantation, the autograft becomes inflamed and necrotic leading to a decrease in graft stiffness and strength (Tohyama and Yasuda, 2000). The graft undergoes revascularization and repopulation with fibroblasts followed by a remodeling period with restructure of collagen fibers and proteoglycans (Arnoczky et al., 1982). Further, bone-to-bone healing and tendon-to-bone healing within the femoral and tibial tunnels revealed that there was complete incorporation of the bone block by 6 weeks, but incomplete incorporation at the tendon–bone interface (Papageorgiou et al., 2001). For the latter, the failure mode of the femur–graft–tibia complex (FGTC) consistently occurred as a pull-out from the tibial tunnel. Over time, however, experimental animal studies show that the FGTC gradually shows improvement (Ballock et al., 1989; Butler et al., 1989; Gerich et al., 1996), but its structural properties failed to be restored to levels of the intact femur–ACL–tibia complex (FATC) even after 12 months (Clancy et al., 1981; Ballock et al., 1989; Butler et al., 1989). It is agreed that accelerating graft incorporation and healing may lead to an earlier return to sports and normal activities, and therefore has become a goal of tissue engineering efforts, as discussed later in this chapter. In addition to the graft selection, other important surgical decisions include tunnel placement, graft tension, and fixation. There has been a substantial amount of research which is focused on the impact of these

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THERAPEUTIC APPLICATIONS: TISSUE THERAPY

variables at time zero and following various periods of healing both in human and animal models (Yagi et al., 2002; Abramowitch et al., 2003b; Loh et al., 2003). All these parameters can lead to various degrees of graft tunnel motion which may impact graft integration and graft healing. Ultimately, these factors impact the post-operative rehabilitation and the return to normal activities and sports (Tsuda et al., 2002). Thus, it is important to take into account the large changes in properties of the graft after implantation.

Combined Ligamentous Injuries Combined ACL/MCL injuries occur frequently and the best methods for treatment are still controversial. Some surgeons elect to surgically reconstruct the ACL without addressing the MCL, while others advocate reconstruction of the ACL with repair of the MCL. Regardless of the treatment modality, clinical and basic science studies continue to show that the outcome of this injury is worse than for an isolated MCL injury. Our research center has elucidated the effects of ACL deficiency on the healing of the injured MCL using canine, rabbit, and goat model (Woo et al., 1990; Ohno et al., 1995; Yamaji et al., 1996; Abramowitch et al., 2003c), thus ACL reconstruction has been suggested. Further, repairing the MCL in combination with ACL reconstruction resulted in reduced valgus laxity and improved the structural properties of the FMTC in early stages, but the long-term effect became minimal (Ohno et al., 1995; Yamaji et al., 1996). These data further suggest that only reconstruction of the ACL is necessary for successful healing of the MCL after a combined ACL/MCL injury. Recently, a larger animal model (i.e. the goat knee) was used in our research center to examine the function of the knee and quality of the healing MCL after a combined ACL/MCL injury treated with ACL reconstruction (Abramowitch et al., 2003c). These results confirmed that valgus rotation was twice that for an isolated MCL injury. Moreover, the structural properties of the FMTC and tangent modulus of the MCL substance are all substantially lower than that for the isolated MCL injury (Scheffler et al., 2001; Abramowitch et al., 2003c). These results demonstrate a clear need to enhance ligament healing after such a severe knee injury, requiring improved treatment strategies.

APPLICATION OF FTE FTE emphasizes the importance of biomechanical considerations in the design and development of cell and matrix-based implants for soft and hard tissue repair. Musculoskeletal tissues, especially ligaments and tendons, are accustomed to being mechanically challenged, therefore tissue engineered constructs used to replace these tissues after injury or disease must meet these requirements. By combining the fields of molecular biology, biochemistry, and biomechanics, novel therapeutic approaches (e.g. growth factors, gene transfer/gene therapy, cell therapy, and biological scaffolds) offer the possibilities for improvement of the treatment of ligament and tendon injuries. The following will be a brief review of the current available approaches to enhance ligament and tendon healing.

Growth Factors Growth factors can induce wide ranging effects on cell function including migration, proliferation, and protein synthesis. The application of exogenous growth factors is based on the premise that they can promote ligament regeneration that will lead to a biologically and biomechanically superior healed ligament substance. Many studies, in vitro and in vivo, have tried to define the role of growth factors in ligament and tendon healing and to determine appropriate strategies for the use of growth factors for these structures (Steenfos, 1994; Duffy et al., 1995; Panossian et al., 1997; Sciore et al., 1998).

Functional Tissue Engineering of Ligament and Tendon Injuries 1215

In Vitro Studies Cell culture or tissue explant methodologies involving the addition of exogenous growth factors have been the major study designs. Measured responses include cell proliferation, synthesis of ECM proteins such as collagen, proteoglycans, tissue remodeling enzymes, and cell migration or chemotaxis. In our research center, the effects of eight different growth factors on the MCL and ACL fibroblast culture were determined for proliferation and ECM production (Ohno et al., 1995; Deie et al., 1997; Marui et al., 1997; Scherping et al., 1997). In terms of cell proliferation, PDGF-BB, epidermal growth factor (EGF), basic fibroblast growth factor (bFGF) have been found to have a significant effect on cell proliferation and caused greater proliferation in MCL fibroblasts versus ACL fibroblasts (Scherping et al., 1997). We found that the proliferation of MCL and ACL fibroblasts from skeletally immature rabbits increased by 7.6 times in response to EGF and 5.6 times in response to bFGF (Ohno et al., 1995). The same study in skeletally mature rabbits showed that insulin-like growth factor IGF and bFGF also had significant effects on fibroblast proliferation in both cell types, but the difference was less pronounced (Scherping et al., 1997). The biological role of TGF-β1 in ligament healing has also been recently addressed. Studies in our research center on fibroblast proliferation in skeletally immature rabbits demonstrated that TGF-β1 stimulated proliferation of MCL fibroblasts 1.3–1.4 times greater than in ACL fibroblasts (Ohno et al., 1995). Subsequent studies on skeletally mature rabbits showed little effect of TGF-β1 on cell proliferation for either fibroblast type. Comparison of these results suggests that age has a significant effect on the ability of growth factors to stimulate fibroblast proliferation (Scherping et al., 1997). The effect of TGF-β1 on canine ACL fibroblast proliferation was shown to be dose-dependent because smaller doses acted synergistically with PDGF whereas higher concentrations inhibited the stimulatory effect of PDGF (Desrosiers et al., 1995). These findings show the complex interactions of growth factors to enhance proliferation of fibroblasts. In terms of in vitro protein synthesis in MCL and ACL fibroblasts, collagen synthesis increased 160% over controls in both MCL and ACL fibroblasts treated with TGF-β1, and the majority of this increase was for type I collagen (Marui et al., 1997). The relative increase in protein production was similar for both cell types, but the absolute increase in protein synthesis was twice as much for MCL fibroblasts as for ACL fibroblasts. These data suggest that TGF-β1 may improve ligament healing by increasing matrix synthesis during the proliferative and remodeling phases (Marui et al., 1997). Similar results have also been found by other investigators (Desrosiers et al., 1995). These studies illustrate the ability of TGF-β1 to increase the production of ECM by fibroblasts in vitro. In vitro models, however, are limited in the extent that they cannot reproduce the complex interplay of signals affected by growth factors in the intricate process of ligament or tendon healing. Differences in the effects of different growth factors on cell proliferation and matrix synthesis suggest that wound healing depends on a highly integrated biochemical network of cell signaling events with intrinsic stimulatory and inhibitory feedback loops. Thus, in addition to providing a better physiological model, in vivo studies are critical to defining the interaction of biology and biomechanics and the degree to which healed ligament or tendon substance restores the structural and mechanical properties of the native tissue. In Vivo Studies In vitro studies showed that EGF and PDGF-BB have the greatest effect on ligament fibroblast proliferation, whereas TGF-β1 superiorly promotes ECM production. These growth factors were then applied in vivo at different dosages, in isolation and in combination, for an MCL injury in the rabbit model. It was found that a higher dose of PDGF-BB improved the structural properties of the FMTC compared to a lower dose of PDGFBB, demonstrating that the effects of PDGF-BB were dose-dependent (Woo et al., 1998). However, the mechanical properties of the ligament substance remained unchanged from untreated controls demonstrating that the improved structural properties resulted from a larger quantity of tissue instead of tissue with improved

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quality. In contrast, the combination of EGF or PDGF-BB plus TGF-β1 did not lead to additional improvements in MCL healing compared with PDGF-BB alone. Other investigators have also found that higher doses of PDGF demonstrated a plateau effect in improving the structural properties of the healed ligament. Furthermore, administration of PDGF for more than 24 h after the injury markedly decreased the efficacy of growth factors in improving MCL healing (Batten et al., 1996). Additionally, administration of TGF-β2 to the healing rabbit MCL improved the mechanical stiffness, but not the load at failure, of the ligament scar (Spindler et al., 2003). These data show that in vivo application of growth factors is indeed more complex than in vitro studies. One possible approach to improve the in vivo application might be to combine growth factors with gene transfer technology. Adenoviral bone morphogenetic protein-2 (AdBMP-2) delivered to the bone–tendon interface using a gene transfer technique has been shown to improve the integration of semitendinosous tendon grafts in rabbits (Martinek et al., 2002). The stiffness (29.0 7.1 N/mm versus 16.7 8.3 N/mm) and the ultimate load (108.8 50.8 N versus 45.0 18.0) were also significantly increased in the specimens with AdBMP-2 compared to untreated controls. Hence, the enhancement of tendon to bone healing is a promising approach to accelerate the patient’s return to activity. Based on these studies, an optimal therapy of introducing growth factors to injury sites is still an open question. While promising, timing of application, mode of delivery, and dosage, remain as major hurdles that need to be crossed before success can be achieved in vivo. Gene Therapy Gene therapy is a potential approach to improve ligament and tendon healing. Foreign nucleic acids, gene transfer, can be introduced into cells to alter protein synthesis or induce the expression of therapeutic proteins. Modern gene therapy relies on mammalian viruses and cationic liposomes as delivery vectors, and both have been developed to deliver genes into host tissue via the direct and indirect methods. Direct gene transfer involves in vivo injection of the delivery vector into the host tissue. Indirect gene transfer involves in vitro transduction of host cells with the desired gene, and subsequent replantation of these cultured cells in vivo. Studies have shown that PT fibroblasts can be transduced with the LacZ marker gene both directly using an adenovirus liposomal vector, and directly using a retrovirus, with the expression of the transferred genes persisting for 6 weeks following the application (Gerich et al., 1996). PT fibroblasts staining positively for β-galactosidase were subsequently found to migrate and incorporate into the tendon tissue following injection. In our research center, we sought to determine if genes could be transduced into MCL and ACL fibroblasts and if ligament injury affected gene transfer and expression (Hildebrand et al., 1999). When both the direct and indirect methods were employed using adenovirus and BAG retrovirus, respectively, it was found that both techniques resulted in expression of the LacZ marker gene by fibroblasts from intact as well as injured ligaments. Gene expression lasted longer (6 weeks) with the direct method as compared to the indirect technique (3 weeks). Fibroblasts from injured ligaments showed transduction both in the wound site and in the ligament substance as well. There was no difference in the duration of gene expression by fibroblasts from intact and injured ligaments suggesting that injury does not affect gene transfer or expression (Hildebrand et al., 1999). Newer techniques for gene transfer have recently been reported. Gene transfer using liposomal vectors may reduce the adverse immune responses seen when using viral vectors. Antisense gene therapy involving blocking the transcription or translation of specific genes which may be excessively expressed within healing tissue has been proposed. By the binding of antisense oligodeoxynucleotides (ODN) to target DNA investigators have performed direct transfer of an HVJ–liposome complex containing a labeled ODN for the protein decorin, which has been shown to inhibit type I collagen fibril formation in prior in vitro studies (Nakamura et al., 2000).

Functional Tissue Engineering of Ligament and Tendon Injuries 1217

Histological analysis was performed 24 h after direct injection into rabbit MCL specimens which were 2 weeks post-injury. Quantitative analysis revealed a transduction efficiency of 62% at 1 day and 23% at 7 days. Significant suppression of decorin mRNA expression was seen at both 2 days (42.7%) and 2 weeks (60.3%) (Nakamura et al., 2000). In our research center, we have evaluated the efficacy of utilizing ODNs to regulate the overproduction of collagens III and V (Shimomura et al., 2003; Woo et al., 2004). Normal human patellar tendon fibroblasts (HPTFs) were transfected with antisense collagen III or V ODNs by mixing with lipofectamine. The uptake of the ODNs was detected as early as 1 h and as late as 3 days after delivery. The relative expression of collagen V mRNA was reduced to 67.8 5.1% of missense levels. Also, preliminary reverse transcription-polymerase chain reaction (RT-PCR) results showed that the inhibitory effects of the collagen III antisense ODNs were most dominant at 1 day as the type III collagen mRNA level was 38.9 19.6% of missense controls. At 3 and 7 days, differences could not be observed. These results suggested that antisense gene therapy can indeed be a potential FTE approach to enhance the quality of ligaments and tendons. Despite these promising results, several obstacles currently impede the practical implementation of gene transfer as a biological intervention in ligament healing. The immune reaction against these antigens decreases the expression of the introduced gene (Tripathy et al., 1996). In addition, retroviral infection of fibroblasts often leads to shut-off of the promoter region, which adversely affects expression of the incorporated gene (Krall et al., 1994). Thus, delivering the ODNs to the appropriate target and reproducibility of the results remains a great challenge. Newer strategies in the evolving field of gene transfer include the search for more effective and less immunogenic vectors, modification of promoters to ensure gene expression after incorporation, and temporary and self-limiting gene expression regulation tailored to the changing environment of the healing ligament. As the complex steps involved in gene expression and regulation are further elucidated, the potential therapeutic efficacy of gene transfer is likely to enjoy practical application. Cell Therapy Cell therapy is another potential strategy to enhance ligament and tendon healing. Studies have focused on the application of mesenchymal stem cells (MSCs), bone marrow-derived cells (BMDCs) and synovial tissuederived fibroblasts into the healing site (Young et al., 1998; Watanabe et al., 2002). BMDCs have been shown to play an important role in wound healing (Badiavas et al., 2003; Galiano et al., 2004; Mathews et al., 2004) and can be obtained in high numbers with relative ease (Awad et al., 2003; Juncosa-Melvin et al., 2005). In one study, autologous marrow-derived progenitor cells were seeded on a collagen gel, and subsequently contracted onto a pretensioned suture (Young et al., 1998). The resulting tissue prosthesis was then implanted into the rabbit Achilles tendon gap defect. Significantly greater structural and mechanical properties were seen after the implantation. The treated tissues had a significantly larger CSA, and their collagen fibers appeared to be better aligned than those in the matched controls. Recently, new methods geared toward PT healing have tried to fill the central third PT defect with collagen gels filled with BMDCs, in which BMDCs were expanded in vitro. These MSC–collagen composites were implanted into full thickness, full length, central defects created in the PTs of rabbits. The healing PTs treated with MSC–collagen gels were one-fourth of the maximum stress of the normal central portion of the PT. The modulus and maximum stress of the repair tissues grafted with MSC–collagen gels increased at significantly faster rates than did natural repairs over time. Thus, overall improved mechanical properties were seen when compared to non-treated defects (Awad et al., 1999, 2003; Butler, 2005; Juncosa-Melvin et al., 2005). This particular cell therapy is attractive because the use of autogenous cells would minimize the immune response at the injury site.

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Alternatively, fibroblasts, myoblasts, and bone marrow cells have also been transplanted into injured ligaments following the induction of marker genes or stimulation by growth factors in vitro. These results show great potential in vivo (Day et al., 1997; Caplan et al., 1998; Hildebrand et al., 1999; Watanabe et al., 2002). However, issues remain on cell therapy. MSCs from bone marrow are relatively few in number and their numbers decrease further after transplantation. Thus, it is essential to develop in vitro techniques to expand MSCs without altering their differentiation potential. Further, scaffolds may help to prevent the loss of these cells and might further enhance the effect already provided by this strategy. Scaffolding An ideal scaffold should provide a suitable mechanical and biological environment for cells to migrate into, guide the formation of the newly synthesized ECM, and then slowly degrade such that the new matrix begins to bear the mechanical loads. Both synthetic and naturally occurring biological scaffolds are commonly utilized in FTE of ligaments and tendons (Aragona et al., 1983; Dunn et al., 1992; Badylak et al., 1999; Bourke et al., 2004). The major advantage of using synthetic polymers as scaffolds is their ease of fabrication and reproducibility. A structure can be created that mimics the structure of a ligament or tendon, and appropriate proteins (e.g. growth factors) can be incorporated into the scaffold during manufacturing. However, performance of current synthetic grafts in vivo have been disappointing (Bellincampi et al., 1998; Guidoin et al., 2000); thus, current work has focused on the use of synthetic scaffolds seeded with fibroblasts (Cao et al., 1994; Lin et al., 1999), or alternatively, a number of naturally occurring biological scaffolds, including bovine pericardium (Integra Life Science), human dermal collagen (Alloderm), and porcine small intestinal submucosa (SIS) have been used. The SIS has shown most promising results in enhancing healing of both ligaments and tendons (Badylak et al., 1995, 1999; Dejardin et al., 1999, 2001; Musahl et al., 2004; Liang et al., 2006), as it possesses a structural hierarchy that is naturally arranged, and it is mostly comprised of collagen type I. Further, 40% of the SIS degrades within 1 month in vivo (Record et al., 2001) and its byproducts have been shown to be chemoattractants for cells (including BMDCs) (Badylak et al., 2001; Li et al., 2004; Zantop et al., 2005). Moreover, it contains many bioactive agents (growth factors, fibronectin, and so on) (Voytik-Harbin et al., 1997; McPherson et al., 2000; Hodde et al., 2002) and causes a limited inflammatory reaction (Allman et al., 2002). Based on these positive findings, our research center conducted multidisciplinary studies to determine the effect of SIS treatment on MCL in the short- and long- term (12 weeks and 26 weeks). The mechanistic hypothesis is that SIS would act as a guidance of neo-ligament tissue formation, limiting the cross-sectional growth of the healing tissue, thereby increasing the mechanical demand. The tissue would respond by decreasing the production of collagen type V which in turn would lead to an increase in collagen fibril diameters and improved collagen orientation in the healing ligament. Finally, these changes would result in an improvement in the mechanical properties of SIS-treated MCLs (Figure 71.5). It was observed that the histomorphological appearance and biochemical composition of the healing ligament were indeed closer to the normal ligament when compared to the non-treated ligament. Biochemical analysis revealed that collagen type V/I ratio decreased with SIS treatment, while TEM (transmission electron microscopy) showed a heterogeneous distribution of large collagen fibril diameters (Figure 71.6), and ultimately mechanical properties of the healing MCL were improved. Most importantly, the effects of SIS treatment persisted even up to 26 weeks (Liang et al., 2006). In addition, CSA in the SIS treated group decreased by 28% and tangent modulus increased by 33% compared to non-treated group, also stress at failure was 49% higher than non-treated. These findings demonstrate that a layer of SIS can act as a guide for neoligament formation by inducing better organization and limiting cross-sectional growth of the healing ligament. This, in turn, requires the mechanical properties of the SIS treated ligament to improve through the formation of larger collagen fibrils resulting from a lower collagen type V/I ratio.

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35 SIS treated

30 Stress (MPa)

Non-treated 25 20 15 10 5 0 0

5

10 Strain (%)

15

20

Figure 71.5 Typical stress–strain curves for SIS-treated and non-treated groups at 12 weeks post-injury (permission granted by Interscience-Wiley).

(a)

(b)

Figure 71.6 Transmission electron micrographs (70,000) of collagen fibrils in (a) sham operated MCL (I), SIS-treatedMCL (II) and non-treatedMCL( III) at 26 weeks post-injury. The arrow indicates the appearance of large fibrils between cells in the SIS-treated MCL. (b) The TEM appearance of both large and small fibrils (heterogeneity) in the pericellular area in the SIS-treated MCL (I) and non-treated MCL (II). The arrow indicates the large fibrils surrounding a cell process. F indicates fibroblast (permission granted by Interscience-Wiley).

We have extended the use of SIS on PT healing after the central third was harvested for ACL reconstruction (Figure 71.7). Since SIS has a preferred collagen alignment (Sacks and Gloeckner, 1999), it has the potential of contact guidance and of promoting cells to produce a newly more aligned deposited matrix (Brunette, 1986; Clark et al., 1990; Chen et al., 1998; Walboomers et al., 1999); and as a result, a concomitant set of improved

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SIS

Type V collagen

XSA

Appropriate mechanical environment

Increase fibril diameter

Improved mechanical properties

Figure 71.7 Schematic of possible mechanism of action of SIS on MCL healing response. mechanical and viscoelastic properties. Additionally, the chemoattractant degradation products (Li et al., 2004) and bioactive agents of SIS could enhance the rate of healing. Finally, the SIS scaffold could also form a barrier between the healing PT and the underlying fat pad, limiting adhesion formation to permit motion between them to take place. The maintenance of stress and motion would help the homeostasis of the remaining PT (Woo et al., 1981c, 1982). Together, these effects will limit problems associated with poor healing, such as excessive hypertrophy of the remaining PT, and limiting the deterioration of its mechanical properties. In a preliminary study, the effects of SIS on healing of a central third defect (3 mm width) of the rabbit PT was performed. By 12 weeks, the SIS treated group contained a large number of spindle shaped cells with more organized collagen matrix, while the non-treated group had a sparse distribution of cells with only patches of collagen. After the healing PT tissue was dissected, the CSA was 61% greater in the SIS-treated group compared to those in the non-treated group (5.0 2.0 mm2 versus 3.1 1.2 mm2, respectively). SIS-treatment also showed higher stiffness (33.9 14.0 versus 24.3 14.9 N/mm, or 38%) and ultimate load (67.7 25.8 versus 43.8 27.4 N, or 58%) compared to non-treatment. This study demonstrated that SIS-treatment shows the potential to increase the quantity of healing PT tissue and structural properties of the healing central BPTB complex after a surgically created central third PT defect. Thus, the results of morphology, histology, and structural properties are very encouraging for further investigation of this application. Recent studies have also shown the feasibility of enhancing ACL graft integration following reconstruction using a tri-phasic scaffold (Spalazzi et al., 2006). Fibroblast and osteoblasts were seeded on a section of the scaffold that mimicked the native environment of the ligament insertion and bone, respectively. It was found that the specific cell types populated and thrived in their respective phase and also each migrated into the middle phase, while each cell type expressed the appropriate type of genes for its particular matrix. Collectively these approaches demonstrate that scaffolds have many potential applications for improving the treatment for injured ligaments and tendons. Mechanical Factors Progress in FTE also has included the elucidation of the importance of mechanical stimuli on cells and on tissue development and remodeling (Huang et al., 1993; Banes et al., 1995; Eastwood et al., 1998; Hsieh et al., 2000; Altman et al., 2002; Wang et al., 2003b). Cyclic stretching of cells from ligaments and tendons in vitro has been shown to cause increases in collagen synthesis (Desrosiers et al., 1995; Hsieh et al., 2000) and changes in intra-cellular processes (i.e. different regulation of metabolic and inflammatory genes and calcium signaling) (Banes et al., 1995, 1999; Archambault et al., 2002; Ralphs et al., 2002; Wang et al., 2003a; Yang et al., 2004).

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Figure 71.8 Randomly aligned cells cultured on a smooth dish (upper left). Aligned cellsculture on dish etched with microgroove (upper right). Randomly aligned matrix produced by cells cultured on a smooth dish (lower left). Aligned matrix produced by cells culture on dish etched with microgrooves (lower right) (permission granted by Wang et al., 2003a).

Interestingly, fibroblasts alignment can be a result of the external mechanical environment. When fibroblasts are grown in microgrooved silicone surfaces instead of smooth culture surfaces, they become elongated and aligned within the microgrooves through contact guidance. Most importantly, the ECM that the cells produce is also aligned along the microgroove direction (Figure 71.8) (Wang et al., 2003a). Cells also align along the direction of the maximum principal strain in collagen gels (Eastwood et al., 1998). Similarly, in vitro studies have shown that multidimensional mechanical strains applied to BMDCs embedded in a collagen gel upregulated the gene expression of collagen types I and III and tenascin-C, which are typically expressed in fibroblasts (Altman et al., 2002). Recently, our research center developed a uni-axial stretching system to study the effects of the mechanical stimuli on cells seeded on the bio-scaffold (SIS). In order to first understand the effect of elongating the SIS on the alignment of cells, the scaffold was stretched for 24 h to 15% of its original length and was then seeded with cells. After 5 days, collagen fibers of the scaffold were more aligned compared to non-stretched controls. In addition, the cells seeded on the scaffold demonstrated a preferred alignment along the direction of stretch. Preliminary results have also been obtained for continuous cyclic stretching (15% at 1 Hz for 4 h/day for 5 days) of SIS. Again, it was observed that the collagen fiber organization of the SIS improved to a more aligned state when the SIS was both seeded and stretched, and the cells were also aligned along the stretching direction. Gene expression analysis is underway to determine the differences in matrix protein expression between the cyclic group and the constant elongated group. Since tendon and ligament fibroblasts are aligned with collagen fibrils in vivo, it is hoped that mechanical stimuli can align cells within the scaffold and produce a better organized collagen matrix that may further enhance the healing response when implanted in vivo.

SUMMARY AND FUTURE DIRECTIONS In this review, the biomechanical and biological problems facing healing and repair of ligament and tendon injuries were discussed. There have already been tremendous improvements to clinical treatment paradigms

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based on studies that have established a fundamental understanding of healing following ligament or tendon injury and the benefits of controlled mobilization. Nevertheless, many issues still remain. For ligaments and tendons that display healing potential after injury, the major challenges are recovering their normal ultrastructural appearance, biochemical composition, and mechanical properties. Specifically, increasing the fibril diameters of healing tissues by limiting the production of type V collagen and decorin as well as improving the alignment of healing tissue by guiding the organization of newly produced matrix are important steps to be taken. By manipulating the healing response at the molecular and cellular level and guiding tissue formation, the following FTE approaches may offer the potential to restore the properties of healing tissue to normal levels. We are particularly interested in bioscaffolds, such as the porcine SIS. When applied to a healing ligament or tendon in vivo, it serves as a substrate that provides contact guidance for cells to form more aligned collagen fibers with a concomitant improvement in mechanical and viscoelastic properties when compared to non-treated controls. Further, the chemoattractant degradation products and bioactive agents of SIS could enhance the rate of healing (Li et al., 2004), allowing better maintenance of stress and motion dependent homeostasis. More excitingly, the SIS can be modified in vitro by seeding BMDCs on the scaffold and applying cyclic stretching in order to increase its alignment. Hence, when applied in vivo, the tissue engineered scaffold could serve to accelerate the initiation of the healing process by improving the production and orientation of collagen that ultimately will help to make a better neo-ligament or tendon. On the other hand, for ligaments and tendons that do not heal following injury and require surgical reconstruction using replacement grafts (e.g. ACL reconstruction), the major challenge is to promote a remodeling response such that the graft maintains sufficient stiffness and strength to provide functional stability of the joint. Most importantly, enhancing the rate of integration of tendon–bone interfaces during early graft incorporation that may permit an earlier and more aggressive post-operative rehabilitation (Chen et al., 2002). These complex issues may require a combination of approaches including gene and cell therapies as well as biological scaffolds. Indeed, grafts treated with AdBMP-2 has shown some potential (Martinek et al., 2002) in both canine and rabbit models. Additionally, other biological tissues such as periosteum have also been used to enhance the interface between tendon and bone with some success (Chen et al., 2002). All these results suggest an exciting potential for clinical application. Indeed, FTE has generated many exciting developments. Further the development of stem cell-based therapies presents both opportunities and challenges. To translate the knowledge gained about a particular gene, protein, or cell to a clinical application will require that expertise from many disciplines to work in a seamless fashion. One of the roles of biomedical engineers within this framework would be to help link interactions of the functions of molecules to cells, cells to tissues, tissues to organs, and organs to body. When biologists, biomedical engineers, clinicians, as well as experts from other disciplines, work together this would result in better therapies that lead to the injured ligaments and tendons to heal with properties closer to those of normal ligaments and tendons. Efforts of such a team-based approach on the new developments of FTE will bring a bright future to the outcome of healing of ligaments and tendon injuries.

ACKNOWLEDGMENT The financial support provided by the National Institute of Health Grants AR41820 and AR39683.

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72 Tissue Therapy: Implications of Regenerative Medicine for Skeletal Muscle Shen Wei and Johnny Huard

INTRODUCTION Skeletal muscle is the largest tissue mass in the human body, constituting 40%–45% of total body weight (Garrett and Best, 1994). Up to 55% of all injuries sustained in sports are muscle-related injuries (Kaariainen et al., 2000). Therefore, muscle injury is of major concern in traumatology and sports medicine. The treatment of such injuries can pose challenging problems, particularly because skeletal muscle injuries heal slowly and often result in incomplete functional recovery. After skeletal muscle injury, the traumatized muscle undergoes sequential and overlapping phases of healing, including degeneration, inflammation, and regeneration. In cases involving severe injuries, such as those resulting from burns or laceration, the healing muscle also may undergo fibrosis. Each of these phases is unique in skeletal muscle healing, and all are potential targets of efforts to treat skeletal muscle injury. The optimization of each phase of muscle healing could facilitate full functional recovery of injured skeletal muscle. In this chapter, we discuss the natural healing process of injured skeletal muscle and the tissue-level therapeutic approaches targeted toward each phase of the healing process. The utilization of tissue engineering and gene and cell therapy approaches is discussed in other chapters of this book.

THE NATURAL HEALING PROCESS OF SKELETAL MUSCLE Skeletal muscle is a composite structure mainly consisting of muscle cells and myofibers, nerves, blood vessels, and extracellular connective tissue matrix. The skeletal muscle fiber (or myofiber) is a syncytium derived from the fusion of myoblasts; individual myofibers are surrounded by a connective tissue layer called endomysium. Fascicles or bundles of myofibers are grouped together and surrounded by perimysium. The epimysium is the outermost connective tissue layer surrounding the skeletal muscle (Garrett and Best, 1994). After muscle injury, satellite cells, the myogenic precursor cells located between the basal lamina and plasma membrane (Hurme and Kalimo, 1992), are activated to proliferate and differentiate into myoblasts and play a key role in muscle regeneration by fusing with damaged myofibers or forming new myofibers. The appearance of centralized nuclei within myofibers is an indication of muscle regeneration. Upon myofiber maturation, the nuclei move into the subsarcolemmal layer (Mauro, 1961; Campion, 1984) (Figure 72.1). Skeletal muscle injuries are common results of sporting events and athletic endeavors (Kaariainen et al., 2000). However, they also can result from daily life activities or military combat. Skeletal muscle injuries range

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Epimysium

Muscle

Perimysium Endomysium

Satellite cells

Myoblasts Differentiated myotubes

Mature myofiber

Figure 72.1 Schematic of the structure of skeletal muscle. In injured skeletal muscle, satellite cells are released and are activated to become myoblasts. They fuse with damaged myofibers or with each other to regenerate muscle tissue. Regenerating myofibers contain centrally located nuclei that migrate to the periphery of the myofiber upon myofiber maturation. Reprinted with permission from The Journal of Bone and Joint Surgery, Inc. in severity from common and relatively minor muscle strains and contusions to more unusual and particularly devastating muscle lacerations and burns. In addition, muscle injuries also can result indirectly from ischemia and neurologic dysfunction (Campion, 1984). Injury to skeletal muscle induces a healing response comprising various stages. The traumatized muscle sequentially undergoes degeneration, inflammation, regeneration, and, in cases of severe trauma, fibrotic tissue formation. Generally, the degeneration and inflammation phases start within hours after injury, the regeneration phase begins 3–5 days after injury, and fibrotic tissue first appears approximately 2 weeks after injury (Li et al., 2001; Huard et al., 2002) (Figure 72.2). However, the timing of these events can vary depending on the type of injury and the severity of damage. Furthermore, these healing stages are not distinct; they overlap with each other and interrelate mechanistically. Degeneration Trauma to muscle tissue destroys its normal structure. Proteases initiate the autodigestion of disrupted and necrotic myofibers (Ebisui et al., 1995; Mbebi et al., 1999). Cytokines up-regulate vascular permeability and

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Inflammation Degeneration

1 week

Regeneration Fibrosis

2 weeks

3 weeks

4 weeks

Weeks after muscle injury

Figure 72.2 The temporal sequence of four interrelated healing phases: degeneration, inflammation, regeneration, and fibrosis.

blood flow, which in turn lead is to the ensuing inflammatory response. Toxic free radical species develop and impair the already damaged muscle tissue and the healthy tissue nearby (Clanton et al., 1999). Researchers have made little effort to improve muscle healing by limiting muscle degeneration. The use of anti-oxidant or free radical scavengers has failed to elicit promising results (Childs et al., 2001; Beaton et al., 2002). The technical difficulty impeding such approaches is the fact that degeneration begins immediately after injury occurs; therefore, efforts to interrupt this phase are nearly impossible, and preventive techniques may have to be used. Inflammation Inflammation begins shortly after injury and overlaps with both degeneration and regeneration. Within hours after injury, neutrophils begin to infiltrate and migrate to the injury site. Different populations of macrophages shortly follow the neutrophils. Some macrophage populations are mainly involved in phagocytosis of tissue debris and may evoke more tissue damage (Lapointe et al., 2002), whereas other populations are primarily responsible for the production of early growth factors and cytokines (McLennan, 1993; St Pierre and Tidball, 1994). In addition to macrophages, myogenic cells can release growth factors and cytokines (Huard et al., 2002) that influence the subsequent regeneration and fibrosis phases. Limiting inflammation may reduce symptoms like pain and edema induced by cytokines and prostaglandins. However, the reduced production of growth factors and cytokines and prostaglandins may lead to delayed regeneration and overgrowth of fibrotic tissue. Regeneration Although satellite cell activation can occur as soon as 24 h after injury (Rantanen et al., 1995), the regeneration phase, evidenced by the appearance of centrally nucleated regenerating myofibers, usually starts 3–5 days after injury (depending on the injury type and severity). Regeneration is carried out mainly by satellite cells. In uninjured muscle, these cells reside quiescently in their niche between the basal lamina and the sarcolemma (Hurme and Kalimo, 1992). In response to injury-induced disruption of normal muscle structure, quiescent satellite cells are activated and begin to proliferate and differentiate in response to various growth factors, including insulin-like growth factor 1 (IGF-1), hepatocyte growth factor (HGF), basic fibroblast growth factor (bFGF), and transforming growth factor beta 1 (TGF-β1). Finally, the satellite cells fuse with existing myofibers in the injury area or fuse with each other to form regenerating myofibers (Li et al., 2001; Huard

Regenerative Medicine for Skeletal Muscle

et al., 2002). Recent reports indicate that prostaglandins play key roles in satellite cell fusion, a point we will discuss in the next section. Fibrosis Not all growth factors released at the injury site promote muscle regeneration. Some also stimulate cells to differentiate toward the fibroblastic lineage and form fibrotic tissue. As the fibrotic tissue becomes increasingly dense, it both hinders the regeneration of myofibers and prevents new axons from reaching myofibers and forming neuromuscular junctions (Kaariainen et al., 2000). Dense fibrotic tissue also causes decreased muscle contractility and range of movement (Shanmugasundaram, 1980). Most importantly, the presence of such tissue makes the repaired muscle more susceptible to re-injury. Research results have implicated TGF-β1 in pathogenic fibrosis within many tissues (Gaedeke et al., 2001; Ihn, 2002). After muscle injury, TGF-β1 stimulates muscle-derived stem cells to differentiate into fibroblast-like cells (Li and Huard, 2002; Li et al., 2004) and produce collagen type I, the major component of fibrous tissue (Ghosh, 2002). Our previous studies of lacerated skeletal muscle have shown up-regulation of TGF-β1 that persisted from 3 days after injury until 14 days after injury (Li et al., 2004; Shen et al., 2005). Each of these healing phases plays a unique role in skeletal muscle repair. By interfering with any of them, we can alter the progression of natural healing in skeletal muscle. All these healing phases are interrelated and should be regarded as integral components of the skeletal muscle healing process. For example, inhibiting inflammation may delay the regeneration phase and therefore slow the whole healing process (Shen et al., 2005). The promotion of regeneration cannot alone elicit complete recovery of severely injured muscle because such muscle also undergoes fibrosis (Sato et al., 2003). Efforts to improve skeletal muscle healing should incorporate different approaches aimed at each phase and must account for the interaction between the different phases of healing.

THE ROLE OF INFLAMMATION IN SKELETAL MUSCLE HEALING Inflammation phase, symptomized by “Rubor,”“Tumor,”“Calor,” and “Dolor,” is an important phase of natural healing in many injured tissues. During the inflammation phase, the injured tissue and coagulation cascade release various cytokines that increase the permeability of blood vessels and the chemotaxis of inflammatory cells. First neutrophils and then macrophages rapidly invade the injury site. The infiltration of inflammatory cells persists for several days after injury. These cells play pivotal roles in the inflammatory phase and also influence the subsequent phases. Inflammatory cells participate in the phagocytosis of necrotic tissue debris, which is necessary for the growth of regenerating tissue. The cytokines and growth factors released by these cells can initiate regeneration and fibrosis or cause further damage to the muscle (Tidball, 2005). Inflammatory Cells Studies have revealed a complex picture in which inflammatory cells promote both injury and repair through the combined actions of free radicals, growth factors, and cytokines. Neutrophils invade the muscle injury site as early as 1 h after injury and remain for up to 5 days (Fielding et al., 1993). Studies have shown that the invading neutrophils are phagocytic and can help to remove cellular debris (Lowe et al., 1995), a necessary step in skeletal muscle repair. However, most research on this topic has shown that neutrophils exacerbate muscle injuries by releasing reactive oxygen intermediates, including peroxides, hypochlorite, and superoxide (Jolly et al., 1986; Korthuis et al., 1988; Smith et al., 1989; Fielding et al., 1993; Tidball, 2002). It is not yet clear whether neutrophil-mediated damage is a necessary step in muscle repair.

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In contrast, macrophages mediate the regenerative process in several important ways. Macrophages secrete growth factors and cytokines that act in a paracrine fashion to stimulate myoblast-related regenerative events (Cantini et al., 1994; Massimino et al., 1997; Lescaudron et al., 1999), but no study has definitively shown which factors macrophages release in vivo to promote muscle healing. Our research has shown that the administration of NS-398, a COX-2 inhibitor, reduces the infiltration of macrophages, especially at the early time point of 24 h after injury (Shen et al., 2005). However, by reducing the number of infiltrating macrophages, the administration of a COX-2-specific inhibitor may impede the proliferation of satellite cells and reduce the secretion of some growth factors and cytokines necessary for regeneration. We believe that this inhibition of the inflammatory response was at least partially responsible for the delayed myofiber regeneration observed during the healing process in this study. A recent report also has shown that direct contact with macrophages can rescue myogenic precursor cells from apoptosis after muscle injury. The rescued cells can act synergistically with macrophages to amplify chemotaxis and enhance muscle growth (Chazaud et al., 2003). Furthermore, macrophages may fuse with myofibers directly to promote regeneration (Camargo et al., 2003). Cyclooxygenase Pathway and Prostaglandins Because the symptoms of inflammation are uncomfortable, attempts are always made to shorten or eliminate this phase. Non-steroidal anti-inflammatory drugs (NSAIDs) reduce prostaglandin synthesis by selectively or non-selectively inhibiting the cyclooxygenase (COX) enzymes. COX enzyme has three known isoforms: COX-1, COX-2, and COX-3. COX-1 is constitutively expressed and controls homeostasis in many tissues. COX-2 is an inducible isoform of COX that is expressed mostly in pathologic scenarios (McCormack, 1998). COX-3 is involved with the control of fever via the central nervous system (Warner and Mitchell, 2002). Studies have reported favorable effects of using NSAIDs to reduce muscle weakness and loss of function (Almekinders, 1999; Trappe et al., 2001). However, some research suggests that NSAIDs only have short-term beneficial effects on muscle healing and that NSAIDs treatment beyond the first week after injury has either no effect or a detrimental effect on the recovery of muscle strength (Almekinders and Gilbert, 1986; Obremsky et al., 1994; Mishra et al., 1995). Due to the different animal models, NSAIDs, and dosages and administration routes used by different research groups, outcomes of studies focused on NSAID treatment vary making it extremely difficult to compare the results. COX-2-selective inhibitors (such as celecoxib and valdecoxib) are new variants of NSAIDs. They have similar analgesic and anti-inflammatory properties and reduced side-effect profiles when compared with non-selective NSAIDs. By avoiding COX-1 inhibition, clinicians can essentially avoid interfering with homeostatic functions and side effects like platelet inhibition and gastric mucosal injury. Studies have shown that COX-2-selective NSAIDs have detrimental effects on muscle healing (Bondesen et al., 2004; Mendias et al., 2004; Shen et al., 2005). It has been reported that repetitive use of skeletal muscle up-regulates the COX-2 enzyme and prostaglandins. NSAIDs, whether selective or nonselective, can suppress the up-regulated synthesis of prostaglandins (Vandenburgh et al., 1995; Trappe et al., 2001). Thus, it is likely that the COX-2 enzyme and the prostaglandins play a pivotal role in the healing process of skeletal muscle. NSAIDs appear to delay muscle regeneration by blocking the COX-2 enzyme and the production of prostaglandins, thereby impeding the long-term functional recovery of injured muscle. The early functional improvement observed after NSAID administration is likely due to the reduced inflammatory symptoms. Recent studies have clearly demonstrated the involvement of COX-2 and prostaglandins in muscle repair and regeneration (Bondesen et al., 2004; Shen et al., 2005). Prostaglandins are potent mediators of inflammatory response and are involved in multiple aspects of muscle regeneration after injury (Prisk and Huard, 2003). The prostaglandins are synthesized from arachidonic acid released from membrane phospholipids by

Regenerative Medicine for Skeletal Muscle

NSAIDs



COX-2

PGF2, PGE2

Muscle growth and regeneration

Figure 72.3 The COX pathway and its final products, PGE2 and PGF2α, are very important for skeletal muscle regeneration. The use of NSAIDs can delay skeletal muscle healing after injury by blocking the production of PGE2 and PGF2α.

phospholipase. COX enzymes are key to converting arachidonic acid into the prostaglandin H2 (PGH2). From PGH2, specific synthases are responsible for the further conversion into various forms of prostaglandins, including PGE2 and PGF2α (Hochberg, 1989). Generally, PGE2 and PGF2α mediate fever, pain, and smooth muscle contraction. In injured skeletal muscle, PGE2 and PGF2α play their specific receptor-mediated roles in inflammation, regeneration, and nociception (Hatae et al., 2002; Prisk and Huard, 2003; Sakamoto et al., 2004). Furthermore, they help to regulate muscle protein synthesis and degradation (Palmer, 1990) (Figure 72.3). PGF2α, in particular, can promote the growth of skeletal muscle by stimulating the secondary fusion between single muscle cells and nascent myotubes (Horsley and Pavlath, 2003, 2004). PGE2 appears to play multiple roles in the muscle inflammation phase, including chemotaxis of inflammatory cells, increasing vascular permeability and vasodilation, and induction of nitric oxide synthase and pro-inflammatory cytokine expression (Stenson et al., 1986; Tetsuka et al., 1994; Sakamoto et al., 2004). In addition, PGE2 is a potent down-regulator of TGF-β1-stimulated fibroblast proliferation and collagen synthesis (Frungieri et al., 2002; Scheuren et al., 2002). The production of PGE2 also may be important for inhibiting fibrosis, and the increased fibrosis observed after NSAID treatment may be due to the inhibited expression of PGE2 (Shen et al., 2005). Inflammation and Muscular Dystrophy Duchenne muscular dystrophy (DMD) is caused by mutations in the dystrophin gene and leads to loss of the dystrophin-glycoprotein complex (Hoffman et al., 1987). DMD results in membrane instability of myofibers, and persons with DMD develop progressive muscle weakness beginning at birth. Previous work has revealed possible involvement of uncontrolled chronic inflammation in DMD (Porter et al., 2002; Monici et al., 2003; Grounds and Torrisi, 2004), which is one of the main causes of its secondary effects. Acting on this concept, researchers studying DMD have developed therapies targeted at inflammatory components. Recent studies have shown that treatment with corticosteroids significantly slows the progression of DMD. The use of corticosteroids may prolong ambulation and upper limb and pulmonary function (Balaban et al., 2005). Although the administration of corticosteroids is accompanied by negative side effects, intermittent prednisone treatment generally is recognized as an effective way to preserve motor functions in ambulant persons with DMD (Beenakker et al., 2005).

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Regeneration Fibrosis

Inflammation

Skeletal muscle injury

Figure 72.4 It is important to appropriately control inflammation. Too much inflammation results in uncomfortable symptoms and severe tissue damage; too little inflammation may result in delayed regeneration and overgrowth of fibrotic tissue.

Because Tissue Necrosis Factor-α (TNF-α) expression is up-regulated in DMD patients and can promote inflammation, scientists have studied blocking TNF-α as a possible new therapeutic strategy. Research has shown that injection of dystrophic muscle with anti-TNF-α antibody (Remicade) before the onset of muscle necrosis can delay and reduce the breakdown of dystrophic muscle and has no adverse effect on new muscle formation (Grounds and Torrisi, 2004). Furthermore, the diaphragm muscles of MDX mice (a mouse model of DMD) treated with a soluble receptor fusion protein that binds TNF-α showed decreased mRNA for type I collagen and for TGF-β1. This finding suggests that blocking TNF-α may also attenuate fibrosis in DMD patients and therefore improve muscle function (Gosselin and Martinez, 2004). Inflammation is an important phase of skeletal muscle healing. Inflammatory cells, growth factors, and cytokines are all integral components of inflammation and the entire healing process. Therefore, at least a moderate degree of inflammation is necessary for proper healing of injured muscle tissue. Uncontrolled inflammatory responses, like those that occur in the muscles of DMD patients, can cause severe symptoms and destroy muscle tissue. However, the suppression of inflammation by NSAIDs can lead to insufficient regeneration and overgrowth of fibrotic tissue. Further research is required to determine the extent of inflammation that results in ideal muscle healing (Figure 72.4).

PROMOTING REGENERATION AFTER SKELETAL MUSCLE INJURY Growth Factors Muscle regeneration occurs early in the healing process. It usually begins from 3 to 5 days after injury, peaks during the second week after injury, and then rapidly declines (Huard et al., 2002). Various growth factors, including IGF-1, bFGF, and nerve growth factor (NGF), can improve muscle regeneration during this phase of muscle healing (Huard et al., 2002). Among these growth factors, IGF-1 has the greatest beneficial effect on the healing of injured skeletal muscle (Florini et al., 1996). IGF-1 is highly mitogenic for myoblasts and can increase both the efficiency of muscle regeneration and muscle strength in vivo (Engert et al., 1996). Systemic administration of IGF-1 can both increase the synthesis of muscle protein and reduce muscle protein degradation (Zdanowicz et al., 1995). Gene transfer of IGF-1 by an adeno-associated viral (AAV) vector into mouse skeletal muscle can block age-related loss of muscle mass and function (Barton-Davis et al., 1998). However, as a potent mitogen for fibroblasts, IGF-1 also can increase the synthesis of matrix components like collagen and lead to the formation of fibrotic tissue (Jones and Clemmons, 1995). Research has shown that bFGF and NGF also can improve muscle healing. Local injection of IGF-1, bFGF, and, to a lesser extent, NGF after injury increases the number and size of regenerating myofibers in different mouse injury models.

Regenerative Medicine for Skeletal Muscle

Table 72.1 Growth factors can stimulate or inhibit the proliferation and differentiation of muscle cells. Reprinted with permission from the The Journal of Bone and Joint Surgery, Inc. Growth factor

Cell proliferation

Cell differentiation

Hepatocyte growth factor (HGF) Basic fibroblast growth factor (bFGF) Insulin-like growth factor-1 (IGF-1) Nerve growth factor (NGF) Leukemia inhibitory factor (LIF) Acid fibroblast growth factor (aFGF) Platelet-derived growth factor (PDGF-AA) Platelet-derived growth factor (PDGF-BB) Epidermal growth factor (EGF) Transforming growth factor-α (TGF-α) Transforming growth factor-β1 (TGF-β1)

Stimulates Stimulates Stimulates Stimulates Stimulates Inhibits Inhibits Stimulates Inhibits Inhibits Inhibits

Stimulates Stimulates Stimulates Stimulates Stimulates Stimulates Stimulates Inhibits Inhibits Inhibits Inhibits

This histological evidence of improved muscle healing was supported by the observation of improved muscle strength 2 weeks after injury (Kasemkijwattana et al., 1998; Menetrey et al., 2000). Other growth factors, including HGF, platelet-derived growth factor (PDGF), and vascular endothelial growth factor (VEGF), also may be able to enhance muscle regeneration through different mechanisms (Huard et al., 2002) (Table 72.1). Neutralization of Myostatin Table 72.1 lists the numerous stimulatory growth factors that can influence muscle regeneration. Recent research has shown that neutralization of muscle growth inhibitory factors, such as myostatin (Mstn, also known as growth differentiation factor 8 (Gdf-8)), also leads to increased muscle regeneration in MDX mice (Bogdanovich et al., 2002). Myostatin belongs to the TGF superfamily and plays an important role in downregulating skeletal muscle growth (McPherron et al., 1997). Mice lacking myostatin exhibit a dramatic and widespread increase in skeletal muscle mass due to myofiber hypertrophy and hyperplasia (McPherron et al., 1997). Mstn negatively regulates muscle regeneration by controlling satellite cell activation and by regulating the migration of myoblasts and macrophages to the site of injury (McCroskery et al., 2005). Thus, antagonists of Mstn could be useful as pharmacologic agents for the treatment of muscle disorders. Some research suggests that other molecules regulate Mstn. Immunohistochemistry has shown highly co-localized expression patterns of Mstn and TGF-β1 3 days after injury, which suggests that these molecules play related roles in injured skeletal muscle (Shen et al., 2005). A recent in vitro study has shown that exogenous TGF-β1 (2 ng/ml) significantly increased myostatin levels in both proliferating and differentiating C2C12 myoblasts, whereas silencing of the TGF-beta1 receptor II gene significantly lowered myostatin levels in the myoblasts (Budasz-Rwiderska et al., 2005). Follistatin is another glycoprotein that can antagonize numerous members of the TGF-β superfamily, including Mstn. Follistatin antagonizes Mstn by direct protein interaction, which prevents Mstn’s inhibitory effect on muscle development (Amthor et al., 2004). Therapeutic Ultrasound Therapeutic ultrasound (US) is commonly used in the rehabilitative setting to elicit thermal or non-thermal physiologic effects. Clinicians frequently use therapeutic US treatments to enhance repair of tissue injuries

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and to reduce the associated pain. Some studies have shown that continuous therapeutic US can improve force production after contraction-induced muscle injury (Forester et al., 1982; Karnes and Burton, 2002). However, other studies have shown that US has little effect on skeletal muscle regeneration (Rantanen et al., 1999; Wilkin et al., 2004). Treatment with pulsed US can promote satellite cell proliferation but does not seem to have any significant effect on the overall morphologic manifestations of muscle regeneration (Rantanen et al., 1999). Due to the lack of scientific evidence, the use and prescription of therapeutic US to enhance skeletal muscle regeneration is often based on clinicians’ personal opinions and experience. Hyperbaric Oxygen In cases of trauma, circulatory insufficiency can lead to ischemia, which is associated with significant morbidity and mortality. Reperfusion further contributes to the morbidity. Researchers have advocated treatment with hyperbaric oxygen (HBO) for a variety of conditions of which tissue ischemia is the underlying problem. HBO has prevented the development of gangrene in humans and animal studies by reducing skeletal muscle necrosis and edema (Strauss et al., 1983; Skyhar et al., 1986). HBO treatment is believed to enhance the recovery of blood flow and functional capillary density in post-ischemic muscle tissue, which leads to the attenuation of microvascular dysfunction or damage due to ischemia (Sirsjo et al., 1993). Other research has shown that HBO treatment may improve the healing of ischemic muscle by reducing glutathione depletion and improving metabolic restitution (Haapaniemi et al., 1995). More recently, researchers have begun to use models of muscle injuries other than ischemia to study HBO treatment. Seven days after injury, HBO treatment resulted in superior healing (as evaluated by both functional and morphologic measures) in rabbits subjected to muscle stretch injury (Best et al., 1998). HBO also hastened the functional recovery and myofiber regeneration of skeletal muscle in rats modeling myotoxic injury (Gregorevic et al., 2000). However, HBO treatment has not been effective when used as therapy for exercise-induced muscle injuries in human studies (Harrison et al., 2001; Webster et al., 2002; Germain et al., 2003). Exercise and Muscle Regeneration Using muscle transplantation, researchers have studied the effects of pre- and post-transplantation exercise on satellite cell activation and the eventual regeneration of skeletal muscle transplants (Roberts and McGeachie, 1992). The morphological analysis revealed hypertrophy of the muscle fibers of the exercised transplants compared with controls, and an increase in the number of capillaries in the exercised transplants. Other evidence supports the efficacy of using exercise to induce satellite cell proliferation and muscle growth (Darr and Schultz, 1987; Smith et al., 2001). There is also theoretical support in the literature suggesting that exercise assists in promoting normal growth and repair of mammalian skeletal muscle (Esser and White, 1995; Wanek and Snow, 2000). However, results from a recent study challenge the efficacy of using exercise to improve muscle regeneration. Markert et al. (2005) reported that the exercise protocols used in their investigation did not noticeably enhance skeletal muscle regeneration after contusion injury in rats.

FIBROSIS PREVENTION: ANOTHER APPROACH TO IMPROVING SKELETAL MUSCLE HEALING Fibrosis, which begins 2 weeks after skeletal muscle injury and continues over time, hinders muscle regeneration and prevents full recovery of strength in injured skeletal muscle. Although the administration of exogenous growth factors (e.g. IGF-1, bFGF, or NGF) can enhance muscle regeneration, it does not prevent fibrosis in injured muscle. However, innovative approaches designed to prevent or limit fibrosis have had beneficial

Regenerative Medicine for Skeletal Muscle

effects on skeletal muscle healing (Fukushima et al., 2001; Foster et al., 2003; Chan et al., 2005; Negishi et al., 2005). TGF-β1 and Fibrosis TGF-β1 is involved in the development of fibrosis in various tissues (Gaedeke et al., 2001; Ihn, 2002). Of particular relevance to the topic of this chapter, TGF-β1 is expressed at high levels after skeletal muscle injury (Li et al., 2004; Shen et al., 2005) and is associated with fibrosis in the skeletal muscle of persons with DMD (Yamazaki et al., 1994; Gosselin et al., 2004). Muscle biopsy specimens from persons with dermatomyositis also contain excess TGF-β1 (Confalonieri et al., 1997). This excess TGF-β1 leads to chronic inflammation, fibrosis, and accumulation of extracellular matrix. These results support the theory that the expression of TGF-β1 in skeletal muscle may play an important role in the fibrotic cascade observed in diseased or injured muscle. Therefore, it is conceivable that neutralizing TGF-β1 expression in injured muscle could inhibit the formation of scar tissue. Anti-Fibrotic Therapy Based on Blocking Overexpression of TGF-β1 In an effort to minimize the effects of TGF-β1 in injured muscle, researchers have studied the effects of several molecules, including decorin, suramin, interferon-gamma (IFN-γ), and relaxin. Decorin is a small dermatan sulfate proteoglycan that helps to constitute extracellular matrix in collagen-containing tissues (Yamaguchi et al., 1990). By directly binding to TGF-β1, decorin participates in a feedback system that regulates cell growth and could prevent the fibrotic activity of TGF-β1 (Yamaguchi et al., 1990). Direct injection of decorin into lacerated muscles results in nearly complete functional recovery within 2 weeks of injection (Fukushima et al., 2001). Suramin was originally designed to be an anti-parasitic drug. However, it also blocks the effect of TGF-β1 by competing with TGF-β1 receptors. Suramin effectively inhibits the in vitro proliferation of fibroblasts and the expression of fibrotic proteins (i.e., alpha-smooth muscle actin and vimentin). The injection of suramin 2 weeks after injury can efficiently prevent muscle fibrosis, enhance muscle regeneration, and enable improved functional recovery of injured muscles (Chan et al., 2005). IFN-γ is a TGF-β1 pathway inhibitor that can disrupt TGF-β1 signal transduction. IFN-γ blocks the endogenous collagen expression induced by TGF-β1. In addition, IFN-γ inhibits TGF-β1 signaling by inducing the expression of Smad7, an inhibitory Smad (Ulloa et al., 1999). Our studies have shown that IFN-γ treatment prevents muscle fibrosis and improves muscle healing by down-regulating the level of TGF-β1induced fibrotic protein expression (Foster et al., 2003). The polypeptide hormone relaxin is a member of the growing family of IGFs. In vitro studies of stimulated dermal (Unemori and Amento, 1990), lung (Unemori et al., 1996), and hepatic (Williams et al., 2001) fibroblasts have demonstrated that relaxin can reduce type I and type III collagen production and increase pro-collagenase synthesis. In injured muscle, relaxin treatment results in a dosedependent decrease in myofibroblast proliferation, down-regulated expression of the fibrotic protein α-smooth muscle actin, and the proliferation and differentiation of myoblasts in vitro (Negishi et al., 2005). The administration of anti-TGF-β1 agents has shown promise as a means to prevent the formation of fibrotic tissue, and blocking fibrosis is a useful approach to improving muscle healing.(Figure 72.5).

CONCLUSION Skeletal muscle is the largest organ in human beings, and the repair of injured skeletal muscle is one of the major concerns of sports medicine. This chapter describes the different phases of the natural healing process in skeletal muscle, including degeneration, inflammation, regeneration, and fibrosis. Although researchers have not yet determined how to facilitate complete recovery, studies have demonstrated that each phase of skeletal muscle healing has its unique role and that the different phases are interrelated. By blocking degeneration, regulating inflammation, promoting regeneration, and blocking fibrosis, it is possible to improve skeletal

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Myofibers

Decorin, IGF-I -INF, suramin, relaxin

Satellite cells

TGF-B1 Fibrotic cells (scar tissue)

Figure 72.5 TGF-β1 is the key molecule responsible for inducing fibrosis after muscle injury. Molecules like decorin and suramin can neutralize the effect of TGF-β1 and induce satellite cells to differentiate toward the myogenic lineage.

muscle healing. However, when investigating treatments for muscle injuries, researchers must consider the interrelatedness of the different phases and keep them in balance.

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73 Tissue Therapy: Central Nervous System Jordan H. Wosnick, M. Douglas Baumann, and Molly S. Shoichet

INTRODUCTION Traumatic injuries to the central nervous system (CNS) come in a multitude of forms, most of which lead to long-term or permanent reductions in the quality of life. Traumatic brain injury and traumatic spinal cord injury are particularly devastating, both emotionally and financially. In the United States, there are over 250,000 people living with spinal cord injury (SCI) and approximately 11,000 new spinal cord injuries are registered each year, in addition to 1.5 million new cases of traumatic brain injury. Children and young adults are disproportionately affected by both types of injury; a grim statistic that is compounded by the lifelong disability that usually results. The complexity of the CNS makes efforts to repair injuries very difficult. However, progress in tissue engineering in recent years has led to new approaches for the treatment of CNS injuries and renewed hope. This chapter focuses on tissue-engineering strategies – including those based on nerve guidance channels, drug delivery, and cell delivery – aimed at achieving functional recovery after SCI. Tissue engineering strategies for brain injury repair are predominantly focused on drug delivery systems and stem cell therapy, which are described elsewhere in this book. Physiological Events Surrounding Spinal Cord Injury The brain and the spinal cord, which together make up the CNS, are two of the most complex organs in the human body. The spinal cord serves as the communication pathway between the brain and the periphery, in addition to coordinating reflex responses without the involvement of the brain. Like the brain, the spinal cord is composed of gray matter (consisting of neuronal cell bodies) surrounded by white matter (myelinated axons). The blood–spinal cord barrier (BSCB) consists of endothelial cells and astrocytes (one of the major glial cell types in the CNS) and protects the spinal cord from changes in the biological and chemical environment of the periphery, including fluctuations in the concentration of ions, small molecules, and growth factors. The peripheral nervous system (PNS) has a very different composition, with no structure comparable to the BSCB, and has a large number of Schwann cells, the dominant glial cell type of the PNS (Figure 73.1). Neural damage in the PNS is partially reversible, largely due to the nutrients, orientation, and myelindigesting support provided by Schwann cells to regenerating axons. Oligodendrocytes, the myelinating cells of the CNS, do not provide similar support after neural damage in the CNS. While it was long believed that axons in the CNS were intrinsically unable to regenerate, this was disproved in the early 1980s by Aguayo and colleagues, who showed that CNS retinal neurons are able to regenerate within peripheral nerve grafts, but not beyond the graft into the CNS tissue (David and Aguayo, 1981). This significant finding suggested that manipulation of the chemical and biological environment around the injury site could serve as a powerful

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(a)

Cranium

(b)

Cranium

Venous sinus Dura mater

Cerebral hemispheres

Subdural space Arachnoid membrane Subarachnoid space

Cerebellum

Pia mater Cervical spinal nerves

Function: cushion and protect delicate neural tissue (c) Dura mater

Brain

Spinal cord Autonomic ganglion

Arachnoid membrane Thoracic spinal nerves

Lumbar spinal nerves

Pia mater

Ventral root: carries efferent (motor) (d) information to muscles and glands

Sacral spinal nerves Dorsal root ganglion

Caccygeal nerve

Gray matter: consists of cell bodies of interneurons and some efferent neurons

Central canal

Spinal nerve Dorsal root: carries sensory (afferent) information to CNS

White matter: consists of axons carrying information through spinal cord to brain

Figure 73.1 Anatomical overview of the central nervous system. Gross anatomy of the brain and spinal cord. (a) The spinal cord is divided into the cervical, thoracic, lumbar, sacral, and coccygeal regions. (b) The neural tissue of the spinal cord is protected by the vertebrae, dura mater, arachnoid membrane, and pia mater. Innervation of the body occurs through spaces along the sides of the spinal cord. (c) Interior view of the spinal cord. The dorsal and ventral roots carry sensory and motor information, respectively, from and to the periphery. Reprinted with permission from Silverthorn, D. (2001). Human Physiology, p. 256. Upper Saddle River, NJ: Pearson Education, Inc.

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strategy to treat CNS damage. The differences in the regenerative capacity of CNS and PNS neurons can be largely explained by the composition and function of the glial cell populations. After an SCI, breaches in the BSCB enable the migration of macrophages and other cell types to the injury site. Astrocytes normally resident in the spinal cord are activated and form a tightly interpenetrating network known as the reactive glial scar, which constitutes a major physical obstruction to axon regeneration. In contrast to the PNS, where glial cells provide support for neuron survival and axon regeneration, glial cells at the CNS injury site produce a large number of inhibitory factors that severely limit axonal regeneration, including myelin-associated glycoprotein (McKerracher et al., 1994; Mukhopadhyay et al., 1994), chondroitin sulfate proteoglycans (Fawcett and Asher, 1999), and Nogo (Bandtlow and Schwab, 2000). The presence of an inhibitory chemical environment after CNS injury is manifested in the hours following axotomy, during which severed CNS axons develop new growth cones which extend very short distances before stabilizing and then receding (“abortive sprouting”) (Schwab and Bartholdi, 1996). Even PNS neurons abort axon growth when they are placed in a CNS environment (Giftochristos and David, 1988). Tissue engineering strategies for spinal cord repair generally focus on the need to promote axonal regeneration and overcome both the physical (glial scar) and the chemical (inhibitory molecules) barriers to regeneration.

ANIMAL MODELS Tissue engineering and chemotherapeutic strategies for treatment of SCI are most commonly tested in rats and mice. Both are accessible and low cost among mammalian vertebrates. Although rats offer better working size for spinal surgery and device implantation, the possibility of genetic manipulation makes mice models attractive. In both species, therapeutic strategies are tested in one or more of three types of injury models: transection models involve cutting all or part of the spinal cord; contusion models quickly crush the cord with measured force (Bresnahan et al., 1987; Gruner, 1992); and compression models compress the cord for a defined time (Rivlin and Tator, 1978). In all cases, the results can be analyzed histologically, neurophysiologically, biochemically, and functionally with behavioral assays. The most common animal models are shown in Figure 73.2 and are described elaborately below. Transection Models Transection models are most often employed to demonstrate the regenerative potential of a therapeutic or implant, with the latter often requiring excision of one or more cord segments to provide the necessary working space for implantation. With complete transection of the spinal cord, the model provides unambiguous proof of regeneration when observed functional recovery is lost on retransection. If specific neuronal tracts are of interest, hemisection can lead to defined neurological deficits. Surgically accessible tracts include the rubrospinal tract via unilateral hemisection and the corticospinal tract through dorsal hemisection. In lateral hemisection models, the function of both sides of the animal are affected, with the ipsilateral side showing more severe deficits. One distinction between complete and partial transection is that, in the latter, a length of tissue is removed to create a physical gap between the rostral and caudal tissues. In complete transection, this gap forms automatically as tension in the spinal cord is released. The primary criticism of the hemisection model is that it is both difficult to reproduce and to differentiate between regeneration and axonal sprouting. The primary criticism of transection models is that they represent only a small fraction of human SCI, with the bulk being compressive. Although transection does not accurately model the pathophysiology of secondary injury – the processes of glial scar formation and tissue cavitation occur after compressive injury – this procedure has the utility as an early stage model for determining the regenerative effect of experimental interventions. The transection model is also excellent for those with chronic spinal cord injuries, where intervention will require a neuroregenerative strategy versus a neuroprotective strategy, which is pursued in acute injury models.

Tissue Therapy: Central Nervous System

(a) Lateral hemisection

(b) Dorsal hemisection

(c) Complete transection

(d) Compression or contusion cross section

Figure 73.2 Animal models of spinal cord injury. (a–b) Hemisection models cut and remove a part of the spinal cord to produce specific functional deficits. (c) Complete transection creates a gap between rostral and caudal sections of the spinal cord when natural tension in the cord is released. (d) Mimicking the dominant form of human SCI, contusion, and compression models create a fluid-filled cyst, which extends rostrally and caudally from the site of injury. Copyright (2006) by Michael Corrin, reproduced with permission of the artist. Contusion Models Contusive models of SCI model compressive human injuries, including the formation of a fluid-filled cyst through secondary injury. Recent reviews by Park (Park et al., 2004) and Kwon (Kwon et al., 2002) describe how secondary injury increases the severity of SCI in the days post-injury and is the target of neuroprotective strategies. Multiple contusive models are in current use, most notably the New York University impactor (NYU) (Gruner, 1992) and Ohio State University impactor (OSU) (Bresnahan et al., 1987). The NYU device impacts the cord with a defined force, whereas the OSU impactor produces a defined cord displacement. In both cases the contusion is accomplished quickly, with the force applied and removed within 1 s (Bresnahan et al., 1987; Gruner, 1992). In addition to the glial scar, contusive injuries are characterized both by axotomy and demyelination of axons. Evaluation of neuroprotective strategies can involve the measurement of axonal sparing, the degree of myelination, the change in the volume of tissue lost to secondary injury, and behavioral assessment (Tsai and Tator, 2005). Assigning the source of functional recovery in contused animals is less precise than with transection because the distinction between recovery due to tissue sparing, or due to plasticity among surviving neurons versus regeneration, is not always clear (Kwon et al., 2002). Compression Models Compression models of SCI produce a glial scar similar to that seen in contusion models and also mimic the residual displacement of the spinal column seen in human injury (Grill, 2005). In the model developed by Rivlin and Tator (1978) a modified aneurysm clip creates the injury as a function of both the closing force and the duration of compression. The model has previously been used to study the pathology of secondary injury (Fehlings et al., 1989) and the importance of decompression following injury (Guha et al., 1987), the latter of which is now in clinical trials to become a standard of care. More recently, clip compression has been used to

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evaluate neuroprotective agents (Jimenez Hamann et al., 2005) and a novel drug delivery strategy (Jimenez Hamann et al., 2003; Gupta et al., 2006). Variants on balloon compression, a technique in which injury severity is related to both the pressure of an inflated angioplasty balloon and its period of application, are also in current use. In different models, the balloon is inflated subdurally or epidurally, being introduced into the intervertebral canal via laminectomy (Martin et al., 1992) or an intervertebral foramen (Purdy et al., 2003). Distinction is made between those models, which require laminectomy, the removal of the dorsal aspect of one or more vertebrae, and those that do not, because a laminectomy-free model of SCI more closely mimics a human SCI. The least invasive balloon compression models reach the spinal cord by lumbar puncture, causing minimal extraneous tissue damage and facilitating medical imaging by eliminating surgical artifacts (Purdy et al., 2003). Acute versus Chronic Injury To date, most treatment strategies have been directed toward acute SCI repair, where treatment is initiated immediately after injury. This strategy is valid when seeking to limit secondary injury, but is also used to evaluate regenerative strategies. Given that the majority of human SCI patients are chronically injured (Carlson and Gorden, 2002), strategies to treat chronic SCI have begun to be reported (Woerly et al., 2001a). The animal models are very similar to acute injuries except that treatment begins 4 to 8 plus weeks after injury and the injury site is often surgically prepared immediately prior to treatment. In the case of complete transection, laminectomy may be performed on neighboring vertebrae to expose the lesion and facilitate resection and removal of the rostal and caudal stumps. Surgical preparation of chronic compression injuries is centered on debridement of the injury cavity, but may also include exposing the spinal cord via laminectomy if the initial injury was inflicted without one. Animal Model Concluding Remarks When testing new therapeutic strategies or drug delivery systems for neuroregeneration, transection models provide the best evidence for axonal growth. Neuroprotective strategies are best evaluated in a contusive or compressive model because these paradigms include secondary injury. Tissue engineered implants, including scaffolds and nerve guidance channels, provide physical or haptotactic guidance and thus require at least partial transection (Dillow and Lowman, 2002). Cell therapies, which are often injectable strategies (with or without a scaffold), have been tested in all the three model types (Knoller et al., 2005; Kulbatski et al., 2005). When evaluating reports of spinal cord repair, it is thus important to consider not only the degree of neuroprotection or neuroregeneration, but also the animal model in question, and its applicability to treatment to SCI in man.

STRATEGIES FOR REPAIR AND REGENERATION Entubulation The use of tubular scaffolds to assist nerve regeneration has been investigated in animal models to promote regeneration after SCI. The guidance tube concept originates from the use of peripheral nerve tissue to confer regenerative potential on the CNS. In this case, the tube is meant to mimic the peripheral nerve tissue by providing a permissive environment for regeneration while, at the same time, protecting the environment from the inhibitory factors present in the spinal cord. Similar strategies have been pursued in peripheral nerve repair with clinical success. Tube dimensions, wall thickness, porosity, and mechanical strength are all important physical factors that influence the suitability of tubes for CNS repair applications. Matching the mechanical properties, and specifically modulus, of the tube to the tissue has been shown to be important to avoid necrotic tissue at the interface (Millesi et al., 1995; Dalton and Shoichet, 2001). Tube collapse is a significant problem in

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nervous system tissue engineering but can be avoided through the use of internally reinforced, corrugated, or thick-walled tubes. However, tube thickness affects oxygen diffusion, which is critical for viable tissue. Porosity is important for nutrient diffusion, and the use of porous tubes in the PNS is generally correlated to improved nerve regeneration (Dahlin and Lundborg, 2001), suggesting that the same will be true in CNS regeneration. Smooth-walled tubes have been found to be superior to rough-walled tubes for axon regeneration (Aebischer et al., 1990), likely because the latter elicit a greater foreign-body response. Considerable effort has been expended in the optimization of tube fabrication methods, including polymer extrusion methods, polymer casting in molds, immersion of a polymer solution-coated mandrel in a nonsolvent, and centrifugal casting of phase-separated polymerization mixtures. These techniques vary in their ability to produce tubes with uniform properties, with centrifugal casting remaining the most versatile option and molding being the most simple (see Dalton and Shoichet, 2001 and references therein). While some efforts to repair damaged nerves using tubular conduits have been based on allogeneic tissues such as blood vessels (Walton et al., 1989) and muscle (Hall, 1997), their limited availability and suboptimal mechanical properties have led to an overwhelming preference for synthetic materials in the construction of guidance tubes. One of the first synthetic materials to be used in PNS nerve guidance tubes was silicone, and several groups have pursued the use of this (Borgens, 1999) and other non-degradable materials in the design of guidance tubes for SCI treatment. Some of the most popular include poly(acrylonitrile-co-vinyl chloride) (P(AN/VC)) (Bunge, 2002) and hydrogel-derived materials such as poly(2-hydroxyethyl methacrylateco-methyl methacrylate) (P(HEMA/MMA)) (Dalton et al., 2002). P(HEMA/MMA) in particular is highly biocompatible and has shown great promise in nerve guidance tubes in the CNS and PNS (Tsai et al., 2004; Katayama et al., 2006). In addition to these materials, non-degradable fluoropolymers such as poly(vinylidene fluoride) (PVDF) and its co-polymers show piezoelectric properties (electric charges develop in response to mechanical deformation) that enhance PNS nerve regeneration (Fine et al., 1991; Valentini et al., 1992) and may find utility in the CNS. Figure 73.3 illustrates the use of a P(HEMA/MMA) guidance tube in a rat transection model.

(a)

(b)

(c)

Figure 73.3 Nerve guidance channel implanted in a rat. A poly(hydroxyethylmethacrylate-co-methyl methacrylate) nerve guidance tube in use in a rat spinal cord transection model. (a) Schematic view of the placement of the spinal cord stumps within the channel. (b) Intraoperative view of the channel after implantation into the spinal cord (dorsal view). The small arrows indicate the location of the spinal cord stumps, while the large arrow indicates the expanded poly(tetrafluoroethylene) (ePTFE) membrane that is used to cover the implantation site. (c) Spinal cord with nerve guidance channel implant, excised after 4 weeks. The spinal cord stump locations are indicated by the small arrows and the ePTFE membrane by the large arrow. The ruler markings are in mm. Reprinted with permission from Tsai et al. (2004).

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Despite the successes achieved with non-degradable materials, the dominant trend in tissue engineering in the nervous system has been toward the use of biodegradable polymers for guidance tubes. Of these, the best known are the polyesters – polyglycolide (PGA), polylactide (PLA), polycaprolactone (PCL), and their co-polymers. These materials have been used extensively for nerve repair in the PNS, and a number of groups have applied tubular polyester constructs to CNS injury repair model systems. In one study, tubes made of a mixture of high- and low-molecular-weight poly(L-lactide) (PLLA) and seeded with Schwann cells were seen to promote axon regeneration and vascularization. Interestingly, these tubes also showed minimal collapse after 4 months, while comparable tubes made of poly(D,L-lactide) (PDLLA) had completely collapsed after 2 months (Oudega et al., 2001). Tubes made of PDLLA have also been used in conjunction with brain-derived neurotrophic factor (BDNF) to promote axonal regeneration (Patist et al., 2004). The degradation rate of polyester tubes can be controlled to a certain extent through polymer composition, with PLLA and PCL giving materials with extended degradation lifetimes, relative to those containing PGA. As with other biodegradable materials, the structural nature of the nerve guide (hollow versus corrugated, the presence of plasticizers or reinforcing agents) can also have a dramatic effect on the nature of the degradation process (von Recum et al., 1995). Notably, the production of crystalline polymer fragments during degradation may lead to inflammation and other adverse side effects. In addition, polyesters degrade to produce high concentrations of carboxylic acids, which may locally irritate tissues prior to clearance or metabolism. For these reasons, concerns of the degradation rate (which depends on the size of the injury to be treated) must be balanced with the possibility of tube swelling, collapse, or irritation during degradation. In recent years, the use of poly(hydroxybutyrate)s (PHB) has been explored for the construction of nerve guides. These materials are produced naturally by certain species of bacteria when grown in an environment low in non-carbon nutrients, and they biodegrade by the same mechanism as the fully synthetic polyesters described above. PHB fibers containing alginate, fibronectin, and Schwann cells have been shown to enhance axonal regeneration in a rat model (Novikov et al., 2002). Polycarbonates are closely related to polyesters and have also been used to construct nerve conduits for CNS injury treatment. These have taken the form of rolled membranes, which have been coated with celladhesive poly(lysine) and seeded with Schwann cells or astrocytes prior to implantation in the injured spinal cord or in brain lesions (Montgomery and Robson, 1990). Polycarbonates are also advantageous in that they degrade to neutral by-products, avoiding the acid release common to polyesters. A number of biopolymers have been adapted for use in CNS guidance tubes. Collagen, one of the most common materials in the human body, has been used in the form of rolls (Paino and Bunge, 1991; Bunge, 1994), filaments (Yoshii et al., 2003), and non-structured gels (de la Torre, 1982; Marchand et al., 1993; Houweling et al., 1998a, b). The success of collagen in these applications appears to be highly dependent on the physical form of the implant and whether it is used in combination with alternate therapies (growth factors, cell seeding, etc.). Dilute collagen gels have also been embedded in the lumen of guidance tubes to promote axon regeneration in the PNS, with mixed results (Midha et al., 2003). Fibronectin, an important cell-adhesive extracellular matrix (ECM) protein, has been used in the form of mats for growth factor delivery in the SCI, though results to date have not been promising (King et al., 2004). Among naturally occurring polysaccharides, hyaluronic acid, a non-toxic and biodegradable co-polymer of glucuronic acid and N–acetylglucosamine, has shown promise as a scaffold for tissue implantation (Rochkind et al., 2002) and has been used to repair brain defects in rats (Tian et al., 2005). Tubes made of chitosan, a polysaccharide prepared from chitin (derived from the shells of crustaceans) through chemical de-acetylation, have shown great promise in PNS injury repair (Itoh et al., 2003; Rosales-Cortes et al., 2003), and our laboratory is currently testing this material for use in SCI treatment (Freier et al., 2005a, b). Successful application of entubulation for nerve guidance requires a combination strategy in which the tube provides both, a pathway for regeneration and is filled with factors and/or

Tissue Therapy: Central Nervous System

cells that promote regeneration, and reinnervation into the spinal cord tissue. Figure 73.4 summarizes some of the key features required in a nerve guidance channel, which are described in further detail below. Haptotactic Cues Cell-adhesive proteins produced by Schwann cells after nerve injury are important in nerve regeneration in the PNS, and a number of tissue-engineering strategies have made use of this principle to provide a more permissive environment for regeneration in SCI repair. Among the proteins found in the ECM, collagen, laminin, and fibronectin contain domains that bind strongly with integrin receptors on cell surfaces, promoting cell adhesion, and neurite outgrowth by organizing the neural cytoskeleton. A number of studies have made use of these materials, implanted into the lumen of PNS nerve guidance tubes, to assist nerve regeneration. An alternative to this approach is the modification of biomaterial surfaces with short oligopeptides derived from

5b

1

2

4

Concentration gradient of neurotrophic factors

3 5a

Figure 73.4 Properties of a typical nerve guidance channel. Schematic view of a nerve guidance channel making use of chemical cues and gradients. The key features of this strategy are (1) the polymeric tube itself, (2) the presence of haptotactic cues within (3) a cell-invasive scaffold that also contains a gradient of chemotactic cues, and (4) a drug-delivery system that is integrated into the walls of the tube. The stumps of the nerve cable (5a, b) are inserted into the ends of the channel (1). Reprinted with permission from Cao, X. and Shoichet, M.S. (2002). Tissue engineering strategies for axonal regeneration following spinal cord injury. Biomimetic Materials and Design. New York: Marcel Dekker, p. 427.

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ECM proteins, such as the YIGSR and RGD sequences, which control cell adhesion (Pierschbacher and Ruoslahti, 1984; Tashiro et al., 1989; Tong and Shoichet, 2001; Woerly et al., 2001b), or the IKVAV sequence, which promotes neurite outgrowth (Jucker et al., 1991; Bellamkonda et al., 1995). The patterning of haptotactic cues on materials has been widely exploited as a method for achieving nerve guidance in in vitro models. The physical or chemical attachment of proteins such as laminin onto otherwise non-adhesive surfaces has been used to promote axon guidance in several model systems, in which it was found that a clear contrast between adhesive and non-adhesive regions was necessary for guidance effects (Tong and Shoichet, 1998; Saneinejad and Shoichet, 2000). For in vivo use, the incorporation of polymer fibers modified with haptotactic cues into nerve guidance tubes – in effect, mimicking the bands of Büngner seen in the PNS – may provide a useful method of chemically and physically directing axon regeneration in SCI repair. Along similar lines, the incorporation of cell-adhesive gels into nerve guidance tubes has shown promise, particularly in the case of dilute collagen gels (Labrador et al., 1998; Midha et al., 2001) and laminin-modified agarose (Bellamkonda et al., 1995; Yu et al., 1999). Chemotactic Cues The tendency of severed nerves to extend axons toward the distal stump has been known since the time of Santiago Ramon y Cajal at the turn of the 20th century. In recent years, the biochemical principles that underlie this tropism have been elucidated, and it is now widely accepted that axon guidance to specific targets is mediated by gradients of chemo-attractive and repulsive cues (Bregman et al., 1997). Growth cones at the tip of growing axons sample the chemical environment by means of numerous small protrusions (Goodman, 1996; Tessier-Lavigne and Goodman, 1996) and direct neurite extension by the asymmetric remodeling of the cytoskeleton and cell membrane (Keynes and Cook, 1995). While a number of systems for the systematic delivery of neurotrophic agents to the injured spinal cord have been developed (see below), axon guidance in a bioengineered constructs requires the use of gradients that mimic the signals provided to the developing spinal cord. Concentration gradients of chemotactic cues such as nerve growth factor (NGF) (Cao and Shoichet, 2001; Kapur and Shoichet, 2004) and neurotrophin-3 (NT-3) (Cao and Shoichet, 2003) have been exploited to guide axon extension in vitro. The application of concentration gradients of chemotactic factors to nerve guidance scaffolds for SCI repair models will likely require immobilized neurotrophic factors to prevent gradient homogenization and uptake by non-neural cells; however, it is not yet clear how neurons will respond to immobilized factors in vivo. Cell Seeding in Nerve Guidance Scaffolds Some of the most promising approaches for SCI treatment involve bioengineered constructs that use a combination of guidance tubes, non-tubular scaffolds, and/or cells. For example, Schwann cells have long been known to be a key supporting factor in the natural nerve regeneration seen in the PNS. For this reason, significant efforts have been made to implant cultured Schwann cells into nerve guidance channels for SCI treatment. Bunge’s group has found that P(AN/VC) tubes seeded with Schwann cells, when used in combination with neurotrophic factors or hormones, facilitated axon regeneration (Jones et al., 2001; Bunge, 2002), and Kamada et al. saw functional recovery in a rat spinal cord transection model using cell-seeded polyether sulfone tubes (Kamada et al., 2005). Olfactory ensheathing glia (OEG) is another type of glial cell (found exclusively in the olfactory system) and can be induced to myelinate large axons. These cells have been used to promote long-distance axon regeneration in conjunction with Schwann cells in a P(AN/VC) tube (Ramon-Cueto et al., 1998). A great deal of attention has recently been focused on neural stem cells (NSCs). The pluripotent capability of NSCs allows them to be differentiated into neurons, astrocytes, or oligodendrocytes, and it has been shown that the differentiation pathway taken by NSCs depends on the presence of specific

Tissue Therapy: Central Nervous System

growth factors and the chemical nature of the material on which they are seeded (the ECM equivalent). Among naturally occurring materials, collagen (Lin et al., 2004) and alginate (Wu et al., 2001) have been used to induce NSC differentiation. Synthetic polymers such as PGA, PLA, and their co-polymers have also been used in conjunction with NSCs in the form of three-dimensional scaffolds for SCI and brain injury treatment (Lavik et al., 2002; Teng et al., 2002). A synthetic self-assembling peptide nanofiber system was recently described that differentiates NSCs specifically to neurons (Silva et al., 2004). The use of NSCs to aid functional recovery after SCI in bioengineered constructs holds great promise, though it is important to note the risk of uncontrolled proliferation that is inherent in stem cells. In addition, the maintenance of cell viability and function after seeding is a particular challenge for all of the cell-inclusive combination strategies described above. Drug Delivery The use of drug delivery as a reparative and regenerative strategy of SCI has been investigated in animal models clinically. With the notable exceptions of methylprednisolone (Bracken et al., 1998), GM-1 gangliosides (Young, 1989), minocycline (Wells et al., 2003), and cAMP moderators (Pearse et al., 2004), a majority of molecules used to treat SCI are high-molecular-weight proteins. A selection of these molecules are reviewed sections “Neuroprotective agents” and “Neuroregenerative agents.” These biomacromolecules generally possess short serum half-lives, are poorly transported across the BSCB, and are thus impractical to administer systemically. The increased risk of side effects (as with hyperalgesia due to nerve growth factor treatment (Namiki et al., 2000)), and the high cost of systemic delivery of recombinant proteins further recommend local delivery. Many of the drug delivery systems investigated for spinal cord repair incorporate a scaffold or tube strategy for release therefrom and were discussed earlier. Thus, this section is focused on strategies principally designed to deliver soluble molecules to the site of injury when a nerve guidance channel or other implant is not appropriate. Whether the therapeutic agents are neuroprotective, neuroregenerative by enhancing regenerative capacity, or neuroregenerative by suppressing inhibitory factors, delivery strategies for all three approaches are similar. Routes of Spinal Drug Delivery Local spinal drug delivery in humans is described as epidural, intrathecal, or intramedullary according to the delivery site. Epidural and intrathecal delivery are used clinically, most commonly for pain management and control of spasticity (Amar et al., 2005). Intramedullary delivery is the only route, which may result in damage to the nervous tissue within the spinal cord and is not in human use. For this reason, and the observed tissue damage to animals in studies of intramedullary-administered therapies (Namiki et al., 2000), intramedullary drug delivery is not reviewed. Epidural and intrathecal delivery consists either of a single bolus injection or insertion of a catheter for continuous infusion. Bolus injection is preferred for its simplicity, but continuous infusion provides better dose control and is indicated for regenerative strategies, which have shown increased efficacy with extended treatment (Tuszynski et al., 2003). The best-established system of epidural drug delivery is continuous infusion of a drug solution (Amar et al., 2005), with more recent reports of injectable hydrogels (Paavola et al., 1998) and liposomes (Paavola et al., 2000) for analgesic delivery. Epidural delivery of proteins for SCI repair faces the challenge of diffusion/transport across the dura mater, arachnoid, fluid filled intrathecal space, and pia mater before reaching the white matter of the spinal cord. Epidural strategies for neuroregeneration exist (Fernandez et al., 1993), but the diffusive barriers are real and intrathecal injection is more effective in animal models. Intrathecal delivery sees injection into the intrathecal, or subarachnoid, space. The intrathecal space is filled with cerebrospinal fluid (CSF) and functions to support the spinal cord, deliver nutrients, and remove cellular waste. Drugs delivered to the intrathecal space must diffuse across the pia mater before entering the

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spinal cord. As the most local route of spinal drug delivery besides intramedullary diffusion (with attendant tissue damage (Namiki et al., 2000)), intrathecal injection is often the first drug delivery system (DDS) used in the evaluation of novel chemotherapies in contusion or compression models. Both bolus injection and continuous infusion of drug solutions are common. An advantage of minipump-based infusion studies is that implantable pumps are in current clinical use (Ethans et al., 2005), facilitating transition of these therapies to human use. One important qualifier is the evidence that continuous intrathecal infusion from both polyethylene catheters and steel cannulae causes scarring and compression over 14 days (Jones and Tuszynski, 2001). Although the great majority of human SCI is compressive and the indication is that even in neurologically complete injuries a rim of intact white matter persists, transection models are valued as a means to test new regenerative strategies (Carlson and Gorden, 2002). The relationship of epidural and intrathecal injection is lost in partial and complete transection models of SCI. With the rostral and caudal stumps of the spinal cord exposed, the role of the spinal meninges as diffusive barriers is minimal because the transected area affords direct penetration into the white and gray matter of the spinal cord. Drug delivery to the lesion site can take the form of bolus or continuous injection as previously discussed, but the use of drug-loaded foams and sponges is often used to localize an initial drug dosage during surgery. These strategies increase local drug concentration but are not often assessed as a DDS with respect to release rates or period of delivery, both of which are important criteria when formulating clinically acceptable drug delivery strategies. Methods of Spinal Drug Delivery Treatment strategies which rely on direct access to neural tissue are likely to require modification prior to clinical use because many contusion and compression injuries leave the spinal meninges as intact barriers to diffusion (Rivlin and Tator, 1978; Bresnahan et al., 1987; Gruner, 1992). Drug delivery strategies founded on contusion/compression models are thus better suited to spinal drug delivery because this paradigm is prevalent in man. Of the strategies previously discussed, bolus intrathecal injection and continuous infusion are most promising but not without limitations. Although facile, bolus injection of a drug solution provides only short-term delivery. For example, neither epidermal growth factor nor basic fibroblast growth factor were detected in the spinal cord after bolus intrathecal injection (Jimenez Hamann et al., 2005). Conversely, osmotic minipumps can deliver a solution of one or more drugs over a period of months and are currently the method of choice in trials targeting neuroregeneration (Namiki et al., 2000; Coumans et al., 2001). In these cases the minipumps are implanted subcutaneously and deliver a constant drug volume to the epidural fat or intrathecal space (Figure 73.5). One increasingly important consideration is that the presence of a subdural catheter in continuous minipump delivery may cause scarring and compression of the spinal cord, a definite barrier for chronic use in human patients (Jones and Tuszynski, 2001). The invasiveness of implantation and attendant chronic inflammatory response are also drawbacks of minipump use. Strategies to prolong drug release after bolus injection include epidural injection of hydrogels, liposomes, and combinations thereof. Fibrin sealant has also been used in spinal drug delivery, either epidurally or to fill spinal defects (Dergham et al., 2002). The release profile from these systems is either poorly characterized or indicates that a majority of the drug load is released within 24 h (Paavola et al., 2000). Seeking to create a localized release DDS, work within our laboratory has focused on the development of an intrathecal drug delivery system (IDDS) as shown in Figure 73.6 (Jimenez Hamann et al., 2003, 2005; Gupta et al., 2006). The IDDS has a drug-loaded biomaterial injected into the intrathecal space at the site of injury where it gels thereby providing localized release. Early research showed that injected collagen remained localized and released EGF for between 6 and 24 h (Jimenez Hamann et al., 2005). This proof-of-concept also demonstrated that injection of collagen into the intrathecal space was safe as determined by behavioral, histological, and immunohistochemical techniques

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Figure 73.5 Implanted minipump with intraspinal catheter. Infusions pumps may be implanted under the skin in a patient’s abdomen and an intraspinal catheter tunnelled subcutaneously into the epidural or intrathecal space. Models used in animal studies are often powered by osmotic pressure, those in clinical use by batteries or pressurized propellant. Copyright (2006) by Michael Corrin, reproduced with permission of the artist.

Injectable hydrogel Spinal cord Dura mater

CSF filled intrathecal spece CSF flow

Figure 73.6 Injectable drug delivery to the intrathecal space. When injected into the intrathecal space a hydrogel based drug delivery system can localize and modulate drug release at the site of injury. This route is preferred over epidural delivery when the diffusive barrier presented by the dura mater is significant, as is the case when delivering large molecular weight therapeutics. Copyright (2006) by Michael Corrin, reproduced with permission of the artist. for both uninjured and SCI animals (Jimenez Hamann et al., 2003). Subsequent work with a blend of hyaluronic acid and methylcellulose (HAMC) resulted in an IDDS, demonstrating erythropoietin release for 3 days in vitro and in vivo localization of the blend for between 3 and 7 days (Gupta et al., 2006). In Figure 73.7 the IDDS has been fluorescently labeled and visualized in vivo. Both of these delivery systems were designed to release neuroprotective factors over a few days. Extending intrathecal drug delivery over a longer time is

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Dura mater

HAMC

Spinal cord

Figure 73.7 In vivo localization of an intrathecal drug delivery system. Parasagittal section of rat spinal cord rostral to site of injection shows that the fluorescent HAMC is localized in situ in the intrathecal space where it was injected (scale bar  1 mm). Reprinted from Gupta, D. et al. (2006). Copyright (2006), with permission from Elsevier.

also of value because neuroregenerative factors often show increased effect with extended treatment (Grill et al., 1997). Neuroprotective Agents Numerous natural and synthetic molecules are known to provide neuroprotection in animal models of traumatic spinal cord injury, but few have been tested clinically and none have shown conclusive benefit. In the strictest test, neuroprotection is evident as a long-lasting increase in functional abilities, but the term is also applied to cases demonstrating increased axonal sparing and reduced lesion volume. Much attention has been paid to the steroid methylprednisolone and GM-1 ganglioside, although both showed only limited success in clinical trials (Ramer et al., 2005). Methylprednisolone is valued as an anti-inflammatory agent (Bracken et al., 1998), while GM-1 ganglioside is purported to inhibit apoptosis and promote axonal sprouting (Geisler et al., 2001). Other molecules targeting the host inflammatory response include a tetracycline derivative, minocycline (Stirling et al., 2005); and the cytokines interleukin-10 (Bethea et al., 1999) and erythropoietin (Eid and Brines, 2002). Each of these molecules probably acts on a variety of cellular targets: minocyline has been shown to reduce dieback of specific neuronal populations, promote oligodendrocyte survival, and inhibit microglial activation in addition to reducing lesion volume in a rat model (Ramer et al., 2005). Interleukin-10 and erythropoietin attenuate inflammation, but it has been suggested that their neuroprotective effects stem from direct suppression of neuronal apoptosis following injury (Eid and Brines, 2002). In the CNS, activated macrophages, astrocytes, and microglia produce tumor necrosis factor-α (TNF-α) as part of a progressive inflammatory response following injury (Bethea et al., 1999). The role of TNF-α in neuroprotection is controversial; however, as mice lacking TNF-α receptors show increased apoptosis and greater tissue cavitation (Kim et al., 2001), but separate in vivo studies demonstrate neuroprotection following TNF-α blockage with antibodies (Lavine et al., 1998). These seemingly contradictory results may be the result of differences in injury type and location. The neurotrophins, a class of growth factors best known for their regenerative potential, also provide neuroprotection. Nerve growth factor (NGF), BDNF, and NT-3 enhance the survival of either or both of sensory and motor neuron survival (Nakahara et al., 1996). Outside the neurotrophins, ciliary neurotrophic

Tissue Therapy: Central Nervous System

factor (CNTF) and glial cell line-derived neurotrophic factor (GDNF) are also neuroprotective. Both growth factors enhance motor neuron survival (Sendtner et al., 1990), while GDNF also reduces sensory neuron dieback (Matheson et al., 1997). Neuroregenerative Agents Neuroregeneration after SCI comprises elongation of existing axons, sprouting and growth of new axons from neural cell soma, remyelination, plasticity among surviving connections, and functional recovery. The complexity of this process provides multiple cellular and molecular targets for intervention, but the importance of axonal growth as a precursor to functional recovery has focused attention on this goal. Neuroregenerative therapies are divided into two classes by their mode of action: some, like the neurotrophins mentioned earlier, act on regeneration associated genes (RAGs) to stimulate axonal sprouting and growth. Others seek to silence the effects of inhibitory glycoproteins and proteoglycans acting on the Rho-ROCK pathway, thereby allowing neurite outgrowth in the hostile environment present after SCI. In addition to the neurotrophins, CNTF, GDNF, and the fibroblast growth factors (aFGF and bFGF) also exhibit neuritogenic effects. There is evidence for regeneration after NGF delivery (Ramer et al., 2000), but these results are tempered by observations that NGF causes sprouting of uninjured sensory axons (Romero et al., 2001) and hyperalgesia (Christensen and Hulsebosch, 1997). Trials involving BDNF provide conflicting results (Nakahara et al., 1996; Oudega and Hagg, 1999), but differences in dosage, delivery method, and treatment length make direct comparison difficult. It is known, however, that BDNF supports the outgrowth of motor (Braun et al., 1996) and sensory axons (Oudega and Hagg, 1999). In vivo experiments utilizing NT-3 also show regeneration of motor (Braun et al., 1996) and sensory axons (Dijkhuizen et al., 1997). NT-3 also enables PNS axons to cross the PNS–CNS transition zone, the point at which regenerating PNS axons terminate in untreated animals (Ramer et al., 2000). CNTF promotes axonal sprouting from motor neurons as well as growth of existing axons (Siegel et al., 2000). With results similar to NT-3, GDNF promotes axonal growth through the PNS–CNS transition zone (Oyesiku and Wigston, 1996) as well as enhancing the survival and growth of motor (Ramer et al., 2000) and sensory axons (Henderson et al., 1994). The FGFs have also been linked to regeneration, with emphasis on the potential of bFGF to encourage axonal growth across the PNS–CNS transition zone (Matheson et al., 1997). Each of the growth factors mentioned acts on one or more RAGs and is predicted to produce synergies when combinations of factors are delivered to the correct location, in an appropriate order and concentration (Schwab, 2002; Enomoto et al., 2004). In addition, enhanced regeneration has been observed between combinations of growth factors and Rho-ROCK inhibitors (Lu et al., 2004). Activation of the Rho-ROCK pathway after SCI is known to inhibit neurite outgrowth in vivo through a multitude of receptor–ligand interactions, many of which are now therapeutic targets (Mueller et al., 2005). Antibodies raised against the myelin protein NI-35 (Nogo), myelin associated glycoprotein, and oligodendrocytemyelin associated glycoprotein enhance neurite outgrowth, as does saturation of their cell-surface receptor Nogo-R (Schnell and Schwab, 1990). Treatment with chondroitinase ABC, a protease acting on chondroitin sulfate proteoglycans (CSPGs) also improves regeneration and functional recovery (Bradbury et al., 2002). CSPGs are ECM components produced by reactive astrocytes and activate Rho-ROCK via their neuronal receptors (Dow et al., 1994). In an attempt to block Rho mediated signaling, delivery of Rho and ROCK inhibitors result in significant regeneration (Mueller et al., 2005). The best-known examples, C3 transferase (Ellezam et al., 2002) and the small molecule inhibitor Y-27632 (Mueller et al., 2005), substantially block ROCK activation and promote neurite outgrowth on CNS substrates. The cyclic nucleotide cAMP, and Rolipram, an inhibitor of cAMP degradation have also been shown to enhance neuroregeneration in vivo (Pearse et al., 2004). cAMP activates protein kinase A (PKA) which in turn inhibits activation of the Rho pathway.

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Overall, there are numerous factors of interest and of therapeutic potential; however, a single therapeutic molecule will not be able to overcome the limitations of the system. A combination strategy of therapeutic agents for neuroprotection, and regeneration is required and must be combined with a strategy to overcome the inhibitory environment present after injury. Cell Therapy Cell therapy, unlike drug or biological therapy, relies on the continuous and endogenous de novo production of cellular products that have therapeutic benefit. While drug/biological delivery is simpler, the delivery strategies require drug replenishment by reinjection and may lack dose control. Cell therapy relies, instead, on the cells either producing the desired therapeutic molecule over a prolonged time, thereby replacing a lost function, or additionally, replacing lost tissue as in the stereotypical tissue engineering/regenerative medicine approach. There are essentially two methods of cell therapy: (1) cell transplantation, where the cells are derived from autogeneic, allogeneic, or xenogeneic sources and (2) cell stimulation, where the endogenous stem cell population is stimulated to promote repair. Cell transplantation therapy has been investigated for neurodegenerative diseases and disorders in clinical settings for several years, specifically for chronic pain, Parkinson’s disease (PD), ALS, and SCI, among others. Interestingly, lentiviral delivery strategies have also been investigated for neurodegenerative diseases. For example, a lentivirus modified with GDNF was investigated for PD and showed some therapeutic benefit (Kordower et al., 2000). For amyotrophic lateral sclerosis (ALS), baby hamster kidney cells were genetically engineered to produce ciliary neurotrophic factor (CNTF) and then encapsulated in a PAN/VC hollow fiber membrane prior to implantation in the intrathecal space of humans in a phase I/II clinical trial. The PAN/VC membrane immuno-protected cells and allowed nutrients to diffuse through the membrane, thereby maintaining cell viability, while at the same time, allowing cell products to reach the intrathecal space and initiate a therapeutic affect (Aebischer et al., 1996). For PD, a number of cellular approaches have been tested clinically. For example, dopamine-producing pheochromacytoma (PC12) cells were encapsulated in PAN. For chronic pain, bovine adrenal chromaffin cells were encapsulated in a similar PAN/VC hollow fiber membrane prior to implantation in the subarachnoid space of humans in a phase I safety trial (Buchser et al., 1996). The chromaffin cells were dispersed within the hollow fiber membrane in an alginate matrix and shown to survive for an extensive period. These cells produced a combination of molecules that had an analgesic effect in patients. For PD, a number of cellular approaches have been tested. Dopaminergic cells have been implanted in both microcapsules and hollow fiber membranes in several animal models – from rats to non-human primates and the results have been promising. One of the limitations in PD is where to transplant cell tissue – this may be overcome by multiple transplant sites (Ramachandran et al., 2002). Despite some clinical success with amyotrophic lateral sclerosis (ALS), chronic pain and PD, cell-loaded membranes are not used clinically due to inconsistent therapeutic results. One of the major challenges associated with cell transplantation is the host’s immune response to the transplanted cells. Even if cells are immuno-isolated by a polymeric membrane – whether it is a hollow fiber membrane or a macrocapsule – the immune response to shed antigens poses a significant problem to prolonged cell viability (Jones et al., 2004). The capsules often become “walled-off ” from the body and this then leads to greater cell death within the capsule. Use of autogeneic cells overcomes the problems associated with this immune response to foreign cells; however, the source of autogeneic cells is often limited. Autogeneic stem or other cells provide some hope in overcoming the limited cell supply issues with autologous cell therapy. For PD, fetal transplantation and stem cell transplants have been investigated. For example, mesencephalic dopamine fetal cells were transplanted in the caudate putamen in the adult brain (Freed et al., 1990) and fetal

Tissue Therapy: Central Nervous System

midbrain cell suspension was transplanted into the striatum or substantia nigra (Mendez et al., 2005). While both demonstrated some promise, the reproducibility of these methods, incomplete recovery, and the ethical questions surrounding the use of fetal tissue transplants have limited progress in this area. While promising, stem cell therapy is faced with three major hurdles: (1) cell survival – often only 5–20% of implanted stem cells survive; (2) tumors – stem cells may form tumors (Singh et al., 2004); and (3) source – where to harvest the stem cells from is a big question. One potential answer to the source of stem cells is work pioneered by Freda Miller (Toronto) where skin-derived stem cells are being investigated in laboratory animals for a number of neurological disorders, including SCI. Neurospheres, identified by Sam Weiss (Calgary), derived from the subependymal zone in the adult brain, are another potential source of neural stem cells. Another potential answer to the source of autogeneic cells is the work pioneered by Michal Schwartz where macrophages are being implanted into the injured spinal cord, where the inflammatory response is thought to be insufficient to overcome the insult. The hypothesis is that the macrophages will “clean-up” the inhibitory environment due to the glial scar and/or myelin. This strategy is currently being pursued in an on-going clinical trial; the outcome of the trial was not known at the time of writing. Stimulation of endogenous stem cells is a relatively new approach that is being investigated in the laboratory. The idea of stimulating endogenous stem cells to repair injured tissue was demonstrated by Derek van der Kooy’s laboratory in spinal cord tissue (Martens et al., 2002). Here, specific factors (EGF, FGF2, and heparin) were delivered to the ventricle in an attempt to locally stimulate neurons and precursor cells. The IDDS strategy described above may be an appropriate methodology to locally stimulate a specific stem cell population.

CONCLUSIONS AND OUTLOOK The use of tissue-engineered constructs in the treatment of CNS injury holds immense promise, though the field remains in its infancy. Important advances with cell therapy have been made in many neurological disorders, including traumatic SCI. Using advanced combination strategies show great promise in animal models, yet require significant advancements for enhanced recovery after traumatic SCI. Guidance tubes alone are insufficient for significant recovery after CNS injury. For this reason, future efforts in this area will likely see an increased emphasis on combination strategies, in which cells and growth factors are delivered inside, or in conjunction, with biomaterials-based scaffolds. A number of challenges remain in the design of the scaffolds themselves: non-biodegradable polymers, which have historically been the most commonly used materials for nerve guidance tubes, may require a second surgical procedure for removal or may collapse over time, while biodegradable polymers may produce toxic by-products during degradation. In the area of cell therapy, effective strategies for preserving cell viability and function after implantation into scaffolds remain to be developed. Drug delivery strategies in the CNS will also likely benefit from optimizations in the timing, spatial control, and rate of delivery. Notwithstanding these challenges with what has been termed the “final frontier of science” – the brain, tissue engineering holds the promise to provide real solutions to CNS injury repair. ACKNOWLEDGMENTS We gratefully acknowledge financial support from the Natural Sciences and Engineering Research Council of Canada, the Canadian Institutes of Health Research, the Canada Foundation for Innovation, and the Ontario Research Fund.

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74 Peripheral Nerve Regeneration Mahesh C. Dodla and Ravi V. Bellamkonda

INTRODUCTION Problems and Challenges of Peripheral Nerve Injuries Injuries to the peripheral nervous system (PNS) occur frequently, and are a major source of disabilities. PNS injuries impair the ability to move muscles, to feel normal sensations, and result in painful neuropathies. PNS injuries are classified as being traumatic, non-traumatic, or surgical in nature. Traumatic nerve injuries result from collisions, motor vehicle accidents, gunshot wounds, fractures, lacerations, or other forms of penetrating trauma. In 2002, more than 250,000 patients suffered traumatic peripheral nerve injuries in the United States (AxoGen Inc., 2006). Of these patients, only 15% could be treated due to difficulties in treatment. Even among the patients who received treatment, more than 50% did not show any measurable signs of recovery or suffered from drastically reduced muscle strength. Most of the non-traumatic peripheral nerve injuries are attributed to nerve compression and adhesion. In 2002, more than 400,000 repair procedures were done in the United States to correct carpal tunnel syndrome, a non-traumatic nerve injury. Although treatments for non-traumatic nerve injuries have higher efficacy than traumatic nerve injuries, patients still suffer from pain, loss of muscle strength, and dexterity for several weeks to months. Surgical injuries result from procedures, such as prostatectomy, to remove prostate tumors. Prostatectomy procedures most often require sacrificing one or both of the cavernosal nerves, adversely affecting erectile function and bladder control. In 2002, more than 260,000 patients in the United States suffered major injuries to cavernosal nerves due to prostatectomy procedures (AxoGen Inc., 2006). In order to repair these nerve injuries, several techniques have been used. Historical Background The simplest technique for nerve repair, in the case of nerve transection injuries, is coaptation of the two ends of the nerve using sutures or fibrin glue. However, in many cases, there is loss of nerve segment due to injury, or there might be a time lag between injury and surgical repair during which the nerve ends might retract, resulting in a nerve gap. In such cases, end-to-end nerve suturing cannot be done without creating tension in the nerve segment, resulting in a poor regeneration outcome (Terzis et al., 1975). To overcome this problem, the two nerve ends are approximated using grafts, such as nerve autografts/allografts, muscle grafts, vein grafts, muscle–vein grafts, and synthetic nerve guidance conduits/channels (NGCs). During the 19th and early 20th centuries, various materials were used to promote nerve repair, such as bone (Gluck, 1880), metal tubes (Payr, 1900), blood vessels (Weiss and Taylor, 1946), and fat sheaths (Kirk and Lewis, 1915). The use of autologous nerve grafts was also first reported during this time (Albert, 1885). However, due to improper surgical techniques, anatomical repair rarely led to an appreciable return of

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function. During World War II, surges in the number of nerve injuries led to advances in microsurgical techniques and instrumentation. Further refinements in microsurgical techniques, revolutionized by Millesi (Millesi et al., 1972), and drug therapies have had beneficial effects. Significant advances in surgical techniques have been achieved, and now biological factors rather than surgical techniques are limiting improvements in nerve regeneration. Further advances may come from greater understanding of the molecular mechanisms of nerve regeneration, advances in nerve cell culture, development of new biomaterials and improved genetic techniques. Current Strategies for Regeneration At present, nerve autografts are considered the “gold standard” for bridging nerve gaps in the PNS (Lundborg, 1988). Autologous tissue grafts possess several advantages. They are likely to be more biocompatible than artificial materials, are less toxic and provide a support structure to promote cell adhesion and migration (Evans, 2001). However, there are several disadvantages with this technique. Obtaining natural graft leads to loss of function, as well as potential neuroma formation at donor site, needs multiple surgeries, and multiple small grafts are needed for a long nerve gap. There can be a size mismatch between the donor nerve graft and injured nerve. In addition, complete functional recovery is seldom obtained with autografts (Kline et al., 1998). Therefore, there is interest in developing techniques to not only enhance the performance of autografts, but also to synthesize alternatives with better functionality than autografts. Analytically, natural or artificial grafts used to bridge nerve gaps can be thought to have four central components germane to regeneration: (1) scaffold/substrate, (2) growth factors, (3) extracellular matrix (ECM) molecules and (4) cells. A graft may have any combination of the four components. In this analytical framework, the presence of these components and their spatio-temporal distribution determines the efficacy of the graft. In this chapter, based on the distribution of these four components within the graft, the grafts are classified as isotropic or anisotropic. In isotropic grafts, the components are distributed uniformly within the graft, with no directional cues. In anisotropic grafts one or more of these components are distributed anisotropically, usually along the direction of regeneration, to direct the axonal growth toward the distal target (Figure 74.1).

Components of nerve grafts Scaffolds, trophic factors, ecm cues, cells

Uniformly distributed (Isotrophic grafts)

veins veinsmuscle veinscells veinscollagen

NGC NGCgels NGCcells NGCECM cues NGCtropic factors NGCECM cues cellstrophic factors

Directionally oriented (Anisotropic grafts)

Nerve autografts NGCaligned gels NGCaligned filaments Nerve allografts NGCcell-seeded aligned filaments NGCECM cues-coated aligned filaments NGCtrophic factors-coated aligned filaments NGCaligned filamentscellstrophic factorsECM cues

Figure 74.1 Classification of nerve grafts as isotropic or anisotropic. Four basic components make up a nerve graft. All nerve grafts consist of either one or more of these components. In isotropic grafts, the components are distributed uniformly. In anisotropic grafts, the components are aligned longitudinally or in an increasing gradient within the graft.

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ISOTROPIC NERVE GRAFTS FOR REGENERATION This section on isotropic grafts has been divided into four sub-sections, to discuss the four components influencing peripheral nerve regeneration. Natural and Synthetic Scaffolds for Nerve Repair A scaffold can consist of two components. The first is a tubular structure serving as a “guidance channel” and the second consists of scaffold elements that are inside the tubular structure. In general, scaffolds for nerve repair should support axonal proliferation, have low antigenicity, support vascularization, be porous for oxygen diffusion, and avoid long-term compression. The scaffold can be made from natural or synthetic materials. Natural Materials as Scaffolds Isotropic natural materials used as scaffolds include veins, skeletal muscle fibers, and collagen. Although these materials support nerve regeneration, they do not provide any direction to the axons. Autologous vein grafts have been shown to provide a good environment for axonal regeneration in short nerve gaps (Wang et al., 1993; Ferrari et al., 1999). However, use of vein grafts for long nerve gaps has been less successful because of the collapse of veins due to their thin walls, and constriction due to the surrounding scar tissue (Chiu and Strauch, 1990). In order to prevent vein grafts from collapsing and to improve their performance, intraluminal space fillers such as autologous Schwann cells (SCs), collagen, and muscle fibers have been used. Collagen-filled vein grafts were found to promote better axonal growth than empty vein grafts for a 15 mm nerve gap in rabbits (Choi et al., 2005b). Similarly, SC-seeded venous grafts supported axonal growth and performed better than unseeded grafts in repairing 40 mm nerve gaps (Zhang et al., 2002) and 60 mm nerve gaps in rabbits (Strauch et al., 2001). The principal drawback of this approach is that it requires the availability of a relevant amount of live autologous SCs (up to 8 million cells/ml) that are difficult to obtain. Muscle–vein combined grafts, in which the muscle fibers are inserted in veins, were used in 10 mm long nerve defects in rats and found to promote axonal regeneration comparable to that of syngenic nerve grafts (Geuna et al., 2004). Although the muscle–vein grafts were able to promote nerve regeneration in 55 mm long nerve defects in rabbits, they were not comparable to nerve autografts (Geuna et al., 2004). Autologous muscle–vein combined grafts have been used clinically in humans to bridge nerve gaps ranging from 5 to 60 mm. The results were scored as “poor”, “satisfactory”, “good”, and “very good”, based on the recovery of sensory and motor functions. Of the 21 lesions repaired (in 20 patients), 10 were lesions of the sensory nerves and 11 were mixed nerve lesions. All lesions in the sensory nerves, except one greater than 30 mm, showed “good” to “very good” recovery. All lesions in the mixed nerves showed “satisfactory” to “good” recovery of motor and sensory functions (Battiston et al., 2000). Although autogenous/natural materials have shown encouraging results when used for nerve repair, they still have certain drawbacks. Autogenous grafts, require a second surgery, and result in the loss of function at donor site, and neuropathic pain at donor site. Allografts have problems related to preservation and immuno-rejection. In order to avoid these problems, grafts made of artificial/synthetic materials have been used. Synthetic Scaffolds for Nerve Repair Among the artificial materials, synthetic tubular NGCs have shown the most promising results so far (Figure 74.2). Some of the commonly used synthetic scaffolds are given in Table 74.1. The use of NGCs reduces tension at the suture line, protects the regenerating axons from the infiltrating scar tissue, and directs

Peripheral Nerve Regeneration

Nerve guidance channel

Proximal nerve end

Components promoting nerve regeneration

Distal nerve end

Filaments ECM protein Scaffold Cells Neurotrophic factor

Figure 74.2 A schematic of a synthetic NGC. The NGC, sutured to the nerve ends, is filled with hydrogel, filaments, cells, neurotrophic factors, and ECM proteins. For an isotropic graft, there would be no filaments and the other components would be distributed uniformly. For an anisotropic graft, there may be filaments, and the other components would be aligned longitudinally or in increasing concentration from proximal to distal nerve end. Table 74.1 Classification of nerve grafts Classifications Isotropic grafts Have uniform distribution of one or more of the four components A. Scaffolds 1. Natural materials 2. Synthetic materials a. NGCs b. Gels

B. Neurotrophic factors

C. ECM proteins D. Support cells Anisotropic grafts Have directional distribution of one or more of the four components A. Scaffolds 1. Aligned filaments

Examples

Veins (Wang et al., 1993 and Ferrari et al., 1999); muscle fibers (Geuna et al., 2004) PLA (Cai et al., 2005), PLLA, PGA; AN/PVC (Uzman and Villegas, 1983) Agarose (Yu and Bellamkonda, 2003); alginate (Suzuki et al., 1999) NGF (Levi-Montalcini, 1987; Thoenen et al., 1987), BDNF (Sendtner et al., 1992), IGF (Glazner et al., 1993); FGF (Gospodarowicz et al., 1987) Laminin (Yu and Bellamkonda, 2003), fibronectin (Chen et al., 2000); collagen (Choi et al., 2005b) SCs (Guenard et al., 1992), fibroblasts (Nakahara, et al., 1996); stem cells (Ansselin et al., 1997; Choi et al., 2005a)

Collagen (Yoshii et al., 2003; Matsumoto et al., 2000); PLLA (Ngo et al., 2003) (Continued )

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Table 74.1 (Continued) Classifications 2. Magnetically aligned gels B. Neurotrophic factors C. ECM proteins D. Support cells

Examples Fibrin (Dubey et al., 2001); collagen (Ceballos et al., 1999; Dubey et al., 1999) NGF (Cao and Shoichet, 2003; Kapur and Shoichet, 2004), BDNF (Cao and Shoichet, 2003); CNTF; FGF Laminin (Kam et al., 2001; Saneinejad and Shoichet, 1998); fibronectin; collagen SCs (Hadlock, et al., 2000; Rutkowski et al., 2004); fibroblasts; stem cells

Autologous nerve grafts (Gospodarowicz et al., 1987; Nichols et al., 2004) Have all the four components: scaffolds, neurotrophic factors, ECM proteins, and cells Nerve allografts (Evans et al., 1999; Midha et al., 2001) Acellular grafts, but structurally similar to autologous nerve graft

the sprouting axons toward their distal targets. The luminal space of NGCs can be filled with growth promoting matrix, growth factors, and/or appropriate cells. In some cases of nerve repair, NGCs have been used to intentionally leave a small gap between the injured nerve ends, to allow accumulation of cytokines, growth factors, and cells (Dahlin and Lundborg, 2001). The NGCs can be used as an excellent experimental tool, to precisely control the distance between the nerve stumps, test the fluid and tissue entering the channel, and vary the properties of the channel. Although NGCs prevent regenerating nerve fibers from wandering, they do not direct axonal growth microscopically. Hence, for the purposes of this chapter, NGCs have been considered as isotropic scaffolds. Nerve regeneration in silicone NGCs has been studied in detail (Williams et al., 1983). Within a few hours of implantation the tube fills with serum exuded by the cut blood vessels in the nerve ends. This fluid contains neurotrophic factors, as well as several cytokines and inflammatory cells, such as macrophages. The macrophages help remove the myelin and axonal debris formed due to injury. The fluid also contains the clotforming protein fibrin. Within days, the fibrin coalesces and forms a longitudinally oriented fibrin cable bridging the two nerve ends. Without the formation of the fibrin cable, axonal regeneration cannot occur, thus making the fibrin cable formation a critical step. The fibrin cable is then invaded by cells migrating from the proximal and distal nerve stumps, including fibroblasts, macrophages, SCs, and endothelial cells (which form capillaries and larger vessels). Axons from the proximal end grow into the fibrin matrix and are engulfed in the cytoplasm of SCs. Some of these axons then reach the distal nerve end and get myelinated. In inert silicone tubes of 10 mm or shorter, these processes occur spontaneously. However, it is generally accepted that impermeable, inert NGCs such as silicone do not support regeneration across defects larger than 10 mm without the presence of exogenous growth factors. The regeneration process can be improved by changing the properties of the tube (permeability, porosity, texture, and electric charge characteristics), and the addition of matrices, neurotrophic factors, ECM molecules, and cells (Valentini and Aebischer, 1997).

Peripheral Nerve Regeneration

Table 74.2 Classification of NGCs based on porosity and degradability Porosity

Degradability

Examples

Impermeable Semipermeable

Non-degradable Non-degradable

Resorbable

Degradable

Silicone (Lundborg et al., 1982) PS (Yu and Bellamkonda, 2003), PAN/PVC (Uzman and Villegas, 1983; Aebischer et al.,1989) PLA (Cai et al., 2005), PGA

Based on the porosity and/or degradability of the material used, NGCs can be classified as impermeable, semipermeable and resorbable (Table 74.2). The silicone tube is an example of an impermeable NGC since it does not permit movement of molecules across the tube walls. Porosity affects the movement of soluble factors, oxygen, and waste products, into and out of the NGCs, which is vital for nerve regeneration. Examples of semipermeable tubes are polysulphone (PS) and polyacrylonitrile/polyvinylchloride (PAN/PVC). Nerves regenerated in semipermeable tubes featured more myelinated axons and less connective tissue (Uzman and Villegas, 1983; Aebischer et al., 1989). PAN/PVC channels with a molecular weight cutoff of 50 kD support regeneration even in the absence of a distal nerve stump (Aebischer et al., 1989). Examples of bioresorbable tubes are polylactic acid (PLA), polyglycolic acid (PGA), poly(L-lactide-co-glycolide) (PLGA), poly(lactide-co-caprolactone) (PLC), and poly(3-hydroxybutyrate) (PHB). The use of bioresorbable tubes negates the need for a second surgery to remove the implant and prevents long-term compression of the nerve. However, it is critical that the degradation of the tube not allow fibroblasts to invade the lumen space before regeneration occurs, as this may prevent axons from regenerating. Inclusion of Hydrogels As Scaffolds NGCs can be filled with gels to support axonal elongation. Here we briefly describe some of the isotropic gels used for nerve regeneration. Agarose Gels

Agarose is a polysaccharide derived from red agar and is widely used in gel electrophoresis and gel chromatography. SeaPrep® agarose hydrogel has been shown to support neurite extension from a variety of neurons in a non-immunogenic manner (Bellamkonda et al., 1995; Labrador et al., 1998; Dillon et al., 2000). Agarose gels also allow molecules to be covalently linked to the gels through functional groups on their polysaccharide chains. For example, laminin protein or fragments of laminin can be covalently coupled to SeaPrep® agarose gels to enhance their ability to support neurite extension (Yu et al., 1999). Although agarose gels support neurite growth on their own, coupling of molecules, such as laminin, significantly enhances the gels’ ability to promote neurite extension. Collagen Gels

Collagen gels and filaments have been used to promote PNS regeneration (scaffolds with collagen filaments will be discussed later in the anisotropic scaffolds section). Collagen gel can be used to fill the intra-luminal space of a vein graft to prevent it from collapsing and improve its nerve repair efficiency. In collagen-filled vein grafts, the number and diameter of myelinated axons was significantly increased compared to vein grafts without collagen gel (Choi et al., 2005b). Nerve repair with silicone tubes can be significantly improved by

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filling them with collagen gel. Collagen tubes filled with collagen gel have promoted more rapid nerve sprouting, and better morphology, than saline-filled collagen tubes (Satou et al., 1986). In some cases, collagen gels have hindered regeneration (Valentini et al., 1987). This negative effect, presumably due to gel remnants blocking diffusion and axonal elongation, might be overcome by reducing the density of the collagen gel (Labrador et al., 1998). Hyaluronic acid, an ECM component, is associated with decreased scarring and improved fibrin matrix formation. It is hypothesized that during the fibrin matrix phase of regeneration, hyaluronic acid organizes the ECM into a hydrated open lattice, thereby facilitating migration of the regenerated axons (Seckel et al., 1995). Hyaluronan-based tubular conduits, used for peripheral nerve regeneration, resulted in more myelinated axons and higher nerve conduction velocities than silicone tubes filled with saline (Wang et al., 1998), with little cytotoxicity (Jansen et al., 2004) upon degradation. Other gels used to promote nerve regeneration in vivo include Matrigel, alginate gels, fibrin gels, and heparin sulfate gels (Madison et al., 1988; Suzuki et al., 1999; Dubey et al., 2001). Neurotrophic Factors for Nerve Regeneration Neurotrophic factors are produced in the target organs and by SCs in response to injury. These neurotrophic factors help maintain the target organ-nerve synapse. A nerve injury usually results in disruption of communication between the target organs and the neuronal cell body, and leads to Wallerian degeneration (breakdown of myelin sheath and axons). Due to cytokines released during Wallerian degeneration, SCs are activated and produce neurotrophins such as nerve growth factor (NGF) and brain-derived neurotrophic factor (BDNF). Although many other trophic factors, including insulin-like growth factor (IGF), fibroblast growth factor (FGF), and ciliary neurotrophic factor (CNTF), have been shown to be involved in the promotion of nerve regeneration (Gospodarowicz et al., 1987; Glazner et al., 1993), it is believed that they are released from SCs following mechanical damage to the cells. NGF is produced in the target organs of sensory and sympathetic nerves in the PNS, and has been shown to stimulate and promote the survival of sensory ganglia and nerves, including spinal sensory nerves and sciatic nerves (Levi-Montalcini, 1987; Thoenen et al., 1987). BDNF is expressed in very low levels in intact adult peripheral nerves, but is upregulated following injury. BDNF is effective in promoting the survival and outgrowth of not only sensory and sympathetic nerves, but also motor nerves (Sendtner et al., 1992). Neurotrophic factors are likely an important part of future clinical therapies for peripheral nerve injuries/diseases. Diseases in which the functions of SCs are severely suppressed (multiple sclerosis, for example) or when acellular grafts (containing no viable SCs) are used, application of neurotrophic factors could be highly effective in facilitating nerve regeneration. Various studies have utilized the functions of NGF to promote nerve regeneration. Hubbell and Sakiyama-Elbert have developed a fibrin matrix that immobilizes heparin molecules by electrostatic interactions, which in turn immobilizes heparin-binding growth factors. The fibrin matrix, when implanted in vivo, releases the bound growth factor due to fibrin degradation. This system was used to deliver NGF (Lee et al., 2003) for peripheral nerve regeneration in vivo, and basic fibroblast growth factor (bFGF) (Sakiyama-Elbert and Hubbell, 2000) for neurite extension from chick dorsal root ganglia (DRG) in vitro. Fibrin–heparin–NGF matrix was observed to promote nerve regeneration comparable to syngenic nerve grafts over a 13 mm nerve gap in rats. Fibrin matrix that released bFGF enhanced neurite extension from DRGs by 100% compared to unmodified fibrin matrix. ECM Molecules for Nerve Regeneration Insoluble ECM molecules, such as laminin, fibronectin, and certain forms of collagen, promote axonal extension and therefore, are excellent candidates for incorporation into the lumen of NGCs. Agarose gels crosslinked

Peripheral Nerve Regeneration

with laminin showed enhanced neurite extension from chick DRG in vitro (Yu et al., 1999). Agarose gels crosslinked with laminin and soluble NGF showed nerve regeneration comparable to autografts over a 10 mm gap in rats (Yu and Bellamkonda, 2003). However, axonal extension in the laminin gels depends on gel density. High concentrations of laminin hinder regeneration (Labrador et al., 1998). Matrigel, a gel containing collagen type IV, laminin, and glycosaminoglycans, supports some degree of regeneration over a long nerve gap in adult rats, when introduced into the lumen of NGCs (Madison et al., 1988). Similarly, a gel mixture containing laminin, collagen, and fibronectin significantly improved nerve regeneration compared to saline-filled silicone channels (Chen et al., 2000). Seeding Neuronal Support Cells for Nerve Regeneration In the PNS, SCs are support cells that wrap around the axons. SCs form a multilamellar sheath of myelin, a phospholipid-containing substance around axons that serves as an insulator and increases nerve conduction velocity. An individual SC may ensheath several unmyelinated axons, but only one myelinated axon within its cytoplasm. In NGCs used for nerve regeneration, formation of fibrin cable, migration of SCs and longitudinal arrangement of SCs (known as bands of Büngner) are necessary processes for axonal regeneration. For nerve gaps less than the critical length (10 mm) these processes occur spontaneously, leading to axonal regeneration. However, for nerve gaps greater than 10 mm, spontaneous nerve regeneration does not occur, due to lack of formation of a fibrin cable and the bands of Büngner (Lundborg, 1988). SCs of uninjured nerves are quiescent. Following nerve axotomy, the SCs become “reactive” and produce a number of neurotrophic factors, including NGF, BDNF, and CNTF (Thanos et al., 1998). They also synthesize and secrete ECM molecules, such as laminin, which is known to modulate neurite outgrowth and express a variety of other cell adhesion molecules. All these components have been known to play roles in supporting neuronal survival and axonal regeneration. Using SCs in NGCs bypasses the fibrin cable formation step, accelerates the formation of bands of Büngner, and introduces a persistent source of neurotrophic factors, leading to more efficient nerve repair. This could decrease the time required by the axons to reconnect to the target organ, as well as increase the distance over which regeneration occurs. SCs isolated from the peripheral nerve of a patient, and expanded in vitro, could be used to treat the patient’s nerve injuries. Addition of SCs has been shown to significantly improve the performance of various scaffolds, such as empty NGCs, collagen gels, venous nerve grafts, and muscle grafts, as compared to control scaffolds without SCs (Ansselin et al., 1997; Strauch et al., 2001; Keilhoff et al., 2005). The ability of SC-seeded NGCs to promote regeneration was found to be dependent on the SCs seeding density, and the immunocompatibility between donors and host (Guenard et al., 1992). For syngenic SCs, it was observed that increasing the seeding density improves the nerve regeneration outcome. Heterologous SCs elicited a strong immune reaction, impeding the nerve regeneration. The performance of SC-seeded NGCs was further improved by designing longitudinally aligned channels in the tube to resemble acellular nerve grafts. This will be discussed in more detail in the section on anisotropic scaffolds. As an alternative to SCs, other cells could be used, as they are genetically modified, to produce desired levels of neurotrophic factors, or to express specific ECM molecules. Fibroblasts, genetically modified to produce NGF, BDNF, neurotrophin 3 (NT-3), and bFGF, showed promising results in central nervous system (CNS) regeneration (Nakahara et al., 1996). Olfactory ensheathing cells have been shown to promote regeneration of cut nerves in the adult rat spinal cord (Li et al., 2003). Although these are examples of CNS regeneration, genetically modified cells can also be used for PNS regeneration. The addition of bone marrow stromal cells to the NGCs has shown improved regeneration over empty NGCs (Choi et al., 2005a). Similarly, pluripotent stem cells derived from hair follicles have shown improvements in rats (Amoh et al., 2005). However, the

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(a)

(b)

(c)

DRG

Figure 74.3 Nanofilaments for contact guidance mediated growth. (a) SEM image of PAN-MA nanofilaments, with diameter of 400–700 nm. (b) Three-dimensional nanofilaments-based scaffold, along with agarose gel, embedded in a polysulfone NGC can be used to direct neurite growth in vitro, or axonal growth in vivo. (c) Chick DRG extending neurites along horizontally oriented PAN-MA filaments, in vitro (Scale bar  1 mm). (Images courtesy: Dr. Young-tae Kim, Georgia Institute of Technology, Atlanta, GA.)

difficulties of isolating and culturing these cells from the patient prior to surgery could limit this approach for some surgical procedures.

ANISOTROPIC NERVE GRAFTS The four essential elements of nerve grafts, scaffolds, neurotrophic factors, ECM molecules, and cells, can be presented in an aligned fashion so as to orient the regenerating axons toward their distal targets. In this section, studies involving nerve grafts that provide directional guidance are discussed. Aligned Anisotropic Scaffolds We hypothesize that the superior performance of autologous nerve grafts is due to its cellular components and its longitudinally aligned structure. The longitudinally aligned structure of the degenerating nerves in the autografts provides contact guidance and direction to the regenerating nerves. In an attempt to mimic autografts, longitudinally patterned or oriented gels and filaments to guide and accelerate the regenerating axons have been designed (Figure 74.3). It has been seen that a poly (acrylonitrile-co-methylacrylate) (PAN-MA) nanofilament-based scaffold by itself can facilitate regeneration across a 17 mm nerve gap in rats (Kim et al.). Many other combinations of materials have been used, such as collagen filaments embedded in collagen tubes (Yoshii et al., 2003), laminin-coated collagen fibers in collagen tubes (Matsumoto et al., 2000), laminin–fibronectin double-coated collagen fibers in collagen tubes (Tong et al., 1994), poly(L-lactide) (PLLA) filaments in silicone tubes and PLA tubes (Ngo et al., 2003; Cai et al., 2005), and PGA fibers in

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chitosan tubes (Wang et al., 2005). All have been found to significantly improve regeneration compared to saline-filled tubes. In addition to synthetic filaments, magnetically aligned fibrin and collagen type I gels have also been used to provide directional guidance to neurites in vitro (Dubey et al., 1999, 2001) and axons in vivo (Ceballos et al., 1999). Neurotrophic Factors Neurotrophic factors have been delivered in vivo mostly in an isotropic manner. However, in vitro studies have suggested that gradients of neurotrophic factors can direct growth cones toward the source of neurotrophic factor (Gundersen and Barrett, 1979). Insoluble and soluble gradients of NGF, NT-3, and BDNF have been shown to direct the growth of neurites from PC12 cells toward increasing concentrations of neurotrophic factors (Cao and Shoichet, 2003; Kapur and Shoichet, 2004). Therefore, anisotropic scaffolds having gradients of neurotrophic factors, along with other components, might be important tools for PNS regeneration. ECM Molecules Gels containing ECM molecules, such as laminin, collagen, fibronectin, and glycosaminoglycans, have been widely used to make isotropic scaffolds for nerve regeneration (discussed earlier in “Current Strategies for Regeneration”). ECM molecules promote axonal growth by the mechanism of differential adhesion, wherein axons preferentially grow on substrates of ECM molecules due to the presence of specific cell surface receptors. In vitro, experiments with spatial patterns of whole ECM molecules (Kam et al., 2001) or their peptide derivates (Saneinejad and Soichet, 1998) have been used to direct the growth of neurites, as well as enhance neurite extension. In vivo, ECM protein-coated fibers have been used to enhance nerve regeneration, where the fibers provide the contact guidance for regenerating axons and the ECM protein provides the adhesive substrate (Tong et al., 1994; Matsumoto et al., 2000). Experiments performed in vitro have demonstrated that gradients of ECM proteins could orient and enhance neurite outgrowth toward increasing concentrations of ECM molecules (Dertinger et al., 2002; Adams et al., 2005). However, this technique is yet to be utilized to enhance nerve regeneration in vivo, due to the difficulties in making gradients of proteins in three-dimensional scaffolds. Cell-Seeded Longitudinally Aligned NGCs Neuronal growth supporting cells can be incorporated with longitudinally aligned filaments and gels in NGCs to enhance nerve regeneration. Since support cells synthesize ECM proteins and neurotrophic factors, aligned cells often result in directionally aligned ECM. Biodegradable conduits of a copolymer of lactic and PLGA with longitudinally aligned channels have been used for nerve regeneration (Hadlock et al., 2000). The channels, with the lumen coated with laminin and seeded with SCs, showed regeneration comparable to nerve autografts over a 7 mm nerve gap in rats. PLA tubes, with a micropatterned inner lumen and seeded with SCs, showed better nerve regeneration compared to unpatterned tubes with SCs (Rutkowski et al., 2004). The disadvantages of cell-seeded NGCs include the need for prolonged isolation and cell culture to prepare cells for implantation, high cell yield, and high cellular morbidity.

NATURAL NERVE GRAFTS A common source of nerve grafts is the sural nerve, which is easy to obtain, has the appropriate diameter for most grafting needs, and is relatively dispensable. Other graft sources include the anterior branch of the medial ante-brachial cutaneous nerve, the lateral femoral cutaneous nerve, and the superficial radial sensory

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nerve (Sunderland, 1991). However, a motor nerve has a preference for a motor pathway (i.e. motor nerve graft) and shows inferior regeneration if a sensory nerve graft, such as sural nerve, is used. Similarly, a mixed nerve shows superior regeneration with either a mixed nerve or a motor nerve graft as compared to a sensory nerve graft (Nichols et al., 2004). Therefore, clinical outcomes might be improved by using alternatives to sensory nerve grafts in the reconstruction of a mixed nerve. However, there are relatively few expendable motor/mixed nerves in the human body that could be used as graft materials. Therefore, a more feasible alternative would be to use nerve allografts or biosynthetic graft materials. Cadavers are a source of graft materials and avoid the complications of harvesting autografts. However, cadaveric nerve allografts require maintenance and can be used only with immunosuppressive therapy. The withdrawal of the immunosuppressant leads to profound loss of axons in the allografts. The axonal loss is most profound in mixed nerve allografts as compared to motor nerve allografts, followed by sensory nerve allografts (Midha et al., 2001). Allografts, cold-preserved and/or freeze-thawed to prevent immuno-rejection by the host body, perform better than fresh allografts in terms of axon density, fiber diameter, and nerve conduction velocity (Evans et al., 1999). Using natural materials (nerve grafts) for regeneration is ideal. However, it has been shown that if autografts or allografts are preserved for too long, their ability to support regeneration is compromised (Gulati, 1996). Also, the pre-treated allografts do not perform as well as autografts (Evans et al., 1999). Although nerve autografts are used as a “gold standard,” the lack of functional recovery even with autografts remains an important clinical problem. Techniques utilized to improve the performance of the nerve autografts include treatments to either remove inhibitory molecules like chondroitin sulfate proteoglycans (CSPGs), or provide factors for axonal growth, such as neurotrophins 4/5 (NT-4/5) or BDNF. CSPG molecules have a core protein structure with glycosaminoglycan (GAG) side chains composed of chondroitin sulfate. Due to their large size and negative charge, the GAGs of CSPGs are thought to hinder neurite access to growth-promoting matrix molecules and also to repel the axons, thereby inhibiting their growth (Properzi et al., 2003). It has been shown that CSPGs are upregulated almost seven-fold in the distal segment of peripheral nerve gaps following transection nerve injury (Zuo et al., 1998). The upregulated CSPGs contribute significantly to inhibition of neurite sprouting and, consequently, growth into the distal nerve segment. Treatment of the injury site with chondroitinase ABC, which digests away the inhibitory CSPGs, increases the neurite ingrowth into the distal nerve segment several fold, as compared to untreated controls without any chondroitinase ABC treatment (Zuo et al., 2002). Treatment with chondroitinase ABC, however, did not improve neurite ingrowth in a crush injury model, suggesting that CSPGs are not upregulated in a crush injury. Syngenic nerve grafts treated with chondroitinase ABC, heparinase I, heparinase III, or keratanase enzymes have shown significantly improved axonal ingrowth from the proximal nerve end into the nerve graft, as compared to untreated controls (Groves et al., 2005). Autografts treated with a combination of all these four enzymes showed the most significant neurite growth into the graft. However, the combination treatment was not significantly different from the arithmetic sum of the individual treatments. This suggests that molecules such as heparan sulfate proteoglycan and keratan sulfate proteoglycan also contribute to inhibition of neurite growth apart from CSPGs, and the pathways/mechanisms of inhibition for each of these molecules might be independent of each other. In another study, an exogenous supply of BDNF and NT-4/5 was found to increase the number of axons regenerating into the nerve graft as compared to untreated nerve grafts (English et al., 2005). These techniques, used in clinical applications, could lead to better results with autografts/allografts.

ANIMAL MODELS Traditionally, nerve regeneration studies have involved the use of various animal models, such as mouse, rat, swine, canine, sheep, and non-human primates. Rat or mouse models are used initially to determine the

Peripheral Nerve Regeneration

efficacy of the various treatments. If the results are encouraging, they are followed by experiments with larger animal models. For PNS regeneration studies, the most commonly studied nerve models are the sciatic nerve and its branches, the tibial and the peroneal nerves. Other models include the cavernosal nerve and the facial nerve. The most common nerve injury model is the single-anastomosis model, where the injury and repair are done on one sciatic nerve and the contralateral sciatic nerve is used as a control. This model is useful when the nerve gap is less than 20 mm. The second version is the cross (double) anastomosis model, where both contralateral sciatic nerves are transected; the proximal end of the right sciatic nerve is then sutured to one end of an implanted tube, and the distal end of the left sciatic nerve is inserted into the other end of the tube (Lundborg et al., 1982). This model allows the study of gaps in excess of 25 mm. Although very convenient, the rat and mouse models suffer from the serious drawback that they present only short nerve gaps for regeneration studies. In order for a regeneration technique to be successfully applied in clinical trials, the nerve gap model has to be more than 40 mm in length. To create a long nerve gap model, rabbits (Geuna et al., 2004), cats (Suzuki et al., 1999), dogs (Matsumoto et al., 2000), sheep (Lawson and Glasby, 1998), and non-human primates (Ahmed et al., 1999) have been used. The large animal models are an important intermediary step before clinical application of experimental therapeutic approaches.

SUMMARY In spite of significant advances in research to the development of synthetic NGCs, nerve autografts are still considered the first-choice strategy for nerve repair, especially in the case of long nerve gaps. However, even the performance of autografts has been unsatisfactory. Using autografts generally results in a good recovery of sensory functions, but negligible return of motor functions. Hence, there will be continued interest in ideas to further enhance the performance of autografts by various treatments, such as chondroitinase ABC, NT-4/5, and BDNF. However, shortage of autografts and allografts is a hindrance to their usage. This shortage can be overcome only by developing synthetic alternatives to autografts. Modulating the spatio-temporal distribution of the four components of grafts germane to regeneration can potentially improve the potential outcomes with these grafts. Ongoing rapid advances in cell biology, cell culture techniques, genetic engineering, and biomaterials research are likely to provide new tools to improve regeneration using NGCs, and the day an engineered construct performs as well as autografts may be near.

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Bellamkonda, R.V., Ranieri, J.P., Bouche, N. and Aebischer, P. (1995). Hydrogel-based three-dimensional matrix for neural cells. J. Biomed. Mater. Res. 29: 663–671. Cai, J., Peng, X., Nelson, K.D., Eberhart, R. and Smith, G.M. (2005). Permeable guidance channels containing microfilament scaffolds enhance axon growth and maturation. J. Biomed. Mater. Res. A. 75(2): 374–386. Cao, X. and Shoichet, M.S. (2003). Investigating the synergistic effect of combined neurotrophic factor concentration gradients to guide axonal growth. Neuroscience 122(2): 381–389. Ceballos, D., Navarro, X., Dubey, N., Wendelschafer-Crabb, G., Kennedy, W.R. and Tranquillo, R.T. (1999). Magnetically aligned collagen gel filling a collagen nerve guide improves peripheral nerve regeneration. Exp. Neurol. 158(2): 290–300. Chen, Y.S., Hsieh, C.L., Tsai, C.C., Chen, T.H., Cheng, W.C., Hu, C.L. and Yao, C.H. (2000). Peripheral nerve regeneration using silicone rubber chambers filled with collagen, laminin and fibronectin. Biomaterials 21(15): 1541–1547. Chiu, D.T. and Strauch, B. (1990). A prospective clinical evaluation of autogenous vein grafts used as a nerve conduit for distal sensory nerve defects of 3 cm or less. Plast. Reconstr. Surg. 86(5): 928–934. Choi, B.H., Zhu, S.J., Kim, B.Y., Huh, J.Y., Lee, S.H. and Jung, J.H. (2005a). Transplantation of cultured bone marrow stromal cells to improve peripheral nerve regeneration. Int. J. Oral Maxillofac. Surg. 34(5): 537–542. Choi, B.H., Zhu, S.J., Kim, S.H., Kim, B.Y., Huh, J.H., Lee, S.H. and Jung, J.H. (2005b). Nerve repair using a vein graft filled with collagen gel. J. Reconstr. Microsurg. 21(4): 267–272. Dahlin, L.B. and Lundborg, G. (2001). Use of tubes in peripheral nerve repair. Neurosurg. Clin. N. Am. 12(2): 341–352. Dertinger, S.K., Jiang, X., Li, Z., Murthy, V.N. and Whitesides, G.M. (2002). Gradients of substrate-bound laminin orient axonal specification of neurons. Proc. Natl Acad. Sci. USA 99(20): 12542–12547. Dillon, G.P., Yu, X. and Bellamkonda, R.V. (2000). The polarity and magnitude of ambient charge influences threedimensional neurite extension from DRGs. J. Biomed. Mater. Res. 51(3): 510–519. Dubey, N., Letourneau, P.C. and Tranquillo, R.T. (1999). Guided neurite elongation and Schwann cell invasion into magnetically aligned collagen in simulated peripheral nerve regeneration. Exp. Neurol. 158(2): 338–350. Dubey, N., Letourneau, P.C. and Tranquillo, R.T. (2001). Neuronal contact guidance in magnetically aligned fibrin gels: effect of variation in gel mechano-structural properties. Biomaterials 22(10): 1065–1075. English, A.W., Meador, W. and Carrasco, D.I. (2005). Neurotrophin-4/5 is required for the early growth of regenerating axons in peripheral nerves. Eur. J. Neurosci. 21(10): 2624–2634. Evans, G.R. (2001). Peripheral nerve injury: a review and approach to tissue engineered constructs. Anat. Rec. 263(4): 396–404. Evans, P.J., MacKinnon, S.E., Midha, R., Wade, J.A., Hunter, D.A., Nakao, Y. and Hare, G.M. (1999). Regeneration across cold preserved peripheral nerve allografts. Microsurgery 19(3): 115–127. Ferrari, F., De Castro Rodrigues, A., Malvezzi, C.K., Dal Pai Silava, M. and Padvoni, C.R. (1999). Inside-out vs. standard vein graft to repair a sensory nerve in rats. Anat. Rec. 256: 227–232. Geuna, S., Tos, P., Battiston, B. and Giacobini-Robecchi, M.G. (2004). Bridging peripheral nerve defects with muscle–vein combined guides. Neurol Res. 26(2): 139–144. Glazner, G.W., Lupien, S., Miller, J.A. and Ishii, D.N. (1993). Insulin-like growth factor II increases the rate of sciatic nerve regeneration in rats. Neuroscience 54(3): 791–797. Gluck, T. (1880). Ueber Neuroplastik auf dem Wege der Transplantation. Arch. Klin. Chir. 25: 606–616. Gospodarowicz, D., Ferrara, N., Schweigerer, L. and Neufeld, G. (1987). Structural characterization and biological functions of fibroblast growth factor. Endocr. Rev. 8(2): 95–114. Groves, M.L., McKeon, R., Werner, E., Nagarsheth, M., Meador, W. and English, A.W. (2005). Axon regeneration in peripheral nerves is enhanced by proteoglycan degradation. Exp. Neurol. 195: 278–292. Guenard, V., Kleitman, N., Morrissey, T.K., Bunge, R.P. and Aebischer, P. (1992). Syngeneic Schwann cells derived from adult nerves seeded in semipermeable guidance channels enhance peripheral nerve regeneration. J. Neurosci. 12(9): 3310–3320. Gulati, A.K. (1996). Peripheral nerve regeneration through short- and long-term degenerated nerve transplants. Brain Res. 742(1–2): 265–270. Gundersen, R.W. and Barrett, J.N. (1979). Neuronal chemotaxis: chick dorsal-root axons turn toward high concentrations of nerve growth factor. Science 206(4422): 1079–1080.

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Hadlock, T., Sundback, C., Hunter, D., Cheney, M. and Vacanti, J.P. (2000). A polymer foam conduit seeded with Schwann cells promotes guided peripheral nerve regeneration. Tissue Eng. 6(2): 119–127. Jansen, K., van der Werff, J.F., van Wachem, P.B., Nicolai, J.P., de Leij, L.F. and van Luyn, M.J. (2004). A hyaluronan-based nerve guide: in vitro cytotoxicity, subcutaneous tissue reactions, and degradation in the rat. Biomaterials 25(3): 483–489. Kam, L., Shain, W., Turner, J.N. and Bizios, R. (2001). Axonal outgrowth of hippocampal neurons on micro-scale networks of polylysine-conjugated laminin. Biomaterials 22(10): 1049–1054. Kapur, T.A. and Shoichet, M.S. (2004). Immobilized concentration gradients of nerve growth factor guide neurite outgrowth. J Biomed. Mater. Res. A 68(2): 235–243. Keilhoff, G., Pratsch, F., Wolf, G. and Fansa, H. (2005). Bridging extra large defects of peripheral nerves: possibilities and limitations of alternative biological grafts from a cellular muscle and Schwann cells. Tissue Eng. 11(7–8): 1004–1114. Kim, Y., Haftel, V.K., Kumar, S. and Bellamkonda, R.V. Oriented nanoscaffolds match the performance of autografts in facilitating regeneration across long peripheral nerve gaps. Nat. Biotechnol. (in review). Kirk, E.G. and Lewis, D. (1915). Fascial tubulization in the repair of nerve defects. J. Am. Med. Assoc. 65: 486–492. Kline, D.G., Kim, D., Midha, R., Harsh, C. and Tiel, R. (1998). Management and results of sciatic nerve injuries: a 24-year experience. J. Neurosurgery 89: 13–23. Labrador, R.O., Buti, M. and Navarro, X. (1998). Influence of collagen and laminin gels concentration on nerve regeneration after resection and tube repair. Exp. Neurol. 149(1): 243–252. Lawson, G.M. and Glasby, M.A. (1998). Peripheral nerve reconstruction using freeze-thawed muscle grafts: a comparison with group fascicular nerve grafts in a large animal model. J. Roy. Coll. Surg. Edin. 43(5): 295–302. Lee, A.C., Yu, V.M., Lowe III, J.B., Brenner, M.J., Hunter, D.A., Mackinnon, S.E. and Sakiyama-Elbert, S.E. (2003). Controlled release of nerve growth factor enhances sciatic nerve regeneration. Exp. Neurol. 184(1): 295–303. Levi-Montalcini, R. (1987). The nerve growth factor 35 years later. Science 237(4819): 1154–1162. Li, Y., Decherchi, P. and Raisman, G. (2003). Transplantation of olfactory ensheathing cells into spinal cord lesions restores breathing and climbing. J. Neurosci. 23(3): 727–731. Lundborg, G. (1988). Nerve Injury and Repair. New York: Longman Group UK Ltd. Lundborg, G., Dahlin, L.B., Danielsen, N., Gelberman, R.H., Longo, F.M., Powell, H.C. and Varon, S. (1982). Nerve regeneration in silicone chambers: influence of gap length and of distal stump components. Exp. Neurol. 76(2): 361–375. Madison, R.D., Da Silva, C.F. and Dikkes, P. (1988). Entubulation repair with protein additives increases the maximum nerve gap distance successfully bridged with tubular prostheses. Brain Res. 447(2): 325–334. Matsumoto, K., Ohnishi, K., Kiyotani, T., Sekine, T., Ueda, H., Nakamura, T., Endo, K. and Shimizu, Y. (2000). Peripheral nerve regeneration across an 80-mm gap bridged by a polyglycolic acid (PGA)-collagen tube filled with laminincoated collagen fibers: a histological and electrophysiological evaluation of regenerated nerves. Brain Res. 868(2): 315–328. Midha, R., Nag, S., Munro, C.A. and Ang, L C. (2001). Differential response of sensory and motor axons in nerve allografts after withdrawal of immunosuppressive therapy. J. Neurosurg. 94(1): 102–110. Millesi, H., Meissl, G. and Berger, A. (1972). The interfascicular nerve-grafting of the median and ulnar nerves. J. Bone Joint Surg. Am. 54: 7727–7750. Nakahara, Y., Gage, F.H. and Tuszynski, M.H. (1996). Grafts of fibroblasts genetically modified to secrete NGF, BDNF, NT-3, or basic FGF elicit differential responses in the adult spinal cord. Cell Transplant. 5(2): 191–204. Ngo, T.T., Waggoner, P.J., Romero, A.A., Nelson, K.D., Eberhart, R.C. and Smith, G.M. (2003). Poly(L-lactide) microfilaments enhance peripheral nerve regeneration across extended nerve lesions. J. Neurosci. Res. 72(2): 227–238. Nichols, C.M., Brenner, M.J., Fox, I.K., Tung, T.H., Hunter, D.A., Rickman, S.R. and Mackinnon, S.E. (2004). Effects of motor versus sensory nerve grafts on peripheral nerve regeneration. Exp. Neurol. 190(2): 347–355. Payr, E. (1900). Beitrage zur Technik der Blutgefass und Nervennaht nebst Mittheilungen uber die Vervendung eines resorbibaren Metalles in der Chirurgie. Arch. Klin. Chir. 62: 67. Properzi, F., Asher, R.A. and Fawcett, J.W. (2003). Chondroitin sulphate proteoglycans in the central nervous system: changes and synthesis after injury. Biochem. Soc. Trans. 31(2): 335–336.

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Rutkowski, G.E., Miller, C.A., Jeftinija, S. and Mallapragada, S.K. (2004). Synergistic effects of micropatterned biodegradable conduits and Schwann cells on sciatic nerve regeneration. J. Neural. Eng. 1(3): 151–157. Sakiyama-Elbert, S.E. and Hubbell, J.A. (2000). Development of fibrin derivatives for controlled release of heparin-binding growth factors. J. Control. Release 65(3): 389–402. Saneinejad, S. and Shoichet, M.S. (1998). Patterned glass surfaces direct cell adhesion and process outgrowth of primary neurons of the central nervous system. J. Biomed. Mater. Res. 42(1): 13–19. Satou, T., Nishida, S., Hiruma, S., Tanji, K., Takahashi, M., Fujita, S., Mizuhara, Y., Akai, F. and Hashimoto, S. (1986). A morphological study on the effects of collagen gel matrix on regeneration of severed rat sciatic nerve in silicone tubes. Acta. Pathol. Jpn. 36(2): 199–208. Seckel, B.R., Jones, D., Hekimian, K.J., Wang, K.K., Chakalis, D.P. and Costas, P.D. (1995). Hyaluronic acid through a new injectable nerve guide delivery system enhances peripheral nerve regeneration in the rat. J. Neurosci. Res. 40(3): 318–324. Sendtner, M., Holtmann, B., Kolbeck, R., Thoenen, H. and Barde, Y.A. (1992). Brain-derived neurotrophic factor prevents the death of motoneurons in newborn rats after nerve section. Nature 360(6406): 757–759. Strauch, B., Rodriguez, D.M., Diaz, J., Yu, H.L., Kaplan, G. and Weinstein, D.E. (2001). Autologous Schwann cells drive regeneration through a 6-cm autogenous venous nerve conduit. J. Reconstr. Microsurg. 17: 589–595. Sunderland, S. (1991). Nerve Injuries and Their Repair: A Critical Appraisal. New York: Churchill Livingstone. Suzuki, Y., Tanihara, M., Ohnishi, K., Suzuki, K., Endo, K. and Nishimura, Y. (1999). Cat peripheral nerve regeneration across 50 mm gap repaired with a novel nerve guide composed of freeze-dried alginate gel. Neurosci. Lett. 259(2): 75–78. Terzis, J., Faibisoff, B. and Williams, B. (1975). The nerve gap: suture under tension vs. graft. Plast. Reconstr. Surg. 56(2): 166–170. Thanos, P.K., Okajima, S. and Terzis, J.K. (1998). Ultrastructure and cellular biology of nerve regeneration. J. Reconstr. Microsurg. 14: 423–436. Thoenen, H., Barde, Y.A., Davies, A.M. and Johnson, J.E. (1987). Neurotrophic factors and neuronal death. Ciba Found. Symp. 126: 82–95. Tong, X.J., Hirai, K., Shimada, H., Mizutani, Y., Izumi, T., Toda, N. and Yu, P. (1994). Sciatic nerve regeneration navigated by laminin–fibronectin double coated biodegradable collagen grafts in rats. Brain Res. 663(1): 155–162. Uzman, B.G. and Villegas, G.M. (1983). Mouse sciatic nerve regeneration through semi-permeable tubes: a quantitative model. J. Neurosci. 9: 325–338. Valentini, R.F. and Aebischer, P. (1997). Strategies for the engineering of peripheral nervous tissue regeneration. In: Lanza, R.P., Langer, R., Chick, W.L. (eds.), Principles of Tissue Engineering. Austin, TX: R.G. Landes Company, pp. 671–684. Valentini, R.F., Aebischer, P., Winn, S.R. and Galletti, P.M. (1987). Collagen- and laminin-containing gels impede peripheral nerve regeneration through semipermeable nerve guidance channels. Exp. Neurol. 98(2): 350–356. Wang, K.K., Costas, P.D., Bryan, D.J., Jones, D.S. and Seckel, B.R. (1993). Inside-out vein graft promotes improved nerve regeneration in rats. J. Reconstr. Microsurg. 14: 608–618. Wang, K.K., Nemeth, I.R., Seckel, B.R., Chakalis-Haley, D.P., Swann, D.A., Kuo, J.W., Bryan, D.J. and Cetrulo Jr., C.L. (1998). Hyaluronic acid enhances peripheral nerve regeneration in vivo. Microsurgery 18(4): 270–275. Wang, X., Hu, W., Cao, Y., Yao, J., Wu, J. and Gu, X. (2005). Dog sciatic nerve regeneration across a 30-mm defect bridged by a chitosan/PGA artificial nerve graft. Brain 128(Pt 8): 1897–1910. Weiss, P. and Taylor, A.C. (1946). Guides for nerve regeneration across nerve gaps. J. Neurosurg. 3: 275–282. Williams, L.R., Longo, F.M., Powell, H.C., Lundborg, G. and Varon, S. (1983). Spatial–temporal progress of peripheral nerve regeneration within a silicone chamber: parameters for a bioassay. J. Comp. Neurol. 218(4): 460–470. Yoshii, S., Oka, M., Shima, M., Taniguchi, A. and Akagi, M. (2003). Bridging a 30-mm nerve defect using collagen filaments. J. Biomed. Mater. Res. A 67(2): 467–474. Yu, X. and Bellamkonda, R.V. (2003). Tissue-engineered scaffolds are effective alternatives to autografts for bridging peripheral nerve gaps. Tissue Eng. 9(3): 421–430. Yu, X., Dillon, G.P. and Bellamkonda, R.B. (1999). A laminin and nerve growth factor-laden three-dimensional scaffold for enhanced neurite extension. Tissue Eng. 5(4): 291–304.

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75 Dental Tissue Engineering Yan Lin and Pamela C. Yelick INTRODUCTION It has long been appreciated that the oral cavity functions as an important barrier to microbial infections that can adversely affect the overall health and well-being of an individual. Diseased teeth and resulting compromised oral health present real threats to healthy lifestyles. This threat becomes more significant over time as teeth wear and/or become damaged after years of prolonged use and as the potential susceptibility to disease increases with advancing age. The need for improved dental tissue repair/regeneration methods has been recognized for many centuries, and a variety of approaches have been explored to restore teeth loss, including tooth autotransplantation, allotransplantation, and dental implant methods (Yelick and Vacanti, 2004; Yen and Sharpe, 2006). Although the current availability and use of prosthetic dental implants contributes to an improved quality of life for many individuals, certain limitations associated with those procedures, including limited tooth autograft available and the possibility of tooth allograft or foreign body rejection by the patient makes this method unsuitable for a number of potential recipients. Thus, the development of new methods to repair dental tissues or to replace whole teeth is in great demand. Recent advancements in stem cell biology and material science have highlighted tissue engineering as an emerging science with the potential, in the relatively near future, to facilitate the successful development of replacement tissues and organs. In ongoing preclinical studies, cell- and gene-based therapies are being developed for a variety of tissues and organs, including bone, heart, liver, and kidney. Methods to deliver stem/progenitor cells, biodegradable scaffolds, and growth factors and/or morphogen gradients to tissue injury sites are being used to accelerate and/or induce natural biological regeneration. With respect to dental applications, stem cell-based dental tissue engineering provides the opportunity to consider biologically based reparative and/or replacement tooth therapies, with the potential to regenerate dental tissues exhibiting physical and esthetic properties that are equal to, or better than, existing counterparts. Because of the rapid growth in this field, this chapter will present and discuss some of the most recent developments in various subfields of dental tissue engineering, which collectively highlight the powerful impact that biologically based dental tissue regeneration strategies will have on the field of dentistry in the foreseeable future. Natural Tooth Tissue Development and Repair Teeth consist of crown, neck, and root structures, and are composed of a variety of hard and soft dental tissues (Figure 75.1). Tooth crowns consist of an outer mineralized enamel layer, the subjacent mineralized dentin layer, and an inner dental pulp tissue. Tooth roots consist of a small root canal containing dental pulp and nerves, surrounded by mineralized dentin and cementum layers and integrated periodontal ligament (PDL) tissue, which secures the tooth to the underlying alveolar bone (Nanci, 2003). The enamel surface of the tooth is produced by specialized dental epithelial cells, called ameloblasts. Enamel is primarily mineral and contains enamel-specific proteins including amelogenin (Snead et al., 1983), ameloblastin, and enamelin (Krebsbach et al., 1996; Paine et al., 1998). Enamel is extremely hard and durable

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Enamel Crown

Dentin Pulp

Neck Cementum Root

Periodontal ligament Root canal

Figure 75.1 Erupted tooth structures.

but is not self-regenerative, since ameloblasts are no longer present in postnatal or adult teeth once they have erupted. With the exception of enamel, all of the other tooth tissues are derived from neural crest cell (NCC)derived dental mesenchyme. The tooth layer underlying the enamel is composed of dentin, which is approximately 75% mineralized and contains dentin-specific gene products, including dentin sialoprotein (DSP), dentin phosphoprotein (DPP), and dentin matrix protein-1 (DMP-1) (Begue-Kirn et al., 1998; Butler et al., 2002). Dentin is produced by NCC-derived dental mesenchymal cells called odontoblasts, which exhibit limited regenerative capacities in response to injury or disease. The dental pulp at the center of the tooth is composed of dental mesenchymal cells, including putative dental stem cells (DSCs), nerves, and blood vessels that thread through the root canal and support the tooth organ. The periodontium, which supports and attaches teeth to the jaw, primarily consists of four connective tissues: mineralized cementum, fibrous PDL, alveolar bone, and gingival tissue (Beertsen et al., 1997). Cementum is a bone-like mineralized tissue lining the dentin of the root that protects the root and also serves as an attachment surface to anchor the PDL to the tooth (Diekwisch, 2001). The PDL functions to secure the tooth to the underlying alveolar bone and is essential for the long-term survival of the tooth. Sub-functioning PDL tissue can result in tooth root damage and eventual ankylosis of the tooth. Cells of the PDL include fibroblasts, cementoblasts on the root surface, osteoblasts and osteoclasts present on the alveolar bone face, and undifferentiated mesenchymal tissues in the body of the ligament (Taba et al., 2005; Nanci and Bosshardt, 2006). The periodontium also possesses some regenerative capability, as tissue loss during the early phases of periodontal diseases can be restored to a certain degree. However, once periodontitis becomes established only therapeutic intervention has the potential to induce regeneration (Shi et al., 2005). When considering methods to regenerate teeth, it is first necessary to have a thorough understanding of de novo tooth development. Like many other organs, such as hair follicles, salivary glands, intestines, and kidneys, tooth development is the cumulative result of reciprocal signaling events between the ectoderm-derived dental epithelial and the NCC-derived mesenchyme (Jernvall and Thesleff, 2000). Molecular signals initiated in the dental epithelium induce gene expression in the subjacent dental mesenchyme, and the acquired odontogenic potential in the mesenchyme triggers further epithelial morphogenesis and cytodifferentiation. As tooth development progresses, dental mesenchyme differentiates into dentin, pulp, cementum, PDL, and alveolar bone, and the epithelial tissues produce dental enamel. The cellular and molecular natures of enamel,

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dentin, cementum, and root development have been extensively characterized (Linde and Goldberg, 1993; Zeichner-David et al., 1995; Bartlett and Simmer, 1999; Chai et al., 2000; Grzesik et al., 2000; Jernvall and Thesleff, 2000; Saygin et al., 2000; Diekwisch, 2001; Bosshardt, 2005; Luan et al., 2006). The morphological stages of natural tooth development consist of early bud and cap stages, and later bell stage where tooth histodifferentiation and morphodifferentiation occur (Thesleff et al., 2001). Tooth development begins as a localized thickening of the dental epithelium, which proliferates and invaginates into the underlying dental mesenchyme, appearing as a small epithelial bud. The primary “enamel knot,” a signaling center that forms at the tip of the epithelial tooth bud, and subsequently formed enamel knots, all of which mark the site of future cusp formation, are considered to be central regulators of tooth development, as they link dental cell differentiation to tooth morphogenesis (Thesleff et al., 2001). The primary enamel knot expresses at least 10 different signaling molecules belonging to the bone morphogenetic protein (BMP), fibroblast growth factor (FGF), hedgehog (Hh), and Wnt (an amalgam of wingless- and int-related proteins) families, which are responsible for the transition from bud to cap stage (Thesleff and Sharpe, 1997; Jernvall and Thesleff, 2000). Epithelial signals emanating from the primary enamel knot induce the formation of the dental papilla, the mesenchymal cell layer underlying the dental epithelium, at the transition from bud to cap stage. Later on, cells of the dental papilla differentiate into odontoblasts, which subsequently differentiate and produce dentin (Thesleff et al., 2001). In bell stage teeth, so named for their characteristic bell-shaped appearance, odontoblasts induce ameloblast formation, and the ameloblasts differentiate and synthesize enamel. Also, at the bell stage the dental lamina connecting the tooth germ to the oral epithelium disappears, isolating the tooth from the oral cavity. Dental Tissue Engineering Realizing the full potential of tissue regenerative treatments for the oral complex will require the integration of three key elements: (1) dental stem/progenitor cells; (2)inductive morphogenetic signals (morphogens); and (3) scaffold material upon which progenitor cells attach and elaborate an extracellular matrix (Nakashima and Reddi, 2003). DSCs One way to regenerate whole teeth would be to mimic the process of natural tooth development, either in vitro or in vivo, using DSCs. DSCs are a somewhat elusive population of self-renewing cells that exhibit the potential to form biological replacement tooth tissues. A variety of DSC populations have been identified for potential use in tooth tissue engineering strategies, as recently reviewed (Krebsbach and Robey, 2002; Murray and Garcia-Godoy, 2004; Yelick and Vacanti, 2004; Risbud and Shapiro, 2005; Shi et al., 2005; Bartold et al., 2006; Yen and Sharpe, 2006). Because tooth development is dependent on epithelial and mesenchymal cell interactions, DSCs necessarily consist of two types: epithelial DSCs, which form ameloblasts, enamel, stellate reticulum, and stratum intermedium, and mesenchymal DSCs, which form the dental papilla, odontoblasts, predentin, dentin, cementum, PDL, and alveolar bone (Nanci, 2003). The exclusive environment in which stem cells reside, called the “stem cell niche,” is thought to support the maintenance and self-renewal of stem cells (Spradling et al., 2001). In the remaining paragraphs, we will focus on the applications of stem cells in dental tissue engineering. Significant efforts currently focus on the development of methods to control the differentiation and proliferation of DSCs, to facilitate the creation of stem cell-based therapies that can be used more effectively than current synthetic material-based dental treatment regimes (Yelick and Vacanti, 2006). For instance, Miura

Dental Tissue Engineering

et al. (2003) isolated stem cells from human exfoliated deciduous teeth, which were found to be capable of differentiating into a variety of cell types, including neural cells, adipocytes, and odontoblasts. The reparative potential of these progenitor cells is now being carefully scrutinized. It has long been appreciated that the PDL harbors progenitor cells that can differentiate into fibroblasts, osteoblasts, and cementoblasts (McCulloch, 1995; Bartold et al., 2006). Many investigators, including Seo and coworkers (2004), have isolated multipotent postnatal human PDL stem cells and demonstrated their capacity to generate cementum/PDL-like tissue after in vivo transplantation. However, at the present time, limited knowledge exists about the manner in which the specialized periodontal tissues are organized and the identity of the cells that contribute to the formation of each type of periodontal tissue (Shi et al., 2005). What has become increasingly clear is the fact that as knowledge of stem cell biology continues to advance, including methods to identify, isolate, expand, and differentiate DSCs, biologically based dental repair and whole-tooth regeneration therapies will become realistic possibilities. Growth Factors and Morphogen-Based Cell Signaling Gradients A major focus of contemporary developmental biology has been to delineate the biological cues that drive stem cell proliferation and differentiation (Thesleff and Sharpe, 1997; Nakao et al., 2004). Knowledge of the responsiveness of DSCs to various morphogenetic signals is of potentially significant importance for improving dental regenerative therapies. Continued efforts to elucidate signaling pathways regulating natural tooth development focus on four families of growth factors, the BMP, FGF, Hh, and Wnt, which govern tissue-specific patterning and morphogenesis during odontogenesis (Yen and Sharpe, 2006). Growth factor ligands from each of these families are currently being evaluated for their utility in guided, stem cell-based dental tissue engineering efforts. It has been demonstrated that delivery of signaling molecules, including BMP (Ripamonti and Reddi, 1997; Nakashima and Reddi, 2003; Iohara et al., 2004) and platelet-derived growth factor (PDGF) (Rutherford et al., 1993; Giannobile et al., 2001; Jin et al., 2004) via gene therapy vectors, resulted in partial dental tissue regeneration, for example, dentin and periodontal tissues. The results suggest that such factors, alone or in combination with other agents, may eventually be used to promote dental tissue regeneration. Scaffolding Materials In the context of tissue engineering, scaffolds made of natural or synthetic materials play a central role in supporting cells during the formation of functional tissues. Therefore, a suitable three-dimensional carrier with tailored macroscopic properties, a well-tuned degradation profile, and specific biological cues are necessary to promote successful tissue growth (Kleinman et al., 2003; Abukawa et al., 2006; Ramseier et al., 2006). Polymers and ceramics are the two major families of biomaterials currently used in tissue engineering. To date, various synthetic biodegradable polymer scaffolds, such as polyglycolic acid (PGA) (Sumita et al., 2006) and poly-L-lactate-co-glycolate (PLGA) (Young et al., 2002, 2005) have been used in dental-related tissue regeneration. Jin et al. (2003) and Zhao et al. (2004) have also used PLGA polymer sponges to deliver cementoblasts, fibroblasts, and dental follicle cells to facilitate periodontal regeneration. These synthetic polymers can be fabricated into desirable sizes and shapes, with readily adjustable degradation profile and pore features. However, they exhibit poor biocompatibility due to release of acidic degradation products and lose mechanical properties very early during degradation (Gunatillake and Adhikari, 2003). Natural polymers, like collagen, have been explored as scaffolds in whole-tooth (Sumita et al., 2006) or periodontal (Rutherford et al.,

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Epithelium

Mesenchyme

Scaffold

Epithelium/mesenchyme interaction

Yellow oval: ameloblast Green oval: odontoblast Ameloblasts Enamel Dentin Odontoblasts Pulp

Figure 75.2 Scheme of whole-tooth tissue engineering.

1993) bioengineering. Properties that can compromise the efficiency of collagen-based scaffolds are rapid absorption and weak mechanical strength (Chevallay and Herbage, 2000). Ceramics, including hydroxyapatite (HA) and tricalcium phosphate (TCP) (Gronthos et al., 2000, Grzesik et al., 2000, Seo et al., 2004; Saito et al., 2005; Shi et al., 2005; Zhang et al., 2005), are also being examined for dental tissue regeneration applications, due to their biocompatibility, osteoconductivity, and structural similarity to the inorganic component of bone. However, the predictability of ceramic degradation is poor, as it is dependent on many factors, including material crystallinity, porosity, density, and host response (Theiss et al., 2005). Tissue regeneration is a highly coordinated process, involving stem cells, scaffolds, and regulatory signals. Scaffold chemistry, morphology, and structure directly impact cell–cell interactions and signaling cascades. For dental tissue regenerative purposes, studies of scaffolding materials are being conducted to identify compatible matches between materials and biological functions. In particular, since tooth development depends on interactions between dental epithelial and mesenchymal cells, to successfully bioengineer teeth of predetermined size and shape, biodegradable scaffolds are desired that can initially facilitate proper epithelial and mesenchymal cell orientation, to guide the subsequent formation of highly mineralized tooth tissues (Figure 75.2). Prior and Current Dental Tissue Regeneration Research The foundation of dental tissue bioengineering has ancient roots. Therapies to replace missing teeth can be traced back at least 2,500 years, when the Etruscans learned to substitute missing teeth with bridges made from

Dental Tissue Engineering

artificial teeth carved from the bones of oxen (Ring, 1995). More recent efforts to regenerate teeth from ectopically grafted embryonic tooth buds have been previously reviewed (Yelick and Vacanti, 2004; Yen and Sharpe, 2006), demonstrating the feasibility of growing teeth in an appropriate environment. State-of-the-art dental tissue engineering methodologies, which employ in vitro expanded presumptive DSCs that are then seeded onto three-dimensional biodegradable scaffolds and implanted back into individuals, are rapidly advancing. Reparative Pulp and Dentin Dentin, a highly calcified connective tissue, is similar to, but distinct from bone (Linde and Goldberg, 1993). Dental pulp tissue exhibits the potential to regenerate dentin in response to noxious stimuli, such as caries, and it was hypothesized that the stem/progenitor cells present in the dental pulp differentiate into odontoblasts in response to BMP signaling (Nakashima and Akamine, 2005). BMP-driven stem cell therapies have shown considerable promise in dentin–pulp complex regeneration (Shi et al., 2005; Murray and GarciaGodoy, 2006). Iohara et al. (2004) discovered that when stimulated by the morphogenetic signal BMP2 porcine pulp cells differentiate into odontoblasts, as indicated by the expression of dentin sialophosphoprotein (DSPP) and enamelysin/matrix metalloproteinase 20 (MMP20). The autogenous transplantation of the BMP2-treated pulp cells into an amputated pulp cavity resulted in reparative dentin formation in dogs. Zhang et al. (2005) isolated dental pulp cells from maxillary incisors of rats and found that these cells, cultured on either titanium or TCP scaffolds, differentiated into odontoblast-like cells and produced calcified nodules similar to dentin. Furthermore, Gronthos and collaborators (2000) have isolated highly proliferative progenitor cells from adult human dental pulp, which produced densely calcified nodules in culture. After in vivo transplantation of cultured cells into immune compromised mice, a dentin/pulp-like complex was generated, again demonstrating the potential use of this approach for successful dental tissue regeneration. Periodontal Tissue Regeneration The wide prevalence of periodontal disease, the limited regenerative capability exhibited by the PDL, and the critical role of the PDL in maintaining tooth health and function have made periodontal tissue engineering an extremely active area of research. As mentioned above, the PDL complex contains putative stem cells that can commit to a variety of cell fates, such as cementum, ligament, and bone. Periodontal tissue engineering offers a powerful approach to supplement existing treatment regimens for periodontal disease, as recently reviewed (Taba et al., 2005; Bartold et al., 2006; Ramseier et al., 2006). A considerable amount of research has explored periodontal regeneration capability of progenitor cells isolated from the PDL of humans (Grzesik et al., 2000; Seo et al., 2004; Shi et al., 2005), bovine (Saito et al., 2005), and mice (Jin et al., 2003; Zhao et al., 2004). According to the findings of Grzesik and colleagues (2000), human cementum-derived cells (HCDCs), expanded in vitro, formed a mineralized matrix when attached to a ceramic carrier and transplanted subcutaneously into immunodeficient mice. The mineralized matrix exhibited features identical to cementum, with typical organized bundles of collagen fibers (Sodek et al. 1977). Saito and coworkers (2005) transplanted bovine cementoblast progenitor cells subcutaneously into nude mice on an HA/TCP scaffold, also resulting in the formation of cementum-like tissue. In addition, Zhao et al. (2004) reported that rat cementoblasts, delivered via PLGA polymer sponges, have a marked ability to promote periodontal regeneration by inducing mineralization and PDL formation in periodontal wounds. Also in this study, the demonstrated healing of alveolar bone defects was found to be limited only by the size of the transplanted PLGA carrier.

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Interestingly, Kawaguchi et al. (2004) used autologous bone marrow stem cells (BMSCs), in combination with atelocollagen (2% type I collagen), to regenerate periodontal tissues in beagle dogs. The repaired experimental Class III defects consisted of cementum, PDL, and alveolar bone, suggesting that autotransplantation of BMSCs, embedded in the appropriate environment niche, is a novel option for the regeneration of complex tissues, including periodontium. Growth factors or morphogens modulate the cellular activities of, and induce cell differentiation and extracellular matrix production in, developing tissues. The effects of several growth factors, including BMPs (Ripamonti and Reddi, 1997; Giannobile and Somerman, 2003; Wikesjo et al., 2004; Xu et al., 2004; Dunn et al., 2005; Miranda et al., 2005), insulin-like growth factor-I (IGF-I) (Saygin et al., 2000), PDGF (Giannobile et al., 2001; Jin et al., 2004; Nevins et al., 2005; Sarment et al., 2006), and transforming growth factor-β (TGF-β) (Tatakis et al., 2000), on periodontal regeneration of various animal models, including baboons (Miranda et al., 2005), dogs (Tatakis et al., 2000; Wikesjo et al., 2004), rats (Jin et al., 2004; Dunn et al., 2005), and humans (Xu et al., 2004; Nevins et al., 2005; Sarment et al., 2006), have been evaluated. For instance, Saygin et al. (2000) reported that several growth factors, including IGF-I, PDGF-BB, and TGF-β, influence the mitogenetic potential, phenotypic gene expression profile, and biomineralization potential of cementoblasts. Even though preclinical and early clinical data for these growth factors appear promising in stimulating periodontal regeneration, they are not sufficient for definitive conclusions at this time (Ramseier et al., 2006). Furthermore, the methods to deliver growth factors, and the kinetics of release must also be considered to optimize the dosage and exposure time to the cells, and to eventually make periodontal tissue engineering a widely practiced, clinically available therapy. Based on Melcher’s (1976) proposition that cells that repopulate the periodontal wound would determine the type of new attachment onto the root surface, a surgical procedure called guided tissue regeneration (GTR) has demonstrated tremendous potential for periodontal tissue regeneration, by using a barrier membrane to selectively encourage progenitor cell populations to remain at the wound site (Quinones et al., 1996; Laurell and Gottlow, 1998). For example, Aukhil et al. (1986) reported the use of a cell-occlusive barrier to prevent gingival epithelium and connective tissue from growing into the periodontal space and to provide a favorable niche to promote maximal functional PDL regeneration in beagle dogs. More recently, Wikesjo et al. (2003) have shown that the combined use of BMP-2 with a physical barrier, such as a bioabsorbable, space-providing polymer membrane, significantly enhanced periodontal bone and soft tissue regeneration in dogs. However, clinical results using this method to achieve complete restoration of periodontal defects are often unpredictable and vary significantly, possibly due to variations in defect morphology among individuals. Whole-Tooth Tissue Engineering Tooth loss due to periodontal disease, dental caries, trauma, or a variety of genetic disorders continue to adversely affect most individuals at some time in their lives. Recent achievements in tooth tissue engineering promise to provide viable alternatives to currently available human tooth replacement therapies. Young et al. (2002) and Duailibi et al. (2004) have demonstrated the potential to bioengineer complex tooth structures from pig or rat tooth bud tissues, respectively. Using a tissue engineering approach, cells from tooth buds of either 6-month-old porcine third molar or 4-day postnatal (dpn) rats were seeded onto biodegradable polymers. After a period of growth in the omenta of adult rat hosts, a conducive environment to promote vascularization and tissue growth, recognizable tooth structures formed that contained dentin, odontoblasts, a well-defined pulp chamber, putative Hertwig’s root sheath epithelia, putative cementoblasts, and a morphologically correct enamel organ containing fully formed enamel. Both of these studies demonstrated that continuous,

Dental Tissue Engineering

de novo tooth development had occured in the implants, suggesting the potential utility of this method for the regeneration of mammalian dental tissues. As mentioned above, tooth development originates from interactions between dental epithelial and mesenchymal cells, with the epithelium providing the instructive signals for tooth initiation and shape determination. Based on these properties, Ohazama et al. (2004) examined the odontogenic potential of embryonic stem cells, neural stem cells, or adult bone marrow-derived cells, combined with oral epithelium. The oral epithelium was found to stimulate odontogenic responses in each of these three non-dental mesenchymal cell types. Transfer of the tissue recombinants into adult mice renal capsules resulted in the formation of tooth structures and associated bone. In addition, transfer of embryonic tooth primordia into the adult jaw led to the development of tooth structures, indicating that an embryonic primordium can develop in an adult environment. The ability for heterogeneous postnatal and adult cell populations obtained from rodents to form bone and teeth in tissue-engineered rudiments may have important implications for the further development of these procedures on humans. The ability to bioengineer whole teeth inspires additional important therapeutic strategies. For instance, a logical therapeutic approach for treatment of edentulism and accompanying alveolar bone loss would be to combine whole-tooth bioengineering with bone tissue engineering approaches (Young et al., 2005; Yelick and Vacanti, 2006). The combined usage of these methods would result in the generation of both jaw and tooth root structures, to provide teeth and supporting periodontal tissues for individuals born without, or to replace structures lost to disease or injury. The widespread availability of such coordinated therapeutic treatments would provide potentially powerful tools for dental tissue therapeutic strategies and dramatically alter the current landscape of treatments for a variety of craniofacial anomalies.

CONCLUSIONS Dental tissue engineering is a newly emerging field, with just over a decade of creation and development. Although still largely experimental in nature, recent progress in this field promises that tissue engineering approaches will be used to bioengineer replacement teeth in the foreseeable future. The successful development of methodologies for dental tissue repair/regeneration from autologous adult tissues will change the face of modern dentistry and reduce the need for synthetic materials in clinical dental practice. The pressing challenge confronting dental tissue engineering is how to perfect the current techniques, such that bioengineered dental tissues or whole teeth are integrated physically and functionally with preexisting dental tissues (Yelick and Vacanti, 2006). Ideally, the bioengineered teeth would be fabricated to occlude properly with opposing and adjacent teeth and be anchored to the underlying alveolar bone via the PDL to transmit mechanical signals when necessary. Biological teeth would also exhibit proper proprioception, facilitating the life of the substitute itself, as well as surrounding teeth. To accomplish these goals, some important hurdles need to be overcome, including: the identification and maintenance of multipotential DSCs in vitro; the establishment of conditions that foster DSC growth and differentiation; and the development of suitable scaffolds and inductive factors that help DSC-scaffold constructs integrate into the surrounding environment for the reconstruction of dental tissues. With parallel and synergistic advancements in stem cell biology and material science, it is possible to envision that the day may soon come when a patient can visit his or her dentist to obtain a biological, living tooth, grown from the patient’s own cells. ACKNOWLEDGMENTS The authors would like to thank Dan McCloskey and Susan Orlando for library science expertise.

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Saito, M., Handa, K., Kiyono, T., Hattori, S., Yokoi, T., Tsubakimoto, T., Harada, H., Noguchi, T., Toyoda, M., Sato, S. and Teranaka, T. (2005). Immortalization of cementoblast progenitor cells with Bmi-1 and TERT. J. Bone Miner. Res. 20: 50–57. Sarment, D.P., Cooke, J.W., Miller, S.E., Jin, Q., McGuire, M.K., Kao, R.T., McClain, P.K., McAllister, B.S., Lynch, S.E. and Giannobile, W.V. (2006). Effect of rhPDGF-BB on bone turnover during periodontal repair. J. Clin. Periodontol. 33: 135–140. Saygin, N.E., Giannobile, W.V. and Somerman, M.J. (2000). Molecular and cell biology of cementum. Periodontol. 2000 24: 73–98. Seo, B.M., Miura, M., Gronthos, S., Bartold, P.M., Batouli, S., Brahim, J., Young, M., Robey, P.G., Wang, C.Y. and Shi, S. (2004). Investigation of multipotent postnatal stem cells from human periodontal ligament. Lancet 364: 149–155. Shi, S., Bartold, P.M., Miura, M., Seo, B.M., Robey, P.G. and Gronthos, S. (2005) The efficacy of mesenchymal stem cells to regenerate and repair dental structures. Orthod. Craniofac. Res. 8: 191–199. Snead, M.L., Zeichner-David, M., Chandra, T., Robson, K.J., Woo, S.L. and Slavkin, H.C. (1983). Construction and identification of mouse amelogenin cDNA clones. Proc. Natl. Acad. Sci. USA 80: 7254–7258. Sodek, J., Brunette, D.M., Feng, J., Heersche, J.N., Limeback, H.F., Melcher, A.H. and Ng, B. (1977). Collagen synthesis is a major component of protein synthesis in the periodontal ligament in various species. Arch. Oral Biol. 22: 647–653. Spradling, A., Drummond-Barbosa, D. and Kai, T. (2001). Stem cells find their niche. Nature 414: 98–104. Sumita, Y., Honda, M.J., Ohara, T., Tsuchiya, S., Sagara, H., Kagami, H. and Ueda, M. (2006). Performance of collagen sponge as a 3-D scaffold for tooth-tissue engineering. Biomaterials 27: 3238–3248. Taba Jr., M., Jin, Q., Sugai, J.V. and Giannobile, W.V. (2005). Current concepts in periodontal bioengineering. Orthod. Craniofac. Res. 8: 292–302. Tatakis, D.N., Wikesjo, U.M., Razi, S.S., Sigurdsson, T.J., Lee, M.B., Nguyen, T., Ongpipattanakul, B. and Hardwick, R. (2000). Periodontal repair in dogs: effect of transforming growth factor-beta 1 on alveolar bone and cementum regeneration. J. Clin. Periodontol. 27: 698–704. Theiss, F., Apelt, D., Brand, B., Kutter, A., Zlinszky, K., Bohner, M., Matter, S., Frei, C., Auer, J.A. and von, R.B. (2005). Biocompatibility and resorption of a brushite calcium phosphate cement. Biomaterials 26: 4383–4394. Thesleff, I. and Sharpe, P. (1997). Signalling networks regulating dental development. Mech. Dev. 67: 111–123. Thesleff, I., Keranen, S. and Jernvall, J. (2001). Enamel knots as signaling centers linking tooth morphogenesis and odontoblast differentiation. Adv. Dent. Res. 15: 14–18. Wikesjo, U.M., Lim, W.H., Thomson, R.C., Cook, A.D., Wozney, J.M. and Hardwick, W.R. (2003). Periodontal repair in dogs: evaluation of a bioabsorbable space-providing macroporous membrane with recombinant human bone morphogenetic protein-2. J. Periodontol. 74: 635–647. Wikesjo, U.M., Sorensen, R.G., Kinoshita, A., Jian, L. and Wozney, J.M. (2004). Periodontal repair in dogs: effect of recombinant human bone morphogenetic protein-12 (rhBMP-12) on regeneration of alveolar bone and periodontal attachment. J. Clin. Periodontol. 31: 662–670. Xu, W.P., Shiba, H., Mizuno, N., Uchida, Y., Mouri, Y., Kawaguchi, H. and Kurihara, H. (2004). Effect of bone morphogenetic proteins-4, -5 and -6 on DNA synthesis and expression of bone-related proteins in cultured human periodontal ligament cells. Cell Biol. Int. 28: 675–682. Yelick, P.C. and Vacanti, J.P. (2004). Handbook of Stem Cells, Vol. 2. Academic Press, pp. 279–292. Yelick, P.C. and Vacanti, J.P. (2006). Bioengineered teeth from tooth bud cells. Dent. Clin. N. Am. 50: 191–203. Yen, A.H. and Sharpe, P.T. (2006). Regeneration of teeth using stem cell-based tissue engineering. Expert Opin. Biol. Ther. 6: 9–16. Young, C.S., Terada, S., Vacanti, J.P., Honda, M., Bartlett, J.D. and Yelick, P.C. (2002). Tissue engineering of complex tooth structures on biodegradable polymer scaffolds. J. Dent. Res. 81: 695–700. Young, C.S., Abukawa, H., Asrican, R., Ravens, M., Troulis, M.J., Kaban, L.B., Vacanti, J.P. and Yelick, P.C. (2005). Tissueengineered hybrid tooth and bone. Tissue Eng. 11: 1599–1610.

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76 Innovative Regenerative Medicine Approaches to Skin Cell-Based Therapy for Patients with Burn Injuries Jörg C. Gerlach, Steven E. Wolf, Christa Johnen, and Bernd Hartmann

INTRODUCTION The clinical need for improved therapy of burn victims is evident; mortality is still high (Caldwell et al., 1996; Raff et al., 1996; Germann et al., 1997). Regenerative medicine, which has achieved initial success in the utilization of human cells for tissue repair and regeneration, represents a great potential for the treatment of patients with severe burns. Technologies are developing rapidly, with the ultimate goal to deliver advanced skin cell-based therapies as safely and efficiently as possible. Developments in the area of technical devices, cell biological methods, and methods in medical practice have to be considered. Research in this field includes surgical techniques, skin cell procurement, cell culture, stem cell biology, cell application, wound healing support, and wound dressings. This requires an interdisciplinary cooperation of cell biologists, surgeons, and bioengineers. Our chapter attempts to give an overview of the application of skin cells after burn injuries and focuses on utilizing autologous skin cells from a healthy area of the patient’s skin. Our topics are current commercial products, problems in the field, skin cell isolation and culture, skin progenitor cells, cell spray transplantation methods, planning clinical studies, case reports, and regulatory issues. CONVENTIONAL THERAPY AND CURRENT COMMERCIAL PRODUCTS Wound healing in victims suffering from full-thickness III° burns is a challenge. The therapy of choice, surgical skin auto grafting, is limited by the availability of healthy skin cell area and donor site. The standard treatment is the split-skin transplantation method, whereby the use of autologous split-skin (Stark and Kaiser, 1994) is superior to allogenic (Stark and Kaiser, 1994) split-skin application, as rejection eventually occurs. If autologous split skin is not sufficiently available, however, allogeneic skin is indicated to support the wound until an autograft is available. During the past two decades, various commercial products for wound treatment have been applied in clinical studies, whereby one can differentiate between cell-free products and products applied with cell cultures. Integra Life Sciences developed Integra™, which consists of a collagen layer in a matrix with a silicon overlay is a widely accepted product. Indications are described as deep II° and III° burns. Application is performed

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after excising necrotic tissue to temporarily cover the wound while the dermal part engrafts. Tissue engineering aspects were published (Michaeli et al., 1990; Stern et al., 1990; Grzesiak et al., 1997) and initial clinical studies described (Stern et al., 1990; Kinner et al., 1992; Cameron, 1997; Helvig, 1997; King et al., 1997; Lorenz et al., 1997; Besner et al., 1998; Clayton et al., 1998; Pandya et al., 1998; Senior, 1999). This product is useful clinically for temporary coverage of excised wounds when autograft is not available. Furthermore, it has the potential to act as a dermal equivalent that could presumably decrease scar contracture and improve functional and cosmetic outcomes after grafting. Mylan Lab introduced Biobrane™, which is a nylon film with a silicone layer that is coated with a porcine xenogeneic type I collagen. Indications include temporary coverage of superficial burns. Clinical studies were described for the treatment of burns (Bradley et al., 1995; Erdmann et al., 1996; Jones, 1998; Ou et al., 1998; Still et al., 1998), wound coverage of transplanted areas (Kirwan,1995; Wang et al., 1996; Vander et al., 1997; Levy and Salomon, 1998), and applications in plastic surgery (Sakai et al., 1996). This product is used for the temporary coverage of superficial wounds or donor sites, but cannot be construed as a dermal equivalent. If used in full-thickness wounds, the wound must be eventually covered with autograft. LifeCell Corporation developed Alloderm™, which is a cell-free human allogeneic dermal substitute with basement membrane preservation. Indications include dermal substitution for skin defects. Clinical studies are described for deep burns (Wainwright, 1995; Lattari et al., 1997) and plastic tissue replacement (Kridel et al., 1998; Tobin and Karas, 1998). It has the potential benefit of replacing the dermis, and thus inhibiting wound contracture after full wound closure. However, this product is only a dermal equivalent without epidermis; therefore, it cannot be used as an artificial skin. It must be used as a component with an epidermal graft. Initial commercial products that utilize allogeneic cell cultures for the generation of the products were clinically introduced several years ago. Advanced Tissue Sciences developed Trans Cyte™/Dermagraft-TC™. Smith & Nephew introduced this into clinical studies. Trans Cyte was developed for acute burns. Dermagraft was developed for chronic skin wounds, including diabetic ulcers. These dressings have a “dermal” layer of allogenic fibroblasts on polyglycol acid and polyglactin tissue. Clinical studies were published for the treatment of diabetic ulcers (Gentzkow et al., 1996; Edmonds et al., 1997; Naughton et al., 1997; Grey et al., 1998; Mansbridge et al., 1998), and burns (Hansbrough et al., 1992, 1997; Economou et al., 1995; Hansbrough, 1997; Herndon, 1997; Purdue et al., 1997; Spielvogel, 1997). The reimbursement situation was described (Parente, 1997; Purdue, 1997), and tissue engineering aspects were discussed (Rennekam et al., 1996; Advanced Tissue Sciences and Smith & Nephew, 1997; Sacks et al., 1997; Jiang et al., 1998). These products, like Biobrane listed above, can only be used for temporary coverage, but the inclusion of live fibroblasts is thought to accelerate wound healing by the production of growth factors. Organogenesis developed Apligraft™, which was introduced into clinical studies by Novartis. It is characterized as a double-layered composite graft of bovine collagen with allogenic fibroblasts and allogenic keratinocytes. It was developed for venous and arterial diabetic ulcer treatment. Tissue engineering aspects (Eaglstein and Falanga, 1997, 1998; Trent and Kirsner, 1998) and studies for the treatment of venous ulcers were described (Alvarez et al., 1998; Fahey, 1998). Although these products have been shown to improve clinical results in specific clinical situations, a generally accepted wound treatment product for severe burns is not available. Some have used Integra successfully even in massive burns, but this product has probably not been tested side by side with autograft for utility in wound closure or for the prevention of long-term contractures. The same can be said for AlloDerm in terms of utility of the product. However, the purported benefits of accelerated wound closure and decreased scarring have not been verified in clinical trials, which then continue to promote further research efforts toward innovative therapies in the field of skin cell-based therapies in regenerative medicine (Leigh and Watt, 1994; Boyce, 2001).

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Table 76.1 Skin cell isolation methods Keywords

Author

Year

Dissocation dermis/ epidermis dispase Collagenase-based isolation

Stenn et al. Green Sugihara et al. Dunnwald et al. Sugihara et al. Johnen et al.

1989 1991 2001 2001 2001 2006

Trypsin-based isolation

Table 76.2 Methods for skin cell in vitro culture Keywords

Author

Year

Standard keratinocyte culture methods

Rheinwald and Green Green Boyce and Ham Andreassi O’Connor et al. Gallico et al. Tenchini et al. De Luca and Cancedda Boyce and Ham Tenchini et al. Johnen et al Sebök et al. Kaiser et al. Harris et al. Horch et al. Bannasch et al. Hoeller et al. Johnen et al. Hohlfield et al.

1975a, b; 1977 1991 1983 1992 1981 1984 1992 1992 1983 1992 2006 1990 1994 1998 1998 2000 2001 2006 2005

Keratinocyte culture as sheets with serum

Keratinocytes as monolayer serum free culture

Keratinocytes as suspension

Fibroblast culture Human derived fetal skin cells

Methods for the isolation and proliferation of autologous keratinocytes were established around 30 years ago; the isolation of fibroblasts in parallel followed. A literature overview is given in Tables 76.1 and 76.2. Studies on culture media, critical for the expansion of autologous skin cells, are summarized in Table 76.3. Wound covering procedures through the application of keratinocytes, multi-layered epidermal carrier, or fibrin-glue immobilized transplants are described. Genzyme Biosurgery introduced Epicel™, which is an autologous keratinocyte expansion service under industrial Good Manufacturing Practice (GMP) conditions. Cultured epidermal autografts are delivered in the form of sheets for patients with severe burns. There have been many reports of its use (Williamson et al., 1995; Munster,1996; Carsin et al., 2000) mostly in those with burns over 60% of the total body surface area (TBSA). One study compared clinical outcomes of patients with greater than 90% TBSA burns treated with either cultured epithelial autografts or standard skin grafting methods, and found no differences in mortality but increased costs with Epicel mostly related to increased hospital stay and increased numbers of reconstructive procedures. However, long-term cosmetic outcomes were superior in the Epicel group (Barret et al., 2000) giving the promise that such methods may actually give the best long-term outcomes if they could be refined.

Regenerative Medicine for Burn Injury

Table 76.3 Culture media for in vitro culture of skin cells Keywords

Author

Year

Standard culture medium

Rheinwald and Green Boyce and Ham Paini et al. Huang et al. Wang et al. González-Castro et al. Finch et al. Hammar et al. Girolomoni et al. Tavakkol et al. Gibbs et al. Marchese et al. Swope et al.

1975, 1977 1983 1997 2006 1995 1997 1989 1990 1993 1999 2000 2001 2001

Culture in conditioned medium Medium with pituitary extract Medium with growth factors

PROBLEMS IN THE FIELD The speed of wound healing in extensively burned skin represents one of the major factors contributing to the patient’s survival. Patients with larger areas of burns often die during the critical time period of 4 weeks after trauma due to delays in wound closure. To cover wounds, semi-synthetic biomaterials, for example bovine collagen engineered with aminopolysaccharides, such as chondroitin sulfate and chitosan have been used, either alone or in combination with a skin graft (Bell et al., 1981a, b; Burke et al., 1981; Bell et al., 1983; Murphy et al., 1990; Yannas and Burke, 1990; Yannas et al., 1990). At the same time, many groups focus on using the above-mentioned in vitro proliferation of skin keratinocytes to provide an adequate cell number for autologous cell transplantation. Cultured keratinocyte grafting was established as a treatment option for severe burn injuries and in the clinical management of chronic venous ulcers (Srivastava et al., 1990; Phillips and Pachas, 1994; Sabolinski et al., 1996). The cells are taken from the patient’s healthy skin areas and expanded as 3–5 cell layers to cell sheets in Petri dishes, then transferred as confluent and stratified sheets. The functional and cosmetic results, however, require further improvement and innovative methods are of interest. Disadvantages of keratinocyte sheets include long in vitro expansion times of 2–4 weeks due to high passages to reach the required cell number, and the differentiation of basal keratinocytes with increasing culture time. Confluent growing cells in sheets in comparison to single cell cultures are divisionally less active, start proliferation later after transfer, and show a reduced migration activity. These available cell-based methods are not yet advanced enough as a routine treatment to help patients in more severe cases of burn disease. Problems include serous fluid blisters between the transplanted tissue sheets and the wound, and the subsequent loss of cell areas, non-optimal cell nutrition in the early phase after cell transfer, water and electrolyte derangements in the wound, toxin accumulation, pH derangements, and infections. In fact, it is thought that the eventual amount of skin that remains after grafting is approximately 25% (Barret et al., 2000). Consequently, the biomatrix for the transplanted cells in the wound is non-physiological, and the cell sheet grafting and cell proliferation are not optimal (Figure 76.1). Furthermore, the transplanted keratinocytes only produce a very thin neo-skin with low protective quality. Therefore, further improvements are of interest. A clinical problem of wound healing is scar tissue formation. In connective tissue, fibroblast remodeling and contraction of granulation tissue of wounds lead to scar tissue formation (Figure 76.2). This imperfect

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No blister formation Larger treated area

Figure 76.1 Using sprayed cells result in the need of fewer cells while in a larger treatment surface can be enabled for therapy. Blister formation is avoided. Reducing the cell number speeds up application time, reduces in vitro differentiation and therefore better preserves basal keratinocyte progenitor cells.

(a)

(b)

Figure 76.2 Clinical example for large, deep facial burns. (a) Hypertrophic facial scaring after deep IIb° dermal burn. (b) Hypertrophic scaring on the neck after deep IIb° dermal burn.

repair is characterized by disorganized collagen, which does not completely recapitulate the natural structure and function of the tissue (Yamaguchi and Yoshikawa, 2001; Hollander et al., 2003). Excess collagen deposition and altered contraction and remodeling lead to aberrant scarring outcomes, such as hypertrophic scars and keloids, in the dermis. The degree of scarring has been linked to early stages of wound healing in which inflammatory mediators play a crucial role in guiding fibroblast activity in the wound bed. In addition, the new scar tissue contracts for 1–2 years after wound closure, producing functional abnormalities when the scar extends over joints or facial tissues. Most burn scar reconstruction occurs for these reasons, making the inhibition of scar contracture an even bigger clinical issue.

CELL APPLICATION INTO THE WOUND BY SPRAYING Distributing single cells over a skin wound area via an aerosol is of interest to maximize the transfer of viable cells onto the wound surface. Currently, cells are laid down in a sheet layer; however, it is conceivable that delivering cells by spray has significant advantages, including distributing the same number of cells over a larger space, and thus decreasing the necessary time to gain enough cells for wound closure. Table 76.4 summarizes literature describing methods of spraying cells into a skin wound.

Regenerative Medicine for Burn Injury

Table 76.4 Methods of spraying cells into a skin wound Keywords

Author

Year

With fibrin-glue

Kaiser et al. Bannasch et al. Currie et al. Navarro et al. Wood Wood Currie et al.

1994 2000 2003 2000 2001 2003 2003

Without fibrin-glue

The simplest way to distribute liquid medium into a surrounding gaseous phase without generating a jet is through dripping down from a pipe. Additional force and utilization of a nozzle generate sputtering into a discontinuous stream. When scattering, a liquid stream executes wave-like oscillations that cause its breakdown into numerous droplets in the spray. When nebulizing into a spray, the fluid distribution occurs through the nozzle/nebulizer in a stream at high speed into regular individual drops of comparable size. Scattering, spraying, or atomization can occur by generating a liquid stream and applying pressure through a pipe and the subsequent breakdown of the stream in or behind a nozzle/nebulizer. Here, the diffused fluid immerges into a latent gaseous phase. Alternatively, a dual substance spray head with a fluid flow and a gas flow can be applied, in which the fluid in the nozzle is induced into the spray head at a 90o angle to the speed direction of the air stream. Depending on the kind of air compression used, different types of flow occur and the drop formation occurs inside or directly outside the nozzle. Disadvantages of keratinocyte-sheet transplantation can be addressed by using a sprayed suspension of subconfluently cultivated keratinocytes. Reduced hypertrophic scaring was observed in pre-clinical studies (Navarro et al., 2000) and clinical applications (Dietch et al., 1983; Wood, 2003). Spraying cells has several advantages over sheet application. In comparison to cell sheet transplantation, blister formation can be avoided by using sprayed cells. The use of a cell spray method for cell distribution enables a larger wound treatment area (Figure 76.3). Thus, using sprayed cells results in the need for fewer cells while a larger treatment surface can be enabled for therapy. Reducing the cell number during cell expansion speeds up the application time. When applying primary cell suspensions that are isolated from autologous skin and immediately applied onto the patient, natural pigmentation of the skin area is possible because primary isolated skin cells contain all cell types occurring in the dermis (Navarro et al., 2000; Stoner and Wood, 2000). Also, it is conceivable that the proper phenotype of keratinocytes is more available because of decreased passaging of cells which will favor the non-differentiated proliferative type over the differentiating phenotype that would presumably have greater tendency to adhere, and thus provide the basis for better and more secure keratinocyte expansion.

PATIENT GROUPS AND SELECTION CRITERIA FOR CLINICAL STUDIES IN THE FIELD OF SKIN CELL-BASED THERAPY DEVELOPMENT IN BURN DISEASE PATIENTS Planning clinical studies for the development of regenerative therapy methods for severe burns is problematic. The question arises of how to determine selection criteria for burn patients when testing innovative skin cell therapies, including an ethical approach. Based on the frequently occurring various degrees of burns on different areas of the body, it is possible to apply different therapeutic methods on different parts of the body independently from one another. Clinically, we were interested in testing the application of keratinocytes via

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(a)

(b)

(c)

Figure 76.3 Keratinocyte spray transplantation. Case report, male, age 34. (a) Keratinocyte spray transplantation after 14 days post burn injury with not healed IIb° facial wound (14 mio cells in 10 ml suspension sprayed after 6 days expansion time). Also, a split-skin transplant was performed ventral on the left neck. (b) Post-operative result 3 months after trauma. A split-skin transplant was performed ventral on the left neck. (c) Post-operative result 3 months after trauma.

a spray transplantation method to open wounds (Wood, 2003; Wood and Allen, 2003). First, we had to define selection criteria/patient groups. In general, category IIa° superficial burn injuries in the face heal within 8–14 days with good results. Category IIb°, deep dermal burns on the neck and face present an indication for excision followed by autologous split-skin transplantation, usually as a sheet-graft. For III° burns that exhibit complete loss of dermis, skin autograft as a sheet with a thicker dermal portion is the preferred choice of therapy. Functionally, the results are usually satisfactory, but esthetically disappointing. However, it is difficult to recommend therapy methods for medium-to-deep dermal burn injuries of category IIb°. After split-skin transplantation, in cases of “overgrafting,” the cosmetic results are often unsatisfactory, leaving the transplanted skin with clearly visible color and texture variations. When leaving these wounds to spontaneous healing, they almost always result in esthetic and functional hypertrophic scaring (Figure 76.4). Therefore, an increasingly conservative wound treatment approach has prevailed for II° dermal facial burns, first awaiting the results of spontaneous healing. Thereafter, only the IIb° areas that have not adequately healed within 7–12 days are surgically treated. However, due to above-described disadvantages,

Regenerative Medicine for Burn Injury

(a)

(b)

Figure 76.4 Keratinocyte spray transplantation. Case report, female, age 47. (a) Keratinocyte spray transplantation after 14 days post burn injury with not healed IIb° wound and split-skin transplant ventral left neck (2 mio cells sprayed after 10 days expansion time). (b) Post-operative result, 9 months after trauma.

making the decision to treat small areas with split-skin transplantation remains difficult. From our perspective, the clinical situation for IIb° burn injuries on the neck and face that have not healed within a prescribed time, appears to be the most suitable for study with innovative cell-based therapy methods. Vitally critical situations are precluded, and the conventional therapy method is not satisfactory; however, the patient can be given hope for possibly improved results. The search for applicable areas and patient groups for the study of innovative therapy methods is also dependent on the technology. The above-described method of spray transplantation using keratinocytes in suspension allows combination of spraying skin cells with the standard therapy of split-skin transplantation, because the isolated and expanded skin cells in the cell suspension can be sprayed between the mesh of the transplant. In this case, such a combination method allows a gradual clinical introduction of a new technology. Patients with III° burns over 10–30% of the body surface area are another potential population in which to test these technologies because using the standard therapy method with autologous mesh-graft split-skin transplantation combined with an additional spraying of cells between the mesh may allow for more rapid wound closure while keeping within the current standard of care which would be skin grafting. Even though this combination method can only provide initial evidence of the effect of the cell spray method, it allows the introduction of the technology under ethical desired conditions prior to introducing this technology into clinical studies as an alternative to split-skin transplantation. In our clinical proceedings, the keratinocyte spray transplantation was applied in both clinical situations and patient groups.

INITIAL OWN CASE REPORTS SHOW THAT THE SPRAY TRANSPLANTATION METHOD WITH IN VITRO EXPANDED KERATINOCYTES YIELDS AN IMPROVED OUTCOME OF DEEP DERMAL BURN INJURIES ON THE NECK AND FACE For 2 years we have been applying skin cells by spray transplantation in our Berlin burn center similar to that mentioned above (Johnen et al., 2006; Hartmann et al., 2007). This work focuses on autologous expanded keratinocytes on patients in the patient group with III° medium-and-deep dermal burn injuries on the neck and face. These patients exhibited III° burns in other areas of the body that required grafting. The decision for cell-based therapy on IIb° on the neck and face was not made until the second treatment week.

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The goal of this therapy is to provide cosmetically satisfactory results, reduce the delay of spontaneous healing, and reduce the possibility of pronounced scarring when applying the otherwise necessary split-skin transplantation method on IIb° on the neck and face. There was no randomization. Initial treatments were collected as case reports of our Berlin center. The applications were in agreement with the local authorities and institutions; the patient or his/her next of kin gave informed consent. All patients that consented to the treatment were included. Initially, the primary therapy included wound cleaning and gentle debridement. Following, the wound was conservatively treated with silicone wound dressings and local antiseptic ointment (Lavasept®-gel, Fresenius, Norderstedt, Germany) and keeping the wound moist. The rationale was to first support the reepithelialization process of superficial wound areas. After approximately 1 week the depth of the injury and the expected course of healing were better assessed, and the indication for skin cell treatment was made. As an earlier step, a small 4  4 cm split-skin biopsy was taken from another part of the body either under general anesthesia or local anesthesia, depending on the type of subsequent surgery to treat another III° area. The biopsy was taken to our laboratory for keratinocyte isolation. The obtained keratinocytes were cultured and expanded as described by Johnen et al. (Johnen et al., 2006). After 10 days, a suspension with expanded autologous patient cells was available for cell spray transplantation. Now, 14 days after injury, superficial dermal facial wounds had healed and deep dermal wounds had not, at which point the final decision for surgery was made. Under general anesthesia all areas that were not adequately healed were cleaned of necrotic tissue and treated with keratinocyte spray transplantation. After the spray transplantation was completed, the treated areas were covered with fine mesh fatty gauze or an adhering cell compatible polylactide membrane (Suprathel®, Asklepios, Germany) followed by a gauze-cotton dressing. The first dressing change occurred after 5 days, and then every 2 days thereafter until the transplanted surface was completely epithelialized and mechanically stabile enough that an open wound treatment with fatty ointment can be applied. The treatment/healing success was examined within the scope of the ambulatory post-surgical check-ups and 1-year post trauma. The Vancouver Scare Scale and the German Donnersmark Scale were used to evaluate the healing success. We compared the quality of the scars and the subjective discomfort of the patients treated with the spray transplantation with those treated with the conventional therapy. The spray transplantation method with in vitro expanded autologous keratinocytes was applied on 16 patients, with 8 having IIb° burns on the neck and 14 with IIb° burns on the face. The range of IIb° affected area on the face was 45–300 cm2 (mean 195 cm2), and on the neck was 31–160 cm2 (mean 58 cm2). Six were female and 10 were male ranging between 18 and 66 years of age (mean 39.5 years). The mean of the burned skin area of the patients was 16.2%; the ABSI score (Tobiasen) was 6.7. The amount of cells in the suspension was between 900,000 and 2,100,000, and the expansion time was between 6 and 9 days (mean 8.2 days). All cases exhibited excellent healing results. After 5–9 days (average 7.3 days) complete epithelialization of the treated wounds was present. The delayed healing process through the formation of granulated tissue in these deep dermal wounds and the subsequent hypertrophic scaring appeared to be stopped when applying this method. We believe that the autologous cells function as immediate wound closure as well as facilitators to accelerate wound healing, which is indicated by the absence of, or significantly reduced disruption of pigmentation that is based on the activation and promotion of self-healing of the dermal cells that remained deep in the dermis, including melanocytes. The spray transplantation of autologous cells presents a treatment option with significantly improved outcome for middle-to-deep neck and facial wounds. This method enables wound healing with minimized hypertrophic scaring and good cosmetic and functional results for wounds that do not exhibit reepithelialization after

Regenerative Medicine for Burn Injury

14 days. The typical results of split-skin transplantation with overgrafting, significant discoloration and textural variations of the grafted skin that can be exceedingly noticeable and disfiguring on the face, are omitted; likewise, hypertrophic scaring after significantly delayed wound healing beyond 14 days’ epithelialization time. Based on the results of the described method, we applied at our Berlin center autologous keratinocyte suspension through spray transplantation, in addition to mesh-graft split-skin transplantation, to patients in another patient group, with complete dermal burns (III°) on the neck and face. The goal of this application was to determine whether the often unsatisfactory cosmetic results in the transplanted areas of the skin with subsequent significant scarring can be improved. Initial treatments are collected as case reports in our center. There is no randomization; the allocation occurs consecutively after the patient’s consent, and all patients that consent are included. III° areas are covered with thin split skin in mesh-technique, and the mesh-graft areas are sprayed with cells. So far, only a few cases can be reported. There were two female patients and four male patients; their age was between 27 and 52 years with a mean of 38.7 years. The III° area treated was between 140 and 580 cm2 with a mean of 392 cm2. The amount of cells in the cell suspension was between 950,000 and 2,500,000; the expansion time was between 6 and 10 days, and the mean was 8.4 days. The assessment of and the criteria for the healing success occur as described above. The initial clinical results of the combined mesh-graft and skin cell transplantation method on deep dermal burns exhibit impressive, favorable results. As also described by the Australian group of Fiona Wood, swift wound closure, reduced scaring, and less visible transplant pattern are apparent after mesh transplantation. From an esthetic perspective, the combination of mesh-graft and spray transplantation of autologous proliferated cells presents as an advantageous treatment option, especially for the neck and face. Hypertrophic scaring can be minimized, and wounds can heal with good cosmetic and functional results. Typical results of mesh-graft transplantation, significant discoloration and textural variations that can often be exceedingly noticeable and disfiguring on the face, are omitted. By distributing the cells over a larger wound area, utilizing cell spray transplantation can reduce the required cell number for therapeutic applications. Wood et al. therefore use cell spraying without in vitro cell expansion (Wood, 2003) with a donor site taken at the same procedure. In a kit devised by these investigators, keratinocytes are separated from the split thickness donor skin and are suspended in a liquid form. These are then sprayed onto the wound area with an approximately 100:1 expansion ratio. Using this cell spray method as well as cultured cells that are sprayed onto the wound, Dr. Wood and her colleagues have shown a decrease in the number of operative procedures and length of hospital stay (Wood et al., 2006). This practice avoids in vitro culture at all and thus in vitro differentiation of the cells. Consequently the basal skin progenitor cells are better preserved for the clinical application.

SKIN PROGENITOR CELLS Several groups focus on addressing disadvantages of using adult skin cells. To address the logistics of adult cell procurement, promoting cell growth (Rheinwald and Green, 1977) during expansion was introduced. To address availability of cells, continuous growing cell clones (Rheinwald and Green, 1975) or the development of continuous growing cell lines from tumor tissue (Rheinwaid and Green, 1975) were discussed. The practical problems of using autologous adult cells also suggest consideration of skin progenitor cells, including basal keratinocytes, and stem cells, as alternative to the use of autologous adult keratinocytes. This is detailed in Tables 76.5 and 76.6. Studying the mechanics of fetal skin is of interest to further develop therapeutic strategies. Fetal wound healing in the dermis (Dostal and Gamelli, 1993) is distinguished by minimal inflammation and contraction

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Table 76.5 Literature sources on skin progenitor cell studies Keywords

Author

Year

Epidermal stem cells

Bickenbach and Chism Watt Fuchs and Raghavan Janes et al. Brouard and Barrandon Webb et al. Morasso and Tomic-Canic Kobayashi et al. Oshima et al. Claudinot et al. Schlabe et al. Li et al. Tani et al. Dunnwald et al. Häkkinen et al. Toma et al. Michel et al. Lyle et al. Levy et al. Pellegrini et al. Trempus et al. Klima et al. Ohyama et al. Watt and Hogan Christiano et al. Tumbar et al. Rizvi and Wong

1998 2002 2002 2002 2003 2004 2005 1993 2001 2005 2007 1998 2000 2001 2002 2005 1996 1998 2000 2001 2003 2005 2006 2000 2004 2004 2005

Hair follicle stem cells

Isolation and cultivation

Marker

Niche

Table 76.6 Use of hair follicle cells in tissue engineering Keywords

Author

Year

Cultivation

Limat and Noser Lenoir et al. Limat et al. Kurata et al. Limat and Hunziker Limat et al. Tausche et al. Navsaria et al. Reynolds and Jahoda Krugluger et al. Randall et al. Jahoda and Reynolds Jahoda Jahoda et al. McElwee et al. Richardson et al. Wu et al.

1986 1988 1991 1994 2002 2003 2003 2004 1994 2005 1996 2001 2003 2003 2003 2005 2005

Clinical use

Reconstruction of hair follicle in vitro Dermal papilla (DP) dermal sheath cells (DSC) tissue

Regenerative Medicine for Burn Injury

leading to complete regeneration of wounds characterized by the organization of quickly deposited collagen that is indistinguishable from non-injured tissue once healing is complete (Nemeth, 1993; Clark, 1996). The altered inflammatory response is thought to be a primary component of this scar-less repair. Another key component of this regenerative wound healing is the intrinsic characteristics of the fetal fibroblasts, which are thought to be the key effectors of scar-less fetal wound healing (Lorenz et al., 1995). Fetal skin-derived cells may even present an interesting cell source for therapy development (Hohlfeld, 2005).

REGULATORY ISSUES Before ending this overview, some regulatory issues are summarized for the United States that could be of interest in practical therapy development. There are several potential “products” in the field; from a regulatory standpoint one has to consider: Devices, for example active skin wound dressings. Drugs, for example skin wound treatment solutions. Biologics, for example expanded autologous skin cells from a GLP/GMP laboratory. Addressing regulatory affairs for the clinical translation of research and development in the United States has to consider the federal Food and Drug Administration (FDA) and specific regulations prior to entering into clinical studies. Developments and clinical studies in the field have to consider three FDA centers with associated and varied regulatory requirements: The Center for Drug Evaluation and Research (CDER): If a development in the field considers a drug, an Investigational New Drug Exemption (IND) may be required prior to clinical studies. Center for Biologics Evaluation and Research (CBER): If a development in the field considers a isolated and expanded cells, these will be considered as biologics. GLP, GMP requirements may apply. A pre-clinical Investigational New Drug Exemption may be required prior to clinical studies. The Center for Devices and Radiological Health (CDRH): If a development in the field considers a device. The main set of regulations that are applicable to medical devices are contained in Title 21 Code of Federal Regulations Parts 800–1200 (21 CFR 800–1200). These controls apply to all medical devices: control of design process, clinical trials, marketing approval, and manufacturing. Some devices may be regulated by the CDRH with input from CBER. The following briefly summarizes some of the FDA processes and requirements that are relevant to products mentioned in this overview: FDA clearance or approval of medical devices: Section 510(k) of the FD&C Act requires a manufacturer who intends to market a medical device to submit a pre-market notification (510(k)) to the Agency at least 90 days prior to introducing the device to the market. If the device is determined to be “not substantially” equivalent (NSE) it must have an approved pre-market approval (PMA) application (or be reclassified into Class I or Class II) before being marketed. The final determination of whether or not a device is substantially equivalent resides with the FDA. There are several routines: PMA: Should a PMA be required, the application must present scientific evidence that the device is safe and effective for its intended use, typically requiring more in-depth supporting data than a 510(k).

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510(k) submission: Must demonstrate that the device to be marketed is as safe and effective as an existing, legally marketed (predicate) device, for example in the FDA’s terminology, that the new device is “substantially equivalent” to the predicate device. GMP Requirement: Production of a device must meet the FDA’s good manufacturing requirements as set forth in the Quality System Regulations (21 CFR 820). Using materials for production coming from an FDA registered manufacturer may allow reference to a corresponding master file, if one exists and permission is obtained from the Master File Holder. Clinical trials: An investigational device exemption (IDE) allows a device to be used in a clinical study for the purpose of collecting safety and effectiveness data required to support a PMA application or a Pre-market Notification (510(k)) submission to the FDA. Clinical studies are most often conducted to support a PMA. Only a small percentage of 510(k)’s require clinical data to support the application. Humanitarian use device (HUD) medical device: A HUD can be developed and marketed by a company in a more streamlined process than a typical new medical device, if it is intended to benefit patients in the treatment of a disease affecting fewer than 4,000 patients per year in the United States. It is required that the patient population is not being addressed by an approved device or therapy. This avenue can be followed if the manufacturer is not making a profit on the HDE; he can only recover R&D cost, manufacturing, and handling. The HUD must demonstrate safety and probable benefit (not efficacy standard for normal products). Compassionate use of a medical device: A compassionate use allows clinicians to use an “unapproved” device on a patient or patients with a life-threatening disease that does not qualify for an existing clinical trial and has no comparable or satisfactory alternative therapy available. Usually the therapy is in a clinical trial and has demonstrated acceptable performance in its intended patient population. For example, a drug that has completed Phase II or Phase III IND testing may qualify, but a drug in Phase I will not qualify because it has not demonstrated safety or any efficacy yet. The device, therefore, must be entered into a clinical trial via an Investigational Device Exemption (IDE) for a device or an IND for a drug. The Committee for Human Research and the FDA will have to approve the application for compassionate use. In well-defined cases of intra-operative cell harvest without cell alteration and immediate application to the same patient no FDA-specific regulations apply (see below); consulting the FDA is required. Autologous cell transplantation may be considered as unregulated in cases where cells are taken from the patient and transplanted with minimal cell manipulation at the point of care, for example in the same operation by the responsible surgeon. Most commercial applications of skin cells or tissues, however, may be considered as biologics. The FDA’s Center for Biologics Evaluation and Research regulates biological products under Sections 351 and 361 of the Public Health Service (PHS) Act and under specific sections of the Federal Food, Drug and Cosmetic Act (FD&C Act). Under these authorities, CBER is responsible for ensuring:

• • • • •

the safety of the US blood supply and the products derived from it; the production and approval of safe and effective childhood vaccines, including any future AIDS vaccines; the oversight of human tissue for transplantation; an adequate and safe supply of allergenic materials and anti-toxins; the safety and efficacy of biological therapeutics.

Licensing of Biologics The PHS Act requires individuals or companies who manufacture biologics for introduction into interstate commerce to hold a license for the products. These licenses are issued by CBER. Licensing of biologic products

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under the PHS Act is very similar to the new drug approval process for human drugs. Following initial laboratory and animal testing, a biological product is studied in clinical trials in humans under an investigational new drug application (IND). If the data generated by the studies demonstrate that the product is safe and effective for its intended use, the data are submitted to CBER as part of a biologics license application (BLA) for review and approval for marketing. Regulation of HCT/Ps Human cells, tissue, and cellular and tissue-based products (HCT/Ps) that are intended for implantation, transplantation, infusion, or transfer into a human recipient are regulated by CBER under 21 CFR Parts 1270 and 1271. Examples of such tissues are bone, skin, corneas, ligaments, tendons, dura mater, heart valves, hematopoietic stem/progenitor cells derived from peripheral and cord blood, oocytes, and semen. Note that vascularized organs, bone marrow, blood, and non-human cells are not considered HCT/Ps, because they are regulated under other laws/agencies. If the HCT/Ps are for homologous use, are only minimally manipulated, and meet the additional criteria listed below in 21 CFR 1271.10, marketing may proceed so long as the requirements of the 21 CFR Parts 1270 and 1271 are met. The FDA has defined “minimal manipulation” (21 CFR 1271.3(f)) as follows: 1. For structural tissue – processing that does not alter the original relevant characteristics of the tissue relat-

ing to the tissue’s utility for reconstruction, repair, or replacement. 2. For cells or non-structural tissues – processing that does not alter the relevant biological characteristics of the cells or tissues. The FDA has defined “homologous use” (21 CFR 1271.3(c)) as follows: 1. The repair, reconstruction, replacement, or supplementation of a recipient’s cells or tissues with an HCT/P

that performs the same basic function or functions in the recipient as in the donor. 21 CFR Parts 1270 (Human Tissue Intended for Transplantation) and 1271 (HCT/Ps): The FDA requires tissue establishments to meet the requirements of 21 CFR 1270 and 1271, which include:

• • • • •

screen and test donors; prepare and follow written procedures for the prevention of the spread of communicable disease; maintain records; register and list their HCT/Ps with FDA; follow current Good Tissue Practices (GTP) for HCT/Ps.

The following criteria in 21 CFR 1271.10 are referenced in determining whether an HCT/P is subject to Part 351 or Part 361 of the PHS Act, as described in the following. 21 CFR 1271.10(a): 1. The HCT/P is minimally manipulated (see examples in Table 76.7). 2. The HCT/P is intended for homologous use only, as reflected by the labeling, advertising, or other indica-

tions of the manufacturer’s objective intent (see examples in Table 76.8). 3. The manufacture of the HCT/P does not involve the combination of the cells or tissues with another arti-

cle, except for water, crystalloids, or a sterilizing, preserving, or storage agent, provided that the addition of water, crystalloids, or the sterilizing, preserving, or storage agent does not raise new clinical safety concerns with respect to the HCT/P (see examples in Table 76.9).

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Table 76.7 HCT/Pss “351” HCT/P

“361” HCT/P

An HCT/P regulated under Section 351 of the PHS Act is defined as a product that does not meet one or more of the criteria in 21 CFR 1271.10 (listed above). “351” HCT/Ps will be subject to premarket review and approval, as appropriate based on the product: – Biologics License Application (BLA), if a biological product. Clinical study – Investigational New Drug (IND) process. – New Drug Application (NDA), if a new drug. Clinical study – Investigational New Drug (IND) Process. – Premarket Approval (PMA) Application, or Premarket Notification (510(k)), if a new device. Clinical study – Investigational Device Exemption (IDE) process. Pre-license approval inspection will be conducted by FDA. Both investigational and licensed “351” HCT/Ps must comply with 21 CFR 1271 (Subparts A, B, C, and D (although if in compliance with cGMPs in 21 CFR 210/211, compliance with GTPs is likely). Investigational “351” HCT/Ps also must comply with: – 21 CFR 312 – IND Application – 21 CFR 210/211 – cGMP (as defined) – 21 CFR 50 – Protection of Human Subjects (Informed Consent) – 21 CFR 56 – Institutional Review Boards Licensed “351” HCT/Ps also must comply with: – 21 CFR 201 – Labeling – 21 CFR 202 – Advertising – 21 CFR 210/211 – cGMP – 21 CFR 600 – Biological Products (includes reporting of adverse experiences and biological deviations) – 21 CFR 601 – Licensing (BLA) – 21 CFR 610 – General Biologics Standards.

An HCT/P is regulated solely under Section 361 of the PHS Act and 21 CFR 1271 if it meets all of the criteria noted above in 21 CFR 1271.10. “361” HCT/Ps are not subject to premarket review. No application to FDA is required. FDA inspections will be conducted to confirm compliance. Compliance with 21 CFR Part 1271 is required. This regulation includes the following subparts: A – General Provisions and Definitions B – Registration and Listing C – Donor Eligibility D – Current GTP (as defined) E – Additional Requirements (Reporting and Labeling) F – Inspection and Enforcement

•

• •

• •

•

•

Table 76.8 Defining manipulations HCT/P category

Minimal manipulation

More than minimal manipulation

Structural tissue

1. Fascia or dermis processed into particulate form. 2. Allogeneic dehydrated and decellularized amniotic membrane intended for wound covering. 3. Cutting, grinding, shaping of bone.

1. Allogeneic, decellularized human arteries, veins, heart valves, or valve conduits intended to replace dysfunctional cardiovascular tissue. (Continued )

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Table 76.8 (Continued) HCT/P category

Minimal manipulation

More than minimal manipulation

Cell

1. CD34+ selection of allogeneic peripheral blood 1. Autologous cultured (expanded hematopoietic stem or progenitor cells. ex vivo) epithelial cells isolated from 2. Density gradient separation to remove a particular skin biopsies and intended to cover type of cell from a mixture of cells. burns. 2. Gene therapy

HCT/P category

Homologous use

Non-homologous use

Structural tissue

1. Demineralized bone matrix used as a bone void filler during orthopedic surgery. 2. Bone recovered from a limb, used as a bone dowel for spinal surgery.

1. Allogeneic veins or arteries intended for use as arterio-venous access (A-V shunts) for hemodialysis. 2. Cartilage tissue used in the bladder for treatment of reflux.

Cell

1. Allogeneic placental/umbilical hematopoietic stem/progenitor cells used for hematopoietic reconstitution. 2. Pancreatic islet cells used for treatment of type 1 diabetes.

1. Autologous bone marrow hematopoietic stem/progenitor cells used for myocardial repair. 2. Nasal mucosal cells used to regenerate nerve tissue.

Table 76.9 Defining combinations “Not combined with”

“Combined with”

1. Lyophilized pericardium and a vial of saline packaged together in a kit, for reconstitution by the physician. 2. Any HCT/P to which a sterilizing (e.g. antibiotic), preserving (e.g. Optisol), or storage (e.g. DMSO) agent is added.

1. Demineralized bone matrix combined with another article to create a paste or putty used to fill bone defects. 2. Tendon allograft combined with a suture for use in ligament reconstruction.

4. Either:

(i) the HCT/P does not have a systemic effect and is not dependent upon the metabolic activity of living cells for its primary function or (ii) the HCT/P has a systemic effect or is dependent upon the metabolic activity of living cells for its primary function, and: (a) is for autologous use; (b) is for allogeneic use in a first-degree or second-degree blood relative; or (c) is for reproductive use. “351” HCT/P or “361” HCT/P HCT/Ps are described as either “351” HCT/P or “361” HCT/P, in reference to the PHS Act sections to which these products must comply, as outlined in Table 76.7. Table 76.8 provides examples to assist in defining and distinguishing “minimal manipulation” with “more than minimal manipulation,” and “homologous” with “non-homologous” for the two categories of HCT/P tissues (structural and cells/non-structural).

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These examples were provided by the FDA in a presentation on February 8, 2006, at the 2nd Annual FDA and the Changing Paradigm for Tissue Regulation in Las Vegas, Nevada (obtained from the FDA’s website). The examples in Table 76.9 illustrate the FDA’s definition of “combined with” or “not combined with” as referenced in 21 CFR 1271.10(a).

DISCUSSION The transition from whole-skin grafts to meshed split-skin grafts quadrupled the potential area for surgical treatment because the basal keratinocytes in the graft are able to regenerate the gaps in the mesh. In vitro expanded keratinocytes could further enlarge the surface area for treatment but leave patients untreated for the critical initial time phase of 2 weeks after injury. The extended in vitro time also results in the loss of basal keratinocyte cell fraction. The enormous regenerative capacity of the basal keratinocytes in a mesh graft can be compared with that of seeded grass leading to a complete new lawn if appropriately nursed, for example with fluid and nutrition supply. Comparably the potential of basal keratinocytes would be of better use, if they would be more thinly distributed as single cell transplant in the wound and be optimally supported in the early phase after “seeding” in the wound (Figure 76.5). Developments in regenerative medicine aim to combine methods synergistically. Innovative skin cellbased therapies should involve considerations according to the therapeutic application phases: cell procurement from donation sites, cell application to wound, and cell support in the wound. Consequently one could consider:

• •

Some groups are already using isolated but not expanded autologous skin cells, containing the basal keratinocyte fraction, to provide a progenitor cell component to the wound. This should result in higher quality of the neo-skin formation. Application of a cell spray method for cell distribution for islet-like distribution of single cells in the wound without in vitro expansion was also introduced. This enables a larger treatment area, protects the basal keratinocytes from differentiation during expansion, reduces the time to begin therapy, and avoids the formation of cell sheet blisters.

(a)

(b)

Figure 76.5 Case report. Male, age 38. (a) Keratinocyte spray transplantation after 14 days post burn injury with not healed IIb° wound (1.7 mio cells sprayed after 9 days expansion time). (b) Post-operative result, 21 months after trauma.

Regenerative Medicine for Burn Injury

• •

The clinical use of an active skin wound dressing to support wound healing, by providing a more physiological wound environment. The use of a perfused membrane in the wound dressing may also be used to continuously supply growth factors to support cell proliferation and cell migration in the wound (Chen et al., 1993; Hinz et al., 1999). Co-use of autologous skin cells, containing the basal keratinocyte fraction with artificial membranes in an active wound healing for wound support.

The combination of using the progenitor cell component, distributing the progenitor cells over a larger surface using cell spray application, and an active artificial wound membrane for initial cell support should enable the transplanted cells to perform “tissue engineering” in the wound. Preserving the basal keratinocytes, reducing blister formation, and providing a more physiologic environment after cell transplantation should improve the wound healing. Figures 76.6 and 76.7 show the concept of using a perfused flat sheet membrane bag in an active wound dressing. The membrane is temporarily placed above the sprayed cells and below the outer wound dressing. Such a membrane-based wound dressing can provide nutrition, oxygenation, pH regulation, electrolyte balance, and detoxification of wound debris. The proposed therapy combination may improve the clinical outcome by reducing the time of wound healing while enabling larger treatment areas, and thus reducing the mortality rate in patients with large surface burns. Basal keratinocytes, skin fibroblasts, single cells

Figure 76.6 Tissue engineering in the wound by a temporary artificial wound flat sheet membrane, manufactured to a perfuseable “bag”: active wound dressing with means to support proliferation, cell migration, and organ restructuring.

Fresh medium Medium drain Membrane bag Medium

Cells

Dressing

Figure 76.7 Temporary artificial flat sheet membrane “bag” placed on a wound and perfused with medium.

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ACKNOWLEDGMENT The information, support, and co-editing of this text by Patsy J. Trisler, J.D., RAC, Regulatory Consultant – Medical Devices ([email protected]) is greatly appreciated.

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Hohlfield, J., Hirt-Burri, N., Gerber, S., Scaletta, C., Hohlfield, P. and Applegate, L.A. (2005). Tissue engineered fetal skin constructs for paediatric burns. Lancet 366: 840–842. Hollander, D.A., Erli. H.J., Theisen, A., et al. (2003). Standardized qualitative evaluation of scar tissue properties in an animal wound healing model. Wound Repair Regen. 11(2): 150–157. Horch, R.E., Bannasch, H., Kopp, J., Andree, C. and Stark, G.B. (1998). Single-cell suspension of cultured human keratinocytes in fibrin-glue reconstitute the epidermis. Cell Transplant. 7: 309–317. Huang, Y.C., Wang, T.W., Sun, J.S. and Lin, F.H. (2006). Investigation of mitomycin-C treated fibroblasts in 3-D collagen gel and conditioned medium for keratinocytes proliferation. Artifi. Organs 30(3): 150–159. Jahoda, C.A.B. (2003). Cell movement in the hair follicle dermis-more than a two-way street? J. Investig. Dermatol. 121: Jahoda, C.A.B. and Reynolds, A.J. (2001). Hair follicle dermal sheath cells: unsung participants in wound healing. Lancet 358: 1445–1448. Jahoda, C.A.B., Whitehouse, J., Reynolds, A.J. and Hole, N. (2003). Hair follicle dermal cells differentiate into adipogenic and osteogenic lineages. Exp. Dermatol. 12: 849–859. Janes, S.M., Lowell, S. and Hutter, C. (2002). Epidermal stem cells. J. Pathol. 197: 479–491. Jiang, W.G., et al. (1998). Enhancement of wound tissue expansion and angiogenesis by matrix-embedded fibroblast (Dermagraft), a role of hepatocyte growth factor/scatter factor. Int. J. Mol. Med. 2(2): 203–210. Johnen, C., Steffen, I., Beichelt, D., Bräutlgam, K., Witascheck, T., Toman, N., Moser, V., Ottomann, C., Hartmann, B. and Gerlach, J.C. (2007). Biodegradable and digestable transfer-membrane-based culture of human fibroblasts and keratinocytes. (Accepted) Burns. Johnen, C., Hartmann, B., Steffen, I., Bräutigam, K., Witascheck, T., Toman, N., Küntscher, M.V. and Gerlach, J.C. (2006). Skin cell isolation and expansion for cell transplantation is limited in patients using tobacco, alcohol, or are exhibiting diabetes mellitus. Burns 32(2): 194–200. Jones, L.M. (1998). The Biobrane stent. J. Burn Care Rehabil. 19(4): 352–353. Kaiser, H.W., Stark, G.B., Kopp, J., Balcerkiewicz, A., Spilker, G. and Kreysel, H.W. (1994). Culture autologous keratinocytes in fibrin glue suspension. Exclusively and combined with STS-allograft (preliminary clinical and histological report of a new technique). Burns 20: 23–29. King, WW., et al. (1997). Evaluation of artificial skin (Integra) in a rodent model. Burns March 23 (Suppl 1): S30–S32. Kinner, M.A., et al. (1992). Skin transplantation. Crit. Care Nurs. Clin. N. Am. 4(2): 173–178 (review). Kirwan, L. (1995). Management of difficult wounds with Biobrane. Conn. Med. 59(9): 523–529. Klima, J., Smetana Jr., K., Motlik J., et al. (2005). Comparative phenotypic characterization of keratinocytes originating from hair follicles. J. Mol. Histol. 36: 89–96. Kobayashi, K., Rochat, A. and Barrandon, Y. (1993). Segregation of keratinocyte colony-forming cells in the bulge of the rat vibrissa. Proc. Natl Acad. Sci. USA 90: 7391–7395. Kridel, R.W., Foda, H. and Lunde, K.C. (1998). Septal perforation repair with acellular human dermal allograft. Arch. Otolaryngol. Head Neck Surg. 124(1): 73–78. Krugluger, W., Rohrbacher, W., Laciak, K., et al. (2005). Reorganization of hair follicles in human skin organ culture induced by cultured human follicle-derived cells. Exp. Dermatol. 14: 580–585. Kurata, S., Itami, S., Terashi, H. and Takayasu, S. (1994). Successful transplantation of cultured human outer root sheath cells as epithelium. Ann. Plast. Surg. 33: 290–294. Lattari, V., Jones, L.M., Varcelotti, J.R., Latenser, B.A., Sherman, H.F. and Barrette, R.R. (1997). The use of a permanent dermal allograft in full-thickness burns of the hand and foot: a report of three cases. J. Burn Care Rehabil. 18(2): 147–155. Leigh, I.M. and Watt, F.M. (1994). The culture of human epidermal keratinocytes. The Keratinocyte Handbook. Cambridge, MA: University Press Cambridge, pp. 43–51. Lenoir, M.C., Bernard, B.A., Pautrat, G., Darmon, M. and Shroot, B. (1998). Outer root sheath cells of human hair follicle are able to regenerate a fully differentiated epidermis in vitro. Dev. Biol. 130: 610–620. Levy, L., Broad, S., Diekmann, D., Evans, R.D. and Watt, F.M. (2000). Beta1 integrins regulate keratinocyte adhesion and differentiation by distinct mechanisms. Mol. Biol. Cell 11: 453–466. Levy, P.M. and Salomon. D. (1998). Use of Biobrane after laser resurfacing. Dermatol. Surg. 24(7): 729–734. Li, A., Simmons, P.J. and Kaur, P. (1998). Identification and isolation of candidate human keratinocyte stem cells based on cell surface phenotype. Proc. Natl Acad. Sci. USA 95: 3902–3907.

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Stoner, M.L. and Wood, F.M. (2000). The treatment of hypopigmented lesions with cultured epithelial autograft. J. Burn Care Rehabil. 21(1): 41–46. Sugihara, H., Toda, S., Yonemitsu, N. and Watanabe, K. (2001). Effects of fat cells on keratinocytes and fibroblasts in a reconstructed rat model using collagen gel matrix. Br. J. Dermatol. 144: 244–253. Swope, V.B., Supp, A.P., Greenhalgh, D.G., Warden, G.D. and Boyce, S.T. (2001). Expression of insulin-like growth factor I by cultured skin substitutes does not replace the physiologic requirement for insulin in vitro. J. Investig. Dermatol. 116(5): 650–657. Tani, H., Morris, R.J. and Kaur, P. (2000). Enrichment for murine keratinocyte stem cells based on cell surface phenotype. Proc. Natl Acad. Sci. USA 97: 10960–10965. Tausche, A.K., Skaria, M., Bohlen, L., et al. (2003). An autologous epidermal equivalent tissue-engineered from follicular outer root sheath keratinocytes is as effective as split-thickness skin autograft in recalcitrant vascular leg ulcers. Wound Repair Regen. 11: 248–252. Tavakkol, A., Varani, J., Elder, J.T. and Zouboulis, C.C. (1999). Maintenance of human skin in organ culture: role for insulin-like growth factor receptor. Arch. Dermatol. Res. 291: 643–651. Tenchini, M.L., Ranzati, C. and Malcovati, M. (1992). Culture techniques for human keratinocytes. Burns 18: 11–15. Trempus, C.S., Morris, R.J., Bortner, C.D., et al. (2003). Enrichment for living murine keratinocytes from the hair follicle bulge with the cell surface marker CD34. J. Investig. Dermatol. 120: 501–511. Trent, J.F. and Kirsner, R.S. (1998). Tissue engineered skin: Apligraf, a bi-layered living skin equivalent. Int. J. Clin. Pract. 52(6): 408–413. Tobin, H.A. and Karas, N.D. (1998). Lip augmentation using an alloderm graft. J. Oral Maxillofac. Surg. 56(6): 722–727. Toma, J.G., Akhavan, M., Fernandes, K.J., et al. (2001). Isolation of multipotent adult stem cells from the dermis of mammalian skin. Nat. Cell Biol. 3: 778–784. Tumbar, T., Guasch, G., Greco, V., Blanpain, C., Lowry, W.E., Rendl, M. and Fuchs, E. (2004). Defining the epithelial stem cell niche in skin. Science 303: 359–363. VanderKam, V.M., Achauer, B.M. and Finnie, G. (1997). Use of a semipermeable dressing (Biobrane) following laser resurfacing of the face. Plast. Surg. Nurs. 17(3): 177–179. Wainwright, D.J. (1995). Use of an acellular allograft dermal matrix (AlloDerm) in the management of full-thickness burns. Burns 21(4): 243–248. Wang, H.J., Chen. T.M., Cheng, L.F., Cheng, T.Y. and Tung, Y.M. (1993). Human keratinocyte culture using porcine pituitary extract in serum-free medium. Burns 21: 503–506. Wang, H.J., Wan, H.L., Yang, T.S., Wang, D.S., Chen, T.M. and Chang, D.M. (1996). Acceleration of skin graft healing by growth factors. Burns 22(1): 10–14. Watt, F.M. (2002). Role of integrins in regulating epidermal adhesion, growth and differentiation. Embo J. 21: 3919–3926. Watt, F.M. and Hogan, B.L. (2000). Out of Eden: stem cells and their niches. Science 287: 1427–1430. Webb, A., Li, A. and Kaur, P. (2004). Location and phenotype of human adult keratinocyte stem cells of the skin. Differentiation 72: 387–395. Williamson, J.S., Snelling, C.F., Clugston, P., Macdonald, I.B. and Germann, E. (1995). Cultured epithelial autograft: five years of clinical experience with twenty-eight patients. J. Trauma 39(2): 309–319. Wood, F.M. (2001). The first seven years of the west Australian skin culture laboratory. Cultured Human Keratinocytes and Tissue Engineered Skin Substitutes. pp. 275–283. Wood, F.M. (2003). Clinical potential of autologous epithelial suspension. Wounds 15(1): 16–22. Wood, F.M. and Allen, P. (2003). The use of cultured epidermal autograft in the treatment of major burn injuries. J. Burn Care Rehabil. 13(1): 154–157. Wood, F.M., Kolybaba, M.L. and Allen, P. (2006). The use of cultured epithelial autograft in the treatment of major burn wounds: eleven years of clinical experience. Burns 32(5): 538–544 (Epub June 14). Wu, J.J., Liu, R.Q., Lu, Y.G., et al. (2005). Enzyme digestion to isolate and culture human scalp dermal papilla cells: a more efficient method. Arch. Dermatol. Res. 297: 60–67. Yamaguchi, Y. and Yoshikawa, K. (2001). Cutaneous wound healing: an update. J. Dermatol. 28(10): 521–534. Yannas, I.V. and Burke, J.F. (1990). Design of an artificial skin. I. Basic design principles. J. Biomed. Mater. Res. 14: 65–69. Yannas, I.V., Burke, J.F., Gordon, P.L., Huang, C. and Rubenstein, R.H. (1990). Design of an artificial skin. II. Control of chemical composition. J. Biomed. Mater. Res. 14: 107–132.

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77 Military Needs and Solutions in Regenerative Medicine Sara Wargo, Alan J. Russell and Colonel John B. Holcomb INTRODUCTION The human body has the inherent ability to heal a wide variety of its organs and tissues following damage from both acute traumatic injury and long-term disease processes. Throughout history, the art and science of medicine has sought to facilitate this intrinsic self-renewing ability by relieving damaged tissues from their functional burden and providing what was empirically perceived to be the ideal environment for tissue healing. As medical science has blossomed in recent decades, an increasing emphasis has been placed upon the replacement of diseased tissues with synthetic, and more recently, transplanted or engineered tissues. The notion of off-the-shelf replacement parts has captured the imagination of clinicians, scientists, and engineers alike and is routinely presented in the lay press with diagrams of the human body surrounded by an array of replacement parts generated from synthetic materials of the past or tissue engineered constructs of the future. Researchers around the world are investigating the promise of replacement tissues and organs. We define regenerative medicine as the utilization of therapeutic techniques to reestablish function lost in diseased or damaged tissues or organs. Regenerative medicine and tissue engineering utilize a variety of approaches to address tissue or organ insufficiency, including:

• • • •

The replacement of tissue function with devices (such as the case with artificial organs). Functional restoration with constructs that comprise both devices and cellular components (such as in biohybrid organs). The combination of temporary biodegradable scaffolds with cellular components (such as in conventional tissue engineering). Cellular therapies, including those involving differentiated cells, stem cells, and genetically manipulated cells (such as for the repair of damaged tissue and muscle).

These approaches are illustrated in Figure 77.1. The defense-related needs for regenerative therapies are clear, and it is vital to develop short- and longterm strategies to deliver on the promise of regenerative rehabilitation.

THE APPLICATION OF TISSUE ENGINEERING TO COMBAT CASUALTY CARE Musculoskeletal Tissue Engineering Casualty and trauma care today do a superb job of saving lives, but the therapies that are used to save lives do not always enhance the chances of tissue and functional recovery as the patient heals. The preponderance of

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Figure 77.1 Examples of the approaches used in regenerative medicine. Box A shows damaged tissue being replaced entirely by an artificial retina. Box B illustrates a device used to harvest cells to heal a facial burn. A conventional tissue engineering approach is illustrated in box C, where the jaw bone defect is healed with a biodegradable scaffold. Box D shows the injection of stem cells into skeletal muscle, demonstrating cellular therapies. soldier injuries are musculoskeletal, and in the civilian population it is well known that tissue engineering approaches to accelerating the pace of skin, bone, and cartilage healing are in clinical implementation. Cardiothoracic and Vascular Tissue Engineering Chest trauma in the military and general population has a crippling effect on productivity and sustainability. The cost of care can be enormous, and new technologies can have a real and lasting impact. Widespread success in the development of a biological vascular graft would depend upon performance superior to currently utilized bypass vessels (saphenous vein, radial artery), and may require genetic engineering to augment native endothelial antithrombotic activity if small vessels (Teebken and Haverich, 2002; Heyligers et al., 2006) (4 mm) are to be obtained. Attempts to develop such a substitute vessel have faced numerous obstacles, including immunologic barriers associated with allogenic or xenogenic tissue, supplying the engineered vessel with a capillary blood supply to maintain mural cell viability, and achieving physiologic biomechanical properties. Wound Healing The civilian population struggles with how to heal incalcitrant wounds. Diabetes is resulting in more than 60,000 amputations per year in the United States alone (Meltzer et al., 2002) as a result of the progression and infection of non-healing wounds. Although soldiers are generally healthy and resilient the conditions that they experience often lead to small and large wounds that can become infected. The issues that must be

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addressed in dealing with these wounds include bleeding, infection, pain, fluid loss, and accelerating repair. One can envision tissue engineers developing a toothpaste tube-like dispenser for self-treating of wounds with polymers that prevent infection, induce regeneration, stop blood loss, and deliver pain relief. Military Needs Wounded soldiers are treated in three phases: pre-hospital battlefield treatment, in-hospital care, and longterm therapy. Each phase of treatment has unique challenges, and in each case tissue engineering and regenerative medicine can play important roles in advancing therapies. Naturally, it is in the long-term restorative therapy arena where regenerative medicine is likely to have its greatest impact. Pre-hospital Care Wound healing is the primary focus of improved pre-hospital care for injured soldiers. An analysis of battlefield statistics explains why this is of dire importance to the future health and well-being of the injured. Since the beginning of the American Civil War in 1861 to the present time, approximately 20% (Bellamy, 1984) of all US combat casualties have resulted in death prior to reaching a medical treatment facility. Of those deaths, 50% were due to hemorrhage (Alam et al., 2005). Casualty analysis from the Vietnam War revealed that 40% of all battlefield exsanguination deaths could have been prevented with hemorrhage control (Alam et al., 2005). What is surprising is that the treatment for battlefield hemorrhaging has not changed significantly in nearly 2,000 years (Zimmerman and Veith, 1993). Hemorrhage prevention is a prerequisite to the retention of cellular function, which is necessary for optimal future regenerative therapies. Interestingly, a biodegradable matrix has been the basis of the most significant advance in hemorrhage control in many decades. The need for advanced battlefield hemorrhage technology has been recognized by the Department of Defense. The US Army and the US Marine Corp are currently using different FDA-approved products for hemorrhage control. Marine Polymer Technologies and HemCon, Inc. both produce a hemostatic agent derived from poly-N-acetyl glucosamine (p-NAG). This complex polysaccharide has been processed into either thin sheets, as shown in Figure 77.2, or a liquid, and is the hemostatic agent of choice by the US Army. The product gained approval by the FDA in November 2002. The mechanism through which the dressing stops blood loss is not entirely elucidated, but is likely to be due to an increase in the concentration of clotting factors at the site of injury (Sondeen et al., 2003; Alam et al., 2005). Multiple studies have been conducted to demonstrate the

(a)

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Figure 77.2 The p-NAG patch is used as a quick means for clotting blood. Placing the patch over a wound temporarily induces clotting on the battlefield while the injured soldier is transported to the hospital for further care.

Military Needs and Solutions in Regenerative Medicine

product’s effectiveness in several injury models, including peritoneal injury, liver and spleen injuries, and brain hemorrhage (Brandenberg et al., 1984; Klokkevold et al., 1991; Fukasawa et al., 1992; Chan et al., 2000; Jewelewicz et al., 2003). All these studies demonstrate the ability of a p-NAG bandage to stop hemorrhage rapidly. Detailed data regarding the operational success of this dressing are yet to be reported (Alam et al., 2005). The US Marine Corp has used a mineral zeolite hemostatic agent. This product works via a chemical reaction, adsorbing the water from the site of injury, thereby increasing the natural concentrations of key clotting factors in the coagulation cascade (Alam et al., 2003). The zeolite is produced by Z-Medica, LLC, and is known by the trade name QuickClot (Alam et al., 2005). Very few published studies exist on the effectiveness of this product, but those that exist describe rapid clotting (Alam et al., 2003, 2004; Wright et al., 2004). The mechanism by which the mineral zeolite operates is purely chemical resulting in the exothermic adsorption of water, thereby increasing the temperatures of a wound for brief time periods. Temperatures can range from 40°C to 60°C, and the dressing can cause significant pain while preventing hemorrhage (Alam et al., 2005). A natural question that emerges is whether biologically inspired principles of regenerative medicine could add further to the suite of products available to prevent blood loss. Consider the development of a lightweight liquid polymer that would solidify when applied to a wound, filling the wound space. It would need to be a nontoxic, sterile (pathogen-free) product that could be placed into the wound by the soldier. It would also need to be stable for prolonged periods at extremes of temperature and simple enough in design to be self-administered. It could serve as a wound covering to reduce complications and stimulate healing while still on the field. Once the soldier returns from the mission, it could then be removed if necessary. At the University of Pittsburgh, Professor Eric Beckman has developed a polymer that is amorphous for easy application to wounds of varied size and shape, but forms a gel when placed within the wound space. It does so at extremes of temperature, and the gel solidifies in seconds within the wound space allowing contact between the gel and the injured tissues. Thus, the gel serves as a delivery platform supplying factors to the wound to facilitate healing. The polymer is nontoxic and does not inhibit wound healing. Figure 77.3 illustrates this innovation. Dr. Beckman’s group has successfully created a liquid prepolymer system that cures in times as short as 10 s, where both the prepolymer and the curing system are composed of benign building blocks. The prepolymer is composed of three basic building blocks: (a) a biocompatible polyhydroxy core material, (b) a lysine linker, and (c) terminal L-DOPA moieties. L-DOPA is known to polymerize in the presence of free radicals,

Figure 77.3 A developing technology for controlling battlefield hemorrhage. The polymers mix and cure in the wound in a matter of seconds, stopping the injured from continued bleeding.

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and hence a biocompatible curing system would need to include free radical initiators. They chose the twopart system silver (Ag2)/peroxydiphosphate, as this redox pair is known to rapidly phosphate radical anions at mild temperatures. Silver catalyzes the formation of radicals and also exhibits anti-bacterial properties, which makes it an ideal part of the cure system. They have found that using appropriate amounts of both the diphosphate and the silver, curing is accomplished in times as short as 10 s, and that the cure time can be adjusted from seconds to hours by varying silver and diphosphate concentration. Changing the average functionality of the polymer network (through mixing of various prepolymers) allows one to vary the physical properties of the cured material from elastomeric to more rigid. In-hospital Care Hemorrhage is still the leading cause of death of those survivors receiving hospital care. Wound healing is also an important phase of in-hospital care. Wound healing is generally understood to occur in three basic phases: the inflammatory phase, the matrix deposition phase, and the remodeling phase (Witte and Brabul, 1997; Goldman, 2004). Tissue-specific healing processes vary considerably but the three phases are a uniting principle (Hildebrand et al., 2005). A variety of proteins and growth factors have been identified as being vital for the wound healing process to occur. Until we more fully understand how all these factors work with each other to orchestrate natural wound healing we will not be in a position to guide the process in the direction that we want. That said, remarkable progress has been made in the use of natural biodegradable matrices that deliver all these factors as they degrade. Bioscaffolds derived from xenogeneic (porcine) extracellular matrices (ECMs) have been evaluated in numerous pre-clinical animal studies and in human clinical applications for the repair and reconstruction of injured or missing tissues. A family of such materials has now been used in more than 500,000 human clinical procedures (Bell et al., 1981; Lantz et al., 1993; Prevel et al., 1994; Kropp et al., 1996; Rickey et al., 2000). Over a decade of clinical experience and pre-clinical studies have now demonstrated that acellular ECMs induce an unnatural but highly beneficial wound healing response that is different from the default, scar-forming method of wound repair in adult mammals. Common features of the ECM-induced response include the recruitment of bone marrow-derived circulating progenitor cell population, angiogenesis, and cytokine profile that possesses many of the features of the developmental biology process of tissues and organs. Stated differently, tissue reconstruction rather than simple tissue repair by scar tissue occurs. ECM-based bioscaffolds have particular utility in a military setting, specifically the application of this bioscaffold to traumatic soft tissue injury. Recent studies by Badylak and colleagues have shown that ECM bioscaffolds have hemostatic properties, inherent antibacterial properties, and angiogenic properties (Sarikaya et al., 2002; Li et al., 2004). The use of multilaminate ECM bioscaffolds in severe acute abdominal and thoracic injuries is a particular area of interest. In this setting the ECM bioscaffold must be strong enough to withstand rough handling and tensile loads that exceed those of normal musculotendonous structures of the abdominal and thoracic wall structures, and yet be light enough to fit easily into a backpack for easy transport. Material-induced wound healing works by providing appropriate signals to cells. Naturally, as we learn more about the cells that are involved we are in a position to consider how to use those cells directly in therapies. Hart et al. (2005) demonstrated a potential role for mesenchymal stem cells in the facilitation of wound healing. The plasticity of stem cells would make them ideal for wound healing as they could be applied to multiple tissue types. Although the use of multipotent adult and embryonic-derived stem cells has attracted a great deal of media attention, relatively little is known about how these cells might elicit and then successfully manage a wound healing response. Indeed, there is an ongoing debate about the balance of advantages and disadvantages in allogeneic versus autologous cell sources (Ahsan and Nerem, 2005) and whether or not embryonic (Health, 2000) or adult (Rubio et al., 2005) stem cells provide more effective putative therapies. It is likely that stem cell therapy

Military Needs and Solutions in Regenerative Medicine

will be a key element of in-hospital combat casualty care in the future. Throughout this book one can find examples of how cells will be used in this manner. In a hospital setting we can envisage the use of regenerative medicine to induce healing of the skin, the musculoskeletal system, the internal organs, and the vasculature. Several living skin equivalents are commercially available, including epithelial replacements, dermal replacements, and full-thickness replacements. As suggested by each category name, the epithelial replacements promote the growth and the replacement of the epithelium, the dermal replacements promote the growth of the dermal layer, and the full-thickness replacements promote both epithelial and dermal growth (Seal et al., 2001; Atiyeh et al., 2005). Dermal skin grafts were made by Advanced Tissue Sciences (Dermagraft) and Life Cell Corporation (AlloDerm). Advanced Tissue Sciences developed a polymeric graft that is now in widespread use (Naughton et al., 1997; Buinewicz and Rosen, 2004). Epithelial skin replacements are not as abundant, but Genzyme has developed Epicel, a product that has demonstrated with promising. Epicel is an autograft of human keratinocytes and murine cells backed by petrolatum gauze (Genzyme Biosurgery, 2005a, b). Almost all treatments for burns today revolve around the concept of replacing lost tissue with some natural or manufactured tissue. Rather like sodding a lawn, the approaches only differ in the makeup of the sod. A radically different approach would be to consider how to seed and fertilize a wound bed. The idea of a “seeding” approach is now reality because “wearable bioreacators” are being combined with cell sprayers in a clinical trial in Europe (Gerlach, 2005; Bornemann, 2007; Hartmann et al., 2007). Musculoskeletal tissue engineering can also begin in the hospital. One of the most promising and near-term advances in tissue engineering of relevance to injured soldiers is regeneration of bone. The use of proteins, such as bone morphogenic protein, to stimulate bone deposition is now commonplace in the clinic. The question is now focusing on which materials and cells can be used to most effectively deliver the necessary signals that will regenerate non-union fractures. Key requirements for these materials include mechanical strength and degradability. Also, the starting material and the degradation products must be biocompatible and non-immunogenic. A common approach to regenerating bone is to combine a synthetic polymer with a ceramic material (Haney et al., 1993; Jansen et al., 1995; Kikuchi et al., 1999; Kikuchi et al., 2002). Ceramic materials offer the advantage of bonding directly to bone (Kikuchi et al., 1999; Kikuchi et al., 2002), while the polymer serves as the support for in-growing tissues (Jansen et al., 1995). Commonly chosen materials for the polymer support system are lactic acid derivatives, and ceramic materials include hydroxyapatite and β-tricalcium phosphate (Seal et al., 2001). In general, these materials have a mechanical strength that mimics that of native tissue and successfully supports bone growth. DegraPol (Saad et al., 1998) and Polyactive (Sakkers et al., 2000) are two new materials that have been developed and are currently under investigation. Published data indicate that these are suitable supports for bone growth. Dr. Charles Sfeir and colleagues are using mineralized proteins as scaffolds for bone and tooth regeneration. More specifically, they are combining calcium phosphate with proteins that serve as direct nucleators of mineralization and cell differentiation inducers to generate bone both in vitro and in vivo (Pola et al., 2004). This therapy could be employed on the injured to direct and accelerate the healing of bone tissue. Muscle regeneration is also necessary. Riboldi et al. have investigated the use of DegraPol in an attempt to regenerate muscle tissue in vitro (Riboldi et al., 2005). They have electrospun the polymer and grew skeletal muscle cells on the construct. They were able to demonstrate satisfactory elasticity and cell growth. Other materials have also been investigated for muscle regeneration. The key property in these scaffold materials is elasticity. Research groups have investigated the properties of in vitro muscle tissue grown in stationary culture (Swasdison and Mayne, 1992; Shansky et al., 1997; Dennis and Kosnik, 2000), on collagen constructs (Van Wachem et al., 1996; Shansky et al., 1997; Chromiak et al., 1998) and on elastic membranes (Vandenburgh et al., 1991; Shansky et al., 1997; Chromiak et al., 1998). As with bone, the future is promising for muscle regeneration; however, an artificial muscle has not been made that can be deemed comparable to native tissue.

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Dr. Johnny Huard is investigating the regeneration of muscle tissue without a scaffold. His group has been able to successfully isolate muscle-derived stem cells and show that they have an improved transplantation capacity in both skeletal and cardiac muscle than satellite cells do. They have also been able to isolate pericytes and show that this population of cells has the strongest myogenic potential. Dr. Huard’s group is working to identify the regenerative capacity of various human muscle-derived cell populations after implantation into skeletal muscle (Cao et al., 2005; Zheng et al., 2007). This would be ideal for the military as it would eliminate the need for scaffold and the complications that arise from using a scaffold. Other groups are focusing on using stem cells for stimulating the regeneration of blood vessels ischemic in limbs. A Japanese group has reported that injection of CD34 stem cells could enhance vascularization in the limbs. In order to qualify for this study, patients had to have chronic limb ischemia and could not be candidates for surgical vascularization. After injection, the ankle brachial pressure index, the transcutaneous oxygen pressure, and the rest pain were monitored with improvements seen as early as 4 weeks after injection (Tateishi-Yuyama, 2002). A group from Northwestern University will be conducting a Phase I clinical trial and will be recruiting patients with peripheral vascular disease (PVD) or coronary artery disease (CAD) that are not candidates for surgery. This study seeks to determine the affect of various doses of autologous stem cells in patients and determine how these injections change the patient's overall health (Burt et al., 2003; Losordo, 2007). Results of these studies have significant ramifications. Long-Term Therapy The final phase of care for the wounded soldier is long-term care. The hope of tissue engineering is that regenerative medicine and long-term rehabilitation science and engineering will fuse as fields. The resultant discipline of regenerative rehabilitation provides an exciting opportunity to no longer accept that rehabilitation is about how to adjust to a new reality. Instead, in the future we will influence the achievable reality by using tissue engineering to restore lost function and esthetics. Relatively simple tissue engineering therapies are being delivered to restore skin, bone, and cartilage. The excitement, however, will come when we can combine tissue engineered tissue types and replace more complex structures, such as organs, digits, or even limbs. It is no coincidence that the US Defense Advanced Research Project Agency has taken the lead on a large limb regeneration initiative. The field has made great strides with regard to repairing damaged tissue, as previously described. However, regenerating whole organs and limbs is much more difficult. Limbs, for example, are composed of bone, muscle, and nerves. They have a complex capillary network and a biological structure that may not be completely understood. Similarly, organs are composed of a variety of cell types and vary greatly both in form and function. In order to construct a whole limb or organ, the aspects of how all of the tissue types and cells interact to make the limb functional need to be understood. The key in developing these complex structures is really in understanding the fundamental biology. Some lower level organisms have the ability to regenerate limbs, and scientists are studying these animals as models for limb regeneration in humans (Endo et al., 2004). It is these mechanisms that will allow scientists to build an interactive system of cells that can act functionally as one unit, hence successfully constructing a limb. Mammalian fetuses in the first two trimesters have the full capacity to re-grow a functional amputated limb. Children below the age of 18 months can regenerate the tip of an amputated finger. Even as adults we retain the ability to perfectly reproduce parts of our bodies on a regular basis, such as the lining of the intestines, the inner surface of the uterus, and the entirety of the bone marrow. However, as we “mature” past 2 years of age we lose the capacity to regenerate entire limbs and digits, and instead, we typically repair tissues by forming scar tissue. Why do we change our mechanism of healing? Our genetic makeup does not change, so the capacity to regenerate tissues and organs must still reside within our cellular and extracellular environments. Surely, with the appropriate stimuli and signals, this regenerative capacity can be re-awakened. These stimuli

Military Needs and Solutions in Regenerative Medicine

and signals are inextricably linked not only to the immediate and acute surgical interventions that we employ, but also to the rehabilitation protocols that follow. These principles will apply not only to digits and limbs, but also to tissues and organs, such as damaged heart tissue, diseased livers, and non-functional kidneys.

CONCLUSION Millions of people suffer injuries and debilitating conditions with devastating consequences. Almost every treatment paradigm ameliorates symptoms but does not provide a “cure” for the underlying disease or injury. Rehabilitative approaches often seek to adjust a patient to a new reality rather than boldly participating in changing that reality. Regenerative medicine seeks to harness the innate ability of the body to heal itself and drive regenerative processes. The reality of regenerative medicine is that its success will depend heavily on a completely new paradigm in the approach to patient rehabilitation. We must seek to induce regeneration via rehabilitation and therapies through vertical integration from conception, innovation, translation, implementation, and follow-on care. As the graphic illustrates, the opportunity for participation at all stages in the evolution of regenerative medicine technology is important in the outcome and in the rate at which it is achieved. This concept of vertically integrated regenerative therapies, illustrated in Figure 77.4, is a dramatic departure from traditional therapeutic development, especially in a biological science world that is dominated by cell-derived therapies. The human body has the inherent ability to heal a wide variety of its organs and tissues following damage from both long-term disease processes and acute traumatic injury. Throughout history, the art and science of medicine has sought to facilitate this intrinsic self-renewing ability by relieving damaged

Traditional approach

Follow-on care

Fusion and vertical integration

Follow-on care

Rehabilitation

Implementation Implementation

Translation

on

ration Regen e

Innovation

bilitati

Reha

Innovation

Regeneration

Translation

Conception

Conception

Figure 77.4 Schematic comparison between the traditional approach to regenerative medicine and the vertically integrated regenerative therapies.

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tissues from their functional burden and providing what was empirically perceived to be the ideal environment for tissue healing. In our excitement to develop and optimize complete replacement therapies for organ failure, the underlying regenerative potential of the organs that we seek to excise and replace can be overlooked. While a significant number of tissues may indeed be irreparably damaged, a large fraction of patients burdened with a wide variety of organ and tissue insufficiencies might be better served by developing treatment modalities that embrace a fundamental principle of medicine: optimization of the regenerative potential intrinsic to many organ systems. War is hell, and the civilian and military casualties that survive that hell need regenerative medicine.

ACKNOWLEDGMENTS Much of the work described here is being performed at the National Tissue Engineering Center under DOD contratcts (NTEC Grant DAMD17-02-1-0717, STRAC Grant W81XWH-04-1-0848, ARM Grant W81XWH07-0415 and SWH Grant W81XWH-04-1-0848). Figures 77.1–77.3 were illustrated by Randal McKenzie ([email protected]). Figure 77.4 was developed in collaboration with Dr. Clifford Brubaker. We thank Drs. C. Brubaker, R. Cooper, E. Beckman, S. Badylak, J. Huard, and J. Gerlach for their assistance in the preparation of this chapter.

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Riboldi, S.A., Sampaolesi, M., Neuenschwander, P., Cossu, G. and Mantero, S. (2005). Electrospun degradable polyesterurethane membranes: potential scaffolds for skeletal muscle tissue engineering. Biomaterials 26: 4606–4615. Rickey, F.A., Elmore, D., Hillegonds, D., Badylak, S., Record, R. and Simmons-Byrd, A. (2000). Re-generation of tissue about an animal-based scaffold: AMS studies of the fate of the scaffold. Nucl. Instrum. Methods Phys. Res., Sect. B. 172: 904–909. Rubio, D., Garcia-Castro, J., Martín, M.C., de la Fuente, R., Cigudosa, J.C., Lloyd, A.C. and Bernad, A. (2005). Spontaneous human adult stem cell transformation. Cancer Res. 65: 3035–3039. Saad, B., Ciardelli, G., Matter, S., Welti, M., Uhlschmid, G.K., Neuenschwander, P. and Suter, U.W. (1998). Degradable and highly porous polyesterurethane foam as biomaterial: effects and phagocytosis of degradation production in osteoblasts. J. Biomed. Mater. Res. 39: 594–602. Sakkers, R.J.B., Dalmeyer, R.A.J., de Wijn, J.R. and van Blitterswijk, C.A. (2000). Use of bone bonding hydrogel copolymers in bone: an in vitro and in vivo study expanding PEO–PBT copolymers in goat femora. J. Biomed. Mater. Res. 49: 312–318. Sarikaya, A., Record, R., Wu, C., Tullius, B., Badylak, S. and Ladisch, M. (2002). Antimicrobial activity associated with extracellular matrices. Tissue Eng. 8: 63–71. Seal, B.L., Otero, T.C. and Panitch, A. (2001). Polymeric biomaterials for tissue and organ regeneration. Mater. Sci. Eng. R. Rep. 34: 147–230. Shansky, J., Del Tatto, M., Chromiak, J.A. and Vandenburgh, H.H. (1997). A simplified method for tissue engineering skeletal muscle organoids in vitro. In Vitro Cell Dev. Biol. Anim. 33: 659–661. Sondeen, J.L., Pusateri, A.E. and Coppes, V.G. (2003). Comparison of 10 hemostatic dressings in an aortic injury. J. Trauma 54: 280–285. Stocum, D.L. (2005). Regenerative biology and medicine: an overview. www.cellscience.com. Swasdison, S. and Mayne, R. (1992). Formation of highly organized skeletal fibers in vitro. Comparison with muscle development in vivo. J. Cell Sci. 102: 643–652. Tateishi-Yuyama, E., Matsubara, H., Murohara, T., Ikeda, U., Shintani, S., Masaki, H., Amano, K., Kishimoto, Y., Yoshimoto, K., Akashi, H., et al. (2002). Therapeutic angiogenesis for patients with limb ischemia by autologous transplantation of bone-marrow cells: a pilot study and randomized controlled trial. Lancet 360: 427–435. Teebken, O.E. and Haverich, A. (2002). Tissue engineering of small diameter vascular grafts. Eur. J. Vasc. Endovasc. Surg. 23: 475–485. Van Wachem P.B., van Luyn M.J. and da Costa M.L. (1996). Myoblast seeding in a collagen matrix evaluated in vitro. J. Biomed. Mater. Res. 30: 353–360. Vandenburgh, H.H., Swasdison, S. and Karlisch, P. (1991). Computer-aided mechanogenesis of skeletal muscle organoids for reversible gene therapy. FASEB J. 5: 2860–2867. Witte, M.B. and Brabul, A. (1997). General principles of wound healing. Surg. Clin. N. Am. 77: 509–528. Wright, F.L., Hua, H.T., Velmahos, G., Dave, T., Deemitriades, D. and Rhee, P. (2004). Intracorporeal use of a hemostatic agent QuickClot in a coagulopathic patient with combined thoracoabdominal penetrating trauma. J. Trauma 56: 205–208. Zheng, B., Cao, B., Mihaela, C., Sun, B., Li, G., Logar, A., Yap, S., Pollett, J.B., Drowley, L., Cassino, T., Gharaibeh, B., Deasy, B.M., Huard, J. and Peault, B. (2007). Prospective identification of myogenic endothelial cells in human skeletal muscle. Nat. Biotechnol. 25:1025-1034. Zimmerman, L. and Veith, I. (1993). Great Ideas in the History of Surgery. San Francisco: Norman Publishing.

Part VII Regulations and Ethics

78 Ethical Considerations Louis M. Guenin As natural phenomena are to the scientist, so are arguments to the philosopher. The philosopher Richard Hare once said, “I like to give arguments for my position. They come in handy when people don’t agree with me.” Consider then the inclination of many among us—perhaps by virtue of being busy, or for other reasons—to pronounce a verdict, when a moral controversy comes along, by consulting aphorisms or slogans. By following this inclination, earnest people may unwittingly betray the moral views that they aspire to uphold. For they may fail to take account of the depth and subtlety of their respective moral views. This chapter presents arguments offered in support of using human embryos in research and therapy, sorting them into arguments that work and arguments that do not. I first review six arguments that I place in the latter category. Each of these purports to justify research that I happen to support. But inasmuch as a good case is not made better by overstatement, and no case is made by an unsound argument, I am going to disavow those six arguments. I urge other supporters of donated embryo use to disavow them as well, because, as Bernard Williams once said, openness to criticism is the homage that candor pays to truth. To support my own view, I shall go on to sketch arguments that, so I shall suggest, are sound. I shall then say a bit more about cloning in particular and shall close by remarking on the risk of abuses.

ARGUMENTS THAT DO NOT WORK Imminent Death as Justification for Killing Consider the argument that the imminent death of an embryo—for instance, a surplus embryo in a fertility clinic—justifies its consumption in research. A more extreme proposal would have us define a concept of embryo death according to which embryos not destined for intrauterine transfer are dead. This concept of death seems to defy common sense. The embryos that I contemplate as research subjects are alive. About this concept of death I doubt that it is necessary to say more. Let us consider the plain argument that imminent death of an admittedly living embryo justifies killing it in research. In refutation of that argument, consider the following Wild West example. One day the notorious varmint Hatfield is riding about on his horse. Feeling tired, Hatfield elects to dismount beside the railroad tracks. He sits down and eventually dozes off, stretching out between the tracks. Sometime later, as Hatfield lies sound asleep, a train approaches at high speed. Whereupon there happens to ride onto the scene Hatfield’s archenemy, McCoy. Spotting Hatfield, McCoy gallops to the tracks, dismounts, and—just in the nick of time before the train arrives—yanks Hatfield from the tracks. McCoy then immediately pulls out his rifle, trains it on Hatfield, and kills Hatfield. In this case, even though Hatfield would have died under the wheels of the train a moment earlier, we hold McCoy guilty of wrongful killing. In general, the imminent death of a victim does not justify its killing. Embryo research will not be justifiable solely on the ground of imminent embryo death.

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Nonindividuation Argument Against Zygotic Personhood The nonindividuation argument against zygotic personhood runs as follows. Prior to formation of an embryo’s primitive streak at day 14 of development, it can happen that the embryo splits into monozygotic twins. If twinning occurs, the twins can fuse. Hence it has been suggested that in respect of any embryo, one cannot say until day 14 whether there exists one individual or more. If one cannot say how many persons exist, it is untenable to say that any person exists. Another version of the nonindividuation argument begins from the premise that in twinning, an embryo vanishes and leaves no earthly remains. How could an individual person die leaving no earthly remains? If no corpse ever exists, there could not earlier have existed a person. Thus does the nonindividuation argument characterize zygotic personhood as metaphysically impossible. The argument’s biological sophistication has led many scientists to regard it as a clinching argument for embryo use. Despite that sophistication, the nonindividuation argument is susceptible to the following refutation owed to David Oderberg (Oderberg, 1997). Suppose at time t0 a somatic cell x.By t1, x has divided, and there exist x’s daughter cells d1 and d2. The process by which this has occurred, namely mitosis, is routine. Notice that as we look back at the history of x up to t1, we do not have any doubt that x was an individual cell. Plainly x had the capability of dividing, and in fact, x did divide, but it is not incoherent to say, and we unhesitatingly do say that for so long as x existed, x was an individual. As for the apparent puzzle of dying without leaving earthly remains, again a reflection on mitosis sheds light. Necrosis is not the only means by which a life form may cease to exist. Dividing is another means. It happens that after division, there is no corpse. So it is not metaphysically incoherent to say that an embryo capable of dividing is an individual. Or that an embryo that does divide was an individual before it divided. To rehabilitate the nonindividuation argument, a proponent might contend that indivisibility is somehow intrinsic to the individuality of a person as it may not be to the individuality of a cell, so that a being that is divisible cannot be or correspond to a person (Guenin 2006). To buttress this claim, the proponent might offer the example that an adult individual cannot divide into surviving individuals. Or the proponent might contend that even if adult individual x could be split into surviving individuals—say, by a brain split and transplant operation in which, as imagined by philosophers of mind, x’s brain is split and each half is transferred into a new body so that each successor retains memories and otherwise achieves psychological continuity with x—it would not be the case that x is the same individual as either of its successors. One reply to this, owed to the whimsy of Peter King, is that it is possible to survive with only half a brain, though in such case one is restricted to a career in politics. But we may leave aside what adults can do. We may reply to the proponent of the nonindividuation argument with two thoughts. First, what is feasible for an adult ought not constrain our thinking, because we know of the remarkable ability of an early embryo to split into surviving individuals. Second, the individuality of a being does not depend on its being the same individual, if it happens to split, as any of its successors. When an embryo has split, we may simply say that it was one individual until it split, that the individual ceased to exist when it split, and that two individuals have succeeded it. Thus may we render not only the possibility but the actuality of twinning consistent with individuality of the original embryo and with individuality of its twin successors. Given that the case of an embryo that does split resolves in this way, no impediment arises to individuality of an embryo that has not split. According to this analysis, the nonindividuation argument fails to establish that an embryo cannot be a person. When monozygotic twinning occurs, it may be said that two persons succeed one person that ceases. But even if we defeat the contention that an embryo cannot be a person, there remains the question whether, for purposes of the duty not to kill, we should treat every embryo as a person. That a being “is a person” is not an empirical observation or an a priori truth. Calling a being a person is a shorthand reply to the moral question, “How should we treat it for this purpose?” The shorthand signifies our conclusion that we should classify

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the being among those to whom we think all of us should accord a particular treatment. Whereupon we may fairly be asked what argument supports that conclusion. Failure of the nonindividuation argument leaves the door open to introduce, or oppose, arguments that we are obliged to treat all embryos as persons for purposes of the duty not to kill. The Utilitarian Defense of Embryo Use Utilitarianism commands us to maximize aggregate preference satisfaction for the universe of affected sentient beings. A familiar argument is that if, in regenerative medicine, we sacrifice a relatively small number of embryos in order to help millions, perhaps billions, of suffering people, we can achieve higher aggregate preference satisfaction than we would achieve were we to classify every embryo as a person for purposes of the duty not to kill. John Stuart Mill, who with Bentham brought utilitarianism to prominence, learned calculus at the age of 5 years, but Mill did not envision the mathematical defect of his moral theory. The infirmity came to light through the work of economists in the 20th century. Consider that the number of affected sentient beings in respect of many policy issues, embryo use among them, is enormous. Collecting utility data from so many people would be a monumental task. A more fundamental failing is that there obtains no method of measuring utility. Utilitarianism presupposes a utility function for each member of the set of affected beings. A utility function is nowadays understood as a real-valued order homomorphism representing a transitive and connected binary relation defined by an individual on a set of alternatives. This understanding no doubt explains why many scientists have felt comfortable with utilitarianism and with a utilitarian defense of embryo research. Utilitarianism seems empirical, quantitative, precise. For physician scientists, utilitarianism evokes some of the thinking to which outcome studies, comparing benefits and costs, have accustomed them. But if, for two or more individuals, a utilitarian seeks to sum the utilities of a given alternative, there obtains no common unit of measure. There does not exist a standard measure even for a single person. While it is easy to define, as a representation of an individual’s positioning of alternatives, a real-valued order homomorphism, any of infinitely many other functions formed by affine transformations of that order homomorphism will also represent the positioning. Though the problem of interpersonal incommensurability of utilities remains unsolved, on many occasions a utilitarian’s audience either will be unaware of the problem, or willing to overlook it. For example, if a hospital were to propose construction of a new facility, and if that project would require demolition of the homes of 100 people, displacing those residents in exchange for reimbursement of their homes’ fair market value and their moving expenses, a utilitarian might argue that future gains in utility produced by the project for perhaps hundreds of thousands of patients in future decades will exceed the disutility of the 100 who must presently relocate. Listeners will follow an argument such as this without worrying much about whether the utility calculation has been performed. For it may seem in such a case that for any plausible conversion ratio of units of measure, the comparison of utility between alternatives will be lopsided. But when an advocated alternative is the killing of a life form that some people sincerely believe to be a person, not much tolerance will be found for an argument whose proponents cannot produce the calculation on which the argument purportedly rests. The root of the difficulty is the same as in the blinkered attempt to order, on the basis of supposed measures of quality, incommensurable college football teams. Appeal to “Fact-Based Reasoning” Alone The next argument proceeds by asserting that public advocacy concerning embryo use should appeal to “factbased reasoning” alone, and that it therefore follows that objections resting on any other ground must give way to the progress of biomedical science. This argument serves as a euphemism for saying that appeals to religion

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and various “insular” moral views do not have a place in public debate about science. In reply, we must observe that the view that scientific work grounded in fact-based reasoning should go forward without obstruction by moral views is itself a moral view. When an objection lies on grounds of wrongful killing, it is neither appropriate nor feasible to oust religion or moral views from the conversation. Of course we all agree that we should rely on facts as opposed to errors, but given facts, a normative discussion awaits. A more sophisticated cousin of the foregoing argument, the call to “public reason” by John Rawls, would have us employ in public discourse only premises that can be supported without appeal to any particular comprehensive moral or religious view (Rawls 1993). That, in a moment, is what I shall try to do in stating my own view. Noncomplicity Defenses It has been suggested that even if the consumption of embryos in research is wrong, a government could support derivative research by eschewing complicity in the destruction of embryos. As it has been put by proponents of this move, one could distinguish between embryonic stem cell derivation and embryonic stem cell use. The National Institutes of Health (NIH) at one time adopted this view (65 Fed. Reg. 51976—51981 [2000]). It announced that it would fund projects classified as embryonic stem cell use. The notion seemed to be that this would avoid complicity in wrongful embryo-destructive derivation of stem cells. In another scheme for purportedly conducting embryo research without complicity in embryo killing, fertility clinicians would perform the immunosurgery by which embryonic stem cells are derived from embryos. This scheme is probably impractical, as fertility clinicians do not do that kind of work. They do not customarily derive cell lines from embryos; they customarily nurture and transfer embryos so as to achieve pregnancies. Even if the scheme were practical, it shares with the derivation-use distinction the problem that when a chain of supply runs from someone who sacrifices an embryo to someone who experiments with the sacrificed embryo’s derivatives, we seem compelled to say from a moral point of view that the source and the recipient ride in the same boat. It is untenable to say that the experimenter is not complicit in the work of the supplier (Guenin 2005b). Another noncomplicity strategy, recently played out in the United States, might be called “government surprise.” It would begin as a government announces that it will not fund research that effects or is consequent on destruction of embryos. Then, after this policy has become widely known, the government suddenly announces—in the United States, we saw this happen on August 9, 2001—that in the future it will disperse public funds for studies using derivatives of embryos already then sacrificed. The most cogent philosophical defense of this gambit that occurs to me might be to say that the government had not, prior to the second announcement, induced any destruction of embryos. It would say this on the supposition that theretofore, the government had given everyone to believe that it would not support such research. (The supposition may not be true with respect to the recent history of US government policy. Between the first and second announcements came the just mentioned NIH announcement that it would fund research on embryonic derivatives. What I am here calling “the first announcement” was a prohibition enacted by Congress as a rider to an appropriations bill [Pub. L. No. 104–99, Title I, §128, 110 Stat. 26, 34 (1996)]. But apart from this historical contingency, the government surprise scheme succumbs to the same objection that lodges against the two other schemes that I have just mentioned. The government surprise scheme would place those who participate in funded embryonic stem cell studies in the same boat with those who participate in nonfunded embryodestructive stem cell derivation (Guenin 2004b, 2005b). Denying That Clones Are Embryos The last on this roster of arguments that do not work consists in the claim that somatic cell nuclear transfer performed in research is not cloning and does not produce clones or embryos, this because a suggested new

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semantic regimen would withhold the term “cloning” from any instance of somatic cell nuclear transfer in research, would instead call that process “nuclear transplantation,” and would withhold “clone” and “embryo” from the products of that process. For this context, I have elsewhere tried to sort out the relevant entities and events in a manner informed by biological and moral considerations (Guenin 2003). My analysis leads me to reject the proposed semantic regimen as to both process and products. The proposed terminology risks the appearance of trying to smuggle in a morally significant event—initiation of embryogenesis—by not mentioning it. In trying to withhold “cloning” and “clone” from processes and products of research, a proponent of the proposed terminology would contradict the ordinary and morally significant understanding of cloning as a genetic event, an event completed shortly after oocyte activation, regardless whether transfer to a uterus ensues. The goal that has motivated the proposed terminology, the goal of offering the public a sharp distinction between research and producing children, does not require legerdemain to attain. Rather we may implement a simple distinction, a distinction between “procreative cloning” and “nonprocreative cloning.” The distinction turns on a single observable event, intrauterine transfer. As for “embryo,” any being that is of a kind capable of developing into a neonate upon transfer to a uterus is an object of moral concern. We implicitly acknowledge that concern when we classify every prefetal developing organism as an embryo. (In both scientific writing and popular speech, we have abandoned the textbook definition according to which “embryo” applies only to a conceptus older than 2 weeks.) Recognition of a being as an embryo does not end our moral investigation. We may conclude—for reasons that I shall shortly present, I believe that we should conclude—that we are not obliged to treat every embryo in the same way. But in respect of a product of somatic cell nuclear transfer, recognizing its inclusion within a discussant’s universe of moral concern is the place from which to begin our discussion with one who does not agree with us (Guenin 2004a).

ARGUMENTS THAT WORK Argument from Nonenablement To advance the proposition that embryonic stem cell research is virtuous if not obligatory, I have presented an argument that I call “the argument from nonenablement” (Guenin, 2001 Guenin [forthcoming]). I refer to an embryo that will never enter a uterus as an “unenabled” embryo. In the first instance, I have in mind a situation, which often arises with fertility patients, in which the one person in the world who, together with her partner, is empowered to decide about intrauterine transfer of an embryo formed from her oocyte decides that neither does she wish to have that embryo transferred into her nor does she wish to give the embryo to anyone else for intrauterine transfer. There is no moral view of which I know that asserts a duty of intrauterine embryo transfer into oneself. That is to say that there is no moral view that asserts that a woman lies under a duty to undergo a transfer into her of an embryo that lies outside her. About such a procedure, we respect her autonomy. Imagine, therefore, that a woman declines intrauterine transfer, and in fact, forbids it. She, with her partner, donates to medicine either an embryo created during her fertility treatment, or an embryo that will be created by a scientist from their donated cells. Let us assume that this decision is final and that the embryo has left parental control. Such a donation to medicine I call an “epidosembryo.” I take this name from the Greek epidosis for a citizen’s great beneficence to the common weal. A distinction obtains between the developmental potential of an embryo that lies in a petri dish and will remain there, and an embryo that lies in a uterus, however it got there. In consequence of parental instructions

Ethical Considerations

that an epidosembryo shall be used in research or therapy and shall not be transferred to a uterus, there does not obtain any morally significant chance that from such an embryo, an infant will develop. To put the matter in language that I owe to Hare, no possible person corresponds to an epidosembryo (Hare 1993). We also know, and this is purely empirical knowledge, that an embryo is not sentient. And that for lack of a cortex, an embryo cannot form preferences. Nor can an embryo adopt ends. Therefore nothing that we might do to an epidosembryo can cause it discomfort or frustrate it. Under these circumstances, and when we consider the duty of mutual aid asserted within each of the leading moral views, I claim that it is permissible to use some embryos, namely epidosembryos, in medicine. As the preceding discussion makes clear, I rest the permissibility and virtuousness of epidosembryo use on the autonomous decisions of couples from whose cells such embryos originate. The moral analysis flows entirely from what it is that they decide. If it is permissible for those donations to be made, then it is permissible for recipient scientists to use epidosembryos as instructed. Suppose that someone interjects that precisely because epidosembryos cannot form preferences, it is our obligation to act according to their advantage. I reply that we cannot promote any advantage of epidosembryos. Were we to refrain from using epidosembryos in research, we could not gain anything for them. Entry into the only environment by which they could attain birth has been forbidden by the only persons in the world empowered to decide on entry into that environment. The argument from nonenablement is a consensus argument. It commands assent across a wide range of moral views. The argument from nonenablement differs from an appeal to imminent death as a means of justifying a killing. Nonenablement precludes a conceptus from attaining any of the attributes—autonomy, ability to feel pain, preferences, and, according to the traditional Thomist—Aristotelian teaching of Christianity, the attainment of a soul—whose infringement makes killing wrong. Nonenablement entails that there does not even correspond to an epidosembryo a possible person. It is not that death is imminent, but that development is bounded. Replies to Objections Let me now reply to some objections. The first objection asserts that the sacrifice of an embryo violates the second form of Kant’s categorical imperative, the precept that we should treat “humanity. . .always at the same time as an end, never merely as a means” (Kant 4: 429). In this formulation, by “humanity” Kant understands rational nature. For this reason, the objection misapprehends Kant’s view, which applies only to rational beings. The second form of the categorical imperative does not apply to mentally incompetent adults, infants, or embryos. That is not to say that Kant would take a cavalier view of the helpless. Kant would analyze our moral obligations toward these non-rational human beings as he would analyze moral questions in general. (“Human being,” by the way, is not a decisive moral classification; the phrase obviously applies to any being of the species Homo sapiens, including any somatic cell.) Kant would ask whether we could without contradiction of the will adopt as a universal law whatever maxim we propose about how we shall treat such a being. We do not contradict our will by adopting as a universal law that we shall use epidosembryos, at no cost in potential lives, to provide to those who suffer the aid that we would wish were we in their shoes. A second objection is the simple declaration that a zygote is a person. The most influential version of that objection appears in the official teaching issued within the latter part of the 20th century by the Sacred Congregation for the Doctrine of the Faith of the Roman Catholic Church. Its reasoning begins from the premise that all artificial methods of reproduction, including in vitro fertilization, are illicit. In adopting that stance, the magisterium anticipated a situation now before us—there exist surplus embryos whose creators have effectively consigned them, as the magisterium puts it, to an “absurd fate.” Such embryos will either perish as waste or be frozen indefinitely, never entering a womb. So when the magisterium condemns embryo

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destruction, it speaks consistently. It condemns destruction of embryos in research just as it condemns artificial reproductive practices that inexorably lead to destruction of embryos as waste. Others who approve of in vitro fertilization as presently practiced but oppose use of embryos in research fall into inconsistency. They implicitly approve the destruction of surplus embryos as waste while condemning the use of surplus embryos to help others. Although the official Catholic view cannot be accused of inconsistency in the foregoing respect, on what ground does that view rest? Why condemn the destruction of an unenabled embryo in research? As one studies the magisterial instructions and looks for arguments, one finds a single argument. The argument is not scriptural. The ancients did not understand fertilization or embryogenesis and were not thinking of embryos in petri dishes. The argument purports to be biological. The argument is that because fertilization creates a new genome, fertilization creates a person (Sacred Congregation for the Doctrine of the Faith 1987). This is an argument whose premise one must admit—the biology is correct, fertilization does produce a new genome—but whose conclusion does not follow. The argument presupposes a radical version of genetic reductionism. To say that a person is a genome is a view that even a materialist would not venture. It is a view contradicting the belief, held within the religious tradition from within which the argument is offered, that a person is a corpore et anima unus, a union of body and soul. I imagine that one reason that many people have not heard this argument is that it cannot be maintained consistently with the rest of Christian teaching. I suspect that eventually the argument will fade. Another interpretation of Catholicism might lay stress on the notion, also advanced by the magisterium, that because we do not know whether an embryo is a person, we should exercise caution and act as if it were a person. But suppose that we could have a conversation with God. We would report that in 1998, we discovered how to culture embryonic stem cells. We explain that we have plans to relieve human suffering by the use of embryos that will never enter a womb. Is it plausible that He would tell us that He regards embryos that will never enter a womb as persons in the sense that He includes them in the universe of beings that He wishes us not to consume? I do not know of a tenable argument according to which an all-merciful and omniscient God would assert that preference. For He would know that unenabled embryos will never develop into infants. He would know that our efforts to aid actual lives would exact no cost in potential lives. Under these circumstances, there inexorably come to bear Christian social teachings, including the duty to love thy neighbor as thyself and “the law of charity” (Sacred Congregation for the Doctrine of the Faith 1974). Cloning as a Special Case You may have heard some people say that the justification of embryonic stem cell research lies in the circumstance that the embryo donors initially intended procreation when they created the embryos now regarded as surplus. The argument from nonenablement does not rely on any assumption of initial procreative intent. Hence the argument from nonenablement justifies not only embryonic stem cell research, but also nonprocreative cloning. An objection peculiar to cloning might be this. An oocyte is created for a purpose, namely to issue in offspring, and it is wrong to divert an oocyte to any other purpose. This objection rests on an Aristotelian teleology that, since Darwin, does not exert much grip on our thought. We have learned from the history of medicine how mistaken we humans have often been in inferring purposes of various cells and structures of the human body. Our forbears would have said that bones are what hold us up; today we think of the marrow as a blood factory, and think it appropriate to transfer marrow from one patient to another. We have learned the remarkable adaptability of tissues and cells. We now actively engage in directing tissues and cells to serve chosen purposes in aid of sick patients, calling this treatment “conventional” drug therapy. It is unpersuasive to say that an oocyte can or should serve only one purpose.

Ethical Considerations

From a religious point of view—and the teleological objection that I have just described now finds its greatest support in religious traditions—the ultimate arbiter is divine will. Imagine again that we could have a conversation with God. Would He say that oocytes may serve only the purpose of procreation? Such a rule would seem harsh inasmuch as we should have to measure its effect in suffering unrelieved. It would also seem puzzling insofar as every human female possesses from birth a quarter million or more immature oocytes. That is all that I shall say here about the argument from nonenablement in general. I shall say a bit more about procreative cloning, the baggage of which those waiting to do nonprocreative cloning in research would like to rid themselves. About procreative cloning, we have developed a social consensus. We know from animal studies that the odds of producing healthy offspring in mammalian procreative cloning are so dismal that it would be irresponsible to attempt it in humans. We have heard a considerable stir about legislating that consensus. I should only like to point out that the need for legislation is illusory. In 1998, the Food and Drug Administration (FDA) dispatched a two-page letter to fertility clinicians and to others known to have an interest in procreative cloning. The letter declared that, first, if anyone wishes to attempt human procreative cloning, they must file an investigational drug application, and second, if anyone files such an application, the FDA will deny it. This letter evidently scared the daylights out of even mavericks. Those reported by the press to be attempting cloning have not been attempting it within the United States. In consequence of four-fold statutory authority buttressed by an FDA regulation that took effect January 21, 2004 (Guenin 2005a), there is no need for legislation duplicating the FDA’s interdiction. I mention this because concern abounds that legislators, in their zeal to appear vigilant to their constituents, will sweep valuable research within the maw of prohibitions. That would be a shame when the likely incidence of procreative cloning seems nil. We ought to think carefully about imagined situations in which people purportedly would resort to cloning as a means of having children. Among prospective parents usually imagined for the procedure are carriers of alleles for recessive diseases and those affected by them, infertile couples, homosexual couples, and others. For reasons that I present elsewhere (Guenin [forthcoming]), I surmise that cloning will seldom be their first choice. A more likely choice will be in vitro fertilization followed by prenatal genetic diagnosis as a screening technique. Germ line intervention would also draw interest when it is feasible. Perhaps someday we shall even see the use of artificial chromosomes. Cloning does not offer the flexibility, safety, and other advantages of such alternatives. Still there will arise objections to procreative cloning when the day comes, if it does, that the procedure is as safe as natural conception. One objection will be that the motivation of people who want to clone is a kind of unworthy narcissism. About this it must be said that if narcissism motivates some instances of asexual reproduction, it may motivate many more instances of natural reproduction. Perhaps we ought to shelve that objection. There also abounds the prediction that a clone will undergo an identity crisis. As the scenario goes, a clone’s knowledge that it shares the nuclear genome of a source will so burden the clone that the clone will not understand that he or she is a distinct person. I find this objection implausible. In a typical case, a clone will be about 25 years or more younger than the source, will grow up in a different environment, will meet different people, will have different experiences, and so forth. We also hear speculation that asexual reproduction will threaten the nuclear family, and even that copying genomes by cloning will adversely affect the gene pool. Both eventualities seem improbable, the latter highly improbable unless the incidence of cloning becomes very high. In any case, we can respond to all these speculations by saying that if the day arrives that cloning is safe, we should revisit the question of its propriety. At that time, anyone proposing a governmental prohibition of cloning will need to explain why, if every other method of reproduction is private, this one should not be.

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MORALITY AND THE SCOPE OF FUNDED RESEARCH It is possible to state a public policy such that the scope of publicly funded research using embryos is congruent with what is, according to the foregoing reasoning, morally permissible. That policy, suitable for adoption by any government, is the following: The government shall support biomedical research using human embryos that, before or after formation, have been donated to medicine under donor instructions forbidding intrauterine transfer (Guenin 2005b). By virtue of the congruence of policy and morality, this policy manifests its justification. That attribute avails its presentation in public debate. One move that does not avail in public discussion is to withhold “clone” and “embryo” from a product of nuclear transfer. Organismic cloning occurs when an activated oocyte assimilates a somatic cell’s nuclear DNA. This process is complete by the time that the activated oocyte divides. Ensuing development is called “embryogenesis.” The point worthy of emphasis is that no attempt at procreative cloning occurs unless the resultant embryo is transferred to a womb (Guenin, 2003, 2004a). The practice of somatic cell nuclear transfer in regenerative medicine is an instance of nonprocreative cloning. Echoing claims that regenerative medicine could succeed if it eschewed use of embryos and confined itself to stem cells in the developed human, a number of suggestions have been broached about purported nonembryonic sources of pluripotent stem cells. It has been suggested, usually without pausing to state a moral argument, that pursuit of these alternatives would be a moral improvement. I have argued that even if the imagined technical feats could each be accomplished, the supposed moral superiority of pursuing these alternatives is illusory insofar as they would use or produce embryos or require for their defense an argument that justifies use of donated embryos in general (Guenin 2005c). ETHICAL CONSTRAINTS The following presents some thoughts about minimizing the risk of abuses. I first mention how we learn about such risks. We did not learn about the risks of recombinant DNA research, germ line intervention, or embryonic stem cell research and cloning from moralists hovering about scientific laboratories. We learned about those risks from the scientists who made the enabling discoveries. They and their colleagues, in each case envisioning what might come of their discoveries, acted promptly to bring before the public the question whether it would be moral to proceed. We should give credit to these scientists for their well-tuned moral barometers. Embryo Use Only for Humanitarian Ends For the present task, we ought first to define the sorts of research in which donated embryos would be eligible subjects. We, most of us, I suspect, would be appalled if human embryos were used—as were animals in studies of which we learned not so long ago-in testing industrial chemicals or cosmetics. Hence, we may stipulate— and the concept of epidosembryo embeds this condition—that only in humanitarian work may one consume a human embryo. Only epidosembryos, in other words, may be subjects. Ectogenesis We would also sensibly prohibit ectogenesis. That is, we would prohibit the development of an embryo in the laboratory beyond some specified number of days. British law prohibits development of an embryo beyond day 14 (Human Fertilisation and Embryology Act 1990, ch. 37, §3[3] [a]). Some have taken the laying down of the primitive streak as the first stage in the development of a nervous system. But the nonindividuation objection has become the most influential reason for adopting day 14 as a boundary. The demise of the nonindividuation objection would explode that rationale. Still we shall have to set some boundary if we wish to

Ethical Considerations

preclude ectogenesis. As Bernard Williams once remarked, it is not uniquely reasonable that we draw a line at 14 days, but it is reasonable that we draw it (Williams 1986). Oocyte Donation There looms a risk of abuse in obtaining oocytes for use in nonprocreative cloning. Research might create an aggregate demand for substantial numbers of oocytes. This, it has been said, might lead to the exploitation of women. If asked, a circumspect scientist will say that science should use only the fruits of fully voluntary donations by women who want to help others. One could present a libertarian argument that a woman should be permitted to donate oocytes, and to receive compensation, as she wishes (Resnik 2001). An oocyte donor’s considerable time, inconvenience, and discomfort alone would seem to warrant compensation. But availability of compensation could induce some women in need of money to undergo hyperovulation and oocyte recovery to their physical or psychological detriment. Concern is also expressed about “commodification” of oocytes. Hence, paternalistic prohibitions are frequently proposed. A plausible balance between paternalism and respect for autonomy would allow for compensated transfers of oocytes, but only in a regulated market. Rather than categorically prohibit compensation (as recommended in National Research Council and The Institute of Medicine 2005), and thereby risk unfairness to the donor, a bound may be set on the amount of compensation. Such a bound has already become a professional standard for an oocyte donation by one fertility patient to another (American Society for Reproductive Medicine 2000, which recommends a bound of approximately $5,000). A compensation bound could serve to minimize the risk that women will unduly discount the risks of donation. Another source of oocyte donations may be women already undergoing fertility treatment. For them, selection of a recovered oocyte for donation imposes no additional physical burden. Were many such women to donate to research, the invisible hand effect could be a copious supply. One technological innovation that could increase the supply of oocytes available for research consists in the production of oocytes by induced differentiation of embryonic stem cells. Immunity of Parents and Children From Patent Infringement Claims We also have to decide whether to allow patents on embryos. A product of somatic cell nuclear transfer could plausibly be adjudged new, useful, and nonobvious, and not be a product of nature. In such case it would meet the criteria for award of a patent. But a patent on a method of forming a clone or the clone itself could render a human conception or birth assailable as a patent infringement. That would be a perverse result. The result already looms by dint of patents on DNA sequences (Guenin, 2000). In the case of both DNA and embryo patents, we can avoid untoward results by introducing a simple rule. Each country could amend its patent laws to provide as follows: No claim of infringement shall lie against a parent or child as such. This provision would leave commercial patent owners free to pursue each other on claims of infringement, but would render parents and children in their capacities as such immune from claims of infringement. Any new technology poses some risk of abuse. It would be misleading to suggest that the risks of abuse in the case of embryo use are insubstantial. No one could claim to know that. But it would be uncaring of us to neglect the humanitarian work of which we are capable within the protective constraints that we have the power to impose.

ACKNOWLEDGMENT This chapter is adapted from the author’s “The morality of unenabled embryo use—arguments that work and arguments that don’t,” Mayo Clinic Proceedings 79: 801—808 (2004), and reprinted with permission.

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REFERENCES American Society for Reproductive Medicine (2000). Financial incentives in recruitment of oocyte donors. Fertil. Steril. 74: 216–220. Guenin, L.M. (2000). Ethics of patents on human life forms. In: Murray, T.H. and Mehlman, M.J. (eds.), Encyclopedia of Ethical, Legal, and Policy Issues in Biotechnology. New York: Wiley. _______ (2001). Morals and primordials. Science 292: 1659–1660. _______ (2003). The set of embryo subjects. Nat. Biotechnol. 21: 482–483. _______ (2004a). On classifying the developing organism. Conn. Law Rev. 36: 1115–1131. _______ (2004b). A failed noncomplicity scheme. Stem Cell Dev. 13: 456–459. _______ (2005a). Stem cells, cloning, and regulation. Mayo Clin. Proc. 80: 241–250. _______ (2005b). A proposed stem cell research policy. Stem Cells 23: 1023–1027. _______ (2005c). Wishful thinking will not obviate embryo use. Stem Cell Rev. 1: 309–315. _______ (2006). The nonindividuation argument against zygotic personhood. Philosophy 81: 463–503. _______ ([forthcoming]). The Morality of Embryo Use. Cambridge: Cambridge University Press. Hare, R.M. (1993). Essays on Bioethics. Oxford: Clarendon Press. Kant, I. Groundwork of the Metaphysis of Morals. In: Gregor, M.J. (1996), (trans.), The Cambridge Edition of the Works of Immanuel Kant, Practical Philosophy. Cambridge: Cambridge University Press. National Research Council and The Institute of Medicine (2005). Guidelines for Human Embryonic Stem Cell Research. Washington, DC: National Academies Press. Oderberg, D.S. (1997). Modal properties, moral status, and identity. Philos. Publ. Aff. 26: 259–298. Rawls, J. (1993). Political Liberalism. New York: Columbia University Press. Resnik, D. (2001). Regulating the market for human eggs. Bioethics 15: 1–25. Sacred Congregation for the Doctrine of the Faith (1974). Declarato de Abortu Procurato. Vatican City. Sacred Congregation for the Doctrine of the Faith (1987). Donum Vitae. Vatican City. Williams, B. (1986). Types of moral argument against embryo research. In: The Ciba Foundation, Bock, G. and O’Connor, M. (eds.), Human Embryo Research: Yes or No? London: Tavistock.

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79 To Make is to Know: The Ethical Issues in Human Tissue Engineering Laurie Zoloth

CASE STUDY The small black box holds a perfectly shaped ear. The scientist at the front of the room explains how it was made. There is a scaffold of nanoparticles that support fibroblast cells that grew over the form, and the ear now looks and feels actual: it can be transplanted to tissue. For burn patients, this represents an enormous chance and change. It is the prototype of a new genre of medicine, one that uses powerful technologies and methods of bioengineering and cellular biology to transform the matter of the world. The ear in the box is not a freak example of a new technique. It is, in fact, one of a number of new devices that utilize the convergent technologies of several different fields of science and engineering to create tissue that can mimic the structure and function of the natural world. Other examples include the creation of skin grafts, corneas, bone, cartilage, and in some pilot studies, bladders. Based on the new technologies of genomics, informatics, nanoscale engineering, molecular biology, and stem cell research, tissue engineering can be said to alter the concept of medicine itself. Instead of treating ailing tissues or organs with drugs intended to repair their structure or function, tissue engineering aims at replacing the diseased or injured parts of aging tissues of the body with entirely new ones, made from component parts of materials in the world, both naturally occurring and synthetic. Such an advance heralds a remarkable ability to heal, a long-awaited solution to several intractable problems, and a serious alternative to cadaveric or living donor whole organ transplants, which have long been an ethically challenged sector of medicine.1 Yet such a remarkable construction of the human body asks a great deal of any social world into which it is introduced, for it is the body that is the place of the self, the location of the acts of the sacred, and the sensory arbiter of the real. In fact, tissue engineering queries two of the very aspects of our humanity that we consider distinctive: our integral embodiment and our finitude. If we are indeed a collection of replaceable and adaptable parts, some reason, what is it that separates us from any other engineered machine? If we can engineer, for example, a synthetic and improved lymphatic system, might we improve our chances to adapt to and overcome infections disease? What other capacities for healing or alteration of our bodies might be prudent? How do we ensure that such changes are indeed ethical? It is this query which has greeted the new biotechnologies of the body, one based, this chapter will argue, in social reactions largely shaped by culture both ancient and contemporary. We will then ask: what are the ethical challenges to the field of tissue engineering? Does tissue engineering raise new ethical issues, or is it a 1

Transplantation Ethics, Caplan, et al.; also, Fox, Renee, “Leaving the field”.

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description of one of the modalities enabled by the convergence of other technologies that have been understood to be individually ethically freighted? In this chapter, we will use an established ethical framework that was suggested in 1999 by a committee of the American Association for the Advancement of Science on inheritable genetic germline modification, and was used far more widely in the field of bioethics to assess new technologies. We will then review the responses given to new technologies in the past from a variety of sources in bioethics, philosophy, and theology, and finally, we will reflect on how the legal and regulatory structure for tissue engineering has an impact on our reflections on ethical norms.2

Research Evaluation Are there reasons in principle why performing the basic research should be impermissible? What contextual factors should be taken into account and do any of these prevent development and use of the research? What purposes, techniques, or applications would be permissible and under what circumstances? What procedures, structures, involving what policies, should be used to decide on appropriate techniques and uses? Adapted from the AAAS Working Group on Human Inheritable Genetic Modifications 1998–2000.3

ARE THERE REASONS IN PRINCIPLE WHY PERFORMING THE BASIC RESEARCH SHOULD BE IMPERMISSIBLE? Principled reasons for objections to basic research are extremely difficult to conceive in research that is, by its very nature, intended to be translational and clinical. Yet, ethical objections to the manipulation, replacement, and engineering of human tissue can be seen as part of a long continuum of dissent about medical technology that began to assume full voice in the 1970s when successful genetic manipulation of bacterial genomes became possible.4 All new technology raises new challenges – in particular, technology that refashions the embodied self, becomes a part of the “self” and the identity of the subject, and seems to raise the deepest anxieties. Even tissue engineering, an emerging field with clear targets, clinical successes, and patient needs will raise familiar concerns. First among these is the argument that humans possess an essential nature and live within an essential natural order that cannot be altered without harm. For C.S. Lewis, this is expressed as a concern that the very acts of rational science: dissection, analysis, and quantification are a violation of the sacred integrity that lies behind all of nature: “Now I take it that when we understand a thing analytically, and then dominate and use it for our own convenience, we reduce it to the level of ‘nature’, we suspend our judgments of value about it, ignore its final cause (if any) and treat it in terms of quantity. This repression of elements in what would otherwise be our total reaction to it is sometimes very noticeable and even painful: something has to be overcome before we can cut up a dead man or a live animal in a dissecting room.”5 2

Chapman, et al., Designing Our Descendants. Ibid. 4 Walters, LeRoy, Enhancing Human Traits. 5 Lewis, C.S “The abolition of man,” On Moral Medicine, edited by Lammers and Vehey, p. 274. 3

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For Lewis, the understanding of the body as replaceable is disturbing: “The real objection,” he says,“is that if man chooses to treat himself as raw material, raw material he will be, not raw material to be manipulated by himself as he fondly imagined, but by mere appetite.”6 (Lewis imagines that new transformative technology will be manipulated by “controllers” who will eventually transform man into mere matter.) Callahan echoes his concern, both in the sense that limits need to be placed on what is decent to do to nature, and in the sense that such action is part of a larger danger – that power in the hands of medicine to heal is really power in the hands of the elite, or the state to manipulate and control. He argues: “The word No perfectly sums up what I mean by a limit – a boundary point beyond which one should no go … There at least two reasons why a science of technological limits is needed. First, limits need to be set to the boundless hopes and expectations, constantly escalating, which technology has engendered. Advanced technology has promised transcendence of the human condition. That is a false promise, incapable of fulfillment … Second … limits (are) necessary in order that the social pathologies resulting from technologies can be controlled … while it can and does care, save, and free, it can also become the vehicle for the introduction of new repressions in society.”7 These objections, made over 30 years ago, are still made (despite, one may note with some irony, that 30 years of medicine have indeed seen rapid and successful advances, without them being used by the state for repression and without any fundamental change in the capacities for intellectual and spiritual self-possession). Nevertheless, powerful arguments of opposition to the manipulation and replacement of tissues and organs continue, with some worried that perfection itself is sought when healing is the goal. Such critics, many from the disability community, raise principled objections to the use of tissue engineering if the goal is to alter the disability. Activists in the Deaf community, for example, defend their disability as a culture and a language exchange, not as a loss of function. Others are concerned that our societies’ focus on “fixing things” will allow a devaluation of the persons that currently bear the broken bodies and parts. Adrienne Asche suggests that there is a “troubling side to every cure, that those of us who are uncured are seen as less valuable, perhaps even expendible.” For Gerry McKinney, a community needs to embrace brokenness, and to “deny that the worth of one’s life is determined by how closely one conforms to societal standards of bodily perfection.” McKinney is also concerned that if medicine is successful, it will create a social and economic system that “virtually demands that we be independent of the need to care for others. …”8

DUTY AND HEALING: NATURAL MAKERS IN A BROKEN WORLD While the opposition to medical technology has indeed been persistent, it has not been unchallenged. For many, the response lies in the nature of brokenness and the human duty to respond to the need of the suffering other.9,10 The principle at stake in the assessment of tissue engineering as an ethical act is not how its use might potentially violate an abstract community in the future, but the actual problem of what one must do as a moral being when one’s neighbor is in need? In this important sense, the duty to heal cannot be overridden by a “sense” of discomfort (as Lewis notes in the quote above). It is the nature, goal, and meaning of science to address the human condition in all its yearning and capacity for defeat and failure. In this sense, there is no principle objection to the science of tissue engineering, and in fact, there may well be a strong moral imperative to develop the technology. 6

Lewis, ibid. Callahan, Daniel, “Science, limits and prohibitions,” The Hastings Center Report, Vol. 3, No. 5, November 1973, pp. 5–7. 8 McKinney, Gerald P., “Bioethics, the body, and the legacy of Bacon,” To Relieve the Human Condition: Bioethics, Technology and the Body, Albany, NY: State University of New York Press. 9 Freedman, Benjamin. The Duty of Healing, Routledge, New York, 1998. 10 Zoloth, Laurie, “Difficult freedoms,” God and the Human Embryo. 7

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Noted Joseph Fletcher, in speaking of an earlier generation of medical technology, and answering the critics of science: “The belief that God is at work directly or indirectly in all natural phenomena is a form of animism or simple pantheism. If we took it really seriously, all science, including medicine would die away because we would be afraid to “dissect God” or tamper with His activity …”11 “Every widening and deepening of our knowledge of reality and of our control of its forces are the ingredients of both freedom and responsibility.”12 McKinney, Childress, and Lewis use an argument that is rooted in Christian moral theology: that since human persons are fallen creatures in a fallen world, we cannot really be counted upon to know the right and the good. God, who is transcendent from this fallenness, has set us in this place, not essentially to alter it toward our own transonic, but to find its meaning and purpose. Yet, other faith traditions differ. For Jewish and Islamic theorists, the world is morally neutral. Humans may – and will – fail in their aspirations, but can be trusted to have the capacity for moral behavior and moral yearnings. Finding meaning in suffering is not the core task – the task is to alleviate suffering, which is understood as chaotic, meaningless, and agonistic. Hence, many of the core objects in principle are rooted in religious constructions and understandings.

TO MAKE IS TO KNOW: NOTES ON AN OLD PROBLEM ABOUT KNOWLEDGE The classic debates of the 1970s are not the only set of problems engendered in the history of ethical responses to the technological gesture at the heart of tissue engineering – at stake as well is the special kind of knowledge that such a making implies. For Aristotle, and the Hellenists, useful knowledge, “practical wisdom,” was phronesis. Phronesis implied actually doing an act, making, in order to know. The act of making, not the act of perception or contemplation alone, what-how wisdom, and indeed, rationality and power, were achieved. Hence, making new tissue is a somewhat different moral gesture than curing the body by altering it with drugs that essentially allow the body to heal itself. Secondly, the use of technology within the body of the patient is a different matter than the use of technology to essentially enhance the body of the practitioner. For all earlier technology, the thing that was changed or enhanced was the sense perceptions of the doctor. Stethoscopes and otoscopes allow the sounds of the body to be more audible. X-rays, CT scans, and MRIs allow the inner vistas of the body to be revealed, EEG and EKG, allow the electrical currents that animate the central and peripheral nervous system to be charted in quantifiable units. Microscopes allow the invasive bacteria to be seen at the microscopic and, increasingly, molecular level. These earlier technologies extended the reach of what Bacon increasingly trusted, and that the Greeks did not – the perception and observation of the phenomena of the world and the perception of the outcome of its deliberate perturbation: “Bacon’s method presupposes a double empirical and rational starting-point. True knowledge is acquired if we proceed from lower certainty to higher liberty and from lower liberty to higher certainty. The rule of certainty and liberty in Bacon converges … For Bacon, making is knowing and knowing is making (cf. Bacon IV [1901], 109–110). Following the maxim “command nature … by obeying her’’ (Sessions, 1999, 136; cf. Gaukroger, 2001, 139 ff.), the exclusion of superstition, imposture, error, and confusion are obligatory. Bacon introduces variations into “the maker’s knowledge tradition” when the discovery of the forms of a given nature provide him with the task of developing his method for acquiring factual and proven knowledge.’13 Thus, the world is known by understanding the parts of the world and from that, theorizing (knowing) by induction to principles or axioms or laws of nature, physics, and chemistry. In contemporary science, 11

Fletcher, Joseph, “Technological devices in medical care,” Who Shall Lie, edited by Kenneth Vaux, 1970, Fortress Press. Ibid. 13 Stanford Encyclopedia of Philosophy, http://plato.stanford.edu/entries/francis-bacon/#5. 12

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knowing is done largely by “unmaking” – by the deconstruction of the component parts in a way scientists of Bacon’s era were unable to imagine. Many of these “unmaking” techniques, such as the splicing of alternative DNA, or the manipulation of cellular structures, allow a sense of inherent interchangeability, as if the real and the person were merely a set of Lego parts, awaiting clever recombination.

WHAT IS A THING: THE PERILS OF DECONSTRUCTION? Making actual tissue in mimesis of the real tissue of the actual body extends the Baconian act in radical ways. Here, the experimental perturbation is the unmaking of tissue and the remaking of tissue, only in a more controllable form. This cannot help but excite concern about the nearly infinite possibilities for technological shaping of the self. Heidegger asks: “what is a thing?” and in so reflecting, understands a thing as an object separate from the self. But what of a made thing, an object that becomes the self? The technology of the alteration of the patient is distinctive. Devices for altering the functioning of the body that become a part of the body, and are actually a tissue of the body, are a step beyond the idea of a device held within the body. Why this is important in any ethical assessment of the technology is that the patient’s consent and participation is needed for the final act of the technology to be completed. Such an event only happens in a specific context, for technologies, patients, and practioners operate in a social, religious, and economic contexts. Here we turn to the second ethical consideration. WHAT CONTEXTUAL FACTORS SURROUNDING THE TECHNOLOGY HAVE TO BE TAKEN INTO ACCOUNT AND DO ANY OF THESE PREVENT THE DEVELOPMENT AND USE OF THE TECHNOLOGY? Tissue engineering is a complex procedure still in the experimental stages. Yet, to be an ethical technology, it must be directed toward accessibility, just distribution, and efficacy. Hence, the troubling context of widespread healthcare disparity is a problem not only for this advanced technology, but for all newly emerging technologies. Emergence into an unjust world asks certain moral questions of new technological advances. First among these is the query about burdensomeness versus benefit in a context in which the vast majority of the world’s people suffer from easily treatable infectious diseases, tuberculosis, malaria, AIDS, and infant diarrhea. How can tissue engineering be justly promoted in the face of other, pressing needs? This objection can be typically met by noting that it would be deeply inappropriate to withhold medical knowledge until the world is entirely perfected, and that applications will not only be increasingly available to the poor (as in vaccines, once rare, now increasingly available) but also that the very process of research will typically uncover new and useful ways to understand disease. The goal of tissue engineering is the widespread use of the technique. Unlike solid organ transplants, which would always require significant resources far outside the capacities of developing world clinics, the use of tissue replacement, stem cell therapies, and other transportable therapies are designed for widespread use. The possibility to create a method for allografts, that uses the patient’s own cells, and the possibility for allologous cells to allow for an “off the shelf ” source of tissue may allow for the basis for access – but only if research priorities are discussed in advance of the design process – a process, as we will describe later, that will need careful support and monitoring. The question of how to achieve this and how to enable a more just use of each technology has not yet been solved. The second contextual factor for tissue engineering is that all human tissue is marked by its genomic identity. It is the very nature of cells, that thing which allows them to copy and reproduced, to carry identifying markers linked to some person somewhere. In the past, such use has raised serious objections. Such tissue can be traced and known, which may have implications for the person who is the source of the tissue, raising significant new issues in genetic privacy for the donor.

The Ethical Issues in Human Tissue Engineering

Further, whose is the tissue that is derived from the cells of a particular body? Who should have the rights to, and a fair share of, the profits derived from its use? In the seminal case in the field, Moore versus the Regents of the University of California (51 Cal 3d 120 271 Cal Reporter, 146, P2d 479, 1990), the issue of ownership was addressed. In this case, Mr. Moore had his T-cell lymphocytes taken from his spleen, during the course of his treatment for hairy-cell leukemia, cells that proved effective in deriving resistant cell cultures. Patented after manipulation to make a new “product”, the cells were indeed profitable. Mr. Moore’s complaint was that he was not informed of, much less, a part of the scientific enterprise and the lucrative payout for his cells. The case was decided in favor of the research laboratories, but in the insuring decades, alert patients with unique cell types or unusual cancers sought for research are selling their materials as personal possessions to the laboratory that wishes to procure them. Ownership is limited, however, by the constraints of the common law of the US and the EU which limits the ability to claim tissue as property. The goals of such restraints were put into place to prohibit the buying and selling of human tissue and organs for fear that, given the desperation of the poor, selling the bodies of the poor would become permissible and lead to their exploitation. Thus, the entire process that allows for the derivation of tissue sources needs to be noted. The current context for tissue donation is a mixed system. Organs, tissues such as blood, corneas, and marrow are donated or exchanged without compensation. Gametes, however, are another matter entirely. Because the use of human sperm and eggs emerged in the context of fertility treatment and because this treatment was largely conducted in stand-alone, private clinics that functioned without public oversight or regulation, the marketplace standards prevailed. What originally began as a compassionate exchange between family members of gametes when an infertile couple could not conceive, quickly changed into a robust marketplace in human gametes. As of this writing, the international standards prohibit the use of marketplace incentives for gametes, or embryos.14 The final context for the debate about the ethics of tissue engineering in general is the special case of human stem cells to make tissues. Because some applications of tissue engineering use stem cells as a part of the method of treatment (Egan, 2006), the debate about the ethics of the use of human stem cells is directly adjacent to this technology. For many, the origins of tissue matter a great deal. For some Christians, many Roman Catholics, and some Hindu sects, the destruction of the human embryo, even at the blastocyst stage, is tantamount to killing. For these faith traditions, the derivation of stem cells from embryos is always impermissible. For many other faith traditions, such as Judaism, Islam, Jainism, Buddhism, Confucianism, and Daoism, the use of these cells is permissible within certain constraints, as we will see below. For all faith traditions, however, the manipulation of adult somatic cells, in their precursor form, is completely sanctioned. Precursor cells are not as flexible as pluripotent cells, and it is that very pluripotency and immortality that are important in tissue engineering. These factors raise concern. Yet the contextual factors alone do not entirely prohibit the use of this technology, for justice in distribution, the possibility of the loss of genetic privacy, and the controversy over stem cell research when pluripotent embryonic cells are used affects many aspects of the new techniques in medicine. Hence, we turn to the third major issue.

WHAT PURPOSES, TECHNIQUES, OR APPLICATIONS WOULD BE PERMISSIBLE AND UNDER WHAT CIRCUMSTANCES? Many of the salient, justifying arguments for the use of tissue engineering hinge on the telos, or goal, of the treatment: if the goal is to cure or treat human disease, then the benefits will outweigh the burdens of the work – controversy, cost, and difficulty. Clearly, then, tissue engineering ought not to be used in a trivial or wasteful fashion. Human tissue is understood by many as deserving a special sort of “respect.”15 14 15

International Society for Stem Cell Research, Draft Guidelines, 2006, www. Isscr.org/ task force report. Geron Ethics Advisory Board, “The ethics of human stem cell research,” The Hastings Center Report, January 1999; see also Nelson, Larry, “The elusive nature of respect,” The Hastings Center Report, 2002.

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This proviso may not be so simple, for a core problem in genetic engineering has been the use of the technique for “enhancement” of human characteristics or traits. The initial ethical discussions about therapeutic uses of medicine versus cosmetic ones imagined ethical bright lines that would define the boundary between the use of such technology to restore “species normal functions”16 for each tissue and for the person as a whole. Yet, medical practice has long gone beyond these lines, using surgery, for example, for cosmetic purpose. Will it be possible to restrict tissue replacement to burn victims, spinal cord injury victims, and diabetics? How can such a distinction be made? Some tissue replacement therapies, such as the use of skin grafts for full face transplant, may also raise questions about the nature of identity. Indeed, the notion of a full face transplant alerted us to the depth of resistance to identity altering tissue replacements. (Could persons use any face? What if persons in need of facial transplants wished to change ethnicity? Should faces “match” and why?) Like many other aspects of this technology, this tension about identity was not new, only heightened. For example, the first years of organ transplant raised the same issues for recipients of hearts – a key aspect of identity in many cultures. If the face is the key determinant of the self in modernity, and even more so, if the brain is such, then how are we to understand the use of tissue engineering to transform identity? Hence, linked inexorably to this technology are larger considerations of the use of tissue engineering for neuroscience – both for therapy and for enhancement. The applications of tissue transplant in Parkinson’s are important – yet will there be concern about this use of the neurons of a stranger in the brain of the self? Of all the possible uses of tissue engineering, the ones that may alter consciousness and memory are the most troubling (what capacities or memories could neurons store?). Here, the need for restrictions on applications may be the clearest, yet it is not clear who ought to decide and who ought to ensure that the restrictions on unethical applications are maintained. By what criteria will such limits be set? New research possibilities also offer applications to engineer gametes for use and storage. Engineered follicles may now be saved, frozen, matured, and used in animal models to create the possibility of human fertility after cancer chemotherapy, or other environmental risk.17 With this, as with all such technology, there will have to be careful attention to how the market may drive technology toward specific research goals rather than others, or that research goals will be framed only by the values of profit and efficacy, and not ones of more general interest, compassion, healing, and solidarity. The powerful applications, and the potential for widespread use itself creates the possibility for serious conflicts of interest, as serious market forces may be the core drivers of technology, especially in an aging population with increasing needs for all manner of new tissues and organs. This turns us toward the consideration of our final set of issues.

WHAT PROCEDURES, STRUCTURES, INVOLVING WHAT POLICIES, SHOULD BE USED TO DECIDE ON APPROPRIATE TECHNIQUES AND USES? Much of the first reviews of the ethical issues in tissue engineering have in fact focused on the issues of policy – safety, patents, and gating. Products and drugs are typically controlled via four levels of restraints. The first is elaborate pre-market gating, first involving animal models, then typically done for pharmaceuticals in a decade-long series of tests, phased to test the drug on an increasing, but controllable number of human subjects. Such subjects must be gender balanced, must have full informed consent, and must be able to leave the trial at any time (which may be difficult for implanted tissues). The next gating is the system of intellectual property. Patents and licensing control the use of the products, even the replication of the experiments. The next gating is that of financial backing, To do the enormous 16

Used first by Norman Daniels in the debates about health care justice and the reasonable amount of medical care a person would be entitled to. See Daniels, Norman. 17 Woodruff, et al. Mouse follicles matured after tissue,” 2006.

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clinical trials, to do pre-market investigation, and of course, to actually make and sell the product requires a productive apparatus, which must be assembled and supported. Finally, each drug or device must be approved for use by the insurers.18 As David Smith notes, tissue engineering faces a gamut of issues and a “new order of magnitude in interactions and science patents.” Additionally, notes Smith, the “things” engineered are hybrids of two jurisdictions – that of drugs and that of devices. Are genetically engineered insulin cells a drug like insulin, a device like a stent, or a biologic? Unlike stents, which are entirely synthetic, tissue engineering uses actual human cells – only manipulated in de novo ways. Standards will need to be set for safety, efficacy, and fair use – standard for clinical use, standards for clinical trial, and standards for tissue stability and purity will be needed for the research and application to be safe. Getting informed consent in this case will present significant challenges. Patients in need of organs, for example, are particularly desperate and their consent may be deeply affected by their utter lack of options. Of the medical system, 8% is already devoted to organ transplantation and the lack of organs is an overwhelming problem for nearly half of the patients hoping for transplants.19 Yet, the first year of the use of engineered tissue will be experimental and will need to be conducted under the strongest possible set of National Institutes of Health (NIH) guidelines. How the first trials of engineered tissue are conducted will set the tone and the future for all subsequent use. The question of policy and the regulation of policy are manifested in many of the first documents that evaluate the ethical and legal implications of tissue engineering. While, as Smith notes, the US faces a complex regulatory system. The European Union has regulated such research products as medical products, and these will fall under the regulatory gaze of the European Medical Evaluation Agency (EMEA). In both , the synthetic nature of tissue engineering, the very de novo quality of the work and the uneasy greeting that met genetically modified food has created serious political opposition. Policies need to be crafted with transparency and full public participation for such research needs not only public funding but public understanding of the complex theory and practice of tissue engineering – what promises it can hold, and what cautions need to be applied prior to use. Policy makers will need to attend to calls for justice in distribution, as was noted above, and will need to set in place structures for regulation. How can new technologies best be regulated? I would contend that a full array of regulatory structures can be employed. First among these are local committees, IRBs and local review boards. The National Academies have played a large role in policy writing for both recombinant DNA and for stem cell research, and, in both instances, called for special, national ongoing oversight on such research. It would be prudent to reflect on the need for such a process for tissue engineering, for established structures largely address issues involving the use of donated tissue, not engineered tissue. Structures that protect human subjects also need strong enforcement, as noted above, both for the donors and the recipients of tissues. But regulation, government oversight, and market forces can only go so far in shaping just research goals and commitments. The goal of ethics is to develop moral agents who are aware of a constancy of duty toward subjects and to humanity, who not only follow rules correctly, but who, given the chance and grace to work at the frontiers of science, act with courage and decency in their research.

CONCLUSION Tissue engineering suggests that an old dream – the replacement of human body parts – may be realized. While any sober and reflective scientist understand the long way to success of this idea, the science described in this volume clearly suggests that our society is on the road to the enactment of the possibility. 18 19

Smith, David, “Legal and regulatory issues in tissue engineering.” National Science Foundation Report on Tissue Engineering, 2004. Lysaught, M.J. and O’ Leagh, J.A., The growth of tissue engineering. Tissue Eng. 7(15): 485–493.

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80 US Stem Cell Research Policy Josephine Johnston

INTRODUCTION Since James A. Thomson and colleagues reported the isolation of pluripotent stem cells from human embryos in 1998 (Thomson et al., 1998), stem cell research has seldom been out of US headlines. Because isolating embryonic stem cells involves destroying embryos, some groups and individuals have opposed some or all of the research on moral grounds. This opposition has been translated into a number of policies and laws, including at a federal level. As a result, much embryonic stem cell research is currently ineligible for federal funding. Nevertheless, embryonic stem cell research is progressing in the United States using monies supplied by individual donors, charitable organizations, and states. After briefly discussing ethical and policy issues in adult and fetal stem cell research, this chapter will survey the current policy issues in embryonic stem cell research, beginning with the 2001 federal funding policy, which will be compared to regulation strategies adopted in other nations active in the research, before moving on to consider oversight, donor, and commercialization issues that arise as the research moves forward. SOURCES OF STEM CELLS Stem cells are special kinds of cells that can regenerate themselves and make new, more specialized cells. For the purposes of this discussion, stem cells can be divided into three kinds based on the source of the cells: adult stem cells, fetal stem cells, and embryonic stem cells. In terms of ethics, politics, policy, and law, much depends on the source of the cells. Adult stem cells are derived from the cells of adults and children. Although this source is not without its ethical issues, they are similar to those raised in other human subject research, the most important of which is the requirement for free and informed consent (President’s Council on Bioethics (PCB), 2004). Because the ability of competent adults to consent to research enjoys wide acceptance, adult stem cell research has not been a major focus of ethical or political debate. It does, however, enter public consciousness as a lesscontroversial alternative to embryonic stem cell research. For example, Catherine Verfaillie and colleagues announced in 2002 that they had isolated multipotent adult progenitor cells from bone marrow (Schwartz et al., 2002). Opponents of embryonic stem cell research have cited research like Verfaillie’s as evidence that research using embryonic cells is unnecessary (Orr, 2002). Verfaillie, however, along with many other adult stem cell scientists, insists on the importance of pursuing research on both kinds of cell (Verfaillie et al., 2002). Pluripotent stem cells have also been extracted from the primordial reproductive tissue of aborted fetuses (Shamblott et al., 1999). Any source of stem cells that relies on women undergoing elective terminations is likely to be controversial in the United States simply because of the ongoing debate over the morality and

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legality of abortion. Nevertheless, researchers have used tissue from aborted fetuses since as early as the 1930s and federal funds are currently available for this kind of stem cell research. It is unclear whether pre-existing federal regulations and legislation apply to fetal stem cell research. Fetal research has been regulated in the United States since allegations of experiments on in and ex utero fetuses emerged in the early 1970s (NBAC, 1999), around the same time as the Supreme Court’s famous decision in the abortion case of Roe v. Wade (1973). In response to the allegations, human subject regulations were promulgated in 1975, requiring extra protections where federally funded research involves pregnant women, fetuses, and human in vitro fertilization.Among other things, these regulations forbid researchers from having any involvement in the Woman’s decision to terminate their pregnancy or from offering them financial inducements (45 CFR §46.204(h–i)). Although it has always been clear that these regulations applied to in utero fetal research, there has been some uncertainly as to whether the regulations apply also to research that, like stem cell research, uses cadaveric fetal tissue (NBAC, 1999). In March 2002, the Office for Human Research Protections issued a guidance for research involving fetal stem cells in which it states that the research will only be subject to federal human subject protections where it involves “a living individual,” which would seem to exclude most fetal stem cell research (OHRP, 2001).1 Likewise, the 1993 federal legislation on fetal tissue transplantation research likely does not apply to stem cell research using fetal cells unless the research also involves transplanting the cells into humans. Similar to the restrictions imposed by the federal human subject regulations, this legislation stipulates that no alteration in the timing, method, or procedures used to terminate the pregnancy be made solely for the purposes of obtaining the tissue (498A of the Public Health Service Act). Even though both the federal regulations and the fetal tissue transplantation legislation likely do not apply to most fetal stem cell research, practices similar to those required by these laws will likely be conditions for Institutional Review Board (IRB) approval of much fetal stem cell research. Various states also have laws affecting fetal stem cell research, which will apply to all researchers in those states regardless of their funding source. For example, six states ban research involving aborted fetuses2 and some states ban paying for fetal remains.3 Despite fetal stem cell research’s intimate connection with the controversial practice of induced abortion, this kind of stem cell research has seldom been the subject of public debate. Instead, debate has consistently focused on research that uses stem cells extracted from 6 to 7 day old human embryos. Because any one cell in the very early human embryo can develop into a whole fetus, it is thought that embryonic stem cells will one day be able to repair damaged or diseased parts of the human body. The therapeutic potential of embryonic stem cells, therefore, is thought to be enormous, but so is the moral peril. Extracting stem cells generally necessitates destroying the embryo.4 For this reason, the research has been vigorously opposed by many individuals and groups, including (but not limited to) those who consider the early embryo to be a person or, if not a full person, an entity of sufficient moral significance that it should not be created for, or destroyed in, research. This opposition to destroying embryos in stem cell research is reflected in the current US policy on the federal funding of embryonic stem cell research.

EMBRYONIC STEM CELL RESEARCH: CURRENT US LAW AND POLICY At the time of writing, US regulation of embryonic stem cell research is limited to the federal funding policy as declared by President Bush on August 9, 2001 and any applicable state laws (Andrews, 2004). The 2001 policy 1

See also 45 CFR §46.206. Arizona, Indiana, Ohio, Oklahoma, South Dakota, and North Dakota. 3 See, for example, Tennessee and Arkansas, as well as Section 10(a) and (b) of the Uniform Anatomical Gift Act. 4 Although a method of extraction has been proposed where one cell is extracted by biopsy, leaving an embryo that could theoretically implant and develop just as a biopsied embryo does following pre-implantation genetic diagnosis (Chung et al., 2006). 2

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establishes the criteria for embryonic stem cell research that uses federal funds – it does not ban embryonic stem cell research altogether because it does not apply to research conducted using non-federal monies, such as might be supplied by an individual donor, a company, a charitable organization or a state government. Nevertheless, because much basic biomedical research is traditionally funded by the federal government, the policy has received considerable attention. The 2001 policy declares that federal funds can only be used to study embryonic stem cells that were already in existence as of 9 pm eastern standard time on August 9, 2001. In addition, the cells must have been extracted from embryos that were created for reproductive purposes but were not needed for those purposes (sometimes referred to as spare, surplus, or left-over embryos) and were donated with the informed consent of the donors, to whom no financial inducements may have been offered (The White House, 2001). To facilitate research on existing embryonic stem cells that met these criteria, the National Institutes of Health (NIH) set up a stem cell registry shortly after the policy was announced, in which the NIH lists embryonic stem cell lines eligible for use in federally funded research and available for shipping. In his 2001 address, the President stated that more than 60 embryonic stem cell lines met his criteria, but for a variety of reasons only 22 lines are currently listed in the registry as available for shipping. The 2001 policy announced by President Bush superceded a policy announced by President Clinton a year earlier, in which he permitted the use of federal funds to study embryonic stem cells subject to a number of conditions, including that the cells had been extracted without using federal monies (PCB, 2004). That is, under President Clinton’s policy, which was not able to be implemented before his presidency ended, federal money could be used to study embryonic stem cells, but not to extract them from embryos (PCB, 2004). In opposing President Clinton’s policy, some members of Congress complained that although it did not pay for embryo destruction, it nevertheless encouraged research that required the destruction of human embryos (PCB, 2004). In formulating their respective policies, presidents Clinton and Bush were both bound by a long-standing policy against federal funding of any research involving the destruction of human embryos (PCB, 2004). Since 1996, this policy has appeared as a provision in the annual allocation acts for the Departments of Labor, Education, and Health and Human Services. Until Congress removes this provision, known as the Dickey Amendment, from the annual allocation acts, or until embryonic stem cells can be extracted without destroying or harming the embryo (Chung et al., 2006), federal funding cannot be made available for the critical extraction step in the research. In his remarks on the 2001 policy, President Bush explicitly referred both to scientists’ beliefs in the enormous therapeutic potential of embryonic stem cell research and to his own belief in the value of embryonic life. Of embryonic stem cell research, he noted: “At its core, this issue forces us to confront fundamental questions about the beginnings of life and the ends of science. It lies at a difficult moral intersection, juxtaposing the need to protect life in all its phases with the prospect of saving and improving life in all its stages.” He called himself “a strong supporter of science and technology” but also noted that he believes “that human life is a sacred gift from our Creator” (Bush, 2001). If the President intended his policy as a compromise between the value of scientific research and the value of embryonic life, it was one that left many scientists, disease groups, and others unsatisfied. Substantial and ongoing criticism has been directed at the 2001 policy. Those who oppose any research in which human embryos are destroyed have argued that federal funding restrictions need to be supplemented by a nationwide ban on creating embryos for use in research, including by cloning (NRLC, 2001; PCB, 2004). A more vocal opposition has called the 2001 policy overly restrictive. Their arguments have included that the President over-valued embryonic life (that it is not more important than embryonic stem cell research) (PCB, 2004), that the policy is arbitrary and inconsistent (PCB, 2004), that the cell lines in the NIH registry are of poor quality and inappropriate for long-term use (Dawson et al., 2003), and that the policy is harming American

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science by encouraging scientists to focus on research for which federal funding is available or to move to other countries to conduct their research (PCB, 2004). In addition to critique and analysis from scholars and scientists, the 2001 policy and stem cell research generally have become focal points for political debate. For example, both the research and federal policies were debated by the 2004 presidential candidates Senator John Kerry (the Democratic nominee) and President Bush (the Republican nominee) (Commission on Presidential Debates, 2004). Stem cell research was also a major issue at the Democratic national convention in 2004, where Ronald Regan Jr., son of former Republican President Ronald Regan, gave a speech calling for greater government support of embryonic stem cell research. In so far as they are motivated by disagreement with President Bush’s 2001 policy, actions at a state level to fund research not eligible for federal funding (described below) can also be interpreted as political. To some extent, then, embryonic stem cell research has been a partisan political issue. Other political activity, however, has been bipartisan, including two 2004 letters – one signed by over 200 congressional representatives and the other signed by nearly 60 senators – asking the President to relax federal funding restrictions on embryonic stem cell research. Similar bipartisan support has been expressed for the proposed Stem Cell Research Enhancement Act, which would increase federal support of the research (AAAS, 2004). Republican politicians who support changes in federal policy include Senator Orrin Hatch, who is in the unusual position of mixing opposition to abortion with support of embryonic stem cell research provided it uses only spare embryos (Hatch, 2003).

INTERNATIONAL COMPARISONS Biomedical research is an international affair and embryonic stem cell research is no exception. Amongst those nations in which embryonic stem cell research is active, governments employ a variety of regulation strategies, most of which subject the research to some restrictions and oversight whilst still allowing it to move forward. No one method of regulation has prevailed internationally, although some patterns have emerged. For instance, many nations use national legislation to regulate stem cell research, often requiring oversight by a national stem cell research committee or licensing authority. Substantively, bans on creating embryos by cloning are common, although not universal, as are bans on creating embryos by fertilization except as part of fertility treatment, which means that research in many, although certainly not all, nations is limited to spare embryos. The federally funded/non-federally funded split currently employed in the United States is highly unusual by international standards. In the United Kingdom, comprehensive legislation governing all research and medical use of human gametes and human embryos has existed since 1990 (Human Fertilisation and Embryology Act 1990 (UK)). A major feature of the legislation is that it institutes few bans, but requires that all collection and use of embryos and gametes be licensed and overseen by an independent body, called the Human Fertilisation and Embryology Authority (HFEA). The existence of this legislation and the HFEA before human embryonic stem cells were first isolated meant that the research already had a regulatory system into which it could immediately be slotted, obviating the need for significant new legislative action by the Government in response to the research (although regulations were promulgated in 1991 to add three new purposes for which research on embryos is permitted, including increasing knowledge about serious disease) (Human Fertilisation and Embryology (Research Purposes) Regulations 1991). Some scholars have urged the United States to adopt a similar regulatory system, under which very few activities are banned outright but instead require a license and are subject to oversight (Parens and Knowles, 2003). In neighboring Canada, national funding guidelines similar to the policy formulated by President Clinton were released in 2002 by the Canadian Institutes of Health Research (CIHR). Significant terms of these guidelines include that all research using CIHR funds is subject to national oversight by the Stem Cell

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Oversight Committee, that all embryonic stem cell lines generated using CIHR funding must be recorded in a national electronic registry and made available to other Canadian academic researchers at cost, and that stem cells can only be extracted from spare embryos (embryos cannot be created for research purposes, including by cloning). The CIHR guidelines were followed in 2004 by the Assisted Human Reproduction Act, which, like the legislation in the United Kingdom, regulates much more than just embryonic stem cell research. Several provision of the Act, however, directly impact embryonic stem cell research. In particular, the Act prohibits creating a cloned embryo and creating an embryo for “any purpose other than creating a human being or improving or providing instruction in assisted reproduction procedures.” Therefore, the Act limits embryonic stem cell research in Canada to embryos originally created in the course of fertility treatment, although it does not specify that such embryos be surplus to the fertility needs of the donors (i.e. the embryos need not be “spare”) (Johnston, 2006). This distinction between stem cell research involving embryos created for use in research (including by cloning) and research involving embryos originally created for use in fertility treatment also appears in Australian legislation. Australia’s Research Involving Human Embryo’s Act 2002 limits research to “excess ART embryos,” which are defined as embryos created for use in the assisted reproductive technology treatment that are now excess to the needs of the woman or couple for whom they were created (Research Involving Human Embryos Act 2002 (Commonwealth of Australia)). Similar conditions are attached to research in other nations, including France, Belgium, Denmark, Finland, Greece, the Netherlands, Sweden, and Japan. Korea, a country active in stem cell research, is in the unusual position of permitting the creation of cloned embryos for research into rare or incurable diseases, but prohibiting the creation of embryos for research by in vitro fertilization. Under a 2004 law, in vitro fertilization may only be used to create embryos for reproductive purposes, which can later be donated to research if not used (Act on Bioethics and Biosafety (Republic of Korea) 2004). Cloned embryos may only be created for use in research into the treatment of rare or incurable diseases (Act on Bioethics and Biosafety (Republic of Korea) 2004)). In Israel, another country active in stem cell research, no legislation regulates the field, although there is a law against implanting a cloned embryo (Prohibition of Genetic Intervention (Human cloning and Genetic Manipulation of Reproductive Cells) Law, 5759-1999 (Israel)). Therefore, derivation of stem cells from embryos created for research use, including by cloning, is allowed (although to date there are no reports of Israeli scientists creating cloned embryos) (Walters, 2004). Singapore currently has no legal restrictions on embryonic stem cell research, but has legislation in progress based on recommendations of Singapore’s government-appointed Bioethics Advisory Committee. The legislation will likely allow for research use of spare embryos and embryos created for research, including by cloning (Bioethics Advisory Committee, Singapore, 2006). In this international climate, current US policy is rather unusual since only federal funding rules and some state level restrictions apply to US embryonic stem cell research (Knowles, 2004).

STATE AND PRIVATE FUNDING IN THE UNITED STATES In response to the limitations currently placed on the use of federal funds in embryonic stem cell research, private funders have stepped up their support and advocates have lobbied state governments to provide funds for the research. As a result of private funding, Harvard University’s Douglas Melton and colleagues announced in 2004 that they had derived 17 new embryonic stem cells lines with support provided by the Juvenile Diabetes Research Foundation, the Howard Hughes Medical Institute, and Harvard University (Cowan et al., 2004). According to a special report in Scientific American, about $200 million of private money is spent on US stem cell research annually, which compares to $550 million in federal funds for stem cell research, of which, however, only $24 million is available for embryonic stem cell research meeting the 2001 policy (Beardsley, 2005).

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States are also moving to fund stem cell research. In 2004, New Jersey announced that it would begin funding local stem cell researchers through grants and by creating the Stem Cell Institute of New Jersey, toward which the state has thus far contributed $11.5 million. The current Governor plans to seek public support for a further $380 million of support (Office of the Acting Governor, 2005). Funding has also been proposed in other states, including Connecticut, Illinois, and Wisconsin. Thus far, however, the largest state initiative has been in California. In November 2004, voters in California supported a proposition to allocate $3 billion over 10 years to embryonic stem cell research. The initiative, known as Proposition 71, authorized the state of California to sell $3 billion in general obligation bonds to provide funding for stem cell research and research facilities in California. Under the proposition, the funds are to be distributed as grants and loans to California-based institutions by the newly created Californian Institute for Regenerative Medicine, which is also required to establish regulatory standards for the research (Attorney General of California, 2004). Priority for funding is to go to research that would not be eligible for federal funding, which currently translates into most embryonic stem cell research. Critics of the initiative called it fiscally irresponsible given the state’s economy and other health and research needs. They also argued that the institute as structured lacked accountability (Attorney General of California, 2004). Nevertheless, 59% of voters approved the measure.

POLICY AS EMBRYONIC STEM CELL RESEARCH MOVES FORWARD As embryonic stem cell research moves forward in the United States, even with limited federal support, various ethical and policy questions arise. For example, should researchers create embryos in the laboratory by fertilization or cloning or should they only use spare embryos? Either way, they will need to interact with fertility clinics or with egg, sperm, or embryo donors, raising questions about how those interactions should be conducted. Should researches pay fertility clinics for procuring gametes and embryos for stem cell research? How should gamete and embryo donors be approached to donate, precisely whose consent should be required, and should the donors be compensated? Once researchers extract cell lines, are they obliged to make those lines available to other researchers? Ought researchers to patent new cell lines or new stem cell related processes? If they do obtain patents, what practices should they follow in licensing those lines or processes? Should researchers be allowed to mix human cells and animal cells in the creation of chimeras or hybrids? GUIDELINES FROM THE NATIONAL ACADEMIES Many of these issues were taken up by a panel convened in 2004 by the National Academies, a private, nonprofit organization whose mission is to advise the nation on issues in science, engineering, and medicine. In general, the National Academies’ reports and recommendations are very influential and its Guidelines for Human Embryonic Stem Cell Research (Committee on Guidelines for Human Embryonic Stem Cell Research, 2005) received significant media attention when they were released in April 2005. Although these guidelines do not carry the force of law they are likely to be very influential. The guidelines are not the first document to speak to the conduct of embryonic stem cell research in the United States but at the time they were formulated it was not clear whether any of the previous guidance applied to most contemporary embryonic stem cell research because it was formulated either only for federally funded research or under a previous administration (and so could be considered out of date), or because it was simply too restrictive to meaningfully guide institutions and researchers that had made the decision to move forward with the research (Johnston, 2005). The guidelines were, therefore, received as filling a policy vacuum. The committee that drafted the guidelines was asked to consider the use and derivation of stem cells from embryos originally created during fertility treatment, embryos created using donated eggs and sperm, and

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cloned embryos. It therefore did not engage in the debate over whether it is morally permissible to destroy human embryos in research. It also did not consider whether there is a moral difference between research that uses spare embryos and research that uses embryos created specifically for research purposes, including by cloning. Creating embryos for research use is often opposed on the grounds that it is wrong to create human life for the purposes of destroying it (PCB, 2002). Using cloning techniques to generate embryos for use in research is often opposed on the same ground and on the ground that “cloning-for-research” could lead to “cloning-to-produce-children,” which is opposed by many stem cell scientists and policy makers (PCB, 2002). Research use of spare embryos, on the other hand, has received more support on the ground that the embryos would be destroyed anyway. Whether the intention of the original creator is sufficient reason for permitting research on spare embryos but not on embryos created for research use has been questioned (Baylis et al., 2003; Parens, 2001). Even if this moral issue can be resolved, spare embryos may not be satisfactory as the sole source of embryos for stem cell research because there may not be sufficient numbers available to meet demand. A 2003 survey of US fertility clinics reports that of the more than 400,000 embryos in frozen storage in the United States, only 2.8% have been donated to research (Hoffman et al., 2003). Other researchers have argued that these frozen embryos will not be genetically diverse enough for therapeutic purposes (Faden et al., 2003). Whether for these reasons or others, the National Academies committee proceeded on the basis that researchers could use spare embryos and embryos created for research use, including by cloning. Overall, the National Academies committee recommended banning very little scientific activity. Instead, it recommended institutional review of protocols, oversight of the involvement of egg, sperm, and embryo donors, the establishment of stem cell banks, and the documentation of research activity. Two recommendations were particularly significant. First, the committee recommended that much embryonic stem cell research be subject to a mixture of local and national oversight. Local oversight would occur at each institution engaged in embryonic stem cell research, which would establish an Embryonic Stem Cell Research Oversight (ESCRO) committee to oversee all issues related to the derivation and use of embryonic stem cells, review all proposals for scientific merit, maintain records of research that takes place at the institution, including registries of new cell lines, and educate investigators. Local IRBs would provide additional oversight. Even though much embryonic stem cell research will not strictly speaking need to go before an IRB,5 the committee recommended that the procurement of egg, sperm, and embryos always be reviewed by an IRB, regardless of the applicability of federal regulations, and that IRBs never waive the requirement for informed consent from a person donating cells, eggs, sperm, or embryos to research, even where the federal human subject research regulations provide for such a waiver. The guidelines also called for the establishment of a national oversight body to consider issues of practice and policy on an ongoing basis. Such a body would be similar to the UK’s HFEA and Canada’s Stem Cell Oversight Committee, although these national bodies also conduct some of the local review that the guidelines suggested be carried out by ESCRO committees and, to some extent, IRBs.6 The other significant set of recommendations addressed the involvement of egg, sperm, and embryo donors. In line with much guidance, law, and regulation around the world, the guidelines recommended requiring consent from donors for research use. They went beyond previous US guidance, however, by extending this requirement to the women and men who contributed egg and sperm to an embryo originally

5

OHRP told the committee that merely asking couples whether they wish to donate their surplus embryos to research does not render them “human subjects of research,” if no data on them is being gathered and if there is no substantive interaction with them other than gaining their consent. 6 The UK Authority is also charged with overseeing assisted human reproduction, which is an area largely free from national regulation or oversight in the United States and is not addressed in these guidelines.

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created for fertility purposes. These donors are currently asked to consent to reproductive use of their gametes, but are not, apparently, generally asked to consent also to research use. Accepting this requirement, the committee noted, might rule out the use of some embryos already created for fertility purposes that are now in frozen storage. On the issue of compensating egg, sperm, and embryo donors, the guidelines noted the arguments in favor of compensation: paying egg and sperm donors is routine in the US fertility context, and many Americans participating in other kinds of research are offered financial inducements to secure their participation. They acknowledged that arguments for compensating egg donors are particularly strong: “the invasiveness and risks of the procedure suggest that financial remuneration is most deserved, but at the same time there is a greater likelihood of enticing potential donors to do something that poses some risk to themselves.” Ultimately, the National Academies committee followed previous US guidelines and guidelines and laws from many other nations. It recommended that egg donors be reimbursed only for “direct expenses,” and that no payment whatsoever be offered to sperm or embryo donors. They did allow, however, for reimbursement of fertility clinics for costs, including professional services, associated with obtaining consent and collecting eggs, sperm, or embryos. In addition to cash payment, the guidelines recommended against compensation in kind. Donors are not to receive any benefit from their donation, including “personal medical benefit” (excepting autologous transplantation, where the donor receives stem cells derived from his or her donation). This rule would prevent a kind of egg or embryo sharing arrangement whereby women or couples receive cheaper or free fertility treatment in exchange for donating a portion of their eggs or embryos to stem cell research. Similar arrangements exist in the fertility context, where women or couples receive a discount if they donate some of their eggs to others undergoing fertility treatment.7 This arrangement would help make fertility treatment available at a lower cost, but it would also more quickly exhaust the women or couple’s supply of eggs or embryos for their own reproductive use, thereby possibly reducing their chances of achieving pregnancy (Nisker and White, 2005).

COMPENSATING EGG DONORS: THE ARGUMENTS After the guidelines were issued, a controversy over the procuring of eggs in South Korea for use in cloning research brought additional attention to egg donation. In early 2006, it emerged that in research led by WooSuk Hwang of Seoul National University some egg donors were members of the research team, some egg donors were paid to donate, and researchers accompanied some egg donors as they underwent the extraction procedure (Johnston, 2006). These facts about the donation process raised concerns about whether the women who donated eggs to Hwang’s research did so completely voluntarily. Voluntariness is a core commitment of modern research ethics (WMA, 2000), which generally translates into requirements that no one be pressured to participate in research and that each participant be able to withdraw from the research at any time without endangering ongoing medical care or the care-giving relationship with the researchers (WMA, 2000). Researchers usually avoid enrolling family members and employees because they might reasonably feel significant pressure to participate. The commitment to voluntary participation, and specifically the derived right to withdraw from the research at any time, could also be under threat if researchers physically accompany volunteers through procedures as Hwang and his colleagues apparently did. Voluntariness is also the major factor motivating bans on compensating egg donors, the argument being that the need for money could compel participation, especially, though not exclusively, amongst the poor. 7

See, for example, the HFEA 2003 at Appendix A: Guidance for egg sharing arrangements.

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Nevertheless, as the National Academies committee already noted, many other research participants in the United States receive compensation in exchange for their involvement in research (Committee on Guidelines for Human Embryonic Stem Cell Research, 2005). Such payment, particularly where modest, is said to be not only a necessary incentive, but also fair treatment of research subjects (after all, the researchers and their staff will be paid for the time and resources they contribute toward the research). The concern about voluntariness is likely heightened because egg donation is time-consuming, painful, and involves some risk. Egg donors are injected with drugs over weeks so that they super-ovulate (produce many eggs). The eggs are then removed from the woman by either inserting a hollow needle through her vagina or by laparoscopic surgery. Risks of the stimulation and egg-collection process include hot flashes, headaches, sleeplessness, mood alteration, ovarian hyper-stimulation syndrome, nausea, vomiting, pain, bleeding, and infection. There is even a controversy over a possible danger of ovarian cancer from the medications (Gurmankin, 2001). Could payment encourage women to donate eggs even though donation might be painful and pose a risk to their health? The answer is yes. However, whether the risk and pain are unacceptable is another issue. In theory at least, if an IRB approves research involving egg donation it has decided that the risk to donors is reasonable in relation to the importance of the knowledge that may reasonably be expected to result (45 CFR §46.111(a)). Compensation is not supposed to be offered to research participants in order to seduce them to take an unacceptable risk. But voluntariness may not be the only concern about paying egg donors in stem cell research. There is also some opposition to paying anyone for providing bodily materials (rather than solely for their time and effort), for example as expressed in a federal law prohibiting payment for organ donation (although that law expressly does not apply to blood, sperm, or human eggs) (Uniform Anatomical Gift Act, 1987). The stance against sale of bodily materials is well defended in scholarly circles. Bioethicist Thomas Murray argued nearly 20 years ago that all donations of body parts, whether for research or for clinical treatment, should be gifts and not sales (Murray, 1987). Others, however, including Law Professor Lori Andrews, writing around the same time, countered that individuals “have the autonomy to treat their own (body) parts as property,” particularly those parts of the body that they can regenerate (Andrews, 1986). This question about compensating egg, sperm, and embryo donors is far from settled. Despite recommending against compensation, the National Academies’ guidelines left the issue open for revision.

COMMERCIALIZATION AND ACCESS Another strong argument against paid donation in general is that it can add to the costs of conducting research and generating eventual treatments. But this argument works best if the same spirit is adopted by the scientists, institutions, and companies involved in the research and in any eventual treatment. Indeed, a commitment to scientific progress and widely available treatments in stem cell research might entail a commitment by all those involved to, for example, banking and widely distributing new cell lines, participating in international collaboration, and adopting patenting and licensing practices designed to facilitate access (including possibly not patenting some discoveries at all) (DHHS, 1999). These concerns about secrecy, access, and commercialization in stem cell research mirror a larger debate in biomedical research in general (Krimsky, 2004). In terms of access, particular emphasis is often put on patenting and licensing practices, which can have an enormous impact on progress made by other researchers as well as on the availability of eventual treatments (Heller and Eisenberg, 1998). Indeed, a number of patents already attach to embryonic stem cell research, including the only patent to be issued in the world claiming a purified preparation of primate (including human) embryonic stem cells and a method for isolating them (US Patent 5,843,780.). To establish its registry, NIH negotiated with patent holders to issue licenses to non-commercial entities royalty-free so that research

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could move forward. The NIH negotiated licenses, however, explicitly cover only research and do not extend to commercial activities. If embryonic stem cell treatments begin to emerge, significant patent-related disputes will also likely arise (Rabin, 2005). Patenting and licensing issues were to some extent anticipated in California’s Proposition 71, which included a provision requiring the establishment of standards in all grants and loans allowing the state to financially benefit from licenses, patents, and royalties and resulting from the research activities funded under the measure (Attorney General of California, 2004). The inaugural version of these standards requires that a quarter of profits over $500,000 be returned to the state (CIRM, 2006). Further issues related to patenting and commercializing stem cell research in California and in the rest of the United States will very likely develop as the research progresses.

CONCLUSION As this brief introduction shows, embryonic stem cell research has not just generated significant scientific activity, but has also led to the development of numerous policies. In the United States, as well as in some other nations, these policies have responded to the controversial nature of research involving the destruction of human embryos. The US federal funding policy of 2001, which limits the use of federal funds in embryonic stem cell research, has generated a lot of criticism. In response to the 2001 policy’s limits, funds for embryonic stem cell research have come from private donors, charitable organizations, and states. As embryonic stem cell research moves forward in the United States, much of it without federal funding, policies are developing that respond to a number of important issues, including the need for local and national oversight, the role of donors in the research, and the consequences of commercial interests. ACKNOWLEDGMENTS Parts of this chapter are based on two of the author’s previous publications: Paying egg donors: exploring the arguments, Hastings Center Report 2006; 36(1): 28–31 and Stem cell protocols: the NAS guidelines are a useful start, Hastings Center Report 2005; 35(5): 16–17.

REFERENCES Act on Bioethics and Biosafety (Republic of Korea) (2004), English translation available at www.koreabioethics.net/ pds/journal/05-2/7.pdf. American Association for the Advancement of Science (AAAS) (August 26, 2004). AAAS Policy Brief: Stem Cell Research. www.aaas.org/spp/cstc/briefs/stemcells/index.shtml. Andrews, L.B. (1986). My body, my property. Hastings Cent. Rep. 16(5): 28–38. Andrews, L.B. (2004). Legislators as lobbyists: proposed state regulation of embryonic stem cell research, therapeutic cloning and reproductive cloning. In President’s Council on Bioethics (2004). Monitoring Stem Cell Research. Washington, DC: President’s Council on Bioethics. Attorney General of California (2004). State of California: Stem Cell Research. Funding. Bonds. Initiative Constitutional Amendment and Statute (Proposition 71): Official Text and Summary. www.sos.ca.gov/elections/bp_nov04/prop_71_ entire.pdf. Baylis, F.B., Beagan, B., Johnston, J. and Ram, N. (2003). Cryopreserved human embryos in Canada and their availability for research. J. Obstet.Gynaecol. Canada 25(12): 1026–1031. Beardsley, S. (June 27, 2005). A world of approaches to stem cells. Financ. Times Sci. Am. (special report), A20–A21. Bioethics Advisory Committee, Singapore (2006). FAQs – stem cell research. www.bioethics-singapore.org/faq/stem.html. Bush, President George W. (August 9, 2001). Remarks by President George W. Bush on stem cell research. In President’s Council on Bioethics (2004). Monitoring Stem Cell Research. Washington, DC: President’s Council on Bioethics.

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California Institute of Regenerative Medicine (CIRM) (February 6, 2006). Draft Recommended Revisions to the California Code of Regulations, Title 17, Division 4, California Institute of Regenerative Medicine. www.cirm.ca.gov/meetings/pdf/ 2006/02/021006_item_9.pdf. Canadian Institutes of Health Research (March 12, 2002). Human Pluripotent Stem Cell Research: Guidelines for CIHRfunded research. Canadian Institutes of Health Research (Now superceded by: Canadian Institutes of Health Research (June 7, 2005). www.cihr-irsc.gc.ca/e/28216.html. Chung, Y., Klimanskaya, I., Becker, S., Marh, J., Lu, S.J., Johnson, J., Meisner, L. and Lanza, R. (2006), Embryonic and extraembryonic stem cell lines derived from single mouse blastomeres. Nature 439(7073): 216–219. Commission on Presidential Debates (October 8, 2004). The Second Bush–Kerry Presidential Debate (debate transcript). www.debates.org/pages/trans2004c.html. Committee on Guidelines for Human Embryonic Stem Cell Research, National Research Council (2005). Guidelines for Human Embryonic Stem Cell Research. Washington, DC: National Academies Press. Cowan, C.A., Klimanskaya, I., McMahon, J., Atienza, J., Witmyer, J., Zucker, J.P., Wang, S., Morton, C.C., McMahon, A.P., Powers, D. and Melton, D.A. (2004). Derivation of embryonic stem-cell lines from human blastocysts. N. Engl. J. Med. 350(13): 1353–1356. Dawson, L., Bateman-House, A.S., Mueller Agnew, D., Bok, H., Brock, D.W., Chakravarti, A., Greene, M., King, P.A., O’Brien, S.J., Sachs, D.H., et al. (2003). Safety issues in cell-based intervention trials. Fertil. Steril. 80(5): 1077–1085. Department of Health and Human Services (DHHS) (1999). Principles and guidelines for recipients of NIH research grants and contracts on obtaining and disseminating biomedical research resources: final notice. Fed. Reg. 64(246): 72090–72096. Faden, R.R., Dawson, L., Bateman-House, A.S., Agnew, D.M., Bok, H., Brock, D.W., Chakravarti, A., Gao, X.J., Greene, M., Hansen, J.A., et al. (2003). Public stem cell banks: considerations of justice in stem cell research and therapy. Hastings Cent. Rep. 33(6): 13–27. Gurmankin, A.D. (2001). Risk information provided to prospective oocyte donors in a preliminary phone call. Am. J. Bioeth. 1(4): 3–13. Hatch, Senator Orrin, G. (March 19, 2003). Promoting ethical regenerative medicine research and prohibiting immoral human reproductive cloning (Statement before the Senate Judiciary Committee Hearing on promoting ethical regenerative medicine research and prohibiting immoral human reproductive cloning). http://hatch.senate.gov/index.cfm? FuseAction=PressReleases.Detail&PressRelease_id=726. Heller, M.A. and Eisenberg, R.S. (1998). Can patents deter innovation? The anticommons in biomedical research. Science 280(5364): 698–701. Hoffman, D.I., Zellman, G.L., Fair, C.C., Mayer, J.F., Zeitz, J.G., Gibbons, W.E. and Turner Jr.,T.G. (2003). Society for Assisted Reproduction Technology (SART) and RAND. Cryopreserved embryos in the United States and their availability for research. Fertil. Steril. 79(5): 1063–1069. Human Fertilisation and Embryology Act 1990 (UK). Human Fertilisation and Embryology (Research Purposes) Regulations 1991. Human Fertilisation and Embryology Authority (2003). Human Fertilisation and Embryology Authority, Code of Practice, 6th edn. The Human Fertilisation and Embryology Authority, London. Johnston, J. (2005). Stem cell protocols: the NAS guidelines are a useful start. Hastings Cent. Rep. 35(5): 16–17. Johnston, J. (2006). Is research in Canada limited to “spare” embryos? Health Law Rev. 14(3): 3–13. Knowles, L.P. (2004). A regulatory patchwork – human ES cell research oversight. Nat. Biotechnol. 22(2): 157–163. Krimsky, S. (2004). Science in the Private Interest: Has the Lure of Profits Corrupted the Virtue of Biomedical Research? Lanham, MD: Rowman & Littlefield. Murray, T.H. (1987). Gifts of the body and the needs of strangers. Hastings Cent. Rep. 17(2): 30–38. National Bioethics Advisory Commission (NBAC) (1999). Ethical Issues in Human Stem Cell Research, Vol. 1. Rockland, MD: National Bioethics Advisory Commission. National Institutes of Health (2000). Guidelines for research using human pluripotent stem cells. Fed. Reg. 65: 51975–51981. National Right to Life Coalition (NRLC) (August 10, 2001). President Bush’s Statement (press release): NRLC Press Release. www.nrlc.org/press_releases_new/stemcelldno081010.htm. Nisker, J. and White, A. (2005). The CMA code of ethics and the donation of fresh embryos for stem cell research. Can. Med. Assoc. J. 173(6): 621–622.

US Stem Cell Research Policy

Office of Human Research Protections (OHRP) (March 19, 2001). Guidance for investigators and institutional review boards regarding research involving human embryonic stem cells, germ cells and stem cell-derived test articles. www.hhs.gov/ohrp/humansubjects/guidance/stemcell.pdf. Office of the Acting Governor (August 2, 2005). Codey announces funding for stem cell research grants (press release). www.nj.gov/cgi-bin/governor/njnewsline/view_article.pl?id=2659. Orr, R.D. (2002). The moral status of the embryonal stem cell: inherent or imputed? Am. J. Bioeth. 2(1): 57–59. Parens, E. (2001). On the ethics and politics of embryonic stem cell research. In Holland, S., LeBacqz, K. and Zoloth, L. (eds.), The Human Embryonic Stem Cell Debate: Science, Ethics, and Public Policy. Cambridge, MA: MIT Press. Parens, E. and Knowles, L.P. (2003). Reprogenetics and public policy: reflections and recommendations. Hastings Cent. Rep. 33(4): S1–S24. President’s Council on Bioethics (2002). Human Cloning and Human Dignity: An Ethical Inquiry. Washington, DC: President’s Council on Bioethics. President’s Council on Bioethics (2004). Monitoring Stem Cell Research. Washington, DC: President’s Council on Bioethics. Prohibition of Genetic Intervention (human cloning and genetic manipulation of reproductive cells) Law, 5759-1999 (Israel). Rabin, S. (2005). The gatekeepers of hES cell products. Nat. Biotechnol. 23(7): 817–819. Research Involving Human Embryos Act 2002 (Commonwealth of Australia). Roe v. Wade 410 US 113 (1973). Schwartz, R.E., Reyes, M., Koodie, L., Jiang, Y., Blackstad, M., Lund, T., Lenvik, T., Johnson, S., Hu, W.S. and Verfaillie, C.M. (2002). Multipotent adult progenitor cells from bone marrow differentiate into functional hepatocyte-like cells. J. Clin. Invest. 109: 1291–1302. Shamblott, M.J., Axelman, J., Wang, S., Bugg, E.M., Littlefield, J.W., Donovan, P.J., Blumenthal, P.D., Huggins, G.R. and Gearhart, J.D. (1998). Derivation of pluripotent stem cells from cultured human primordial germ cells. Proc. Natl Acad. Sci. USA 95(23): 13726–13731. The White House (August 9, 2001). Fact sheet – embryonic stem cell research. http://www.whitehouse.gov/news/releases/ 2001/08/print/20010809-1.html. Thomson, J.A., Itskovitz-Eldor, J., Shapiro, S.S., Waknitz, M.A., Swiergiel, J.J., Marshall, V.S. and Jones, J.M. (1998). Embryonic stem cell lines derived from human blastocysts. Science 282: 1145–1147. Uniform Anatomical Gift Act, 1987. US Patent 5,843,780. Verfaillie, C.M., Pera, M.F. and Lansdrop, P.M. (2002). Stem cells: hype and reality. Hematology Am Soc Hematol Educ Program 369–391. Walters, L. (2004). Human embryonic stem cell research: an intercultural perspective. Kennedy Inst. Ethics J. 14(1): 3–38. World Medical Association (WMA) (2000). Declaration of Helsinki: ethical principles for medical research involving human subjects (reprinted). J. Am. Med. Assoc. 284(23): 3043–3045. 45 CFR. §46.204(h-i). 45 CFR §46.111(a). 45 CFR. §46.206. 498A of the Public Health Service Act (42 USC 289g-1).

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81 Overview of FDA Regulatory Process Celia Witten, Ashok Batra, Charles N. Durfor, Stephen L. Hilbert, David S. Kaplan, Donald Fink, Deborah Lavoie, Ellen Maher, and Richard McFarland

INTRODUCTION AND CHAPTER OVERVIEW The field of regenerative medicine encompasses a breathtaking array of interdisciplinary scientific approaches that address a broad spectrum of clinical needs. Recent advances in scientific knowledge related to cell biology, gene transfer therapy, biomaterials, immunology, and engineering principles applicable to biological systems place the regenerative medicine community in a position to address a number of challenging and critical health needs. These include treatment of disease conditions resulting from pancreas, liver, and kidney failure; structural cardiac valve repair; skin and wound repair; and orthopedic applications. Scientific challenges confronting this field include expanding the knowledge base in each discipline as well as developing an interdisciplinary approach for identifying and resolving key questions. The Food and Drug Administration’s (FDA) regulatory review process mirrors this scientific challenge with regard to development of review paradigms that cross scientific disciplines. This chapter will provide a brief historical review of FDA and its organizational structure as well as discuss topics pertaining to the regulation of regenerative medicine products including possible regulatory pathways for combination products and relevant jurisdictional issues. Sources of information concerning FDA regulatory policies important to regenerative medicine product developers will also be discussed. It is essential for individuals, institutions, and companies, collectively referred to in FDA regulations as Sponsors (the term Sponsor for drugs and biologics is defined at 21 Code of Federal Regulations (CFR) 312.3(b), while Sponsor is similarly defined at 812.3(n) for devices), responsible for the clinical trials of regenerative medicine products to be aware of FDA regulatory policies and how to obtain this information. Suggestions will also be provided as to how to effectively engage the FDA during the development of a novel regenerative medicine product. BRIEF LEGISLATIVE HISTORY OF FDA Medical products regulated by FDA include human and animal drugs, medical devices, and biological products, such as vaccines, cellular and gene therapies, and blood products. Among the therapeutic agents of biological origin regulated by FDA are cellular therapies, including products derived in whole or part from human tissue and xenotransplantation. In addition to medical products for human use, FDA also regulates food other than meat and poultry, radiation emitting products for consumer, medical, and occupational use, cosmetics, and animal feed. FDA’s role in medical product regulation extends throughout the entirety of the

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product life cycle. Depending on the product category, this may mean oversight, including review and inspection, of clinical trials, of the premarket product approval/clearance process, of manufacturing controls, controls over labeling, and registration and listing requirements. FDA also continues its oversight once a product is marketed in a variety of ways, including inspections and review of adverse events. FDA laws and regulations have developed over time; prompted partly in response to serious medical adverse events or by other public health and safety concerns. Early regulation of biological products was prompted in part by the death of 13 children in 1901 following administration of diphtheria antitoxin prepared from a source contaminated with tetanus. In response, Congress passed the Biologics Control Act in 1902. This act provided for regulation of viruses, serums, toxins, and analogous products; required licensing of manufacturing establishments and manufacturers; and provided the government with inspectional authority. The Act focused on requiring control of manufacturing processes for producing biological products, reflecting the extent to which the starting source material and the manufacturing process defined the final product. In 1906, Congress passed the Federal Food and Drugs Act proposed in part in reaction to the meat packing industry conditions described in Upton Sinclair’s book “The Jungle.” While the primary focus of the Act was on food safety, the law also required that drugs be provided in accordance with standards of strength, quality, and purity unless otherwise specified in the label. Premarket review of new drugs was not required until the passage of the 1938 Food, Drug, and Cosmetic Act (FD&C Act), which repealed the earlier 1906 Federal Food and Drugs Act. In 1937, the sulfa drug, Elixir Sulfanilamide, previously available only in tablet or powder form to treat streptococcal infections, was marketed as a liquid using diethylene glycol, an analog of antifreeze, as a formulating solvent. This change in formulation, made without the requirement for premarket review, resulted in over 100 deaths, many of them were children, prompting the passage of the 1938 FD&C Act. The 1938 Act also put medical devices and cosmetics under FDA authority and authorized factory inspections. The Public Health Service Act (PHS Act), passed in 1944, incorporated the 1902 Biologics Control Act and is the present legal basis for licensing of biological products. Because most biological products also meet the definition of “drugs” under the FD&C Act, they are also subject to regulation under that Act. The requirement for premarket demonstration of efficacy and the authority for FDA oversight of clinical trials were provided by the Kefauver-Harris amendments to the FD&C Act in 1962. These amendments were prompted in part by the tragic adverse events resulting from use of thalidomide as a non-addictive prescription sedative. This drug, not approved as a sedative in the United States, when taken by pregnant women during the first trimester resulted in thousands of birth defects for children born outside this country. The Medical Device Amendments to the FD&C Act were passed in 1976, following reports of safety issues with respect to the Dalkon Shield intrauterine device. The Medical Devices Amendments included risk-based requirements for premarket notification or approval of medical devices. Prior to 1976, FDA authority was limited to taking action against marketed devices found to be unsafe or ineffective.

LAWS, REGULATIONS, AND GUIDANCE The previous section summarized the history of laws that form the underpinning of FDA medical product regulation. This section provides a brief description of how laws are made and implemented, the procedures for promulgating regulations, and a description of how FDA develops and uses guidance documents. Laws are created as an outcome of legislative activity conducted in the United States Senate and House of Representatives resulting in passage of a bill. Once Congress passes a bill, it becomes law if signed by the

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President. If the President vetoes the bill, it becomes law if two-thirds of the Senate and House of Representatives vote in favor of the bill. A federal law also is denoted as a public law and may contain a name, such as the FD&C or PHS Acts. These laws are then incorporated into the United States Code (USC) which is updated every 6 years with supplements published regularly to incorporate changes to statutes between updates. Drugs, biologics, and device laws can be found in the United States Code at:

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Drugs and Devices: Title 21 Chapter 9 Biologics: Title 42 Chapter 6A

When laws are passed, government agencies, such as the FDA, often implement them by promulgating regulations. Sometimes, an agency may elect to promulgate regulations on its own whereas other laws may explicitly require an agency to issue regulation. The process for making regulations must be performed in accordance to the Administrative Procedures Act (Title 5, USC, Chapter 5). This Act generally requires agencies, such as FDA, to provide public notice and opportunity for comment as part of the rule-making process. FDA regulations are contained in the CFR. Regulations for drugs, biologics, devices, and tissues, along with related regulations, may be found in various parts of Title 21 of the CFR. The following is a list of key regulatory provisions:

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Drugs: 21 CFR Parts 200–299, 300–369 Biologics: 21 CFR Parts 600–680 Devices: 21 CFR Parts 800–898 Human Cells, Tissues, and Cellular and Tissue-based Products: 21 CFR Parts1270/1271 Recalls: 21 CFR 7 Informed Consent/Institutional Review Boards: 21 CFR 50/56 Financial Disclosure by Clinical Investigators: 21 CFR Part 54 Good Laboratory Practice for Nonclinical Laboratory Studies: 21 CFR Part 58 Good Guidance Practices: 21 CFR 10

Guidance documents are non-binding publications that describe FDA’s interpretation of policy pertaining to a regulatory issue or set of issues related to:

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The design, production, labeling, promotion, manufacturing, and testing of regulated products. The processing, content, and evaluation or approval of submissions. Inspection and enforcement policies.

Guidance documents, which are developed in accordance with Good Guidance Practices found at 21 CFR §10.115, are intended to clarify FDA’s current thinking related to regulatory issues and procedures. Unlike regulations and laws, guidance documents are not enforceable. Therefore, Sponsors may elect to choose alternate approaches that still comply with existing laws and regulatory requirements. In most cases, guidance documents are issued in draft for public comment before implementation. In cases where prior public participation is not feasible or appropriate, FDA may issue a guidance document for immediate implementation without first seeking public comment. Many of the guidance documents referred to in this chapter, although available to the public, are still in draft form. This reflects FDA’s efforts to convey up-to-date information to those involved in the developing field of regenerative medicine. FDA is currently working to finalize these draft guidance documents. When considering development of a guidance document, FDA may freely discuss related issues with the public. In fact, the FDA may hold a public meeting, advisory committee meeting, or workshop to obtain input on regulatory issues. Finally, after receiving public input, FDA will evaluate submitted comments and finalize

Overview of FDA Regulatory Process

the document. Guidance documents are a very useful way for FDA to communicate current thinking to the public. Within the arena of regenerative medicine, it is of value to be aware of both product-specific and cross-cutting guidance documents. Some of the more pertinent guidance documents to this field, such as those related to preclinical testing, manufacturing practices, and clinical trial design, are discussed in this chapter. In addition to FDA guidance documents, FDA may also refer to guidelines published by the International Conference on Harmonization (ICH). ICH is an international effort to harmonize regulatory requirements. ICH guidelines, similar to FDA guidance documents, are non-binding.

FDA ORGANIZATION AND JURISDICTIONAL ISSUES Scientific development of regenerative medicine products involves extensive testing and planning prior to initiation of clinical trials; therefore, it can be helpful for individuals and organizations involved in product development to engage in early dialog with the appropriate FDA review unit in order to receive and consider FDA comments on the design of the preclinical development plan. This section describes FDA organizational structure and provides basic information regarding jurisdictional decisions made to determine the appropriate regulatory pathway for a broad range of products. The FDA consists of six Centers and the Office of the Commissioner. Three of the Centers are responsible for regulating medical products for humans. The Center for Biologics Evaluation and Research (CBER) regulates a variety of biological products, including blood and blood products, vaccines and allergenic products, and cellular, tissue, and gene therapies, as well as some related devices. The Center for Devices and Radiological Health (CDRH) is responsible for review of diagnostic and therapeutic medical devices, administration of the Mammography Quality Standards Act (MQSA) program, and ensuring safety of radiation emitting products. The Center for Drug Evaluation and Research (CDER) regulates a variety of drug products, including small molecule drugs, and well-characterized biotechnology-derived drug products that include monoclonal antibodies and cytokines. For many medical use products it is clear which Center within FDA shall have primary jurisdiction for the premarket review. For other products, including some technologically novel products under development, determining which Center has jurisdiction for review may be unclear. Important starting points for determining product jurisdiction are the formal regulatory definitions of biological products, drugs, devices, and combination products, as well as contacts with the agency. The formal definitions are as follows:

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Biological Product (42 USC 262(i)): A virus, therapeutic serum, toxin, antitoxin, vaccine, blood, blood component or derivative, allergenic product, or analogous product, or arsphenamine or derivative of arsphenamine (or any other trivalent organic arsenic compound), applicable to the prevention, treatment, or cure of a disease or condition of human beings. Drug (21 USC 321(g)(1)): (A) articles recognized in the official United States Pharmacopeia, official Homeopathic Pharmacopeia of the United States, or official National Formulary, or any supplement to any of them; and (B) articles intended for use in the diagnosis, cure, mitigation, treatment, or prevention of disease in man or other animals; and (C) articles (other than food) intended to affect the structure or any function of the body of man or other animals; and (D) articles intended for use as a component of any articles specified in clause (A), (B), or (C). Device (21 USC 321(h)): An instrument, apparatus, implement, machine, contrivance, implant, in vitro reagent, or other similar or related article, including any component, part, or accessory, which is (1) recognized in the official National Formulary, or the United States Pharmacopeia, or any supplement to them; (2) intended for use in the diagnosis of disease or other conditions, or in the cure, mitigation, treatment, or prevention of disease, in man or other animals; or (3) intended to affect the structure or any function of the

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•

body of man or other animals, and which does not achieve its primary intended purposes through chemical action within or on the body of man or other animals and which is not dependent upon being metabolized for the achievement of its primary intended purposes. Combination Product (21 CFR 3.2(e)): (1) a product comprised of two or more regulated components, that is, drug/device, biologic/device, drug/biologic, or drug/device/biologic, that are physically, chemically, or otherwise combined or mixed and produced as a single entity; (2) two or more separate products packaged together in a single package or as a unit and comprised of drug and device products, device and biological products, or biological and drug products; (3) a drug, device, or biological product packaged separately that according to its investigational plan or proposed labeling is intended for use only with an approved individually specified drug, device, or biological product where both are required to achieve the intended use, indication, or effect and where upon approval of the proposed product the labeling of the approved product would need to be changed, for example, to reflect a change in intended use, dosage form, strength, route of administration, or significant change in dose; or (4) any investigational drug, device, or biological product packaged separately that according to its proposed labeling is for use only with another individually specified investigational drug, device, or biological product where both are required to achieve the intended use, indication, or effect.

FDA’s Office of Combination Products (OCP), located in the Office of the Commissioner, has broad administrative overview responsibilities covering the regulatory life cycle of drug–device, drug–biologic, and device–biologic combination products. When jurisdiction is uncertain, sponsors may contact OCP and OCP may assign primary review responsibility for the oversight of combination and other medical products, following a formal submission process called a Request for Designation (RFD). The appropriate FDA Center jurisdiction is determined by considering the primary mode of action of the product.

APPROVAL MECHANISMS AND CLINICAL STUDIES There are several premarket approval pathways for medical products, depending on whether the product is a drug, biological product, or device. Approval pathways, explained in more detail below, include the Biologics License Application (BLA) for biologics and New Drug Application (NDA) for drugs. The Premarket Approval Application (PMA), Humanitarian Device Exemption (HDE) and 510k clearance mechanism are various regulatory pathways used for medical devices. Clarification on the type of application needed for a particular regenerative medicine product may be helpful to the Sponsor early in development, to enable the Sponsor to discuss the data needed for a marketing application during the planning stage. A BLA is an application for licensure under the PHS Act; the approval standards set forth in the statute are a demonstration that the product is safe, pure, and potent. Further information concerning the licensure of biological products is provided in “Guidance for Industry: Providing Clinical Evidence of Effectiveness for Human Drugs and Biologic Products.” (US FDA, 1998b) A PMA is an application for approval for most Class III medical devices; the Sponsor must show reasonable assurance of safety and effectiveness. (US FDA 2002c) Under medical device regulation a product can also gain approval as an HDE, which is not a full marketing approval but requires demonstration of safety and probable benefit. (US FDA, 2003d) To qualify for this type of application, a Sponsor would need to first receive a designation from the FDA Office of Orphan Products Development that the device is a Humanitarian Use Device (HUD), intended for treatment or diagnosis of a disease or condition that affects or are manifested in fewer than 4,000 individuals per year in the United States. The 510k clearance process applies to products that are “substantially equivalent” to a Class I or II (or in a few cases, a Class III) device already on the market.

Overview of FDA Regulatory Process

Many, but not all, combination products are approved or cleared under one marketing application. For example, depending on the specific facts, including the primary mode of action of the product, a combination biological device could be licensed under the biologics authorities or approved under the medical device authorities. Following approval of a marketing application there are also post-marketing requirements such as reporting (US FDA, 2002d; NDA; Reporting for Biological Products 21 CFR 314.80 and 21 CFR 314.81, 21 CFR 600.14, 21 CFR 600.80, 21 CFR 600.81, 21 CFR 601.28, 21 CFR 601.70, and 21 CFR 601.93). In addition, modifications to the product or labeling may require prior approval. FDA has published regulations and guidance documents that address submission and approval processes for modifications to marketed products (US FDA, 2002d, 2005d; PMA Supplements; Supplements and Changes to an Approved NDA; BLA 21 CFR 814.39, 21 CFR 314.70, 21 CFR 314.71, 21 CFR 314.72). Compliance with manufacturing requirements is also an ongoing Sponsor obligation. FDA has issued a draft guidance document entitled “Draft Guidance for Industry and FDA: Current Good Manufacturing Practice (cGMPs) for Combination Products” which provides direction on applicable manufacturing requirements for combination products (US FDA, 2004c). Due to the relatively new nature of regenerative medicine and its developmental status, post-approval topics will not be further discussed in this chapter. In circumstances when clinical investigation is needed to evaluate the safety and efficacy of an investigational product prior to marketing approval, an Investigational New Drug (IND) application is required for drugs and biologics, and an Investigational Device Exemption (IDE) is generally required for devices (US FDA, 2003c, 2005b, 2006d). For both types of applications the Sponsor needs to submit a description of the product and manufacturing process, preclinical studies, a clinical protocol, information on any other prior investigations such as human clinical studies, and a rationale for the study design. An Institutional Review Board (IRB) and informed consent are also required. The FDA has 30 days to review the application to determine if the study may proceed. The contents are specifically laid out in FDA regulations for each type of application. Requirements for the content of an IND can be found at 21 CFR 312.23 and for an IDE at 21 CFR 812.20. For some products, there may be applicable guidance with respect to developing the manufacturing or the preclinical data to support the study. For example, the “Draft Guidance for Reviewers: Instructions and Template for Chemistry, Manufacturing, and Control (CMC) Reviewers of Human Somatic Cell Therapy IND’s Applications” discussed in the following section provides information on characterization and manufacturing of a cellular product to be submitted in an IND (US FDA, 2003a). Applicable regulations and guidance should be further consulted for information on adverse event reporting, labeling, study conduct and monitoring, and other topics related to requirements for conducting an IND (US FDA, 2005b, 2006d). For information on general clinical study design and conduct issues, FDA has many guidance documents that may be helpful (US FDA, 2001a, b). For some indications there may be guidance documents that apply across technologies, such as the “Guidance for Industry: Chronic Cutaneous Ulcer and Burn Wounds – Developing Products for Treatment” (US FDA, 2006f). In addition, guidance documents not directly on point for a specific product, indication, or technology may be worth consulting, as the documents may provide some insights into general clinical issues such as assessment parameters that may be of value.

MEETINGS WITH INDUSTRY, PROFESSIONAL GROUPS, AND SPONSORS Although the terminology and procedures may vary, all three FDA Centers performing medical product review encourage meetings with Sponsors to address questions prior to a regulatory submission and at specific developmental milestones. When requesting a formal or informal meeting with FDA it is helpful to provide background information as well as specific discussion questions. Further information about formal meetings, such as what to include in a meeting request, and what type of information to include in the

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information package submitted prior to the meeting, is provided in “Guidance for Industry: Formal Meetings with Sponsors and Applicants for PDUFA Products” (US FDA, 2002b). Early stage device meetings are addressed in “Early Collaboration Meetings under the FDA Modernization Act (FDAMA), Final Guidance for Industry and CDRH Staff ” (US FDA, 2001e). FDA also interacts with organizations representing a group of interested parties (e.g. International Society for Cellular Therapy, American Association for Blood Banks, and Pharmaceutical Research and Manufacturers of America), which provides an opportunity to discuss topics of interest to FDA and the organization. These interactions can be very valuable for FDA and stakeholders as they are a way to better understand general issues of concern as opposed to product-specific discussions with individual firms. In addition to such interactions and meetings with individual sponsors, FDA also has various advisory committees that review available data and information, and make recommendations related to a variety of issues, many of which are pertinent to the field of regenerative medicine. Advisory committees will be discussed further in Section “Advisory Committee Meetings.”

REGULATIONS AND GUIDANCE OF SPECIAL INTEREST FOR REGENERATIVE MEDICINE The topics discussed thus far have been of general applicability for medical product regulation: marketing pathways, clinical trial regulation, meetings, guidance development, and related topics. This section will review a few topics of particular interest to the scientific community engaged in development of regenerative medicine products: FDA regulations on human tissue products, product characterization for cellular products, FDA policy and guidance on xenotransplantation, and gene therapy. Regenerative medicine products often face some unique product development challenges because of their complexity. Some of these products contain metabolically active cells and tissue, making the manufacturing, characterization, and study of these products a challenge because even a small variation in manufacturing may impact product safety and effectiveness. Scaffolds themselves may be difficult to characterize; for example, some of the materials used during development have complex three-dimensional structures. For cell–scaffold combination products the challenges are multiplied; as such a product is not defined by the components alone. For example, when a combination product is chemically or physically combined, product assembly is an important step in product manufacture and there may be other processing steps or further cell–scaffold interactions which will in turn further define the characteristics of the final product. Packaging and shelf life are critical concerns. The preclinical development plan needs careful attention. A particular question for cell–scaffold combinations is to determine which tests need to be conducted on individual components prior to assembly and which are most relevant after product assembly. For many innovative products, such as cell–scaffold combinations and other regenerative medicine products, the final product, and instructions for use, can be expected to undergo iterative modifications over time. Consequently, refinement of the product and review of product modification will be an ongoing process for the sponsor and for FDA, respectively. It will be critical for the sponsor to have a good understanding of its product and the key scientific or clinical issues that could affect safety and effectiveness of the product, including establishment of manufacturing controls. Demonstrating comparability of a biological product after a manufacturing change is a question that FDA routinely asks as the manufacturer modifies their processing methods; for this reason product characterization is extremely important and should be considered early during product development. Regulation of Human Cells and Tissues Intended for Transplantation An understanding of the regulations applicable to cells and tissues is important for developers of regenerative medicine products since human cells or tissues comprise the whole, or a key component, of many products. In 1997, noting the fragmented approach to regulation of human cell and tissue-based products the FDA issued the “Proposed Approach to the Regulation of Cellular and Tissue-Based Products” (US FDA, 1997b). This

Overview of FDA Regulatory Process

document proposed a tiered risk-based approach to regulation of these products. According to the proposed approach, products posing less risk would be subject to the rules designed to minimize communicable disease risks, and additional regulatory requirements would be imposed on those products posing additional risk. The proposed approach to regulation of human tissues was implemented in three parts, collectively referred to as the tissue rules: Registration and Listing, Donor Eligibility, and Good Tissue Practices (GTP). These complete set of rules went into effect on May 25, 2005. The tissue rules focus on control of infectious disease in products containing human cells or tissue. Thus the tissue rules apply to all human cellular and tissue-based products. It is important for sponsors of regenerative medicine products to be aware of these rules, as well as the specific requirements for biologics or devices that may apply depending on the particular regulatory pathway applicable to their products. The tissue rules are published as regulations at 21 CFR Parts 1270 and 1271. With some exceptions that are noted in the tissue rules, human cells or tissue intended for implantation, transplantation, infusion, or transfer into a human recipient are regulated as a human cell, tissue, and cellular and tissue-based product (HCT/P). Examples of HCT/Ps are: musculoskeletal tissue, skin, ocular tissue, human heart valves, dura mater, reproductive tissue, and hematopoietic stem/progenitor cells. Tissues specifically excluded are: vascularized organs, minimally manipulated bone marrow, blood products, xenografts, secreted or extracted products such as human milk and collagen, ancillary products, and in vitro diagnostic products. The tissue rules require tissue establishments:

• • •

To register and list their HCT/Ps with FDA (21 CFR 1271 Subparts A and B). To evaluate donors through screening and testing, to reduce the transmission of infectious diseases through tissue transplantation (21 CFR 1270 and 1271 Subpart C). To follow GTP to prevent the spread of communicable disease (21 CFR 1271 Subpart D).

Additional requirements regarding reporting, labeling, inspections, importation, and enforcement are described in 21 CFR 1271 Subparts E and F. The rules also define the circumstances under which a product would be subject to the tissue rules only; and when additional oversight such as the need for a BLA, PMA, or other marketing application would be required (21 CFR 1271.20). Products that meet the following conditions are regulated by FDA solely under the tissue rules: the tissue is not more than minimally manipulated, is intended for homologous use, is not combined with a drug or device (with certain exceptions), and does not have a systemic effect and is not dependent upon the metabolic activity of living cells for its primary function (except for autologous use or allogeneic use in a first or second degree blood relative, or reproductive use). If all four of these conditions are not met, a marketing application is required. Additional information and documents regarding these rules, as well as electronic forms for registration and listing, can be found on the FDA website (US FDA, 2006a). FDA has issued “Draft Guidance for Industry: Eligibility Determination for Donors of Human Cells, Tissues, and Cellular and Tissue-Based Products (HCT/Ps)” which provides guidance for donor testing and screening that is recommended in making donor eligibility determination for donors of HCT/Ps (US FDA, 2004b). FDA has also issued a draft guidance entitled “Guidance for Industry, Preventive Measures to Reduce the Possible Risk of Transmission of Creutzfeldt–Jakob Disease (CJD) and Variant Creutzfeldt–Jakob Disease (vCJD) by Human Cells, Tissues, and Cellular and Tissue-Based Products (HCT/Ps)” (US FDA, 2002a). FDA plans to issue a one final guidance document on eligibility determination to reduce the possibility of infectious disease transmission for HCT/Ps that will incorporate both draft guidance topics. Human Cellular Therapies Many products in development for tissue repair or replacement are comprised of cells or cells combined with a scaffold. The cell or tissue source and manufacturing process may vary greatly for different products.

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Despite the diversity in products, there are regulatory considerations that apply to all cellular preparations being developed as investigational regenerative medicine products intended for early phase clinical studies. Among these considerations are three that will be discussed briefly: control of the source material, demonstrated control of the manufacturing process, and characterizations of the cellular product that results from the manufacturing process. The cell source will vary for different products and may be autologous or allogeneic, undifferentiated stem/progenitor cells, or terminally differentiated cells. Assuring the safety of source cellular materials used during manufacture of an investigational regenerative medicine product begins by determining the eligibility of the donors selected to provide the source material through screening and testing. This screening and testing is part of the tissue rules described earlier in this chapter. Although autologous products are not required to comply with the donor screening and testing requirements in the tissue rules, for autologous tissue that is either positive for specific pathogens, or that has not been screened or tested, it is recommended that manufacturers document if tissue culture methods could propagate or spread viruses or other adventitious agents to persons other than the recipient (US FDA, 2003a). Donor eligibility determination is required for all allogeneic donors of cells and tissues. The cell source may raise substantial concerns in addition to the possibility of infectious disease transmission. In addition to screening and testing donors for communicable disease agents, according to the document entitled “ICH Guidance on Quality of Biotechnological/Biological Products: Derivation and Characterization of Cell Substrates Used for Production of Biotechnological/Biological Products,” FDA has suggested that Sponsors consider the importance of evaluating donor medical history information and the relevance of conducting specified molecular genetic testing as part of an overall comprehensive assessment program to establish the suitability of a specific cellular preparation for use in the manufacture of a regenerative medicine somatic cellular product (US FDA, 1998c). The rationale and feasibility for collecting additional information about molecular genetic testing was discussed in a public meeting of the FDA Biological Response Modifiers Advisory Committee (now known as Cellular, Tissue and Gene Therapies Advisory Committee (CTGTAC)) convened July 13–14, 2000 on the topic of “Human Stem Cells as Cellular Replacement Therapies for Neurological Disorders” (US FDA, 2006c). A description of the physiological source of the cellular material, including tissue of origin and phenotype such as hematopoietic, neuronal, fetal, or embryonic conveys important information about the cells and their critical attributes. Control of the manufacturing process helps provide assurance of the consistent, reproducible production of the cellular component. Often, manufacturing will involve a multi-step process that must be performed using aseptic techniques to prevent introduction of microbial contamination (US FDA, 2004d). Many types of reagents may be used to manufacture the cellular component of a product including those that promote cellular replication, induce differentiation, and those used to select targeted cell populations, specifically, serum, culture medium, peptides, cytokines, and monoclonal antibodies. It is essential that reagents be properly qualified (US FDA, 1993, 1997a, 2003a, 2004a). Demonstration of manufacturing control is evidenced by strict adherence to standard operating procedures and quality control assessment of manufacturing intermediates as well as testing of the final cellular preparation. Due to inherent biological complexity it is unlikely that a unique biomarker or other single analytical test will be sufficient to permit full characterization of a cellular product. Accordingly, as recommended in the “Guidance for Industry: Guidance for Human Somatic Cell Therapy and Gene Therapy,” FDA asks Sponsors to provide documentation that their testing paradigm developed for the final cell product encompasses a multi-parametric approach that may involve biological, biochemical/biophysical, and/or functional characterization (US FDA, 1998a, 2003a, 2004a). Tests developed to demonstrate identity of the cell product (physical and chemical characteristics, identify the product as being what is designated on the label), purity (freedom from contaminants

Overview of FDA Regulatory Process

including residual reagents and unintended cell populations), and potency/biological activity (the specific ability of the cells, as indicated by appropriate laboratory tests, to effect a given result) should be conceived to determine the degree to which the characteristics of the manufactured cell preparation conform to desired and specified criteria (US FDA, 1998a, 2003a, 2004a). This process can be challenging for a number of reasons. For example, the mechanism of action associated with a cell product may be incompletely understood and thus constrains the ability to develop a specific potency assay. Direct assessment of potency for a cellular preparation may not be possible due to a lack of appropriate in vitro or in vivo assay systems. On February 9–10, 2006, the FDA CTGTAC discussed this challenging topic and obtained input on alternative approaches for performing potency assessments of cellular therapy products (US FDA, 2006c). In summary, assuring the safety of cell products that in and of themselves constitute a regenerative medicine product or that constitute a component of a product requires demonstrated control over each facet of the manufacturing process. This assurance begins with acquisition of the source material and is carried forward through manufacturing and characterization of the final cellular preparation using specified analytical tests based in large measure on the intrinsic biological properties of the cell product. Xenotransplantation The success of allogeneic organ transplantation has increased the demand for human cells, tissues, and organs. Scientific advances in the areas of immunology and molecular biology coupled with the growing worldwide shortage of transplantable organs have lead to increased interest in xenotransplantation. In addition to the potential use of xenotransplantation to address the shortage of human organs for transplantation, there are increasing efforts to utilize xenotransplantation in the treatment of other chronic diseases as well as end organ failure. Along with the promise of xenotransplantation are a number of challenges, including the potential risk of transmission of infectious agents from source animals to patients, and the spread of any zoonotic disease to the general public. Agents pathogenic for humans may not be pathogenic or detectable in the source animal host. In addition, the potential exists for recombination or reassortment of source animal infectious agents, such as viruses, with non-pathogenic or endogenous human infectious agents, to form new pathogenic entities. These considerations demonstrate the need to proceed with caution in this area. The United States PHS Agencies including FDA, National Institutes of Health (NIH), Centers for Disease Control and Prevention (CDC), and Health Resources and Services Administration (HRSA) have worked together to address the risk of infectious disease transmission, publishing the “PHS Guideline on Infectious Disease Issues in Xenotransplantation” (US PHS, 2001). This Guideline discusses xenotransplantation protocols, animal source, clinical issues, and public health issues. Following publication of the PHS Guideline, FDA published a Guidance document entitled “Guidance for Industry: Source Animal, Product, Preclinical and Clinical Issues Concerning the Use of Xenotransplantation Products in Humans” to build on the concepts in the PHS Guideline, and provide more specific advice regarding xenotransplantation product development and production, and xenotransplantation clinical trials (US FDA, 2003b). Xenotransplantation is defined in the PHS Guideline and the FDA Guidance as any procedure that involves the transplantation, implantation, or infusion into a human recipient of either live cells, tissues, or organs from a non-human animal source or human body fluids, cells, tissues, or organs that have had ex vivo contact with live non-human animal cells, tissues, or organs (US PHS, 2001 and US FDA, 2003b). Xenotransplantation products are defined as live cells, tissues, or organs used in xenotransplantation. Examples of xenotransplantation products provided in the FDA Guidance are:

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Transplantation of xenogeneic hearts, kidneys, or pancreatic tissue to treat organ failure, implantation of neural cells to ameliorate neurological degenerative diseases.

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• •

Administration of human cells previously cultured ex vivo with live non-human animal antigen-presenting or feeder cells. Extracorporeal perfusion of a patient’s blood or blood component through an intact animal organ or isolated cells contained in a device to treat liver failure.

FDA encourages any potential sponsor of a xenotransplantation product to familiarize themselves with available documents that can be found on the FDA website (US FDA, 2006b). Gene Therapy FDA regulates human gene therapy products as biological products. The field of gene therapy holds great promise for treating a wide array of illnesses; from genetically inherited diseases such as cystic fibrosis or hemophilia, to heart disease, wound healing, AIDS, graft versus host disease, and cancer. In addition, the use of gene therapy in the area of tissue repair and tissue engineering is also being investigated. There are a number of safety issues associated with gene therapy, some of which are unique to this area. Safety issues specific to gene therapy trials include generation of replication competent virus, vector as well as transgene associated immunity, toxicity associated with transgene expression, and inadvertent germline transmission of vector. Two examples of gene therapy-specific risks are instructive: (1) high doses of adenovirus vector particles have been shown to induce toxicity under certain circumstances, and resulted in the death to a study subject in 1999; (2) genomic integration of retroviral vectors has been shown to result in genotoxicity, such that three children developed leukemia, and one died, as a direct result of altered gene expression after vector integration. Detailed recommendations from FDA regarding what type of information to submit in an early phase study of gene therapy products are available in the FDA “Draft Guidance for FDA Review Staff and Sponsors: Content and Review of CMC Information for Human Gene Therapy INDs Applications” (US FDA, 2004a). This draft guidance covers product manufacturing and characterization information (including components and procedures), product testing (including microbiological testing, identity, purity, potency, and other testing), final release testing criteria, and product stability, giving specifics in these areas that are pertinent to gene therapy. Suggested preclinical testing includes tests designed to describe localization, and persistence of gene expression. For vectors intended for direct in vivo administration, demonstration of the extent of dissemination and gonadal distribution is suggested. Gene therapies may differ from conventional drugs in that vector and transgene expression may persist for the lifetime of the subject. In these cases, there is a risk of delayed adverse effects. Indeed, the previously mentioned leukemias in a clinical study of gene therapy for the treatment of X-linked severe combined immunodeficient (SCID) did not occur until approximately 3 years after exposure to the retroviral vector. These events highlight the need for assessment of long-term risk in research subjects. FDA has discussed these issues, noting that the assessment of risk is based on the persistence of vector sequences, integration into the host genome, and transgene-specific effects. FDA has recently published the “Draft Guidance for Industry: Gene Therapy Clinical Trials – Observing Participants for Delayed Adverse Events” which addresses the duration and types of observations to be performed based on the patient population and the risks presented by the gene therapy product, and hopes to finalize such guidance in the near future (US FDA, 2005a). Although regulatory responsibility for gene therapy trials rests with FDA, NIH serves an important complementary role. In addition to funding a number of gene therapy research studies, NIH also provides an important forum for open public deliberation on the scientific, ethical, and legal issues raised by recombinant DNA technologies and its basic and clinical research applications through the Recombinant DNA Advisory Committee (RAC), an expert advisory committee to the NIH Director (US NIH, 2005). Clinical studies discussed in this forum include studies funded by NIH, as well as industry funded studies conducted at clinical sites receiving NIH funding for DNA recombinant research.

Overview of FDA Regulatory Process

The Preclinical Development Plan For device and drug clinical trials, the goal of preclinical development studies is to establish a scientific rationale for the clinical investigation and to demonstrate an acceptable safety profile. Traditional pharmacology/ toxicology safety studies are important to identify potential toxicity in target organs and tissues, and to obtain information on effective safe starting doses in humans as well as establishing a safety profile for dose escalation and/or clinical monitoring. For cellular therapy, gene therapy, and cell–scaffold combination products there are frequently additional product-specific safety questions that might need to be addressed prior to initiation of a clinical trial. For example, what is their potential to undergo unanticipated undesired changes in their characteristics, such as malignant transformation? For cell–scaffold constructs, are there safety issues associated with the implantation procedure or potential construct failure? Animal models have limitations that are confounded by anatomical as well as physiological differences between the animals and man. Many of these products, because of their novelty, do not have an established paradigm for preclinical evaluation and Sponsors are therefore encouraged to discuss their development plan with FDA early in the development process. In several specific clinical applications FDA has either had public discussion at an Advisory Committee meeting or has published Guidance documents, and in either case it is valuable for Sponsors to be aware of these discussions or publications. For example, in March 2005 the CTGTAC discussed the manufacturing, preclinical, and clinical issues in the development of cellular therapies for repair and regeneration of articular joint surfaces (US FDA, 2006c). In the preclinical session of this meeting the committee focused on addressing specific issues raised by the FDA with respect to the range of the subset of products (device, cellular, and tissue engineered (TE) products) that have been proposed for the repair and regeneration of articular cartilage however, the issues can be readily generalized to a broader range of regenerative medicine products. The major issues discussed were:

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The limitations and capabilities of available animal models for predicting safety and clinical activity. Pivotal animal toxicology studies designed to support a clinical trial of a cellular cartilage repair product. Additional safety concerns for allogeneic cellular products (versus autologous products) that should be addressed in an in vivo study prior to clinical trials.

The committee arrived at a consensus that animal studies were needed to evaluate potential products prior to human administration; however, it also was evident that there was no single animal model that is adequate to test all of the hypotheses involved in development of these complex products. For example, the committee acknowledged small animals might be useful in assessing novel biomaterials or mechanisms of action in early stages of product development. Unfortunately, due primarily, to anatomic and biomechanical considerations, large animals are needed to assess the potential clinical activity, therapeutic durability, and safety of a final product prior to initiation of clinical trials (“pivotal” toxicology studies). The committee also discussed the necessary study design features for pivotal toxicology studies and reached a general agreement that the design of such studies should incorporate:

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Assessment of the mechanism of action of the product by either interim sacrifice of subgroups of animals or by non-lethal methods of assessment such as imaging or arthroscopy (particularly if this use in animals can provide data to support use of these modalities in clinical trials by allowing comparison between in-life and necropsy assessments). Adequate duration prior to terminal sacrifice to demonstrate integration of the product with the native tissue. Histopathological examination of the site of implantation and surrounding regions, draining lymph nodes and remote organs with gross pathology at time of necropsy and histopathology of other organs as guided by data from preliminary studies.

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The committee discussed the potential need to use analogous animal cells in lieu of human cells in testing the final product construct and the limitations posed by this approach in terms of extrapolating the data to human trials. In recognition of this concern the committee highlighted the need to understand the physiology of the animal cells relative to their human analogs. The committee also discussed the potential for additional safety concerns posed by use of allogeneic cells in these products. Clinical Development Plan The goal of the clinical development program is to establish product safety and efficacy. In the field of regenerative medicine, variability in the product, as well as the patient, poses unique challenges in clinical trial design and conduct. An additional challenge is the need, with many of these products, to observe their integration into the host over a prolonged period. Specific feedback regarding the adequacy of certain proposed studies and predictors of clinical benefit can be provided to Sponsors through the use of meetings with the FDA at various development time points, and the use of a Special Protocol Assessment (for products regulated as a biologic) or an Agreement Meeting (for products regulated as a device) prior to initiation of their Phase 3 studies (US FDA, 2001e, 2002e). Cell–Scaffold Wound Healing Skin Constructs Some of the earliest attempts at cell–scaffold combination products are skin constructs and those products are worth a special mention in this chapter because there are a number of approved products on the market. Some of these products are composed of keratinocyte and/or fibroblast cells on a scaffold (e.g. animal collagen, glucosaminoglycans, or gauze); cell–scaffold wound dressings that function primarily as physical wound coverings are generally regulated as medical devices by CDRH. In contrast, regulatory oversight of cellular constructs formulated without a scaffold intended specifically to promote wound healing are assigned to CBER. Today, several cell–scaffold designs are commercially available. Dermagraft (a single cell construct of allogeneic neonatal fibroblasts seeded onto a bioresorbable mesh) was approved in 2001 as a medical device for treating full-thickness diabetic foot ulcers (US FDA, 2001c). In addition, in 2003 an HDE was approved for Dermagraft treatment of wounds associated with dystrophic epidermolysis bullosa (DEB). Apligraft and Orcel are two examples of commercially available bilayered co-culture constructs comprised of allogeneic neonatal keratinocyte and fibroblast cells on bovine collagen scaffolds. Apligraft is indicated for treatment of venous insufficiency and diabetic foot ulcers (US FDA, 2000). Orcel is approved for treating split thickness donor site wounds on burn patients (US FDA, 2001d). An HDE for Orcel use in recessive DEB patient (as an adjunct in covering wounds and donor sites after the surgical release of hand contracture and deformities) was approved in 2001. Information about the clinical performance of each product is available in the published literature and product labeling (Green and Rheinwald, 1975; Cazalet et al., 1988; Boyce et al., 1989; Hansbrough et al., 1992; Haeseker et al., 1993; Baird et al., 1998; US FDA, 2000, 2001c, d, 2005c; Currie et al., 2002). Information about the clinical studies, as well as other studies supporting approval, is available on the FDA website (US FDA, 2005c). FDA’s Standards Development Program Since its inception, the development and use of standards has been critical to the mission of FDA. The use of standards in FDA medical product regulation began with the 1906 Federal Food and Drugs Act. Drugs, defined in accordance with the standards of strength, quality, and purity in the United States Pharmacopoeia and the National Formulary, could not be sold in any other condition unless the specific variations from the applicable standards were plainly stated on the label (Federal Food and Drugs Act, 1906). In current times, Federal

Overview of FDA Regulatory Process

government agencies including the FDA are encouraged to use voluntary consensus standards, whether domestic or international, when performing regulatory activities in lieu of government-unique standards which are developed by the government for its own uses, when practical. Standard-setting activities include the development of performance characteristics, testing methodology, manufacturing practices, product standards, scientific protocols, compliance criteria, ingredient specifications, labeling, or other technical or policy criteria. As with guidance document development, in which Good Guidance Practices describes FDA procedures for developing and using guidance documents, there are specific regulations that describe FDA participation in outside standard-setting activities. Regulations governing this participation can be found in 21 CFR 10.95. Constructive FDA participation in organizations responsible for developing standards applicable to the products regulated by the agency is considered essential. The FDAMA of 1997 provides for the recognition of national and international standards in medical device reviews for IDEs, HDEs, PMAs, PDPs, and 510(k)s (Marlowe and Phillips, 1998). A “recognized consensus standard” is a consensus standard that FDA has evaluated and recognized for use in satisfying a regulatory requirement and that FDA has published in a Federal Register notice. A “consensus standard” is a standard developed by a private sector standards body using an open and transparent consensus process. Conformance with recognized consensus standards is strictly voluntary for a medical device manufacturer. A manufacturer may choose to conform to applicable recognized standards or may choose to address relevant issues in another manner. A complete listing of CDRH recognized consensus standards and guidance documents can be found on the CDRH website (US FDA, 2006e). American Society for Testing and Materials International (ASTM International) and International Standards Organization (ISO) exemplify two organizations FDA works with in standard development. ASTM FO4 Division IV is actively engaged in development of standards for tissue engineered medical products (TEMPS). F04 Division IV consists of five subcommittees: (1) Classification and Terminology, (2) Biomaterials and Biomolecules, (3) Cells and Tissue Engineered Constructs, (4) Assessment, and (5) Adventitious Agent Safety. Currently, the ASTM TEMPs group has developed 19 published standards and standard guides, and has approximately 25 draft standards under preparation. The first standards were for substrates, biomaterials, such as collagen, alginate, and chitosan, and for terminology, cells and cell processing, bone morphogenetic protein, and test methods. A standard guide for in vivo repair of articular cartilage has also been approved. FDA is actively engaged in standards development. In the long term, standards development will benefit the regenerative medicine community, and by extension, the public.

ADVISORY COMMITTEE MEETINGS As mentioned in Section “Meetings with Industry, Professional Groups, and Sponsors” above, because of the diversity of innovative technology evaluated by FDA review staff, FDA makes use of expert scientific Advisory Committees or Panels (for medical devices) to complement its internal review process. These advisors provide outside advice to contribute to scientific regulatory decision making. Outside experts can be asked to review data, or make recommendations about study designs across a product or clinical area; outside advisors can also be helpful at earlier stages of product development. Expertise on the advisory committee often includes scientific, statistical, and clinical experts, as well as consumer representation, patient advocates, and industry participation. Most meetings are public and there is an opportunity for public participation in the form of public comment. There are 30 Advisory Committees and 18 Advisory Panels for medical devices (as well as a number of subcommittees and one Department of Health and Human Services (DHHS) Committee administered by CBER). The areas of responsibility for the panels and committees are divided along product lines.

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The Advisory Committee for cellular, tissue, and gene therapy products, as mentioned earlier, is the CTGTAC. This committee has discussed a number of areas in recent years that are of potential interest to product developers in the regenerative medicine area, including:

• • • • •

Hematopoietic stem cells for hematopoietic reconstitution (February 2003). Allogeneic islet cell therapy for diabetes (October 2003). Somatic cell cardiac therapies (March 2004). Somatic cell therapies for joint surfaces (March 2005). Potency measures for cell, tissue, and gene therapies (February 2006).

The presentations for each topic, as well as a transcript of the discussion, are available on the FDA website referenced at the end of this chapter (US FDA, 2006c).

FDA AND CRITICAL PATH SCIENCE FDA recognizes the complexity of the scientific issues related to these products. FDA has introduced the Critical Path Initiative as a way to identify priority research areas that are expected to advance innovation in medical products. The Critical Path Opportunities List was recently published on March 16, 2006, and is available on FDA’s website (US Department of Health and Human Services, 2006). This list “presents specific opportunities that, if implemented, can help speed the development and approval of medical products” (US Department of Health and Human Services, 2006). “Moving Manufacturing into the 21st Century – Manufacturing, Scale-up and Quality Management” was specifically identified as a critical path element (US Department of Health and Human Services, 2006). Tissue engineering was specifically addressed as an area in need of additional critical path scientific development: “A key hurdle holding back innovation in tissue engineering is the difficulty in sufficiently characterizing a finished product to enable development of meaningful quality controls and release specifications. Often, conventional techniques, such as simple cell morphology, used to evaluate cell characteristics cannot be applied to these products because, for example, the engineered product may also include non-biological materials (e.g. support matrix). Consensus on how to assess these products and ensure manufacturing consistency would give product Sponsors the predictability they need to unlock innovation in tissue engineering” (US Department of Health and Human Services, 2006). FDA critical path research is aimed at bridging some of the challenges between product development and product evaluation. For example, following the observation of unexpected toxicity of adenoviral vector gene therapy in a clinical trial, CBER research provided insight into how adenovirus vectors cause toxicity, and developed an animal model for gene therapy in the context of preexisting liver disease (Raper et al., 2003; Smith et al., 2004). CBER researcher/regulators also worked with a consortium from industry and academia to develop reference material for adenoviral vector particles (Simek et al., 2002). An additional example of FDA critical path research is the active intramural and collaborative research program in the area of heart valves. The challenges that overlie the development, manufacturing, and characterization of a cell–scaffold tissue engineered heart valve (TEHV) center about the complexity of native heart valve biology and valvular tissue remodeling in response to a dynamic hemodynamic environment. A TEHV must be functional and durable at the time of implantation. The identification of factors that modulate the in vivo remodeling of a TE construct may be difficult to predict with certainty using in vitro methods (e.g. cell phenotype characterization); identification of cell products such as cytokines and growth factors; and physiological preconditioning in a bioreactor. Numerous biomarkers are available to describe heart valve biology; however, there are no validated surrogate markers of in vivo long-term TEHV performance and durability. New regulatory approaches, based on both sound engineering and biological principles, will have to be developed to assess the preclinical safety of these novel viable tissue products.

Overview of FDA Regulatory Process

CDRH has an integrated intramural and collaborative heart valve research program that to date has been primarily focused on the effects of preimplantation processing on heart valve tissue-derived biomaterials, cryopreservation and decellularization of allograft heart valves, and the identification of valve-related pathology and potential clinical failure modes. Extensive experience has been gained in the evaluation of replacement heart valves (mechanical, bioprosthetic, and polymeric), cryopreserved allograft, and TEHVs (Jones et al., 1982, 1989; Hilbert et al., 1987, 1990, 1992, 1994, 1999, 2004; Crescenzo et al., 1992, 1993; Hilbert and Ferrans, 1992; Schoen et al., 1992; Grehan et al., 2000). The majority of these investigations have involved preclinical in vivo safety and efficacy studies conducted in juvenile sheep. More recently this research program has focused on the identification of evaluative tools having potential utility for assessment of emerging TEHV and blood vessel constructs. (Hilbert et al., 2004; Xing et al., 2004a, b).

CONCLUSION The field of regenerative medicine is an exciting field with scientific advances leading to the promise of future therapies for current unmet medical needs for patients. The FDA regulatory approach to medical products evaluation includes an ongoing assessment of how the science of those products informs regulatory policy. FDA looks to continue ongoing dialog with the scientific community and product Sponsors to continue to develop science-based regulatory review policies that are robust and predictable in order to meet the needs of the challenging array of products that are on the horizon.

REFERENCES Baird, L.G., Christenson, L., David, J., Du Moulin, G., Gentile, F.T., Omstead, D.R., Maxted, D.D. and Tubo, R. (1998). Voluntary guidance for the development of tissue-engineered products. Tissue Eng. Fall 4(3): 239–266. (Review). Boyce, S.T., Cooper, M.L., Foreman, T.J. and Hansbrough, J.F. (1989). Burn wound closure with cultured autologous keratinocytes and fibroblasts attached to a collagen–glycosaminoglycan substrate. JAMA 262: 2125–2130. Cazalet, C., Cherruau, B., Jaffray, P., Marien, M., Schlotterer, M., Toulon, A. and Wassermann, D. (1988). Preliminary clinical studies of a biological skin equivalent in burned patients. Burns Incl. Therm. Inj. 14(4): 326–330. Changes to an Approved Biologics Licensing Application (BLA): 21 CFR 601.12 (2007). Crescenzo, D.G., Hilbert, S.L., Barrick, M.K., Messier Jr., R.H., Wallace, R.B. and Hopkins, R.A. (1992). Donor heart valves: electron microscopic and morphometric assessment of cellular injury induced by warm ischemia. J. Thorac. Cardiovasc. Surg. 103: 253–258. Crescenzo, D.G., Hilbert, S.L., Messier Jr., R.H., Domkowski, P.W., Barrick, M.K., Lange, P.L., Ferrans, V.J., Wallace, R.B. and Hopkins, R.A. (1993). Human cryopreserved allografts: electron microscopic analysis of cellular injury. Ann. Thorac. Surg. 55: 25–31. Currie, L., Jones, I. and Martin, R. (2002). A guide to biological skin substitutes. Br. J. Plast. Surg. 44: 185–193. Federal Food and Drugs Act, 1906, 34 Stat. 768, repealed by Food, Drug, and Cosmetic Act of 1938, 21 U.S.C. Sec. 329(a). Green, H. and Rheinwald, J.G. (1975). Serial cultivation of strains of human epidermal keratinocytes: the formation of keratinizing colonies from single cells. Cell 6: 331–343. Grehan, J.F., Hilbert, S.L., Ferrans, V.J., Salerno, C.T. and Bianco, R.W. (2000). Development and evaluation of a swine model to assess the preclinical safety of mechanical heart valves. J. Heart Valve Dis. 9: 710–720. Haeseker, B., Koch, R. and Teepe, R.G. (1993). Randomized trial comparing cryopreserved cultured epidermal allografts with tulle-gras in the treatment of split-thickness skin graft donor sites. J. Trauma 35(6): 850–854. Hansbrough, J.F., Dore, C. and Hansbrough, W.B. (1992) Clinical trials of a living dermal tissue replacement placed beneath meshed, split-thickness skin grafts on excised burn wounds. J. Burn Care Rehabil. 13(5): 519–529. Hilbert, S.L. and Ferrans, V.J. (1992). Porcine aortic valve bioprostheses: morphologic and functional considerations. J. Long-Term Eff. Med. Impl. 2: 99–112. Hilbert, S.L. and Hopkins, R.A. (1994). The ultimate challenge. In: Cardiac Reconstructions with Allograft Valves, 2nd edn. New York: Springer-Verlag, pp. 612–620.

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Hilbert, S.L. and Hopkins, R.A. Small intestinal submucosa pulmonary monocusp and mitral bi-flap valves implanted in sheep. FDA Cooperative Research and Development Agreement 74-03. Hilbert, S.L., Ferrans, V.J., Tomita, Y. and Jones, M. (1987). The morphologic evaluation of explanted polyurethane trileaflet cardiac valve prostheses. J. Thorac. Cardiovasc. Surg. 94: 419–429. Hilbert, S.L., Barrick, M.K. and Ferrans, V.J. (1990). Porcine aortic valve bioprostheses: A morphologic comparison of the effects of fixation pressure. J. Biomed. Mater. Res. 24, 773–787. Hilbert, S.L., Ferrans, V.J., McAllister, H.A. and Cooley, D.A. (1992). Ionescu–Shiley bovine pericardial bioprostheses: histologic and ultrastructural studies. Am. J. Pathol. 140: 1195–1204. Hilbert, S.L., Luna, R.E., Zhang, J., Wang, Y., Hopkins, R.A., Yu, Z.X. and Ferrans, V.J. (1999). Allograft heart valves: the role of apoptosis-mediated cell loss. J. Thorac. Cardiovasc. Surg. 117, 454–462. Hilbert, S.L., Schoen, F.J., Jones, M. and Ferrans, V.J. Allograft heart valves: morphologic, biomechanical and explant pathology studies. (1994) In: Cardiac Reconstructions with Allograft Valves, 2nd edn. New York: Springer-Verlag, pp. 193–231. Hilbert, S.L., Boerboom, L.E., Livesey, S.A. and Ferrans, V.J. (2004a). An explant pathology study of decellularized carotid artery vascular grafts. J. Biomed. Mater. Res. 69A: 197–204. Hilbert, S.L., Yanagida, R., Souza, J., Wolfinbarger, L., Linthurst-Jones, A., Krueger, P., Stearns, G., Bert, A. and Hopkins, R.A. (2004b). Prototype anionic detergent technique used to decellularize allograft valve conduits evaluated in the right ventricular outflow tract in sheep. J. Heart Valve Dis. 13: 831–840. Jones, M., Barnhart, G.R., Chavez, A.M., Jett, G.K., Rose, D.M., Ishihara, T. and Ferrans, V.J. (1982). Experimental evaluation of bioprosthetic valve implanted in sheep. In: Cohn, L.H. and Gallucci, V. (eds.), Cardiac Bioprostheses. New York: Yorke Medical Books, pp. 275–292. Jones, M., Eidbo, E.E., Hilbert, S.L., Ferrans, V.J. and Clark, R.E. (1989). Anticalcification treatments of bioprosthetic heart valves: in vivo studies in sheep. J. Cardiac Surg. 4: 69–73. Marlowe, D.E. and Phillips, P.J. (1998). FDA recognition of consensus standards in the premarket notification program. In: Biomedical Instrumentation and Technology. Philadelphia: Hanley & Belfus, Inc., pp. 301–304. PMA supplements: 21 CFR 814.39.(2007). Postmarketing Reports for Applications for FDA Approval to Market a New Drug (NDA): 21 CFR 314.80 and 314.81.(2007). Raper, S.E., Chirmule, N., Lee, F.S., Wivel, N.A., Bagg, A., Gao, G., Wilson, J.M. and Batshaw, M.L. (2003). Fatal systemic inflammatory response syndrome in a ornithine transcarbamylase deficient patient following adenoviral gene transfer. Mol. Genet. Metab. 80: 148–158. Reporting for Biological Products: 21 CFR 600.14, 600.80, 600.81, 601.28, 601.70, and 601.93.(2007). Schoen, F.J., Levy, R.J., Hilbert, S.L. and Bianco, R.W. (1992). Antimineralization treatments for bioprosthetic heart valves: assessment of efficacy and safety. J. Thorac. Cardiovasc. Surg. 104: 1285–1288. Simek, S., Byrnes, A. and Bauer, S. (2002). FDA perspectives on the use of the adenovirus reference material. Bioprocessing 1: 40–42. Smith, J.S., Tian, J., Lozier, J.N. and Byrnes, A.P. (2004). Severe pulmonary pathology after intravenous administration of vectors in cirrhotic rats. Mol. Ther. 9: 932–941. Supplements and Changes to an Approved NDA: 21 CFR 314.70, 314.71, 314.72.(2007). US Department of Health and Human Services, Food and Drug Administration: The Critical Path to New Medical Products (June 2006). http://www.fda.gov/oc/initiatives/criticalpath/. US Food and Drug Administration, Center for Biological Evaluation and Research: Tissue (April 2006a). http:// www.fda.gov/cber/tiss.htm US Food and Drug Administration, Center for Biological Evaluation and Research: Xenotransplantation Action Plan, FDA Approach to the Regulation of Xenotransplantation (April 2006b). http://www.fda.gov/cber/xap/xap.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Cellular, Tissue and Gene Therapies Advisory Committee, (Formerly Biological Response Modifiers Advisory Committee) (March 2006c). http:// www.fda.gov/cber/advisory/ctgt/ctgtmain.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Draft Guidance for FDA Review Staff and Sponsors: Content and Review of Chemistry, Manufacturing, and Control (CMC) Information for Human Gene Therapy Investigational New Drug Applications (INDs) (November 2004a). http://www.fda.gov/cber/gdlns/gtindcmc.htm.

Overview of FDA Regulatory Process

US Food and Drug Administration, Center for Biologics Evaluation and Research: Draft Guidance for Industry: Eligibility Determination for Donors of Human Cells, Tissues, and Cellular and Tissue-Based Products (HCT/Ps) (May 2004b). http://www.fda.gov/cber/gdlns/tissdonor.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Draft Guidance for Industry: Gene Therapy Clinical Trails – Observing Participants for Delayed Adverse Events (August 2005a). http://www.fda.gov/ cber/gdlns/gtclin.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Draft Guidance for Industry, Preventive Measures to Reduce the Possible Risk of Transmission of Creutzfeldt–Jakob Disease (CJD) and Variant Creutzfeldt–Jakob Disease (vCJD) by Human Cells, Tissues, and Cellular and Tissue-Based Products (HCT/Ps) (June 2002a). http://www.fda.gov/cber/gdlns/cjdvcjd0602.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Draft Guidance for Reviewers: Instructions and Template for Chemistry, Manufacturing, and Control (CMC) Reviewers of Human Somatic Cell Therapy Investigational New Drug Applications (IND’s) (August 2003a). http://www.fda.gov/cber/gdlns/cmcsomcell.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Good Clinical Practice Program: Guidances and Information Sheets on Good Clinical Practice in FDA-Regulated Clinical (April 2001a). http:// www.fda.gov/oc/gcp/guidance.html. US Food and Drug Administration, Center for Biologics Evaluation and Research: Guidance for Industry: Acceptance of Foreign Clinical Studies (March 2001b). http://www.fda.gov/cder/guidance/fstud.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Guidance for Industry, Formal Meetings with Sponsors and Applicants for PDUFA Products (April 2002b). http://www.fda.gov/cber/gdlns/mtpdufa.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Guidance for Industry: Guidance for Human Somatic Cell Therapy and Gene Therapy (March 1998a). http://www.fda.gov/cber/gdlns/somgene.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Guidance for Industry: Providing Clinical Evidence of Effectiveness for Human Drugs and Biologic Products (May 1998b). http://www.fda.gov/ cber/gdlns/clineff.pdf. US Food and Drug Administration, Center for Biologics Evaluation and Research: Guidance for Industry: Source Animal, Product, Preclinical, and Clinical Issues Concerning the Use of Xenotransplantation Products in Humans (April 2003b). http://www.fda.gov/cber/gdlns/clinxeno.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Guidances for Submission of IND’s (January 2006d). http://www.fda.gov/cber/ind/indpubs.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: ICH Guidance on Quality of Biotechnological/Biological Products: Derivation and Characterization of Cell Substrates Used for Production of Biotechnological/Biological Products (September 1998c). www.fda.gov/cber/gdlns/qualbiot.pdf. US Food and Drug Administration, Center for Biologics Evaluation and Research: Information on Submitting and Investigational New Drug Application for a Biological Product (June 2005b). http://www.fda.gov/cber/ind/ind.htm. US Food and Drug Administration, Center for Biologics Evaluation and Research: Points to Consider in the Characterization of Cell Lines Used to Produce Biologicals (July 1993). http://www.fda.gov/cber/gdlns/ptccell.pdf. US Food and Drug Administration, Center for Biologics Evaluation and Research: Points to Consider in the Manufacture and Testing of Monoclonal Antibody Products for Human Use (February 1997a). http://www.fda.gov/cber/gdlns/ptc_mab.txt. US Food and Drug Administration, Center for Devices and Radiological Health: Apligraft (Graftskin) Product Label (September 2000). http://www.fda.gov/cdrh/pdf/p950032s016.html. US Food and Drug Administration, Center for Devices and Radiological Health: CDRH Databases (December 2005c). http://www.fda.gov/cdrh/databases.html. US Food and Drug Administration, Center for Devices and Radiological Health: Dermagraft Product Label (September 2001c). http://www.fda.gov/cdrh/pdf/p000036.html US Food and Drug Administration, Center for Devices and Radiological Health: Device Advice: Investigational Device Exemptions (IDE) (July 2003c). http://www.fda.gov/cdrh/devadvice/ide/index.shtml. US Food and Drug Administration, Center for Devices and Radiological Health: Device Advice: Premarket Approval (PMA) (November 2002c). http://www.fda.gov/cdrh/devadvice/pma/.

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US Food and Drug Administration, Center for Devices and Radiological Health: Device Advice: Premarket Approval: Postapproval Requirements: General Requirements (November 2002d). http://www.fda.gov/cdrh/devadvice/pma/postapproval.html#general. US Food and Drug Administration, Center for Devices and Radiological Health: Device Advice: Premarket Notification [510(k)]: Is a New 510(k) Required for a Modification to the Device? (February 2005d). http://www.fda.gov/cdrh/ devadvice/3146.html. US Food and Drug Administration, Center for Devices and Radiological Health: Humanitarian Use Devices (November 2003d). http://www.fda.gov/cdrh/ode/hdeinfo.html. US Food and Drug Administration, Center for Devices and Radiological Health: Orcel Product Label (September 2001d). http://www.fda.gov/cdrh/pdf/p010016.html US Food and Drug Administration, Center for Devices and Radiological Health: Search STANDARDS Database (March 2006e). http://www.accessdata.fda.gov/scripts/cdrh/cfdocs/cfStandards/search.cfm US Food and Drug Administration, Draft Guidance for Industry and FDA: Current Good Manufacturing Practice (cGMPs) for Combination Products (September 2004c). http://www.fda.gov/oc/combination/OCLove1dft.html. US Food and Drug Administration, Early Collaboration Meetings under the FDA Modernization Act (FDAMA), Final Guidance for Industry and for CDRH Staff (February 2001e). http://www.fda.gov/cdrh/ode/guidance/310.pdf. US Food and Drug Administration, Guidance for Industry: Chronic Cutaneous Ulcer and Burn Wounds – Developing Products for Treatment (June 2006f). http://www.fda.gov/cber/gdlns/ulcburn.htm. US Food and Drug Administration, Guidance for Industry: Special Protocol Assessment (May 2002e). http:// www.fda.gov/cber/gdlns/protocol.htm. US Food and Drug Administration: Guidance for Industry, Sterile Drug Products Produced by Aseptic Processing – Current Good Manufacturing Practice (September 2004d). http://www.fda.gov/cber/gdlns/steraseptic.htm. US Food and Drug Administration, Proposed Approach to the Regulation of Cellular and Tissue-Based Products (February 28, 1997b). http://www.fda.gov/cber/gdlns/CELLTISSUE.txt. US National Institutes of Health, Office of Biotechnology Activities: Recombinant DNA and Gene Transfer (September 2005). http://www4.od.nih.gov/oba/Rdna.htm. US Public Health Service, PHS Guideline on Infectious Disease Issues in Xenotransplantation (January 2001). http://www.fda.gov/cber/gdlns/xenophs0101.htm. Xing, Y., He, Z., Warnock, J.N., Hilbert, S.L. and Yoganathan, A.P. (2004a). Effects of constant static pressure on the biological properties of porcine aortic valve leaflets. Ann. Biomed. Eng. 32: 555–562. Xing, Y., Warnock, J.N., He, Z., Hilbert, S.L. and Yoganathan, A.J. (2004b). Cyclic pressure affects the biological properties of porcine aortic valve leaflets in a magnitude and frequency dependent manner. Ann. Biomed. Eng. 32: 1461–1470.

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82 Current Issues in US Patent Law Patrea L. Pabst Intellectual property rights provide a means for the owners of technology to recover their investment in the technology and, in some cases, to make a profit. More importantly, intellectual property rights provide a means for financing the incredibly expensive research and development and testing required for commercialization of new products and processes in the medical and biotechnology field. When the intellectual property rights have been lost, many times it is not possible to obtain the money required to see a product or process reach the clinic and benefit those for whom it is intended. It is only by protecting the technology that it can be used to help those who need it the most.

WHAT ARE INTELLECTUAL PROPERTY RIGHTS? Intellectual property is intangible. It is embodied in patents, trademarks, service marks or trade names, copyrights, and trade secrets. Companies may have rights in such intellectual property by virtue of various licensing agreements, joint venture agreements, or by virtue of employment agreements. The Constitution of the United States gives Congress the power to enact laws relating to patents, in Article I, Section 8, which reads “Congress shall have power . . . to promote the progress of science and useful arts, by securing for limited times to authors and inventors the exclusive right to their respective writings and discoveries.” Trademarks and trade secrets are also typically protected under state law. Further protection can also be obtained using employment non-compete agreements, confidentiality agreements, and material transfer agreements. Extensions of patent rights and exclusive rights may also be obtained for delays in obtaining regulatory approval by the Food and Drug Administration (FDA), for Orphan Drug Act and under the WaxmanHatch act. These provisions are not dealt with in detail in this chapter. Patents Patents are grants by a government entity that gives the patent owner the right to exclude competitors from making or using that which is defined by the claims of the patent. Patents have basically the same requirements throughout the world. In the United States, the requirements for obtaining and asserting a patent are defined by Chapter 35 of the United States Code ( “USC”). Patents are governed exclusively by federal law. US Patents are granted exclusively by the US Patent and Trademark Office (PTO). Patents are granted on applications for new, useful, and non-obvious inventions. A patent is issued by the US PTO after it has examined the application for the patent and determined that it meets various criteria. Most other countries have similar laws governing granting of patents. In some cases, patents are obtained for a “region” such as the European Patent Convention countries, through a centralized patent office. The granted patents are then registered in the individual member countries in which they are enforceable.

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A patent for an invention is the grant of a property right to the inventor, issued by the PTO. Currently the term of a new patent is 20 years from the date on which the utility application for the patent was filed in the United States or, in special cases, from the date an earlier related application was filed, subject to the payment of maintenance fees and disclaimers to earlier expiring patents. Effective May 29, 2000, the patent term for subsequently filed utility applications, was extended to compensate for delays in prosecution due to the Patent Office. Design patents are granted for new, ornamental, and non-obvious designs, and have a maximum term of 14 years from the date of grant. US patent grants are effective only within the US, US territories, and US possessions. The right conferred by the patent grant is, in the language of the statute and of the grant itself, the right to exclude others from making, using, offering for sale, or selling the invention in the United States or importing the invention into the United States.1 What is granted is not the right to make, use, offer for sale, sell, or import, but the right to exclude others from making, using, offering for sale, selling, or importing the invention. In general, technology relating to tissue engineering is subject to the same rules as are other compositions and methods of use and manufacture thereof. As long as the “thing itself ” is new, non-obvious, and subject to written description, it is patentable. In the language of the statute, any person who “invents or discovers any new and useful process, machine, manufacture, or composition of matter, or any new and useful improvement thereof, may obtain a patent,”2 subject to the conditions and requirements of the law. The word “process” is defined by law as a process, act or method, and primarily includes industrial or technical processes. The term “machine” used in the statute needs no explanation. The term “manufacture” refers to articles which are made, and includes all manufactured articles. The term “composition of matter” relates to chemical compositions and may include mixtures of ingredients as well as new chemical compounds. These classes of subject matter taken together include practically everything which is made by man and the processes for making the products. The patent law specifies that the subject matter must be “useful.” The term “useful” in this connection refers to the condition that the subject matter has a useful purpose and also includes operativeness, that is, a machine which will not operate to perform the intended purpose would not be called useful, and therefore would not be granted a patent. In general, patentable subject matter includes compositions, methods of manufacture, and methods of use.2 Compositions may include, for example, biodegradable polymeric matrices seeded with cells, or isolated cells for implantation in the body. Methods of manufacture may be directed to, for example, a process for creating a unique prosthetic device. Methods of use may entail methods for administration of therapeutic compositions, or surgical implantation of, for example, a synthetic tissue matrix containing implanted isolated cells that secrete insulin. The most significant problems we observe in patenting in pharmacology and biotechnology are with the issue of written description and enablement. The PTO has provided guidelines for the written description requirement.3 The fundamental factual inquiry is “whether the specification conveys with reasonable clarity to those skilled in the art that, as of the filing date sought, applicant was in possession of the invention as now claimed.”3 Answering this question is not a single, simple determination, but rather is a factual determination reached by considering a number of factors. The PTO guidelines list the following factors which should be considered by an examiner in his written description analysis: (1) the level of skill and knowledge in the art; (2) partial structure, physical, and/or chemical properties disclosed in the specification; (3) functional characteristics disclosed in the specification, alone or coupled with a known or disclosed correlation between structure and function; and (4) the method of making the claimed invention disclosed in the application.

1

35 USC §271 (a) (2005). 35 USC §101 (2005). 3 Manual of Patent Examining Procedure (MPEP) §2163 (8th Edition, August 2001) (revised August 2005). 2

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The PTO has also provided guidelines for the enablement requirement under 35 USC §112 and its relationship to the utility requirement under 35 USC §101.4 The PTO guidelines focus on a three pronged test for determining whether an invention is “useful” within the meaning of the law: Does the invention have a utility that is specific, substantial, and credible? A specific utility is one that is particular to the subject matter claimed. A substantial utility is one that defines a “real world” use. Utilities that require or constitute carrying out further research to identify or reasonably confirm a “real world” context of use are not substantial utilities. A utility is credible unless the logic underlying the assertion is seriously flawed, or the facts upon which the assertion is based are inconsistent with the logic underlying the assertion. Inventions are defined by claims, which are supported by a detailed disclosure in a specification that tells one how to make and use that which is defined by the claims. For product claims that do not recite any utilities, disclosure or assertion in the specification of one specific, substantial, and credible utility meets the criteria of 35 USC §101. If no credible, specific, and substantial utility is asserted in the specification and none is well established, then the claims are rejected under 35 USC §101. Utilities that constitute curing or preventing a condition are sometimes not credible to one of the skill in the art and thus may raise a question under 35 USC §101. However, any rejection based on lack of credible utility must be supported by documentary evidence or sound technical reasoning by the examiner. Since most diseases or conditions can be treated (although not necessarily “cured”), rejections under 35 USC §101 for treatment claims rarely should be made. Each case is decided based on the specific facts. For example, since vaccines are regularly prepared to combat various viruses and organisms, vaccines would have a credible utility to one of the skill in the art. Thus, vaccines should not raise a question under 35 USC §101. Materials to be used for research, or methods of using those materials for research, raise issues of whether the utilities require or constitute carrying out further research to identify or reasonably confirm a “real world” context of use. One of the areas people frequently have questions on is when does something have utility when it is still in a research stage. For example, if one makes a matrix out of a particular experimental polymer, seeds it with cells, and shows the cells attach and proliferate in a petri dish under laboratory conditions, is that sufficient to make a claim to a matrix with cells seeded thereon to form an organ equivalent for transplantation? The key question is whether those skilled in the art would have a reasonable expectation of success – if they thought it more likely than not that the laboratory tests were predictive of success in a human or animal, then one has met the criteria for utility. What about when one clones a portion of a gene, the so-called “expressed sequence tags” or “ESTs.” Since one does not know what is encoded, even using the various analytical programs, one can only claim the isolated sequence, not the protein that may be encoded by the full length gene, nor its use. The ESTs themselves can be claimed for use as probes but this is a very limited use. Another frequent question is with regard to isolated cells, especially stem cells or fetal cells. These are clearly patentable subject matter, if they meet the other requirements of being novel and non-obvious, and one can either make a deposit with an approved depository such as the American Type Culture Collection, or provide an adequate written description of how to obtain the cells. Although the law provides for patenting of compositions, methods of manufacture, and methods of use, biotechnology can present a problem under US patent law when the subject matter moves away from the realm of the artificial or “things engineered by the hand of man” to a blend or chimera of “artificial” and “natural.”5 An example is when one blends cells and a matrix to form a cell matrix structure that is then implanted in a patient. Then, the matrix degrades to leave only implanted cells and/or the patient’s own tissue grows into an implanted matrix structure which then degrades. At what point do these materials become patient and not 4 5

MPEP §2107.02 (8th Edition, August 2001) (revised August 2005). Diamond v. Chakrabarty, 447 US 303, 206 U.S.P.Q. 193 (1980).

Current Issues in US Patent Law

patentable subject matter? Ethical issues may arise due to overlap between patient material and traditional subject matter, particularly in those cases involving dissociated isolated cells, biodegradable matrices for implantation, polymeric materials for altering cell/cell interaction (such as adhesion or restenosis), as well as materials for implantation that are designed to remain in the body, such as stainless steel hip replacements or cryopreserved pig valves. Interpretations of the statute by the courts have defined the limits of the field of subject matter which can be patented. Examples of materials that cannot be patented include the laws of nature, physical phenomena, and abstract ideas.6 Naturally occurring materials cannot be patented unless altered “by the hand of man”. Id. Thus, isolated cells are patentable; while the naturally occurring source – harvested tissue – is not patentable subject matter. A patent cannot be obtained upon a mere idea or suggestion. A complete description of the actual machine or other subject matter for which a patent is sought is required. Outside of the United States and Australia, methods of treatment of humans or other animals are generally not patentable subject matter. For example, although surgical instruments, drugs, or devices used in surgery are patentable, surgical treatments are not patentable subject matter. Therefore, one cannot obtain a patent on a method for surgically treating a patient. Typically, while this subject matter is not patentable, the compositions and methods of manufacture for use in treating patients are patentable subject matter. Claims may be obtained to the composition per se, which is to be implanted. In Europe, claims can be obtained to a first, or even a second, use of the material when the material itself is known. However, the patentability is quite limited in individual countries and in the European Patent Office, for policy and ethical reasons. Generally, Patent Offices in Asian countries are far less flexible than the European Patent Office in this matter. As a result, patent attorneys have adopted a number of strategic approaches to obtain protection equivalent to that which is available in the United States. For example, one may draft claims directed to methods of manufacture of such materials, as well as to methods of use that are defined by the composition rather than the method of use steps. An exception to the statute of what constitutes patent infringement in the United States was created by the legislature to prevent enforcement of patents for methods of surgery not involving a patentable drug or device. The patent cannot be enforced against the doctor performing the surgery or against the patient. The legislation arose as a result of a controversy over enforcement of a patent on a method for cataract surgery, which solely involved the way a surgical incision was made and available lenses inserted into the eye. The patent would have been enforceable if it utilized a patent device for cutting or inserting the lens, or a drug for treating the eye. Novelty Another requirement for patentability is novelty. In order for an invention to be patentable, it must be new, that is, it cannot be patented if: “(a) the invention was known or used by others in this country, or patented or described in a printed publication in this or a foreign country, before the invention thereof by the applicant for patent” or “(b) the invention was patented or described in a printed publication in this or a foreign country or in public use or on sale in this country more than 1 year prior to the application for patent in the United States . . .”7 Novelty means that no one, including the applicant for the patent, has publicly used or described that which is being claimed, prior to filing an application for patent. In the United States, there is an exception when the publication is made less than 1 year prior to filing of the patent application or when an application by a third party filed before the application of interest is published or granted after the application of interest is filed.8 The publication can be “removed” as prior art if the applicants are able to demonstrate that, prior to

6

Diamond, 447 US at 309. 35 USC §102 (2005). 8 35 USC §102 (a) and (e) (2005). 7

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the publication, they conceived and diligently reduced to practice what they are claiming. In the United States, one can also get into a proceeding between multiple parties who have filed patent applications on the same invention, called an interference, to determine who was first to invent, regardless of who was first to file for a patent, which in turn determines who is entitled to the patent. Outside of the United States, rights are awarded to the first to file, and patent rights are lost immediately upon publication of the invention if made prior to filing for the patent. What constitutes a publication? Generally, a publication is any oral, written, or physical description that conveys to the public that which an applicant would like to claim. It may be a talk at the proceedings of a society (including any slides presented), an article in a scientific journal, a grant application that is awarded (effective as prior art at the time of award, not of application), a thesis (effective at the time of cataloging), or even an offer for sale or a press release. A critical requirement is that the publication must be enabling, that is, it must convey to one of the ordinary skill in the art how to make and use that which is being claimed. Public use means more than using the composition or method in one’s laboratory. However, it can include even a one patient study that is reported during clinical rounds or at a presentation at which a drug company or surgical supply representative is present. The courts, in many cases, have had to interpret what it means to be publicly available. A frequent question is when is a student’s thesis available as prior art. Courts have now held that once the thesis is cataloged, it is publicly available, because it has been entered into a computer database that one searching the database will be able to access.9 Accordingly, the publication date of a thesis is the date on which the thesis is cataloged, not the date on which it is defended or signed by the thesis committee. Slides that are not distributed, but that are shown at an oral presentation, are considered to be publications, particularly if the meeting is attended by those skilled in the art who would be able to understand and use the information in the slides. Additionally, a poster presented at a scientific meeting constitutes a “printed publication” under 35 USC §102.10 Disclosures to another party under the terms of a confidentiality agreement are not publications. Uses that are strictly experimental may not be public disclosures, if, among other aspects, they are designed to determine if that which is to be claimed will work, and if any other parties who are involved are clearly informed that the studies are experimental in nature. If an announcement is made publicly, which does not enable one of the ordinary skill in the art to use or make that which is later claimed, then the announcement is not a publication. For example, an announcement could be a statement made to the press that researchers X and Y have discovered a cure for cancer. Since the announcement does not tell one of the ordinary skill in the art how to cure cancer, it is not enabling. However, enablement can be difficult to prove and standards may change over time. A 1995 case involved the question of whether a publication was enabling for the development of a transdermal patch for delivery of nicotine.11 The court found that a prior publication referring to transdermal patches for drug delivery mentioned that the drug in the transdermal patch for treatment of heart disease could be replaced with nicotine for assisting patients in quitting smoking. The court held that the article disclosed or made obvious the transdermal patch for delivery of nicotine claimed by the applicant, because the applicant merely took the transdermal patch described in the article, put nicotine in it, and demonstrate that the nicotine was delivered and would work exactly as predicted based on the delivery of the drug for treatment of heart diseases. Even though there was no information relating how the drug was to be incorporated into the transdermal patch, the publication was enabling because those of ordinary skill in the art would have been able to determine how to put the nicotine in the transdermal patch without undue experimentation.

9

Philips Elec. & Pharmaceutical Indus. Corp. v. Thermal & Elec. Indus., Inc., 450 F.2d 1164, 1169–72, 171 U.S.P.Q. 641 (3d Cir. 1971); Gulliksen v. Halberg, 75 U.S.P.Q. 252 (Pat. Off. Bd. Int’f. 1937). 10 In re Klopfenstein, 380 F.3d 1345, 72 U.S.P.Q.2d 1117 (Fed. Cir. 2004). 11 Ciba-Geigy Corp. v. Alza Corp., 1995 US App. LEXIS 28214, 37 U.S.P.Q.2d 1337 (Fed. Cir. 1995).

Current Issues in US Patent Law

If the invention has been described in a printed publication anywhere in the world, or if it has been in public use or on sale in this country before the date that the applicant made his/her invention or more than 1 year before the date on which an application for patent is filed in this country, a patent cannot be obtained.7 In this connection, it is immaterial when the invention was made, or whether the printed publication or public use was by the inventor himself/herself or by someone else. If the inventor describes the invention in a printed publication or uses the invention publicly, or places it on sale, he/she must apply for a patent before 1 year has gone by, otherwise any right to a patent will be lost. Obviousness Even if the subject matter sought to be patented is not exactly shown by the prior art, and involves one or more differences over the most nearly similar thing already known, a patent may still be refused if the differences would be obvious. The subject matter sought to be patented must be sufficiently different from what has been used or described before that it may be said to be non-obvious to a person having ordinary skill in the area of technology related to the invention. The claimed method or composition must be non-obvious to those of ordinary skill in the art from what is publicly known.12 This is usefully referred to outside of the United States as a requirement for an “inventive step.” In 1960s, the United States Supreme Court carefully analyzed non-obviousness and those factors that are to be considered in determining whether that which is claimed is obvious from the prior art.13 This analysis is a fact-based determination, involving not only the elements which are claimed, but also the level of skill in the art and the expectation or predictability that the claimed method or composition would perform as predicted, actual success in the marketplace, long felt need, and whether there are unexpected results. If one has no better than a 50–50 chance that a particular method may work, and the method works, it is arguably not obvious, although it may be obvious to try. If one tries something and the results are vastly different from what was expected, then the results are not obvious. For example, if one administered two drugs each in the dosage known to yield a particular effect and the combination results in a substantially greater effect than the sum of the individual effects of the two drugs, resulting in the ability to use a much lower dosage of each drug than expected, then one would have unexpected results or “synergy.” If the prior art teaches away from what the applicant has done, this result would support a finding of non-obviousness. For example, if the prior art states that one cannot proliferate implanted hepatocytes, then it may be non-obvious if the applicant for a patent finds that he can administer particular growth factors with his hepatocytes and see proliferation. Many other considerations factor into whether a claimed composition or method is obvious in view of the prior art. Enablement, Written Description, and Best Mode Additional requirements for patentability are defined by 35 USC §112, first and second paragraphs. These include enabling one of the ordinary skill in the art to make and use that which is claimed; a written description that clearly describes that which is claimed; and, in the United States only, disclosure of the best mode for practicing the invention.14 The applicant must describe that which is claimed in sufficient detail, and with appropriate methods and sources of reagents or other materials or equipment, to enable one of the ordinary skill in the art to make and use that which is claimed. This sounds far simpler than it actually is in practice. In many cases, particularly when coming out of a university study or a start-up company, the invention that applicants would like to

12

35 USC §103 (2005). Graham v. John Deere Co., 383 US 1, 148 U.S.P.Q. 459 (1966). 14 35 USC §112 (2005). 13

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claim is that which the applicant intends to develop over the next several years, based on a limited amount of data available at the time of filing. Particularly in the case of universities, where the applicant must publish or has submitted grant applications (which in and of themselves constitute prior art once they are awarded), the difficulty is in describing that which has not yet been done. The application must not only describe a specific limited example, but must describe the various ways in which one intends to practice that which is claimed. “Invention” usually consists of two steps: “conception” and “reduction to practice.” There are two kinds of reduction to practice: actual and constructive. “Constructive reduction to practice” means that the applicant has described in the application for patent how to make and use that which is claimed, but has not actually made and used what is claimed. This may be as simple as stating that although a biodegradable polymer such as polylactic acid–co-glycolic acid is preferred for making a matrix for culturing cells, other biodegradable polymers such as polyorthoesters or polyanhydrides could also be used. It may be less obvious that other cell types or shapes or methods of processing may be used when only one example showing reduction to practice is available. The rule of thumb in this case is the level of predictability. Therefore, in stating what kind of drugs one could deliver using the claimed technology, one might list a wide variety of drugs based on the data available with one type of drug. However, one may not be able to include in a list of drugs to be delivered peptides or very hydrophobic compounds, which usually are viewed as difficult to deliver, based on data obtained with a drug that is “easy” to deliver, such as a sugar or small molecular weight dye. Being too predictive (i.e. engaging in extensive constructive reduction to practice), which includes “non-enabling” or non-enabled technology, may in some cases be a detriment during prosecution of subsequently filed applications, because the examiner may cite the earlier work as making obvious the applicant’s subsequent work. Patent attorneys frequently must play a balancing game in determining how far to go with constructive reduction to practice in order to exclude competitors while not eliminating the applicant’s own ability to obtain additional, subsequent patent protection. In the United States, there is a requirement to disclose the best mode for practicing that which is claimed at the time of filing the application. No similar requirement exists outside of the United States. Because most applicants file the same application in the United States as outside of the United States, US applicants frequently disclose their best mode in foreign-filed applications. Patent Rights The purpose of a patent is to exclude the competition from making and using that which is claimed, not to “protect” a product – a frequent misconception of patents. In order to exclude competition, one must describe and claim not only that which one intends to practice, but that which another party could practice in competition with the patentee. What does this mean in real terms? It means that the applicant for a patent must describe his preferred method, which is known as of the date of filing, the preferred embodiments that he or his company intends to market, as well as any embodiments that a competitor could make and use in competition with the applicant’s product. Patent Term and Patent Term Extension Under the revised US patent law that was enacted as a result of implementation of the General Agreement for Trade and Tariffs (“GATT”), the term of a patent is 20 years from the original date of filing or the filing date of the earliest utility application to which priority is claimed. Applicants therefore have more incentive to prosecute all claims in a single application in order to minimize costs for prosecuting and maintaining the patent. Under the law in effect prior to June 8, 1995, the patent term was 17 years from the date of issue in the United States. Divisional applications were a commonly used method to extend patent protection to encompass different aspects of the technology over a period of time much greater than 17 years. For example, an application would be filed in 1990, and a single inventive concept (e.g. the composition) would be prosecuted

Current Issues in US Patent Law

in the first application. Three years later, when those claims were allowable and a patent was to issue, a divisional application would be filed with another set of the claims that had been restricted out of the original application. This divisional application would be prosecuted for another 2 to 3 years, the claims would be determined to be allowable, the second patent would issue with a 17-year term, and a third divisional application would be filed. The result is that patents on related technology would issue sequentially over several years, increasing the effective term of patent protection beyond 20 years. Under the new law, this mechanism to extend patent protection is not possible. The GATT was signed into law in the United States on December 7, 1994, and the initial provisions affecting US patent practice were implemented on June 8, 1995. The most significant changes arising from enactment of that agreement, now called Uruguay Round Act, were changes in the patent term in the United States, the implementation of provisional patent applications, and the broadening of what constitutes infringement in the United States. The change in patent term has been discussed above. For those applications filed before June 8, 1995, the term of any issuing patent is 17 years from the date of issue or 20 years from the filing date, whichever is longer. The term of any patent issued on an application filed June 8, 1995 or later is 20 years from the earliest claimed non-provisional priority date (i.e. the filing date of the first utility application to which priority is claimed or the filing date of a Patent Cooperative Treaty (PCT) application designating the United States). Extensions of terms are available upon delays in issuance arising from appeals or interferences. Additional extensions of terms are available for delays in obtaining regulatory approval by the Food and Drug Administration (FDA) for a device or a drug. Recent legislation has provided for an increased patent term due to unreasonable delays by the US Patent Office. Under changes implemented in late 2000, the US Patent Office recognized that losses in patent term were resulting from delays in prosecution of applications through no fault of the applicants. Diversion of user fees to other government agencies, in particular, as well as the initial “up time” associated with contracting out the publication and other services at the US Patent Office resulted in some cases in delays of months to even years. Excessive delays may now be the basis for extensions of patent term. Another result of the GATT was the implementation of provisional patent applications. These provide a means for applicants to make a preliminary filing, with limited claims or even no claims, and less stringent requirements as to form (but exactly the same requirements as to enablement, best mode, and written description), to preserve their initial filing date, while collecting additional data or financing or evaluating the technology. As long as a utility application is filed within 1 year claiming priority to the provisional application, one can benefit from the provisional application priority date. Provisional applications are not themselves examined and expire exactly 1 year after filing. They constitute a convenient way to preserve priority dates and provide a particular advantage in the medical and biotechnology fields, since patent term runs from the date of the utility application, not the provisional application, and therefore the effective patent term of an application claiming priority to a provisional application can be 21 years from the earliest priority date, not 20 years. Failure in the provisional application to completely disclose and enable that which is subsequently claimed in an utility application can result in a loss of the claim to priority to the provisional application, if that which is claimed is not enabled. Merely filing as a provisional application an article that will be published or presented in order to avoid loss of foreign rights usually will not comply with the enablement requirements, and therefore will not serve as an adequate basis for priority. It is essential that applicants who file provisional applications based on an article amplify the description to encompass other embodiments and to provide the basis by which one of the ordinary skills in the art can practice that which is claimed. Application sections that are not required for enablement, which are typically included in a utility application, include the background of the invention, the problems that the claimed invention addresses, and the claims. These sections can be omitted from the provisional application, thus saving time and money in preparing the application. In many cases, fairly standard language can be used to expand or broaden the description in an article in

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order to meet the enablement requirements, providing a means for those with limited amounts of time or money to protect that which they are disclosing with minimum risk and expenditure. The World Intellectual Property Organization (“WIPO”), which implements the provisions of the “PCT” and the European Patent Office, has confirmed that US provisional applications serve as an adequate basis for a claim to priority in corresponding foreign file applications. However, under the Patent Convention, all foreign applications that claim priority from an earlier filed application must still be filed within 1 year of the US filing date or the filing date of the country in which the first application is originally filed. Because biotechnology is a complex field, especially in the areas of patentable subject matter and enablement, many of the general observations relating to patents may not be as directly applicable to more “conventional” patentable subject matter. For example, the US Patent Office consistently maintained that the change from a 17-year term from the date of issue to a 20-year term from the earliest priority date would not result in a significant loss of patent term. However, because the Patent Office applies such a stringent examination proceeding under §112 (written description and enablement) in the biotechnology area, issue time in these cases typically has been considerably longer, not uncommonly taking as many as 5–7 years from the original priority date. The result is that these complex biotechnology patents have a substantially shortened term as compared to many other types of patents. This is the case even if one obtains some extension due to delays by the US Patent Office. Because a patent extension can still be obtained for delays due to regulatory issues involving the FDA, as well as for appeals to the Board of Patent Appeals and Interferences, those in the United States who believe that their patent rights will be limited in term due to delays in prosecution should avail themselves of the Patent Extension Act, if at all possible. One must bear in mind, however, that an extension for regulatory delays can only be obtained on one patent for any particular product or process; thus, the inventor or licensee with multiple, related patents clearly should choose the most important patent or the patent subject to the greatest increase in patent term, when facing such a situation. The patent that is to be extended must be brought to the attention of the FDA. Following FDA approval of the claimed product or process, the extension must be applied for in a timely fashion. Trademarks, Service Marks, and Trade Names A trademark is a word, design, shape, number or slogan, or a combination of those elements, that identifies and distinguishes one company’s products from others. It is a symbol of goodwill (e.g. Coca-Cola®, PC Paintbrush® and Budweiser®). A service mark is to services what a trademark is to products (e.g. SAKS® (retail department store services), Holiday Inn® (hotel services), and Orkin® (pest control services)). A trade name is a name used to identify a business, (i.e. a “commercial name”). A trade name can also be used as a trademark to identify products and as a service mark to identify the services (e.g. IBM®). Trademarks and service marks can be registered on the Federal Register with the US PTO if they are used in interstate commerce, or they can be registered in a state if they are only used within that state. Trade names cannot be registered federally. A trademark is created by use, not by filing of an application. Lack of registration is not a bar to enforcing the mark, but when a trademark owner has obtained a registration, it obtains significant procedural and substantive advantages. For example, federal registration provides constructive nationwide notice to all users of a trademark. This eliminates a second user’s defense of “innocent adoption” which might otherwise permit such a second user to continue to use the mark. Federal registration can also protect an internet domain name or email address from a disruptive challenge. “Intent to use” trademark applications allow a trademark owner to obtain “pre-approval” prior to investing money in the use of the mark. The date of filing of the intent to use application is the priority date for the mark and actual use must be shown within a period of not greater than 36 months following the approval of the mark by the US PTO.

Current Issues in US Patent Law

Copyrights A copyright protects original expression. It is totally a creature of statute, 17 USC, the US Copyright Act. The author of a copyrighted work has the exclusive right to use, distribute, modify, display, reproduce, and perform the work. The copyright protects its owner from actual copying or unauthorized use of the protected expression in the copyrighted work. It does not prohibit the independent creation of a similar work by a third party if that third party does not have access to, or knowledge of the work. A copyright also does not grant the copyright owner a monopoly on the ideas contained within the work. Copyright protection does not extend to any “idea, procedure, process, system, method of operation, concept, principal, or discovery” embodied in the work.15 Such concepts may be protected only through patents or trade secrets. Trade Secrets “Trade secret” means information including, but not limited to, technical or non-technical data, a formula, a pattern, a compilation, a program, a device, a method, a technique, a drawing, a process, financial data, financial plans, product plans, or a list of actual or potential customers or suppliers which derives economic value, actual or potential, from not being generally known to, and not being readily ascertainable by proper means by, other persons who can obtain economic value from its disclosure or use; and is the subject of efforts that are reasonable under the circumstances to maintain its secrecy. This definition suggests three main characteristics of a trade secret. First, a trade secret generally relates to some sort of data, formula, device, method, or other similar type of information. Second, a trade secret must be valuable. This quality is generally self-evident. If a trade secret were not valuable, then a company would not consider it to be worth keeping secret. Third, and most importantly, the information must be the subject of efforts that are reasonable under the circumstances to maintain its secrecy. An owner need not take every steps possible to maintain a secret but must take steps that are reasonable under the circumstances to maintain the secrecy. This typically means that employees will have employee agreements with non-compete provisions that the trade secrets are kept in a restricted area or under conditions wherein the documents cannot be downloaded by others and that parties accessing the information are aware that it is both valuable and confidential. Trade secrets are becoming increasingly valuable in the biotechnology and pharmaceutical areas, with the advent of databases containing information that allows one to accumulate and collate information on a regular basis. An example of a trade secret may be a database including information regarding drug design or what amino acids can be substituted or deleted, to achieve a particular activity. It may also be what kinds of genes are in a microarrary that is being marketed or what kind of controls are required to generate useful data. Trade secrets are undervalued assets by many investors who do not always understand that the trade off for a patent is that one must disclose who to make and use the claimed invention. To enforce rights in a patent, one must be prepared to enter into costly and time-consuming litigation. In contrast, trade secrets do not need to be enforced since the competitor never learns how to practice the trade secret. Of course, this only works if the information that is being maintained in secrecy is not subject to “reverse engineering,” that is, one cannot figure out the secret from the product that is being sold. In the case of a database, it is the product that is designed using the information that is being sold, not the information per se. Trade secrets can last much longer than a patent, and do not cost hundreds of thousands of dollars to obtain and maintain.

15

17 USC §102(b) (2005).

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WHY ARE INTELLECTUAL PROPERTY RIGHTS IMPORTANT? Advancements in the medical and biotechnology fields often are achieved only through substantial investment of industrial, academic, and governmental resources. Patenting of these technological advancements frequently is employed to recoup that investment, to create profits which are used in part to develop new or improved products, and to enhance a competitive commercial edge. Other forms of intellectual property protection, such as trade secrets, copyrights, and trademarks, may also be used to further protect and exploit drug delivery processes, products, and services. In the rapidly expanding fields of microarray screening and genomics, where databases are revised on a daily basis and much of the information does not need to be distributed, protection of only the results obtained using the data (i.e. trade secret) is far more effective than obtaining patent protection. One of the most frequently asked questions is why do we need to go to the trouble and expense of patenting a composition or method. The most common reason is that protecting a new composition or method of manufacture or use provides a means for obtaining the revenue required to develop a new drug or medical treatment. With the cost of developing and obtaining regulatory approval for a new drug approaching $200 million dollars in the United States, patent rights are essential to recovering expenses. For small companies that spend more time raising money than selling products, patents and patent applications represent the company’s only tangible assets which it can show to potential investors. For universities and other non-profit research institutions, patents and associated know-how and, in some limited cases, trade secrets can be used to obtain royalties from license agreements, sponsored research funding in many cases, and equity in new companies started for the purpose of exploiting the technology. Patents and other intellectual properties are valued in many different ways. For example, a process for manufacture typically would be licensed for 2–3% of the gross selling price of a product of the process. This price would be decreased if multiple licenses had to be obtained to use the process. Patents claiming compositions tend to have a greater market value, for example, between 5% and 10% of the gross selling price, due to the perception that these patents are easier to enforce than process patents. These numbers of course are affected by the stage in development, the market cap, and the number of competing technologies in the same field. Enforcement is a risky business. A good patent strategy is to obtain patents that claim a product, methods of manufacture, and methods of use, broadly and specifically, so that a patentee is able to assert multiple patents against an alleged infringer. Patents with broad claims generally will be easier to invalidate than more specific patents. Faced with the prospect of fighting several patents, most parties will opt for settlement. The alternative, litigation, is extraordinarily expensive for both parties and can result in the patents being invalidated, or the infringer being liable not only for damages for infringement, but also for attorney’s fees and punitive damages. Patents give the patent owner the right to exclude competition. This is accomplished by asserting the patent against third parties who are marketing a product or service which falls within the scope of the claims. Referred to as “infringement,” the criteria are totally different from the criteria for obtaining a patent, referred to as “patentability.” In simple terms, a patent claim consists of “elements” in a defined relationship. Certain phrases expand or limit the scope of the claim. For example, the term “comprising” can be translated as “including at least,” while “consisting” means “including only.” If a claim reads: Composition comprising:

• • •

A B C

then the claim would cover any composition including A, B, C, and any other component. Use of the term “consisting” in place of “comprising” would restrict the claim to a composition including only A, B, and C.

Current Issues in US Patent Law

In determining infringement, one must look to the claims of the patent. Claims may be clear on their face, or require reference to the specification, or description, of the patent. Claims also may be limited by agreements made during prosecution, a doctrine referred to as “file wrapper estoppel.” For example, if the prosecuting attorney argues that the claims distinguish over the prior art on the basis that the prior art does not disclose a particular feature that the attorney argues is essential to the claims in the patent, then the claims will be construed to require that limitation, even if not explicitly recited in the claims as issued. Asserting a patent allows the alleged infringer to file an action for declaratory judgment in a federal district court, asking the judge to declare the patent invalid or non-infringed. Litigation is very expensive and especially detrimental to small companies, thus providing a great deal of incentive to license the technology on terms favorable to both parties. Since patents are now published as applications, 18 months after the original filing date, the public gets the benefit of this knowledge prior to the period of exclusivity, since at least in the United States, patent rights are not enforceable until the patent is granted, and then only to the extent that the alleged infringing method or composition falls within the issued claims, although it is now possible even in the United States to backdate damages if the claims as published are essentially the same as the issued claims. A major limitation on the ability of a patent owner to enforce his patents lies in his ability to pay for the cost of litigation. In my experience, the cost of patent infringement litigation in the United States has gone from approximately $1 million in the mid-1980s, to $3 million by 1990, to five times or more that amount by 2000. This effectively prevents small companies and individuals from enforcing their patent rights, even with legislation allowing recovery of treble damages and attorneys’ fees in the event of willful infringement. Further, as discussed briefly above, in some cases, the rights to enforce a patent have been limited by legislative or other political action. In the United States, these rights are defined by 35 USC §271. 35 USC §287, which defines the limitations on damages and other remedies, was amended on September 30, 1996, in an apparent effort to prevent issuance of patents directed to new medical treatments. The amendment, which applies only to patents issued after September 29, 1996, deprives patentees of remedies for direct infringement and induced infringement of patents to surgical or medical procedures that do not involve patented drugs or devices.16 Although the exact scope of the exclusion is subject to interpretation of numerous terms in the amendment, it appears that the exclusion applies only to the performance by a medical practitioner (defined as a person licensed to provide medical activity or a person under the direction of such a person) of a medical or surgical procedure not involving the use of a patented machine, manufacture, or composition of matter (e.g. medical devices, implants, and drugs) or the patented use of a composition of matter. Thus, it appears that the use of patented medical devices and patented drugs will still be subjected to infringement remedies. The amendment does not apply to activities of persons engaged in the commercial development, manufacture, sale, importation, or distribution of a machine, or composition of matter or the provision of pharmacy or clinical laboratory services where such activities are (1) directly related to the commercial development, manufacture, sale, importation, or distribution of a machine, manufacture, or composition of matter or the provision of pharmacy or clinical laboratory services, and (2) regulated under the Federal Food, Drug, and Cosmetic Act, the Public Health Service Act, or the Clinical Laboratories Improvement Act. Although it is not clear how this exclusion will be interpreted, and as of 2005, this clause has not been discussed in any published legal opinion, it appears that infringement remedies will not be limited for performance of a medical or surgical procedure that is related to the commercial development or exploitation of a product or service where the medical or surgical procedure is subject to Federal regulation. 16

35 USC §287 (2005).

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From a practical perspective, patents are usually only enforced against manufacturers or distributors, not doctors, since it would be unduly burdensome, and typically not cost effective to pursue actions against many individuals rather than a single large entity. It is therefore doubtful that the change in law will have any detrimental impact on the health care industry in the United States. The absence of literal infringement does not necessarily mean that a process or device does not infringe a patent. According to the judicially created “doctrine of equivalents,” even though the language of a claim cannot be read literally upon a process or device, a claim can be infringed if the process or device “performs substantially the same function in substantially the same way to obtain the same result.”17 In 2002, the Supreme Court rendered its decision in Festo Corp. v. Shoketsu Kinzoku Koygo Kabushiki Co., addressing the relationship between prosecution history estoppel and the doctrine of equivalents.18 In an unanimous opinion, the Court held that prosecution history estoppel may apply to any claim amendment made to satisfy the Patent Act’s requirements (e.g. novelty, obviousness, enablement, written description, etc.); it is not limited to amendments made to avoid the prior art. The Court also held that prosecution history estoppel creates a flexible bar to the doctrine of equivalents. The Court explained that amending a claim to satisfy the Patent Act’s requirements need not bar suit against every equivalent to the amended claim element. By amending a claim, the inventor is deemed to concede that the patent does not extend as far as the original claim. A rebuttable presumption is made that the amendment to the claim limitation bars the patentee from asserting infringement against an equivalent to the narrowed limitation. The patentee can rebut this presumption and show that the amendment does not surrender the equivalent in question. For example, the patentee could establish that the equivalent was unforeseeable at the time the claim was drafted, the amendment did not surrender the particular equivalent in question or there was some reason why the patentee could not have recited the equivalent in the claim. In some cases it is possible to pursue alternative remedies, such as third party reexamination proceedings in the US patent office to have a party’s patent invalidated over art not considered during prosecution, or in Europe or Japan, by third party opposition proceedings. The latter are considerably less expensive than litigation in the United States and decisions are reached typically in less than 2 years, which is also quite quick compared to litigation in the United States. The results of litigation are also quite unpredictable but very risky in the medical and biotechnology fields. Especially in the biotechnology area, the late 1990s were known for the swift and drastic decisions by the Court of Appeals for the Federal Circuit, invalidating patents on the grounds the claims were not enabled by the specifications.

HOW DO YOU VALUE INTELLECTUAL PROPERTY? A common misconception is that one must have a patent in order to license the technology or for investors to be willing to invest in the technology. This is just not the case. In most cases where the technology arises in a university and many times with individuals, licenses are entered into with only a single application having been filed, and no indication of whether or not a patent will actually issue. Numerous start-up companies have been formed based on technology which is not the subject of US patents, and may in fact be better protected solely by trade secret and collaborative agreements. What kind of products is most attractive to investors? Certainly, in the medical area, drugs that are for known indications are attractive because the issues are simpler; one does not need PhD investors in order for the investor to understand the technology or the market. The regulatory issues are straightforward, unlike

17 18

Graver Tank & Mfg. Co., Inc. v. Linde Air Products Co., 339 US 605, 608, 85 U.S.P.Q. 328, 330 (1950). Festo Corp. v. Shoketsu Kinzoku Kogyo Kabushiki Co., 535 US 722, 122 S. Ct. 1831 (US 2002).

Current Issues in US Patent Law

with technologies such as gene therapy and stem cells where the political and ethical issues are in a constant state of flux. The development costs can be calculated with a higher degree of certainty. A single product usually is not enough to start a company with venture or angle funding. A single technology may be sufficient if it is basic, broad, and exclusive. There is a tremendous difference. A single product may be sufficient to start a company if the product is ready to go to market in a relatively short time frame and does not require huge amounts of funding. Certainly there are numerous examples in the e-commerce and software fields. In the medical areas, this is more difficult. An example of a company that is typical of the models that are being used successfully uses the combination of a very broad technology, at least one product in or close to clinical trials within a few months of funding, funding, and experienced investors who evaluate the technology, the scientists and the market, and who then put together the deal with the university or other source of the technology, form the scientific advisory board, and hire the business people to run the company. Absent this essential element, experienced business people who can put together the market plan, the people, and the money, even the best technology will fail. In summary, there are three things one must have to start a successful company: a business plan which describes a technology base, specific products, cost of development, marketing plans, and means of generating income; the individuals who can develop the technology; and the business people who can run the company and obtain funding. Failure to have anyone of the three will cause the company to fail. A frequent question is what are typical licensing terms. Generally speaking, I suggest the following guidelines. For a new drug, which is patented or patentable as a compound per se, percentages range from 5% to 15%, with 8% not being unreasonable. Variables include whether or not there is animal testing data, clinical data, formulation issues, regulatory issues, a crowded market. For example, a new Hepatitis C drug commands a high price – there are few if any good drugs on the market and the market is huge. A new HIV drug that is another variation of known drugs commands less of a price – market is crowded and it does not represent a radically different solution to the problem. For a diagnostic assay, antibody with desirable specificity to difficult antigen, medical device, screening assay with novel reagents, percentages typically range from 3% to 6% with 4–5% being common. For a method of manufacture or screening, the ranges are typically from 1/2% up to about 3–4%. Major variables limiting price include the need for multiple licenses to practice method, particularly if they have to be obtained from multiple third parties. Up front fees and milestones are highly variable. In the pharmaceutical and medical areas, minimum royalties are primarily designed to insure the technology comes back to the owner if it is not developed, but are low enough not to divert the needed resources from development. Highest milestones are on completion of regulatory steps, such as filing of a new drug application or completion of Phase I clinical studies. Initial fees are usually set at a number to help offset or recover patent expenses. These fees are usually the most contentious; if the product is a success, then both licensor and licensee win. In the case of advances, the licensor has put up money for patenting expenses and development which it wants to recover; the licensee wants to defer as much as possible while it evaluates the technology, so it loses as little as possible if the technology does not pan out. Advantages of licensing versus forming a company include:

• • • • •

Less risk to the technology owner. Greater likelihood of having experienced partner developing, testing, and obtaining regulatory approval. More likely to have sales and distribution team in place. No need to provide separate manufacturing and packaging facilities. Partner responsible for insurance and indemnification. Partner assumes liability and cost of defending technology against competitors.

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• •

Partner assumes responsibility for obtaining licenses to practice the technology, if any. Partner assumes cost of patenting, which can be in the tens to hundreds of thousands of dollars, especially if foreign rights are pursued.

Advantages of forming a company include:

• •

The party developing the technology retains control over the development. Some technology is too early stage to license. The company may be able to fund sufficient development of the product to license or partner the technology and both profit from the technology and see technology developed that would not otherwise be developed.

Potential for Higher Profit There are many successful business models. Some of those that I have seen failed because they did not have sufficient initial investment, did not immediately identify products, or focused on products that took too long to get to market. These were not cases where the technology was not good; rather, it was a case of where the business people failed. In contrast, several years after my experiences with “boom and bust” startups, I was privileged to watch a different model developed by Terry McQuire of Polaris Ventures in Cambridge, Massachusetts. Terry helped form a company called AIR (Advanced Inhalation Research), licensing in some very basic technology developed by Bob Langer of MIT and David Edwards of Penn State University. They had realized that pulmonary drug delivery was the next big target for drug delivery – potentially a means of delivering drugs currently deliverable only by injection, such as insulin. A number of other companies were working on other means for delivery such as transdermal drug delivery or ultrasound mediated drug delivery, but technical problems were delaying progress. The company was funded with very little investment money, and Terry, Bob, and David immediately set out to establish corporate deals as a source of revenue. In an extremely short time, they had deals with half a dozen companies to take their drugs and turn them into pulmonary formulations. The attraction of the technology was that it was quick to market – reformulating already approved drugs, using simple technology based on a single discovery – that “aerodynamic diameter” not “actual diameter” was determinative of whether or not the drug particles were delivered by inhalation into the “deep lung.” The funding for all of the development came from the corporate partners, the time to clinicals was less than 1 year, several products were being developed simultaneously, extremely broad, dominating patents were obtained, and the company was positioned for either sale or an IPO in a little over a year. Approximately 1.5 years after formation of the company, AIR was sold to Alkermes for an excellent return on their investment. A different model that I have seen implemented by David Scheer of Connecticut, over the last few years, has also been extremely successful. David identified markets with good growth potential and high market caps, companies that had done extensive work in developing products for that market but had made a business decision to move out of the market, then acquired the technology, intellectual property, and people to form new companies. The two companies I have had the most experience with are Esperion and Orapharma, both of which went public less than 2 years after being funded, and were then acquired shortly thereafter by Pfizer and Johnson & Johnson, respectively. Esperion initially acquired technology based on Apolipoprotein A1 Milano from Pharmacia Upjohn, including an extensive patent portfolio, and the people involved in manufacturing, regulation, and development. Esperion then expanded to obtain rights to some small molecule technology for treatment of atherosclerosis, then rights to liposomal formulations for treatment of atherosclerosis. Phase I data had been obtained by Pharmacia so that at least one product was positioned for Phase II clinical trials within a year of funding.

Current Issues in US Patent Law

Orapharma was set up similarly, with an initial product including extensive patent portfolio and people experienced with the science and regulatory sides of the technology, for treatment of periodontal disease using polymer-microsphere-based technology for administration of minocycline into the periodontal pockets. Phase I clinical trials had already been completed by American Cyanamid when it decided not to pursue the technology. Orapharma acquired all rights to the technology and used this as a basis for its IPO about 1.5 years after formation. The money from the IPO allowed the company to enter into phase II/III trials almost immediately, with regulatory approval and product on sale about 1 year after completion of its IPO. Simultaneously, it acquired rights to university technology which was much earlier stage for recruitment and regrowth of the bone in the periodontal pockets. No matter how the property is protected, or the technology valued, the important issue is that it be developed. The amount of innovation in the medical field has been truly extraordinary, even as the cost of developing it has escalated exponentially. Indeed, we now see companies that appear to conduct their entire research and development based on identifying products such as Eli Lilly’s Prozac with huge market shares and profits, reviewing the relevant patents to identify weaknesses, then spending millions of dollars to invalidate the patents, thereby not only achieving the right to market the product generically, but having a period of 180 days exclusivity under the Waxman-Hatch Act in which to recover the litigation costs. One can only hope that the current spate of litigation by generic companies to knock out patents will not discourage further innovation and development of new technologies.

1401

Index

1-Methyl-4-phenyl-1,2,3,6-tetra-hydropyridine (MPTP), 309, 331 1,3-Bis(2-chloroethyl)-1-nitrosourea (BCNU), 619 1,3-Bis(p-carboxyphenoxy)propane (CPP), 619 1,25-Dihydroxy vitamin-D3, 890, 895 5-Azacytidine, 233, 273, 330, 1040 “351” HCT/P, 1312, 1313–1314 “361” HCT/P, 1312, 1313–1314 510k clearance process, 1370 α -Smooth muscle actin, 76, 77, 79, 730, 972, 1005, 1006, 1081 β -Adregenic receptor signaling, 1041 β -Catenin, 52, 58, 130 β -Cells, 21, 102, 103, 405, 803, 805 differentiation, 404 regeneration insulin-producing cells, source of, 402–403 replacement therapy, 331–332, 398 benefits, 399–400 limitations to, 401–402 transplantation, 399–400 β -Glycerophosphate, 890, 895 ε -Caprolactone, 616, 641 ABA-type tricopolymers, 622 Abnormal offspring syndrome (AOS), 152 Acellular collagen matrix, 1110, 1126–1127, 1129, 1131, 1142, 1143, 1202 Acid citrate dextrose (ACD), 239 Acquired immune system. See Immunotoxicity Actin, 53, 484, 485, 487, 488, 667 Activin, 54, 229, 405, 430, 502, 852, 860, 890 Activin/nodal signaling, 129–130 Acute coronary syndrome (ACS), 818 Acute inflammation, 585, 704, 708–709, 727 Acute liver failure (ALF), 362, 1086 hepatocyte bridge technique, 916–917 hepatocyte transplantation, 917–918 Acute renal failure (ARF), 787, 788, 1117, 1132 cell therapy for, 1117–1118 Acute tubular necrosis (ATN), 1118

1402

Acute versus chronic SCI, 1252 Ad vector, in gene therapy, 449 Adaptive immune system, 469, 714, 715 Adeno-associated viral-mediated gene transfer, 198, 897, 899, 1238 Adenosine, 955 Adenoviral vectors, 198, 325, 1380 Adherens junctions, 51, 57 Adhesive environment, 539 dynamic change, 548–549 spatial patterning, 541–542 Adipogenesis, 233, 895 Adipogenic lineage, induction of of pluripotent stem cells, 232–233 Adipose tissue-derived stem cells (ATSCs), 746, 772 Adult myoblasts, 783, 784 Adult stem cells (ASCs), 5, 7–8, 20, 21, 29, 30, 32–34, 241, 259, 262–263, 302, 319, 347, 1177–1178, 1354 activation and proliferation, 101, 102–103 cardiac muscle, 106 cell-based therapies, 34–35 cell transplant, 110–111 chromatin of, 151 differentiation potential, 33 endothelial stem cells, 105–106 epithelial stem cells, 103–105 hematopoietic stem cells, 105–106 islet neogenesis from, 405–407, 891 mesenchymal stem cells, 105 plasticity of, 1073, 1119 skeletal muscle, 106 Adult tissues, 16, 35, 132, 345, 349, 351, 402 Advanced Cell Technologies, California, 374 Advanced Tissue Sciences, 1299, 1327 Adventitia, 528, 980, 1009 AffylmGUI, 567, 569 Affymetrix, 564 advantage of, 565 probe set, 567, 572 U133a, 569 AffyPLM, 566

Agarose gels, 38, 527, 680, 699, 1187, 1188, 1275, 1276, 1277 Aggrecan, 18, 69, 105, 494, 523, 524, 769, 1180 Agilent, 564 Aging effects, 108, 392, 468, 473, 474–475, 1061 Akt1, 330 Albumin, 248, 249, 261, 322, 357, 586, 668, 938, 1094, 1095, 1159 Aldehyde dehydrogenase, 251 Alginate, 114, 901, 1025, 1043, 1044, 1133, 1151, 1159, 1185 Alimentary tract colon, 1080–1082 esophagus, 1074–1076 immune function, 1073–1074 intestinal development, 1073–1074 small bowel, 1076–1080 stomach, 1080–1082 Aliphatic polyesters, 581–582 Alkaline phosphatase, 132, 230, 232, 494, 499, 875, 895, 1109 Alkanethiolates, 672 Alkanethiols, 661, 671 Allergy. See Hypersensitivity Alloderm®, 1076, 1299 Allogeneic cells, 23, 690, 717, 821, 1378 bone marrow, 770, 873 derived MSCs, 40, 773 for cell-based therapies, 17, 20–22 collagen, 595 ESCs, 111, 461 HSCs, 328 islet, transplantation, 400, 401, 806 MSCs, 22, 323, 873, 882, 892, 895 myogenic cell transplantation, 789 organ transplantation, 1375 SMCs, 1005 Allografts, 868, 882, 897, 899, 956, 1213, 1272, 1280 Allyl alcohol, 248 Alpha-1 antitrypsin deficiency (A1AT), 919 AlphaCor KPro, 1063 Alzheimer’s disease, 2, 243–244, 309, 562, 970 Ameloblasts, 1286–1287, 1288 American Civil War, 1324 American Liver Foundation, 1086 American Society for Testing Materials International (ASTM International), 1379 Amino acid-derived polymers, 617–618

Ammonium bicarbonate salt, 583 Amniocentesis, 226–227 Amniotic fluid, 730 isolation and characterization, 231–234 and placenta mesenchymal cells from, 227–230 pluripotent stem cells from, 230–234 Amniotic fluid stem cells (AFSCs), 8, 260, 571–572 Amniotic fluid-derived mesenchymal cells (AFCs), 227, 229 Amyotrophic lateral sclerosis (ALS), 115, 241, 242, 243–244, 309, 1262 Anagen, 103 Anaphylactic hypersensitivity, 716, 717, 719 Anatomic location, 387, 871, 1176 Androderm, 1151 Androgen replacement modalities, 1134, 1149, 1150 Angiogenesis, 75, 242, 246, 275, 310, 331, 441, 444, 589, 728, 814, 972, 974, 1020–1021, 1027, 1030, 1031, 1042, 1044, 1076, 1167 Angiogenic growth factors, 271, 440, 590, 690, 1023, 1025, 1026 Angiopoietins, 292, 449, 1022, 1023 Anisotropic nerve grafts, 1278 aligned anisotropic scaffolds, 1278–1279 cell-seeded longitudinally aligned NGCs, 1279 ECM molecules, 1279 neurotrophic factors, 1279 Anklylosing spondylitis, 474 Annexin V, 522, 557 Anophthalmia, 419–420 Anterior cruciate ligament (ACL) fibroblasts, 680, 888, 1209 versus medial collateral ligament (MCL) fibroblasts, 1206 reconstruction, 1213–1214 Anterior definitive endoderm, 403 Anterior–posterior axis, 141, 1211 Antiangiogenic treatment, 443–444 Antibody–antigen interactions, 662 Anti-CD49a antibodies, 322 Anti-CD105 antibody, 319, 321 Anti-fibrotic therapy, based on blocking overexpression of TGF-β1, 1241 Antigen presenting cells (APCs), 3, 22, 58, 714, 715, 806 Antigen receptors, 469 Antisense gene therapy, 1216, 1217

1403

1404

INDEX

Antisense oligonucleotides, 56, 193 Anti-Stro-1 antibodies, 321 Aorta-gonad-mesonephros (AGM), 230, 285 Apatite/collagen composite, 639 Apligraft, 3, 1005, 1299, 1378 Apoptosis, 52, 66, 72, 73, 76–78, 176, 194, 245, 272, 326, 557, 1260 associated proteins, 804 cell proliferation, 489–490 fetal wounds, 80 liver regeneration, 82–83 resistance, 471 signal transduction event, 73 smooth muscle cell, 818 wound healing, 76–78 Appendicular skeleton autogenous bone, 1165 autologous bone marrow, 1166 fracture repair, 1165 future directions, 1168 osteoconductive materials beta-tricalcium phosphate, 1167 ceramics, 1166 collagen, 1166–1167 osteoinductive materials, 1168 tissue repair factors, 1167–1168 Arginine–glycine–aspartic acid (RGD), 586, 694, 1067 Argument from nonenablement, 1338–1339 cloning, 1340–1341 objections, 1339–1340 ARPE-19, 854, 860 ArrayExpress, 563 Arteriogenesis, 271, 1022–1023 Arthritis, 770, 888 Articular cartilage, 18, 40, 325, 507, 520, 521, 525–526, 766, 900–901, 1188 cell therapy, 768 articular chondrocytes, 769–770 autologous versus allogeneic, 768 availability, 767–768 intra-operative versus culture expanded, 769 MSCs, 770–774 rationale for, 766–767 synovial cells, 774 cell-scaffold implants, 774–776 clinical outcomes, 776–777 defects, 766 scaffold-free constructs, 776

Articular chondrocytes, 769, 889 culture-expanded cells, 769–770 intra-operative cell therapy, methods, 769 Artificial cornea. See Keratoprostheses Arx gene, 103, 404 Ascl1, 423, 425 Ascorbic acid, 746, 876, 890, 895, 920, 1006, 1066 Asystolic donors, 344, 354 Atheromas, 1000, 1001 Atherosclerosis, 474, 519, 528, 529, 812, 813, 815, 816, 818–820, 826, 1000, 1400 Atomic force microscopy, 325 ATPase, 153, 915 ATryn, 168 Autogenous bone graft, 1165, 1168 Autograft, 868, 869, 870, 878, 879, 881, 883, 897, 1165, 1170, 1190, 1206, 1213, 1278, 1280 Autoimmunity, 402, 718 Autologen®, 595 Autologous bone marrow, 833, 872–873, 970, 985, 1166 in cell isolation methodology, 875 mobilization, 439 with optimized matrices clinical studies, 882–883 preclinical studies, 876–877 Autologous cell-based therapies, 3, 16, 17 Carticel, 18 disadvantage of, 20 Epicel, 17 physiology, restoration of, 20 structure, 18–20 Autologous chondrocyte implantation (ACI), 767–768, 769, 1190 Autologous chondrocyte implantation/transplantation, 900 Autologous chondrocyte transplantation, 1184 Autologous dermal fibroblasts, 904 Autologous human cells, 968, 1120 Autologous islets transplantation, 399–400 Autologous MSCs, 323, 326, 881, 895, 905 Autologous perichondrium, 767 Autologous scaffolds, 1184–1185 Autologous SKMBs, 820–824, 831, 832 Autologous tissue grafts, 1271 Autologous transplantation model, 773 Autologous versus allogeneic stem cells, 768 Autotransplantation. See Autologous islets transplantation AutoXpress Platform (AXP), 239

Index

Avidin–biotin interactions, 662 Avitene®, 800 Axial skeleton, 1168 future directions, 1170 graft options BMP-2, 1169–1170 BMP-7, 1170 DBM, 1169 Baby hamster kidney (BHK), 955, 1262 Bacon, 1349–1350 Balloon angioplasty, 1001 Basal lamina, 50, 51, 76, 387, 389, 391, 392, 666, 669 invasion of, 53 Basement membrane, 75, 76, 106, 487, 688, 729, 854, 894, 1009, 1064 Basic fibroblast growth factor (bFGF), 191, 213, 229, 304, 420, 428, 588, 589, 680, 691, 774, 853, 1001, 1006, 1011, 1215, 1258, 1276 Basic helix-loop-helix (bHLH) transcription factors, 57, 423, 425 Batten disease, 9 Bayesian model, 567 Bcl-2, 287, 804, 1027 Beckman, Eric, 1325 Benign ovarian teratomas, 574, 575 Benjamini and Hochberg’s model, 565 Bestrophin, 853, 860 Beta-casein promoter, 170, 171 Beta-III-tubulin, 233, 242, 244, 245 Beta-tricalcium phosphate (β-TCP), 1167 Bicistronic vector, 196 Biglycans, 769, 1185 Bilayer model, 523 Bile acid, 916, 921 Bioartificial kidney, 1116, 1119, 1123, 1132 Bioartificial liver (BAL), 365–366, 914, 928, 1087 biological issues, 936–937 cell introduction perfusion scaffold systems, 934 spheroid encapsulation, 934 cell sourcing issues, 936–937 cryosectioning, 939–940 designs computer-regulated bioreactors, 934–935 flat plate bioreactors, 930–931 future design, 935–936 hollow fiber bioreactors, 932–934

ex vivo cell maintenance, 928–930 functional analysis, 937–938 MRI analyses, 938–939 NMRS analyses, 939 nutrient inputs, static versus dynamic, 938 Bioartificial organs, 363, 365–366, 368, 369, 373, 1086 Bioartificial tissues, 109–110, 112–113 Biobrane™, 1299 Bioceramic scaffolds, 638 Biocompatibility, 605, 617, 660, 704, 1148 of aliphatic polyester, 581, 613 fibrosis, and fibrous encapsulation, 713–714 and host response, 1048–1050 immunotoxicity, 714–720 inflammation and wound healing, 705–713 Biodegradable polymeric scaffolds, 582, 680, 1006, 1109, 1131, 1145, 1157, 1201, 1289 Biodegradable synthetic polymers, 605, 613, 972, 1108, 1200 aliphatic polyesters, 581–582 amino acid-derived polymers, 617 block copolymers of polyesters with PEG, 618–619 crosslinked polymer networks, 620–621 polyesters, 621–625 peptides, 617–618 poly(amino acids), 617 polyanhydrides, 582, 619 polyesters, 613 poly(α-hydroxy acids), 613–615 polycarbonates, 617 polylactones, 616 polyorthoesters, 616–617 polyphosphazenes, 619–620 polyurethanes, 618 Biological issues, 359, 936 Biological product (42 USC 262(i)), 1369 Biological scaffolds, 580, 582, 595, 598, 1006–1008 macroporous biodegradable scaffolds, design principles of fabrication, 582–584 surface immobilization, 585–587 Biological signaling molecules, 1167, 1178 Biologics Control Act, 1367 Biologics License Application (BLA), 1311, 1370 Bioluminescence imaging (BLI), 328 Biomarkers, 192, 819, 834, 1373, 1380

1405

1406

INDEX

Biomaterials, 585, 1077, 1200 biocompatibility and bioresponse, 704 fibrosis and fibrous encapsulation, 713–714 immunotoxicity, 714–720 inflammation, 705–713 wound healing, 705–713 biodegradable synthetic polymers aliphatic polyesters, 581–582 polyanhydrides, 582 blood substitutes, 757–759 in cartilage repair, 1183–1187 control, 558–559 digit reconstruction, 1200 evolution, 113 genes and proteins, presentation of, 1024–1026 goal, 112, 605 history, 604 role of, 559, 1126–1127 scaffolds, 689, 692, 697, 896 design principles of, 582–590 selection, 581 surface modification biological modification, 662–664 chemical modification, 657–660 overcoating technologies, 660–662 surface chemical patterning, 664 topographical modifications, 660 sustained release, from macroporous scaffolds, 588–589 three-dimensional, 36, 580 Biomimetic biomaterials, scaffolds of, 498, 505–506 Biomineralization bone development, 744 fracture healing, 745 Bioreactors, 38, 697, 930, 992–995, 1029, 1087 in bioartificial liver, 366, 930–940 blood vessel bioreactors, 528–530 bone bioreactors, 493–494, 527–528 for cardiac tissue engineering, 494–495, 1050 cartilage bioreactors, 492, 525–527 constructions in, 1095–1096 animal models, 1088 clinical trials, 1087 culture models, 1089 cultures, 1011–1012 clinical applications, 1093–1094 differentiation, 1051–1052 oxygen supply, 1050–1051

human liver cell in, 1093 in vitro, 1089–1090 for cartilage growth, 1188–1189 key function, 490 modeling, 491 for stem cell culture, 1098–1099 types, 491–492 Biosafe, 239 Bioscaffolds, 1222, 1326 Bladder, 974–975 cell transplantation, 1129 matrices, 1129 seromuscular grafts, 1128–1129 smooth muscle progenitors, 574 tissue expansion, 1128 Bladder-derived acellular collagen matrix, 1127 Blastema, 104, 106, 107–108 Block copolymers, of polyesters with PEG, 618–619 Blood cells, regeneration of, 105 Blood–material interactions, 704, 705–706 Blood–spinal cord barrier (BSCB), 1248 Blood substitutes biomaterial-based, 757–759 blood transfusion, indications, 756–757 cardiovascular biomechanics, 759–761 cellular-based, 761–762 ethical considerations, clinical trials, 762–763 oxygen to non-oxygen carrying plasma expander, 756 Blood urea nitrogen (BUN), 1117, 1132 Blood vessel bioreactors, 528–530 Blood vessel substitutes, 39, 40, 1002, 1006 Bone, 897, 899–900 bioreactors, 527–528 fetal tissues, applications of, 972 formation of, 320, 520–521, 744, 745 fracture healing, 745 Bone-central third of the patellar tendon-bone (BPTB), 1213, 1220 Bone graft, 499, 745, 868, 876, 881, 1164 appendicular skeleton autogenous bone, 1165 autologous bone marrow, 1166 fracture repair, 1165 future directions, 1168 osteoconductive materials, 1166–1167 osteoinductive materials, 1168 tissue repair factors, 1167–1168

Index

axial skeleton, 1168 future directions, 1170 graft options for, 1169–1170 for bone repair, 870 carriers, 877, 882 craniofacial skeleton BMP-2, 1170–1171 BMP-7, 1170–1171 future directions, 1172–1173 PDGF, 1171–1172 Bone induction, 499–500, 747 Bone–ligament–bone complex, 1208–1209 Bone marrow (BM), 21, 308, 397 adherent cell population from identification and therapeutic use of, 774 centrifugation, 876 compartments, 318, 321 derived MSCs, 219, 310, 312, 314, 315, 316, 321, 322, 1107 derived progenitor cells, 820, 1005 derived SMC, 442 derived stem cells, 21, 268, 270, 774 HSCs, 21, 271 MSCs, 21, 273, 442, 444, 1040 and stem cell niches, interaction of, 271–272 transplantation of, 21, 22, 41, 274, 278, 442, 1040 see also Bone marrow stem cells Bone marrow aspiration technique, 873, 883, 1166 Bone marrow mononuclear cells (BMNC), 814, 816, 819–820, 824–826, 833–834, 1040 EPCs, 825–826 HSCs, 825 MSCs, 825 umbilical cord blood cells, 826 Bone marrow stem cells (BMSCs), 21, 111, 247, 268, 832–833, 955, 1120, 1292 and bone regeneration, 274 clinical applications, for cardiac regeneration, 278 clinical trial, 818–819 endothelial progenitor cells, 271 factors regulating homing and differentiation G-CSF, 276 IL-8, 277 morphogenic proteins, 277 SCF, 277 SDF-1, 277 TGF-β , 277@3:VEGF, 275–276 and heart regeneration, 272–274

hematopoietic stem cells, 269–270 in vivo differentiation, 275 and liver regeneration, 274–275 MSCs, 270–271 mobilization of, 439 and nerve cell regeneration, 275 and skeletal muscle regeneration, 274 and tissue regeneration, 272 Bone marrow stromal cells, 198, 277, 504, 505, 825 Bone marrow-derived cells (BMDCs), 1217 Bone morphogenetic protein (BMP), 100, 127, 128–129, 142, 213, 214, 277, 323–325, 498, 499–503, 505–506, 507, 643, 747–748, 872, 889, 901, 1168, 1169–1171, 1182–1183, 1291 BMP-2, 323–325, 502, 505, 878, 899, 905, 1169–1170, 1170–1171, 1183, 1185 BMP-4, 198, 199, 501, 502, 748 BMP-7, 502, 1168 BMP-9, 323, 325 Bone regeneration, 105, 498, 624, 906, 1164, 1165, 1167, 1201 biological ingredients osteogenic cells, delivery of, 870–872 and biomineralization, 744 and BMSCs, 274 bone development, 744 carriers and growth factors, 872–873 clinical needs, for therapeutic solutions, 868–870 fracture healing, 745 tissue engineering ESCs, 746 growth/differentiation factors, 747–748 MSCs, 322–325, 746 principles, 745 scaffolds, 748–752 Bone repair, 318, 323, 746, 815, 870, 871, 873, 874, 882, 1165 Bone–tissue engineering, 897, 899–900 Bone-to-bone healing, 1213 BOOST trial, 278, 832 Bovine IGF2 gene, 154–155 Brachyury, 326, 328 Brain-dead-but-beating-heart donors, 353, 354, 365 Brain-derived neurotrophic factor (BDNF), 244, 647, 1254, 1261, 1276, 1279, 1280 Bridge technique, 916, 917 Bruch’s membrane, 854–855 Burst-forming unit-erythrocyte (BFU-E), 240

1407

1408

INDEX

C-kit, 260, 277, 1040 C-reactive protein (CRP), 818, 1119 C3b, complement-activated fragment, 709, 712, 714 C3H10T1/2, cell lines, 323, 324, 326, 327 Cachexia, 902 Cadaveric allograft, 869, 1280 Cadavers, 1280 Cadherin, 51–52, 1179 Cadherin switching, 51 Caenorhabditis elegans, 192 Calcineurin, 525 inhibitor, 800–801 Calcium oscillation, 179 Calcium precipitation, 232 Calcium signaling, 1041 California, 372, 374, 1359 CAMP, 423, 524, 1261 Canada, 372, 797 stem cell research policy, 1357–1358 Canaliculi, 493, 889 Cancer, 150 applications to, 574–575 and stem cells, 351–352 T cells, 469 Carbon nanotubes, 556–557, 558 Carboxyfluorescein diacetate succinimidyl ester (CFSE), 290 Cardiac arrhythmia, 787, 1120 Cardiac-derived stem cells (CSCs), 826–831 Cardiac differentiation. See Cardiomyocytes Cardiac muscle, 106, 1051 Cardiac patch implantation, 1043–1044, 1052–1053 Cardiac stem cells, 7, 106 Cardiac tissue, engineering of, 273, 492, 1038 bioreactors differentiation, 1051–1052 oxygen supply, 1050–1051 cell therapy, 1039–1041 clinical applications, of BMSCs, 278 gene therapy, 1041–1042 implantation, of cardiac patches, 1052–1053 organ function, of engineered tissue, 1046 electrical conduction, 1047 endothelialization, 1047–1048 host response and biocompatibility, 1048–1050 mechanical elasticity, 1047 strength development, 1047 thrombogenecity, 1047–1048

tissue architecture, 1047 vascularization, 1048 scaffold-based approaches cell-based cardiac patches, 1043–1044 cell-free cardiac patches, 1043 fibrous scaffolds, 1044–1046 Cardiomyoctes, 7, 115, 200, 246, 273–274, 330–331, 826, 836, 1039–1041, 1043, 1051 Cardiothoracic tissue engineering, 1323 Cardiovascular applications, in regenerative medicine, 39–40 Cardiovascular biomechanics, 759–761 Cardiovascular disease (CVD), 268, 812, 1000 cell-based repair, 814–815 cell therapy, 813–814 Carmustine, 619 Carticel, 3, 18, 40, 900 Cartilage, 448, 767, 776, 901, 1176 biomaterials, in cartilage repair, 1183 autologous scaffolds, 1184–1185 natural scaffolds, 1185–1186 synthetic scaffolds, 1186–1187 bioreactors, 494–495, 525–527, 1188–1189 cell response and transduction mechanisms, 494 clinical needs for, 1176–1177 current, and future trends, 1190–1191 loading conditions, 494 mechanical determinants, 1188 morphogenesis, 506 by MSCs, 325–326 regeneration, 900, 901 repair techniques, 40, 597, 776 tissue engineering, requirements, 587, 775 translation, 1189–1190 types of, 889 Cartilage-derived morphogenetic proteins (CDMPs), 506 Catagen, 103 Cationic lipids, 194 Cauchy stress, 517 cbfa1, 232, 1183 Cbfa1/Runx2, 889 CD13, 271, 444 CD31, 199, 249, 260, 320, 1021 CD33, 445 CD34, 198, 201, 239–240, 242, 244, 245, 246, 248, 249, 250, 292, 320, 440, 444, 445 CD36, 70 CD38, 250, 292

Index

CD44, 69–70, 79, 260, 774 CD45, 198–199, 241, 246, 250, 260 CD105, 231, 318, 319, 320, 774 CD133, 244–245, 249, 250, 271, 440, 774 CD146, 246, 440 .CEL file, 565 Cell, as signal receiver and processor growth, differentiation, and apoptosis, 489–490 mechanochemical transduction, 487–489 receptors and sensors, 483 cell–matrix adhesions, 484–486 tensegrity model, 486–487 Cell-based cardiac patches, 1043–1044 Cell-based drug delivery cell encapsulation challenges, 958, 960 parameters, 956–958 cell-based protein factory, 960–961 clinical applications, 962 companies working on, 962–963 definition, 954 drug-loaded tumor cell system, 961–963 engineered cells, 955 primary cells, 954–955 Cell-based human corneal equivalents, 1066–1068 Cell-based protein factory, 960–961 Cell-based repair autologous SKMBs, 820–824 BMNCs, 824–826 cardiac-derived stem cells, 826, 831 cardiovascular regeneration, 812, 814–815 neovascularization, 818 requirements, 834–839 vascular integrity injury versus repair, 815–820 Cell-based therapies, 8, 16, 1072 adult stem cells, 34–35 allogeneic, 20–22 autologous, 17–20 commercialization, 22–23 ESCs, 34 and EGCs, 34 immunosuppressive properties, 22 production and best quality, 23 rationale for, 16 Cell-based tissue repair technique, 767–768, 1140 Cell–bioactive interactions, 672 cell adhesion, 672–673

cell motility, 673 cell proliferation and differentiation, 673–674 Cell biology, in musculoskeletal tissues chondrocytes, 889 muscle cells, 889–890 osteoblasts, 888–889 osteocytes, 888–889 Cell–cell adhesion, 51–52, 55, 57, 69, 544 Cell culture modeling, 470–471, 1029 Cell–ECM adhesion, 52–53, 484–486 Cell–ECM interactions, 66, 666–667, 673, 687, 894 composition and diversity, of ECM, 66–68 receptors for, ECM molecules, 68–70 during regeneration, 78–83 during skin wound healing, 73–78 signal transduction events, 70–73 Cell–ECM reciprocity, 677 Cell encapsulation cells and materials, usage, 957 parameters, 956–958 for testosterone therapy, 1151–1152 therapeutic delivery system, 955 Cell-enriched grafts, bone regeneration, 879–881, 883 Cell fate potential, 126 growth, differentiation, and apoptosis, 489–490 Cell-free cardiac patches, 1043 Cell fusion, 138–139, 302 Cell isolation techniques, for bone regeneration autologous bone marrow, on optimized matrices, 875 bone marrow centrifugation, 876 culture expansion, 876 enzymatic tissue digestion, 875 point-of-care osteogenic cell enrichment, 875 selective cell retention, 876 Cell line development, 163–165 copy number, 172, 173 FISH analysis, integration site, 172–173 neomycin selection, 172 timeline, 171–172 transfection, 172 Cell-mediated delayed hypersensitivity, 716, 717 Cell phenotype, 139–140 Cell receptors and sensors, 483–484 Cell response and transduction mechanisms in bone tissues, 493 in cartilage, 494 in vasculature, 491–492 Cell–scaffold implantation, 774–776

1409

1410

INDEX

Cell–scaffold wound dressing, 1378 Cell screening, of hESC, 573 derived RPE, 573 for genetic variability, 574 Cell-seeded tissue-engineered bladder, 1129, 1130 Cell seeding techniques, 639, 893, 930, 992–995, 1008–1009, 1256–1257 Cell source, 4–10, 955, 1005, 1040, 1374 for cardiac tissue engineering, 1040 ECs, 1003–1004 fibroblasts, 1004 for histogenesis, 689–690 human ECs, 1005 issues, of bioartificial livers, 936–937 for kidney, 1119–1121 liver support, 1091–1093 smooth muscle cells, 1004 tissue engineered vascular grafts, 992 Cell–substrate interactions, 667 biological properties, effect of bioactive surfaces, development of, 671–672 cell–bioactive surface interactions, 672–674 ligands, 671 chemical properties, effect of, 669 cellular response, 670–671 surface charge, 670 surface chemistry alteration, 671 surface wettability, 670 culture, substrates of, 676–677 importance of, 676 substrates, cellular responses, 677 mechanical stimuli, effect of, 677 cartilage/bone engineering, 680 ECM, cellular responses, 678–679 mechanotransduction, 678 tendon/ligament engineering, 680 vascular grafts, 679–680 physical properties, effect of crystallinity, 668–669 morphology, 669 stiffness, 669 compliance, 669 substrate importance, 667 three-dimensionality, effect of topography, effect of cellular responses, 675–676 fabrication techniques, 674–675 Cell surface markers, 441, 774

Cell-support matrices, 1108, 1200 Cell-surface receptors, 68, 70, 714 Cell survival, of myogenic-cell grafts early survival, 788–789 guarantee, 789 long-term survival, 789 Cell therapy, 1322, 1373–1375 of acute renal failure, 1117–1118 articular chondrocytes, 769–770 autologous chondrocyte implantation, 767–768 autologous versus allogenic, 768 for cardiac tissue, 1039–1041 cell isolation techniques, for bone regeneration autologous bone marrow, on optimized matrices, 875 bone marrow centrifugation, 876 culture expansion, 876 enzymatic tissue digestion, 875 point-of-care osteogenic cell enrichment, 875 selective cell retention, 876 clinical trials, 831–834, 837 clinical studies, for bone regeneration autologous bone marrow, with optimized matrices, 882–883 cell-enriched grafts, 883 commercialization, 22–23 future developments, for bone regeneration, 883–884 G-CSF, 833 genetically modified, 1027 hepatic stem cells, 365 intra-operative, methods for, 768 versus culture expanded, 769 for ligament and tendon healing, 1217–1218 mature, 1026–1027 MSCs, 770–774 for myocardial infarction, 246 osteogenic cells, source of allogeneic bone marrow, 873 autologous bone marrow, 872–873 gene therapy, 874–875 novel tissue sources, 873–874 perichondrial transplantation, 767 periosteal transplantation, 767 peripheral arterial disease, 833–834 preclinical studies, for bone regeneration, 876 autologous bone marrow, with optimized matrices, 877–879 cell-enriched grafts, 879–881 progenitor and stem, 1027

Index

rationale for, 766–767 reparative potential autologous SKMBs, 820–824 bone marrow mononuclear cells, 824–826 cardiac-derived stem cells, 826–831 for spinal cord injury, 1262–1263 Cell therapy, reparative potential of autologous SKMBs, 820–824 bone marrow mononuclear cells, 824–826 cardiac-derived stem cells, 826–831 Cell transplant, 913, 916, 918 adult stem cells, 110–111 ESCs, 111–112 fetal cells, 109–110 Cell types in adaptive immune system, 706, 715 in SCNT, 175–176 Cell-associated mineralization, 232 “Cell-painting” technologies, 884 CellbioreactorModule, 1094–1095, 1096 Cellular microenvironments, engineering dynamic changes adhesive environment, 548–549 soluble environment, regulation, 548 microengineered tools micropatterened screening arrays, 539–541 spatial patterning, 541–542 microfluidics, control soluble cues, 537–539 multicellular constructs, organization heterotypic interactions, 546 homotypic interactions, 544–546 three-dimensional patterning, 546–548 two-dimensional patterning, 544 substrate mechanics engineering MEMS devices, 542–543 patterning stiffness, 543–544 Cellular retinaldehyde-binding protein (CRALBP), 853 Cellular, Tissue and Gene Therapies Advisory Committee (CTGTAC), 1374, 1375, 1377, 1380 “Cellular vacuum”, 361–362, 364 Cellular-based blood substitutes, 761–762 Cementum, 1287 Center for Biologics Evaluation and Research (CBER), 1309, 1310, 1369 Center for Devices and Radiological Health (CDRH), 1309, 1369, 1378, 1379, 1381 Center for Drug Evaluation and Research (CDER), 1309, 1369

Centers for Disease Control and Prevention (CDC), 888, 1375 Ceramics, 899, 1327 osteoconductive materials, 1166 scaffold, 323, 641–644, 1290 Cerberus-like, 142 Chang medium, 231, 233 Chemical-field effect transistors (ChemFETs), 1123 Chemokine, 67, 74, 83, 275, 537–538 Chemotactic cues, 1256 Chemotaxis, 499–500, 502 Chest trauma, 1323 Chimeric proteins, 1025 ChiP on ChiP assays, 258 Chitosan, 595–597, 1186 Cholera toxin, 428–429 Chondral lesion, 766 Chondrocytes, 112, 250, 524, 527, 680, 766, 767, 889, 1148–1149, 1180–1183 bioreactor design, 494–495 cell response and transduction mechanisms, 494 cell therapy, 507 transplantation, 19 Chondrogenesis, 231, 502, 889, 895, 1182–1183 Chondroitin sulfate proteoglycans (CSPGs), 306, 1261 Chorionic villus sampling (CVS), 226, 231 Chronic inflammation, 617, 707, 709–710 Chx10, 425, 427, 430 Ciliary epithelium, 427 Ciliary marginal zone (CMZ), 424–425 Ciliary neurotrophic factor (CNTF), 1262 Cincinnati knee rating system, 777 Circulating progenitor cells, 439, 442, 832 Citrate phosphate dextrose (CPD), 239 ckitneg cells, 232, 233 ckitpos cells, 231–233 Claudin, 57 Cloned cattle, transgenic production in, 182–184 Cloned goats, transgenic production in, 181–182 Cobblestone area forming cells (CAFC), 290 Cochet–Bonnet corneal, 1068 Collagen, 351, 594–595, 680, 751, 903, 1142, 1254 osteoconductive material, 1166–1167 scaffolds, 974, 1006–1007, 1185, 1275–1276 sponge, 1046, 1047, 1049, 1052 and TERP5 hydrogels, 1067–1068 types, 1207–1208 Collagen/alginate composite hydrogels, 646

1411

1412

INDEX

Collagen/elastin/glycosaminoglycan/chondroitin scaffolds, 645 Collagen/hyaluronan composite hydrogels, 646 Collagen Meniscal Implant (CMI), 595 Colloid osmotic pressure, 759 Colloids, 757 applications of, 759 Colon, 1081–1082 Colony-forming unit in cell culture (CFU-C), 290 Colony-forming unit-erythrocyte (CFU-E), 240 Colony-forming unit-granulocyte, erythrocyte, macrophage, megakaryocyte (CFU-GEMM), 240 Colony-forming unit-granulocyte, macrophage (CFUGM), 240 Combination product (21 CFR 3.2(e)), 1370–1371 Combined ACL/MCL injuries, 1214 Committed progenitors, 347 Common lymphoid progenitors (CLP), 105, 262, 288 Common myeloid progenitors (CMP), 262, 288 Company formation advantages, 1400 versus licensing, 1399–1400 Compensatory hyperplasia, 101–102 Complement system, 714 Composite scaffolds, 622, 639–648, 749–750, 1182, 1202 Composite tissue structures, creation of, 1201–1203 “Composition of matter”, 1387, 1397 Compression models, of SCI, 1251–1252 Computational fluid dynamic (CFD) software, 491 Computer-regulated bioreactors, 934–935 Congestive heart failure, 2, 39, 40 Connected cellular network (CCN), 493 Connexin-32, 915 Connexin-43, 1046 Constitutive equation, 517–519 Contigen®, 595 Continuous venovenous hemofiltration (CVVH), 1117 Contusion models, of SCI, 1251 Copyrights, 1395 Coral, 899 CorCap, 1043 Cord blood hematopoietic and tissue regeneration, 240–241 neurological regeneration aldehyde dehydrogenase expressing cells, 251 ALS, 243–244 Alzheimer’s disease, 243–244 cardiac treatment, 245, 246–247 chondrocytes, 250

diabetes, treatment of, 247 endothelial progenitors, 249–250 ex vivo expansion, 250–251 hepatocyte-like cells, 248–249 Huntington’s disease, 243–244 myocardial infarction, 245–246 neural cell surface markers, 244–245 Parkinson’s disease, 243–244 spinal cord and CNS injuries, 244 stroke, 241–243 pluripotent cells from, 241 procurement and processing, 239–240 storage, 240 transplantation, 238 see also Umbilical cord blood Cord blood-derived embryonic-like stem cell (CBE), 248 “Core and skirt” concept, 1062 Cornea, in regenerative medicine allograft transplantation, disadvantage of, 1061 cell-based human corneal equivalents, 1066 advancement, in tissue engineering, 1068 collagen–synthetic polymer matrix replacement, 1066–1068 cellular layer of, 1060 design requirements, for human corneal replacement, 1061–1062 keratoprostheses, 1062 development of, 1063–1066 Cornea allograft transplantation, 1060, 1061 Corneal epithelium, 103, 1064, 1067 Corneal ulceration, 1060 Coronary artery bypass graft operation, 813, 1001 Corporal smooth muscle, 1130 Corpus cavernosum reconstruction acellular collagen matrix preparation, 1145–1146 human endothelial cell culture, from foreskin, 1146–1147 smooth muscle cell culture, 1144–1145 Corticosteroids, 800, 1237 Covalent coating method, 661 COX enzymes. See Cyclooxygenase enzymes CpG dinucleotides, 152, 154 Craniofacial MSCs, 322 Craniofacial skeleton, 1167 BMP-2, 1170–1171 BMP-7, 1170–1171 future directions, 1172–1173 PDGF, 1171–1172

Index

Craniofacial tissue, 904–905 Creutzfeldt–Jakob Disease (CJD), 1061 Crigler–Najjar (CN), 919–920 Critical Path Opportunities List, 1380 Critical-sized defects (CSD), 624, 1164 Cross-differentiation, 302 Cross-linked polyesters fumarate-based polymers, 621–624 polymers containing acrylate, methacrylate, or vinylsulfone functionalities, 624–625 Cross-sectional area (CSA), 108, 517, 938, 1208 Cryopreservation stem cells versus mature cells, 362–363 Cryosectioning, of bioartificial livers, 939–940 Crystallinity, 614, 668–669, 1005 CT scans, 949, 1170 Culture-expanded cells in articular chondrocytes, 769–770 in bone regeneration, 881–882 Cultured autologous articular chondrocytes. See Carticel Cultured autologous epidermal keratinocytes. See Epicel Current Good Manufacturing Practice (cGMP), 401, 1371 Cyclin-dependent kinases, 101, 285 Cycloheximide, 178 Cyclooxygenase enzymes, 1236 Cox-2, 114, 525, 974, 1236 Cyclosporine, 947, 951 Cytofluorimetric analysis, 231 Cytokines, 22, 67, 194, 213, 243, 271, 275, 303, 445, 471, 646–648, 716, 819, 833, 1233–1234, 1235, 1276 Cytomegalovirus (CMV), 196 Cytoplasmic calcium, 523 Cytotoxic hypersensitivity, 717, 718 Cytotoxic T cells (CTLs), 469–470, 575 D407, 573, 860 Daclizumab, 800 Dacron tube, 1075 .DAT file, 565 Data and Safety Monitoring Board, 948 DCHIP, 566 Decorin, 68, 76, 80, 1216, 1241 Dedifferentiation, 17, 106–108, 774, 1187 Definition of HSCs, 388 MSCs, 319–320, 888 regeneration, 101, 498, 688 regenerative medicine, 35, 38–39, 1322

Definitive endoderm, 403, 405 Definitive hematopoiesis, 285 Degeneration phase, in skeletal muscle healing, 1233–1234 Degenerative disk disease, 904 DegraPol, 1327 Delta, 108, 140, 392, 404 Delta-crystallin enhancer-binding factor 1 (δ-EF1), 57 Demineralized bone matrix (DBM), 498–500, 617, 747, 899, 1167 Demineralized freeze-dried bone graft (DFDBA), 1172 De novo methylation, 153 De novo methyltransferase, 154 Dental stem cells (DSCs), 1287, 1288–1289 Dental tissue engineering, 1286 dental stem cells, 1286, 1288–1289 dentin and reparative pulp, 1291 morphogen-based cell signaling gradients, 1289 natural tooth tissue development and repair, 1286–1288 periodontal tissue regeneration, 1291–1292 prior and current researches, 1290–1291 scaffolding materials, 1289–1290 whole-tooth tissue engineering, 1292–1293 Dentin, 105, 1287, 1291 Department of Health and Human Services (DHHS) Committee, 1379 Dermagraft, 3, 1299, 1378 Dermal fibroblasts, 904, 1053 Dermal replacement, 1327 Dermal ulcers, 20, 725 Dermal wound model, 76 Desaminotyrosyl-tyrosine ethyl ester (DTE), 617 Design criteria scaffolding, 748 Design parameters, for histogenesis biomolecular factors, 694–695 cell sources, 689–690 degradation, 692–693 microvasculature, 690–691 porosity, 691–692 Design patent, 1387 Desmin, 233, 1021 Desomosomes, 51 Determined stem cells, 347, 349, 354, 363, 365, 367 identification, 348 purification, 348 sourcing issues, 353–354 DetoxModule, 1094, 1095, 1096

1413

1414

INDEX

Developmental mechanisms, of regeneration and strategies, 100 regeneration mechanisms, 101 adult stem cells, activation of, 102–106 aging effects, 108 compensatory hyperplasia, 101–102 dedifferentiation, 106–108 regenerative medicine, strategies of, 109 bioartificial tissues, 112–113 cell transplants, 109–112 chemical/physical induction, of repair and regeneration, 113–117 Developmental neurons, 303 Developmental origin, of hematopoiesis, 285 definitive hematopoiesis, 285 primitive hematopoiesis, 285 Device (21 USC 321(h)), 1369–1370 Dexamethasone, 588 Dextran, 557, 637, 646, 758, 760, 761 Diabetes, 2, 7, 28, 41, 331, 398, 408, 780, 794, 798, 1309 islet transplantation, 801–803 treatment of, 247 wound healing, 1323 Diabetes Control and Complications Trial (DCCT), 794 Diabetic neuropathy, 247 DialysisModule, 1094, 1096 Dialysis vascular access, 978–980 Diapodesis, 726 Diaspirin, 759 Differentially methylated region (DMR), 154–155 Differentiated cell, 8, 19, 31–32, 33, 39, 131, 227, 234, 1177 Differentiation assays, 23 Differentiation potential of MSC, 320, 892 stem cells, 30, 31, 33, 34 Digit, 1199, 1201–1203 Diisocyanate, 611–612 Dipeptidylpeptidase (DPPIV), 915 Diploid adult cells, 349 Diploid adult hepatocytes, 358 Direct perfusion bioreactor, 490–491 Discoidin domain receptors (DDR), 70 Distraction osteogenesis, 897 Divisional applications, role in patent protection, 1392–1393 Dkk1, 142, 430 DNA construct development, 170–171

DNA methyl transferase (Dnmts), 152–153 DNA methylation, 132, 141, 152–153, 156, 161, 196 Dolly, 148, 158, 169, 176, 457 Donor-specific tolerance, 805–806 L-DOPA, 1325–1326 Dopaminergic neurons, 7, 39, 109, 110, 200 Dor procedure, 1043 Double-quantum filtered magnetic resonance (MR) imaging, 327 Doxorubicin, 961 Drosophila melanogaster, 57, 58, 137, 138, 140, 392, 420 Drug (21 USC 321(g)(1)), 1369 Drug-eluting stents, 1001 Drug-loaded tumor cell (DLTC)system, 961 Drug testing, hepatic stem cell, 366 Dual channel, 565 Dubin–Johnson syndrome, 916, 920 Duchenne muscular dystrophy (DMD), 110, 387, 782, 1237 Ductal plate cells, 356, 357–358 Dulbecco modified Eagle medium (DMEM), 227, 320, 772, 1078 DuPont, 359 Dynamic cardiomyoplasty, 820 Dynamic compression, 521, 526, 1188 Dynamic loading, 527 Dynamic mechanical conditioning, 679, 1012, 1025 “Dynamic reciprocity”, 66 Dynamic regulation, of soluble environment, 548 Dynamic seeding method, 1008 Dysferlin, 783 Dystrophic epidermolysis bullosa (DEB), 1378 Dystrophin, 783–784, 788, 903, 1237 E-cadherin, 51–52, 55, 56, 57–58 E-selectin, 726 E2A, 51, 57 EASE, 567, 569, 571 Ectogenesis, 1342–1343 Ectopic eyes, 420–421 Edmonton Protocol, 41, 796–798, 800–801, 806 EEG, 1349 Effector T-cells, 715 Egg donors compensation, arguments, 1361–1362 EKG, 1349 Elastic cartilage, 1176 Elastin, 645, 666, 1208 Elastin-laminin receptor (ELR), 70

Index

Elastography, 483 Electrical conduction, of cardiac tissue, 1047 Electrical pulse method, 179 Electrical stimulation, to cardiac tissue, 1051–1052 Electron beam lithography (EBL), 675 Electropatterning methods, 547 Electroporation, 172, 192, 194 Electrospinning, 4, 491, 584–585, 616, 639, 675, 691, 748, 902, 990, 1045, 1046 Electrospun scaffolds, 990, 1046 Embryo culture, 179–180 potential, 368 Embryogenesis, 29, 136, 137, 141, 142, 349, 423, 430, 439, 441, 974, 1021, 1022, 1042, 1342 Embryoid bodies (EBs), 6, 37, 41, 126, 215, 462, 746, 762, 968 Embryonal carcinoma (EC) cells, 126, 127, 132, 139, 196, 210–211, 230, 1127, 1177 Embryonic blastomeres, 457, 461, 890 Embryonic gene expression patterns and NT, 155–156 Embryonic germ cells (EGCs), 29, 32, 126, 1177 versus ESCs, 34 Embryonic progenitor cells, 1177, 1183 Embryonic stem cells (ESCs), 30–32, 100, 111–112, 305, 345–346, 349, 402, 456, 569, 746, 762, 890–891, 937, 941, 1098, 1200, 1337, 1338, 1340, 1354 culture of, 212–215 derivation of, 20–21, 29, 150–151, 210, 211–212, 1027 differentiation, 215–216 ethical and legal issues, 367–368 isolation methods, 347–348 pluripotency, 136–137, 216, 258, 690 properties, 968 research, 1359 in Asia, 372–373 in Australia, 372 in US, 372, 1355–1357 retinal pigment epithelium from, 855–859 sources, 1354 see also Human embryonic stem cells (hESCs) Embryonic Stem Cell Research Oversight (ESCRO), 1360 Enamel, 1286–1288 Encapsulation system, 663 End-stage renal disease (ESRD), 400, 794, 795, 978, 1114 in United States, 1114 Endochondral ossification, 744, 767, 889 Endocrine gland-derived VEGF (EG-VEGF), 1022

Endocrine specification, 404 Endoderm, 30, 1073 Endoderm/gut endothelium generation, 403–404 pancreatic differentiation, 404 Endogenous NSC, 309, 310 Endogenous tissue repair, 815 Endomysium, 1232 Endostatin, 76, 443, 955, 961 Endothelial cells, 37, 105, 249, 271, 272, 441, 474, 690, 933–934, 1001, 1145 Endothelial–leukocyte adhesion molecules (ELAMs), 709 Endothelial nitric oxide synthase (eNOS), 199, 524, 525, 529 Endothelial phenotype, induction of, 37, 233 Endothelial progenitor cells (EPCs), 33, 245, 249–250, 271, 825–826, 1027 derived endothelial cells, 441, 446 identification and isolation, 440 in vitro expansion, 440–441 as surrogate marker, 443–444 therapeutic applications in tissue engineering, 446–447 in tissue regeneration, 447–448 in tumor growth, 442–443 Endothelial stem cells (EnSCs), 104, 105–106 Endothelial tube formation, 75, 774 Endothelialization, 1047–1048 Engelbreth-Holm-Swarm (EHS) tumors, 75 Engineered cardiac tissue, function of electrical conduction, 1047 endothelialization, 1047–1048 host response, and biocompatibility, 1048–1050 mechanical elasticity, 1047 strength development, 1047 thrombogenecity, 1047–1048 tissue architecture, 1047 vascularization, 1048 Engineered penile prostheses, 1130–1131 Engineering strain, 514, 516, 517 Engineering substrate mechanics MEMS devices, 542–543 patterning substrate stiffness, 543–544 Enhanced green fluorescent protein (eGFP), 196, 248, 274 Enteroendocrine cells, 1077 Entubulation, 1252–1255 chemotactic cues, 1256 haptotatic cues, 1255–1256 nerve guidance scaffolds, cell seeding in, 1256

1415

1416

INDEX

Enzymatic tissue digestion, 875 Enzyme-linked immunosorbent assay (ELISA), 860, 918 Enzyme-mediated hydrolysis, 1006 Ependymal cells, 104 Epiblast, 403 Epicel, 3, 17, 18, 1300, 1327 Epidermal growth factor (EGF), 67–68, 440, 694, 1064 Epidosembryo, 1338–1339, 1342 Epigenetic and environmental regulation, 131–132 Epigenetic reprogramming, 152 basic mechanisms DNA methylation, 152–153 histone modifications, 154 imprinting, 154 post-zygotic phase, 157–158 pre-zygotic phase, 154–157 Epimysium, 1232 Epithelial cell adhesion molecule (EpCAM), 357 Epithelial cells, 50, 51, 53, 67, 487, 675 Epithelial–mesenchymal relationship, 350–351, 361 Epithelial-to-mesenchymal transition (EMT), 50, 104 cellular mechanisms of, 50 basal lamina, invasion of, 53 cell–cell adhesion, changes in, 51–52 cell–ECM adhesion changes, 52–53 cell motility, stimulation of, 53 molecular control of, 53 ECM signaling, 56 signaling molecules, 54–56 transcriptional program, 57–59 Epstein–Barr virus (EBV), 473 eRas, 130 ErbB2/HER-2/Neu receptor, 55 ERK mitogen-activated protein kinase (MAPK), 213 Erythrocytes, 892 Erythropoietin, 329, 1106, 1259 Escherichia coli, 140, 198 Esophagus, 1072, 1074–1076 Esperion, 1400 Estrogen, 475 Ethical considerations, 762–763, 969–970, 1334 arguments do not work, 1334–1338 arguments that work, 1338–1343 Ethical constraints, 1342–1343 ectogenesis, 1342–1343 humanitarian, embryo use for, 1342 oocyte donation, 1343 patent infringement claims, 1343

European and American organ procurement organizations, 1093 European Medical Evaluation Agency (EMEA), 1353 European Patent Office, 1389, 1394 Eurotransplant data, 1093 Ex vivo, measurements of mechanical properties, 483 Ex vivo cell maintenance, 928–930 Ex vivo expansion, 250–251, 445 Ex vivo gene therapy, 874–875, 905, 1170 for bone regeneration, 873 experiments, 459–460 Ex vivo therapy, 769 Expanded poly(tetrafluroethylene) (ePTFE), 712, 1253 Expansion potential stem cells versus mature cells, 363 Expressed sequence tags (ESTs), 1388 Extracellular matrix (ECM), 50, 66, 350–351, 359–361, 491, 536, 554, 1021, 1062, 1065, 1127, 1162 and cell dynamic reciprocity, 667 composition and diversity, 66–68 developmental changes, 351 receptors for, 68–70 as scaffold material, 597–598, 1326 signaling, 56, 70 and soluble signals, 350–351 tissue engineering, 359–361 Extracellular signaling factors, 127–130 Extracorporeal liver support, use of hepatic cells, 1087 Extrinsic controls, 141–142 anterior–posterior axis, 141 BMP, 141 FGF, 142 Nodal, 141–142 retinoic acid, 141 Sonic Hedgehog, 141 Wnt, 141–142 Eye-field transcription factors (EFTFs), 418–421 F04 Division IV, 1379 Fabrication methods for biodegradable polymer scaffolds, 582–585 “Fact-based reasoning”, 1336–1337 Fanconi’s anemia, 238 FDA Biological Response Modifiers Advisory Committee, 1374 FDA Modernization Act (FDAMA), 1372 of 1997, 1379

Index

FDA regulatory process approval mechanism and clinical studies, 1370–1371 center of, 1369 clinical development plan, 1378 and critical path science, 1380–1381 gene therapy, 1376 guidance, 1367–1369, 1372 history of, 1366 human cells and tissues, for transplantation, 1372–1373 human cellular therapies, 1373–1375 laws, 1367–1369 meetings, 1371–1372, 1379–1380 organization and jurisdictional issue, 1369–1370 preclinical development plan, 1377–1378 regulations, 1367–1369, 1372 skin constructs, 1378 standards development program, 1378 xenotransplantation, 1375–1376 Federal Food and Drug Act, 1367, 1378 Federal Register, 1394 Feed-forward loops, 258 Feedback loop, 349, 361, 362, 934 Feeder cells, 6, 191–192, 345, 361 Femoral bone defects, 323 Femur–graft–tibia complex (FGTC), 1213 Femur–MCL–tibia complex (FMTC), 1206, 1214 Fetal bone cells, 972 Fetal bovine serum (FBS), 321, 769, 772 Fetal calf serum (FCS), 319, 320, 762, 974 Fetal cells, in somatic cloning experiments, 151 Fetal fibroblasts, 78–79, 156, 177 Fetal healing, 730 Fetal heart tissue, 970–971 Fetal human retinal progenitors, 428 Fetal livertyrosine kinase 3 ligand (Flt3L), 191, 244, 249 Fetal origin, of stem cells, 402 Fetal pancreatic tissue, 973 Fetal stem cells, 969, 1354–1355 Fetal stem cells from somatic tissue (FSSCs), 262 Fetal thymic organ culture (FTOC), 290 Fetal tissues, 39, 353, 892, 947, 968, 1132–1133 ethical considerations fetus and oocytes, 969–970 immunological considerations, 969 regenerative medical application bladder, 974–975 bone, 972

genitourinary system, 1126, 1127, 1132–1133 heart, 970–971 hematopoietic cells, 973 kidney, 974 lung, 971 muscle, 972 neural tissue, 970 pancreatic islet cells, 973 skin, 971 stem cells derivation, 969 Fetal wound healing adhesion and migration, 78–79 differentiation, 79–80 proliferation, 79 Fibrin, 76, 114, 706, 727, 960, 1006, 1007, 1009, 1013, 1025, 1066, 1274, 1276, 1277 Fibrin-based grafts, 1008 Fibrin–fibronectin meshwork, 73 Fibrin glue, 768, 773, 1078 Fibrin sealant, 1258 Fibroblast alignment, 1221 Fibroblast feeder layers, 213, 430 Fibroblast growth factor (FGF), 55, 76, 128, 130, 213–214, 404, 852, 961 FGF1, 1022, 1023, 1024 FGF2, 55, 213, 427, 852 FGF13, 569, 571 Fibroblast/myofibroblast apoptosis, 77 Fibroblastic colony-forming units (CFU-Fs), 322 Fibrocartilage, 902, 1176 Fibromodulin, 80, 83 Fibronectin (FN), 57, 74, 79, 81, 487, 499, 666, 727, 1108, 1254 Fibronectin adhesion promoting peptide (FAP), 1064 Fibroplasias, 726–727 Fibrosis, 81, 712, 1235 anti-fibrotic therapy, 1241 and fibrous encapsulation, 713–714 prevention, 1240–1241 and TGF-β1, 1241 Fibrous scaffolds, in cardiac tissue engineering, 1044–1046 Ficoll, 444, 876, 892 Fidap, 935 File wrapper estoppel, 1397 Filtration, 1114, 1116, 1117 Filtration barrier, in kidney, 1121 Finite strain theory, 1210

1417

1418

INDEX

First Piola–Kirchoff stress, 517 Flat bed bioreactors, 930–931 Flat plate bioreactors, 930–931 Fletcher, Joseph, 1340 Fluent, 935 Fluid-induced shear stress, 529 cell response and transduction mechanisms, 491–492 Fluorescence activated cell sorter (FACS), 386, 389, 892 Fluorescenct in situ hybridization (FISH) analyses, 171, 248, 272 integration site of, 172–173 Fluorine nuclei (19F), 557 19 FMRI, 557 Focal adhesion kinase (FAK), 484, 485, 486, 667 Focal injuries, 766 Follistatin, 1239 Food and Drug Administration (FDA), 144, 239, 401, 696, 872, 1341, 1366 Food, Drug, and Cosmetic Act (FD&C Act), 1310, 1367 Foreign body giant cell (FBGC), 711 formation and interactions, 712–713 Foxn4, 423 Fracture healing, 745 Fracture repair guide for bone regenerative therapeutics, 1165 Freeze-gelation method, 596 Frizzled-3 (Fz3), 421 Fumarate-based polymers, 621–624 Fumaric acid, 621 Functional capillary density (FCD) model, 759–761 see also Cardiovascular biomechanics Functional renal tissue in vivo, regeneration of, 1109–1111 Functional tissue engineering (FTE), 679, 897, 1206 applications of cell therapy, 1218–1219 future directions, 1222–1223 gene therapy, 1217–1218 growth factors, 1215–1217 mechanical factors, 1221–1222 scaffolding, 1219–1221 Fusion/cleavage, as screening tool, 179 G-protein, 55, 83, 492, 523, 739 G-protein-coupled receptor, 523, 524, 860 G418 selection, 172, 405

G418/Neo, 171 Gal epitope, 719 Galactose modified PLGA, 586 Galanin, 569 Gambit, 935 Gamma-aminobutyric acid (GABA), 191 Ganciclovir, 5 Gas-foaming/salt-leaching method, 583–584, 587, 588 Gastrointestinal (GI) tract, 275, 1072, 1153 Gastrulation, 55, 258–259, 420 in amniotes, 50 definitive endoderm, 403 Drosophila, 57, 59 morphogenetic movements of, 100 sea urchin, 52 GCRMA, 566 GDF3, 571 Gelatin, 127, 231 Gene activity control, 140–141 Gene complementation, 783–784 Gene expression Myf5, 233, 890 Myf6 (MRF4), 233 Gene ontology (GO), 294, 596 Gene regulation, 192–194 knock-down, 192–194 knock-in/knock-out, 192 Gene therapy for bone regeneration, 747, 874–875, 899, 1168, 1170 in cardiac tissue, 1041–1042 in cartilage regeneration, 901, 1179 in hepatocyte transplantation, 918 in ligament and tendon healing, 1216–1217 MSC, role of, 904 with MSC-based vehicle, 885 in musculoskeletal repair, 897 in peripheral blood stem cells, 448–449 regulations, 1376 GeneChips, in regenerative medicine cancer, applications to, 574–576 cell screening, 573–574 databases, 563 interpretation of, 567–569 diseased tissue, re-engineering, 574–575 experimental design, 564 expression analysis, 566–567

Index

platform, 564–565 sample preparation, 565 stem cell differentiation, 569 AFSC, 571–572 hESC, 569–571 meta-analysis, 572–573 PGESC, 571 General Agreement for Trade and Tariffs (GATT), 1392, 1393 Genetic aberrations, 462 Genital tissues, 1129 corporal smooth muscle, 1130 engineered penile prostheses, 1130–1131 female genital tissues, 1131 Genitourinary system biomaterials, role, 1126–1127 fetal tissue engineering, 1132–1133 injectable therapies, 1133–1134 reconstitution strategies, 1126 testicular hormone replacement, 1134 urologic structures tissue engineering, progress of, 1127–1132 vascularization, 1127 GenMapp, 569, 570, 572 Genome Explorations, 564 Genomic imprinting, 154 Genzyme Biosurgery, 3, 768, 769, 1327 Carcitel™, 900 Epicel™, 1300 Germ layer stem cells, 349 Germ line intervention, 1341 Germ-line transmission, 459 German embryo protection regulations, 372 Gibbon ape leukemia virus (GALV), 195, 201 Glaucoma, 418 Glial derived neurotrophic factor (GDNF), 110, 115, 947, 1261, 1262 Glial fibrillary acidic protein (GFAP), 242, 245, 857 Global gene expression profile, 459 Glomerular slit diaphragm, 1121 Glucagon-like peptide-2 (GLP-2), 1079 Glutaraldehyde, 595, 598, 758, 759, 762 Glutathione, 1106, 1114, 1115, 1117, 1118 Gly–Pro–Nleu, 1064 Glycosaminoglycan (GAG), 66, 68, 350, 596, 667, 1186, 1279 Goblet cells, 1077 GOC, 1159

Goretex®, 607 Graft versus host disease (GVHD), 22, 228, 238 Granulation tissue, 77, 597, 710 development of, 707 formation of, 74, 76, 729 Granulocyte colony-stimulating factor (G-CSF), 275 cell therapy and administration of, 833 Graphical user interface (GUI), 566 Green fluorescent protein (GFP), 261, 327, 449, 890 Green strain, 516 Growth differentiation factor 8 (Gdf-8), 1239 Growth/differentiation factors (GDFs), 128–129, 501, 745, 901 in bone tissue engineering, 747–748 combination of, 747–748 delivery, 747 Growth factors, 67, 113, 571, 710 role in ligament and tendon healing, 1214 in vitro studies, 1215 in vivo studies, 1215–1216 SwissProt, 571 therapeutic use, 2 GSK3β inhibition, 130 GTPase proteins, 485 Gut, 1072 Gut-associated lymphoid tissue (GALT), 1074 Hair follicle stem cells (HFSC), 103, 127 Hank’s balanced salt solution (HBSS), 1078 Haptotactic cues, 1255–1256 Hare, Richard, 1334 HCT/Ps, 1311, 1312, 1313, 1373 Healos®, 877, 883 Health Resources and Service Administration (HRSA), 1375 Heart regeneration and BMSCs, 272–274 see also Cardiac tissue, engineering of Heat shock protein (HSP), 156, 1042 Hedgehog, 54, 356, 503, 1073, 1288 Helicobacter pylori, 55 Hemangioblasts, 199, 245, 438, 439, 449 Hematocrit (Hct), 329, 760 Hematopoiesis players, 289–290 Hematopoietic stem cells (HSCs), 21, 33, 105, 127, 194, 228, 269–270, 284–295, 387, 762, 825, 1120, 1380 and autoimmune applications, 41 cell surface molecules, 283

1419

1420

INDEX

Hematopoietic stem cells (HSCs), (continued) compartment, of bone marrow, 321 developmental origin of, 285 differentiation, 284, 289, 290 ex vivo expansion, 445, 894 hematopoiesis players, 289–290 HSC niche, 272, 291–292 identification and isolation, 445 intrinsic regulators, of self-renewal, 285–288 lineages, in vitro differentiation, 284, 290 mobilization of, 439 multi-lineage repopulation, 288–289 plasticity and therapeutics, 294–295 precursor cells, 285, 290, 302, 459 purification and molecular signature of, 292–294 and tissue regeneration, 232–233 Hematopoietic system, 137, 269, 272, 289, 445, 973 Hemodialysis, 979, 980, 1114 Hemodilution, 760, 761 Hemoglobin use of, 758, 759, 763 Hemorrhage, 756, 760, 1078, 1211–1212, 1324, 1326 Hemostatic agent, 1324 Heparan sulfate proteoglycans, 81, 351, 356 Heparin, 67, 76–77, 227, 356, 589, 590, 662, 776, 779, 878, 1094, 1116, 1276 Hepatic angioblasts, 357 Hepatic artery, 355, 356 Hepatic stellate cell precursors, 357, 361 Hepatic stem cell and cancer, 351–352 categories of committed progenitors, 347 determined stem cells, 347 ESCs, 345–346 multipotent stem cells, 346–347 totipotent stem cells, 345 clinical, commercial, and research application of, 362–367 compartments, 356–357 dynamic interaction, 351 future uses, 364 bioartificial livers, 365–366 cell therapies, 365 drug testing and protein manufacturing, 366 gene therapies, 366 research, 367 vaccine production, 366–367

human tissue sourcing determined stem cell, 353–354 ES cells, 352 multipotent stem cell, 353–354 isolation and purification of, 347–348 and liver regeneration, 361–362 and maturational lineage biology, 344, 348 epithelial–mesenchymal relationship, 350–351 liver as, 354–362 and progenitors, 344–348 need for feeders, 361 properties of, 362 sources of precursors, 358 tissue engineering, 358–359 biological issues, 359–361 tissue reconstitution, 363 Hepatoblasts, 356, 357, 358 Hepatocyte allografts, 918 Hepatocyte apoptosis, 82 Hepatocyte growth factor (HGF), 55, 67, 81, 102, 106, 725, 1023, 1042, 1098, 1234 Hepatocytes, 80, 101, 233, 248, 261, 274, 332, 355, 357–358, 403, 407, 536, 540, 546, 672, 674, 698, 929, 930, 933, 937, 940, 1092 phenotype induction of pluripotent stem cells, 233 transplantation in acute liver failure, 917–918 aspects of, 921 bridge technique, 916–917 choice of sites for, 914 integration of, 914–916 for metabolic liver disease, 918–921 for non-organ transplant candidates, 921–922 preventing factors, 922 shunting of, 917 Hepatocyte-like cells, 248–249, 332 Her2/Neu (herceptin), 575 Herceptin®, 55, 575 Herpes simplex virus (HSV), 5, 198, 1063 Herpes zoster ophthalmicus, 1060 hESC cocktail (hESCO), 191 Heterodimers, 68, 71, 484, 523, 666 Heteroplasmy, 462 Heterotypic interactions, of multicellular constructs, 546, 1010 HG-U133 GeneChip, 562, 574 HIF2α, 132

Index

Histidine (HIS6) tags, 737 Histogenesis, in three-dimensional scaffolds, 686 design parameters biomolecular factors, 694–695 cell sources, 689–690 degradation, 692–693 microvasculature, 690–691 porosity, 691–692 future directions three-dimensional microfabrication, 698 human bioreactor, 697 regeneration, of diseased tissue, 688–689 synthetic materials, of new organs hydrogels, 696–697 hydrolytically degradable polymers, 695–696 Histone acetyltransferases (HAT), 141, 154 Histone deacetylases (HDACs), 141, 153, 154, 572 deacetylation function of, 141 Histone methyltransferases (HMTs), 153, 154 Histones, 141, 154 Histopaque™, 444 HIV disease, retarding/preventing, 474–475 HLA-type ESCs, 461 Hoechst 33342, staining dye, 178 Holiday Inn®, 1394 Hollow-fiber bioreactors, 932–933 metabolic characteristics, 1115–1116 successful with liver, 933–934 transport, 1115 Homeobox genes (HOX), 287, 507 Homodimers, 747 “Homologous use”, 1311 Homotypic interactions, of multicellular constructs, 544–546 Hormonally defined media (HDM), 359, 360–361 Host response and biocompatibility, 1048–1050 to engineered cardiac tissue, 1048–1050 Hox and Sox transcription, 1073 HoxB4, 130, 201–202, 287–288, 460 Huard, Johnny, 1328 Human amniotic epithelial cells (HAEC), 229, 230 Human amniotic fluid stem cell (HAFSC), 571, 572 Human amniotic mesenchymal cells (hAMC), 229 Human autologous chondrocytes (HACs), 18 Human bioreactor, 697 Human bone marrow-derived stem cells (hBMSCs), 260, 262, 331

Human cells, tissue, and cellular and tissue-based products (HCT/Ps), 1311, 1373 Human chorionic gonadotropin (hCG), 1152 Human chromosome 12p13, 132 Human Cloning Regulation Act, 372 Human corneal replacement, 1061–1062 Human-embryo-based technologies, 462 Human embryonic research in Asia, 372–373 in Canada, 372 in Europe, 372 in South America, 372 in UK, 369 in US, 372 Human embryonic stem cells (hESCs), 6–7, 8, 30, 126, 137, 190, 456, 569–571, 855, 856, 890, 968–969, 1199 versus adult stem cells, 1119–1120 applications immunogenicity modification, 202–203 lineage tracking and purification, 202 specific lineages, 201–202 cardiomyocyte derivation, 7 culture of, 213–215 condition, 31 derivation of, 212 derived RPE, culture and properties of, 859–861 developmental potential differentiation of, 215–216 pluripotency of, 216–217 differentiation and transplantation, 198 ectodermal derivatives, 200 endodermal derivatives, 200 mesodermal derivatives, 198–200 differentiation potential, 31 genetic approaches in, 190 genetic modifications, 201 in vitro model of, 31, 212 maintainability of, 191–192 manipulation, genetic approaches to chemical transfection, 194 gene regulation, 192–194 physical transfection, 194 viral transduction, 195–198 phenotypes, 31 derivation, 191 screening, for genetic variability, 574 see also Embryonic stem cells (ESCs)

1421

1422

INDEX

Human epidermal keratinocytes (HEKs), 17, 18 Human Fertilization and Embryology Act (HFA Act), 369 Human fetal liver multipotent progenitor cells (HFLMPCs), 259 Human leukocyte antigen (HLA), 329, 461, 470, 802, 918, 969 Human MSCs (hMSCs), 271, 318, 542, 772, 873, 901 Human nerve growth factor (hNGF), 955 Human renal tubule assist device clinical experience with, 1118–1119 Human TElomerase Reverse Transcriptase (hTERT), 475, 1005 Human tissue engineering, ethical issues in, 1346–1353 arguments for, 1347–1348 brokenness nature and human duty, 1348–1349 case study, 1346 debates of, 1349–1350 deconstruction, perils of, 1350 financial backing, 1352–1353 intellectual property, 1352 ownership, 1350–1351 pre-market gating, 1352 principle reasons, 1347–1348 Humanitarian Use Device (HUD), 1310, 1370 Humanitarinan Device Exemption (HDE), 1370 Huntington’s disease, 243–244, 309 Hurler syndrome, 328 Hyaline cartilage, 18, 624, 889, 1176, 1184 Hyaluronan, 56, 327, 771, 772, 1276 based grafts, 1012 Hyaluronic acid (HA), 78, 581, 587, 646, 1186, 1276 Hyaluronic acid and methylcellulose (HAMC), 1259, 1260 Hybrid, 636, 783, 915, 1008, 1066 scaffold biomaterial ceramic hybrid scaffolds, 641–644 natural polymers hybrid scaffold, 644–646 poly(α-hydroxyacid) family, 639–641 Hybridoma cell production, 259 Hydrogels, 113, 547, 605–606, 607, 608, 609, 620, 623, 625, 676, 693, 696–697, 698, 1002, 1025, 1043–1044, 1067, 1187, 1259, 1275 based cardiac tissue engineering, 1043 Hydrolytically degradable polymers, 695–696 Hydrolytically stable polyurethanes, 611–612 Hydrophobic polymers, 619, 620 Hydrophobic polyphosphazenes, 620

Hydroxyapatite (HAP), 505, 739, 746, 869, 878, 879, 1200, 1290 Hydroxyethyl starch (HES), 239, 758 Hydroxyethylated amylopectin, 744 Hyperacute rejection response (HAR), 160 Hyperbaric oxygen (HBO), 1240 Hyperglycemic liver environment, 401 Hyperplasia, 101–102, 1000 Hypersensitivity, 716, 717, 719 HYPO score, 798 Hypoglycemic unawareness, 795, 796, 798, 806 Hypotension, 756 Hypoxanthine phosphoribosyltransferase-1 (HPRT-1), 192 Hypoxia-inducible factor-1α (HIF-1α), 277, 1024 Hypoxic–ischemic injury NSC therapy, 310–311 Hysteresis, 1210 Id (inhibitor of differentiation) proteins, 100, 213 Ideal scaffold for cartilage engineering, 1183 for ligament and tendon healing, 1218 for tissue engineering, 775 IgE antibodies, 717, 719 IgG, 709, 719 IgM, 717, 719 Iliac crest, 322, 869, 883 autograft harvesting, 869 Imminent death, 1334, 1339 Immune compatibility, 8–10, 1109 Immune complex hypersensitivity, 716, 717 Immune function, of alimentary tract, 1073–1074 Immunocytochemistry (ICC), 231, 235, 308, 564, 1110 Immunodeficiency syndrome, 459 Immunofluorescence, 538, 540, 860 Immunofluorescent labeling, 242 Immunogenicity, 9, 202, 240, 364, 719, 746 Immunohistochemical assay, 109, 449, 773, 1031, 1033, 1051, 1109, 1239 Immunomodulation and islet cell transplantation, 805–806 Immunomodulatory effects, of MSCs, 328–330 Immunostimulation, 718 Immunosuppression, 8, 399, 461, 718, 789, 796, 799, 800, 895, 1052 in islet cell transplantation, 800–801

Index

Immunotoxicity, 704, 714–720 cytokines and effects, 716 effector T lymphocytes, 715 hypersensitivity responses, 716–717 persistent chronic inflammation, 717–718 representative tests, 718 tissue-engineered constructs, 717 Impaired dark adaptation, 853 Implantable protein factory (ImPACT™), 961 Imprinting control regions (ICRs), 154 In silico modeling, 482 In silico post hoc method, 563 In situ hybridization technique, 773 In vitro culture systems in cartilage tissue engineering, 1179–1183 In vitro differentiation of hematopoietic lineages, 290 of MSCs, 320, 326, 332 In vitro fertilization (IVF), 155, 212, 345, 369, 890, 1358 In vitro matured (IVM) oocytes, 149–150 In vitro model, of human ESCs, 31, 212 In vitro neurogenesis, analysis of epigenetic strategy, 303–305 genetic strategy, 305 In vitro organ model, 1030 In vitro techniques organ model, 1030 spontaneous tube formation, 1030 sprouting models, 1030 In vivo cartilage tissue generation, 1189–1190 In vivo construct fabrication, 1012 In vivo-derived counterparts, 459 In vivo differentiation of bone marrow stem cells, 275 of unrestricted somatic stem cells, 230 In vivo gene therapy, for bone regeneration, 874, 1170 In vivo measurements mechanical properties, of tissue, 483 In vivo model, of human ESCs, 31 In vivo neurogenesis, analysis of, 302–303 Indium111, 920 Indomethacin, prostaglandin inhibitor, 329 Inducible nitric oxide synthetase (iNOS), 726 Inflammation, 401, 704, 705, 707, 708, 709, 725–726, 817, 818, 1165, 1234, 1235, 1237 and wound healing acute inflammation, 708–709 blood–material interactions, 705

chronic inflammation, 709–710 FBGC formation and interactions, 712–713 granulation tissue, 710 inflammatory response, 705–706 macrophage interactions, 710–712 provisional matrix formation, 706–707 temporal sequence of, 707–708 Inflammatory cell emigration, 709 Inflammatory mediators, 709 Inhibition of non-specific protein adsorption, 736 Initial commitment phase, 744 Injury-induced regeneration, 103, 115 Injury versus repair, 815–820 Inkjet printing, 1043 Innate immune system, 469, 1048–1049 Inner cell mass (ICM), 126, 210, 211, 216, 855 Instant blood-mediated inflammatory reaction (IBMIR), 804, 806 Institutional Review Board, 948, 1355 Instructive scaffold, 734 Insulin-like growth factor (IGF), 612 IGF-1, 105, 113, 694, 890, 901, 1180, 1238, 1276, 1292 Insulin-producing cells, 247 alternative sources, 402–403 Intact extracellular matrix, as scaffold material, 597–598 Integra®, 114, 1298 Integrin–ECM binding, 72 Integrin-linked kinase (ILK), 56, 73 Intellectual property rights copyrights, 1395 importance, 1396–1398 patents, 1386–1389 novelty, 1389–1391 obviousness, 1391 patent rights, 1392 patent term and patent term extension, 1392–1394 requirements, 1391–1392 service marks, 1394 trade marks, 1394 trade names, 1394 trade secrets, 1395 valuation, 1398–1401 Interconnectivity scaffolds, 748 Interferon-gamma (IFNγ), 471, 726, 1117, 1241 Interlamellar spacing, 671 Interleukin IL-1, 249, 474, 716, 1119 IL-2, 329, 667, 716 IL-4, 271, 708, 713

1423

1424

INDEX

Interleukin, (continued) IL-6, 819, 1098, 1118, 1119 IL-8, 74, 75, 83, 277 IL-10, 329, 1118, 1260 IL-13, 713 Internal ribosome entry sites (IRES), 196 International Conference on Harmonization (ICH), 1369 International Society for Cellular Therapy (ISCT), 893 International Standards Organization (ISO), 1379 Interstitial collagens, 77 Interstitial fluid flow (IFF), 493 Interstitial retinol binding protein, 428 Intervertebral disk (IVD) degenerative disk disease, 904 regeneration, 327–328 Intestinal development, of alimentary tract, 1073–1074 Intestinal stem cells (ISC), 127, 1073 Intima, 528, 697, 980, 1000 Intimal thickening, 1000–1001 Intracellular activators, 484 Intracellular adhesion molecule (ICAM), 709 Intracellular mechanics, 488–489 Intracellular multi-molecular proteins, 484 Intracorporeal kidney support end stage renal failure, 1106 functional renal tissue in vivo, regeneration, 1109–1111 kidney tissue regeneration, principles, 1106–1108 renal structure in vivo, creation, 1108–1109 Intramembranous ossification, 744 Intramyocardial injection, 22, 1042 Intra-operative cell therapy, 769 Intra-operative versus culture expanded, 769 Intrinsic stem cells, 427–428 Investigational Device Exemption (IDE), 1310, 1371 Investigational New Drug (IND), 239, 401, 456, 1118, 1371 Ionic charge (pH) orientation, 738 Ischemia–reperfusion injury, 82 Ischemic cardiomyopathy trial, 831 Islet after kidney (IAK), 400 Islet cell therapy β-cell replacement therapy, 399–402 benefits, 399–400 limitations to, 401–402 in future, 408–409 insulin-producing cells alternative sources of, 402–403

islet neogenesis, 404–407 from adult stem cells, 405–407 from ES cells, 404–405 pancreatic development, 403–404 β-cell differentiation, 404 endocrine specification, 404 endoderm/gut endothelium generation, 403–404 and pancreatic stem cells, 398 transdifferentiation, 403, 407–408 Islet cell transplantation. See Islet transplantation Islet engraftment, in post-transplant period, 803–805 Islet neogenesis, 404–407 from adult stem cells, 405–407 from ES cells, 404–405 Islet transplantation, 399–400, 794 clinical aspects patient assessment and selection, 798 procedure, 798–800 Edmonton Protocol, 796–798 future challenges donor-specific tolerance, 805–806 engraftment post-transplant, improvement of, 803–805 living donor islet transplantation, 801–803 stem cell transplantation, 803 xenotransplantation, 803 history of, 795–796 risks immunospressive therapy and complications, 800–801 surgical complications, 800 Islet xenotransplantation, 803 Islets of Langerhans, 398, 625 Isobutyl-1-methylxanthine, 895 Isotropic nerve grafts, 1272 agarose gels, 1275 collagen gels, 1275–1276 ECM molecules, 1276–1277 natural materials, as scaffolds, 1272 neuronal support cells, seeding, 1277–1278 neurotrophic factors, 1276 synthetic scaffolds, for nerve repair, 1272–1275 ITSFn (insulin, transferrin, selenium and fibronectin), 428 Johnson & Johnson, 1400 Joint loading, 520, 526 Joint motion, 1210–1211 Jungle, The, 1367

Index

Kaposi’s sarcoma (KS), 474 KDR, 245, 440, 441 Kefauver–Harris Amendments, 1367 Keloid scars, 77 Keratinocyte grafting, 1301 Keratins, 17, 486, 759 Keratoprostheses (KPro), 1062–1066 development of, 1063–1066 Kernicterus, 920 Kidney, 960 acute renal failure, cell therapy, 1117–1118 bioartificial kidney in ESRD, 1119 bone marrow to, 275 cell sourcing, 1119–1121 filtration barrier, 1121 hollow-fiber bioreactors transport and metabolic characteristics of, 1115–1116 proximal tubule cells isolation and culture of, 1115 renal tubule assist device (RAD) characterization, 1116–1117 clinical experience, 1118–1119 tissue creation, 1133 tissue generation, principles, 1106–1108 ultrafiltration membrane development, 1121–1123 Knee motion, 1210–1211 Knock-down technology, 192–194, 201 Knock-in/knock-out technologies, 192 Lability index, 798 Laboratoire d’Organogenese Experimentale (LOEX), 1066 lac operon, 140 LacZ, 325, 1216 Lamellar keratoplasty (LKP), 1060, 1067 Laminectomy, 1252 Laminin, 36, 52, 67, 75, 76, 80, 81, 191, 666, 1043, 1046, 1064, 1108, 1115, 1255, 1276–1277, 1279 Langmuir–Blodgett deposition method, 661 Large diameter vessels engineering tissue-engineered blood vessels (TEBVs), 980 ECs, 981–982 smooth muscle cells, 982–984 tissue-engineered vascular grafts cell seeding and preconditioning, 992–995 cell source, 992

large diameter TEBV, 985 process, schematic depiction of, 990 scaffolds, 985, 990–992 small diameter TEBV, 984–985 vascular disease dialysis vascular access, need for, 979–980 prevalence and impact, 978–979 Large offspring syndrome (LOS), 151 Laser in situ keratomileusis (LASIK), 1061 Laser microdissection, 565 Laser scanning confocal microscopy (LSCM), 389, 1033 “Late-outgrowth” EPC, 441 Lateral inhibition process of endocrine specification, 404 LBS-neuronal precursor cells, 947 Lef-1, 58 LEF/TCF transcription factor, 52, 58 Lefty1, 142 LeftyA, 430 LeftyB, 569 Lentiviral vector, 196–198, 897 Leukemia inhibitory factor (LIF), 127–128, 191, 212, 213, 271, 304, 348 LIF receptor (LIFR), 128 Leukocyte–endothelial cell interactions, 709 Lewis, C.S., 1347, 1348 Leydig cells encapsulation, 1151–1152 Lhx2, 418, 420–421 Licensing of biologic products, 1310–1311 versus company formation, 1399–1400 LIGA, 660 Ligament and tendon injuries, 1206 biology, 1207–1208 biomechanics joint function, 1210–1211 tensile testing, 1208–1210 FTE applications cell therapy, 1217–1218 gene therapy, 1216–1217 growth factors, 1214–1216 mechanical factors, 1220–1221 scaffolding, 1218–1220 healing of, 1211–1212 future directions, 1221–1222 hemorrhage phase, 1211–1212 inflammatory phase, 1211

1425

1426

INDEX

Ligament and tendon injuries, (continued) healing of, (continued) remodeling phase, 1212 reparative phase, 1212 MCL, of knee, 1212–1214 ACL reconstruction, 1213–1214 combined ligament injuries, 1214 Light microscopy, 232, 389, 392, 1065 LIM mineralization protein-1 (LMP-1), 1170 Limb prosthetics, 1199 Limbal stem cells, 103 Limbs regeneration, 106, 1328 Lim, 420, 826 LIMMA, 567 Linneg HSCs, 269 Lineage-committed cells, 448 Lineage-committed progenitors, 438 Lineage conversion assays, 111 Lineage-dependent gene expression, 366 Lineage-restricted stem cells, 7 Lineage restriction, 346 Lineage-specific markers, 6 Lipid-mediated transfection, 172 Liquid nitrogen, 240, 939 Liquid prepolymer system, 1325 Liver bioartificial design, 930 flat plate bioreactors, 930–931 hollow fiber bioreactors, 932, 933 lineage model, 356 organization of, 354–356 precursors, sources of, 358 regeneration, 80, 361–362 after partial hepatectomy, 361 adhesion and migration, 80–81 apoptosis, 82–83 bone marrow stem cells and, 274 differentiation, 81–82 following toxic injuries, 361 implications, 362 proliferation, 81 stem cell compartment, 356–358 tissue engineering of, 358–359 biological issues for, 359–361 Liver cell-based therapy bioreactor culture of liver cells clinical applications, 1093–1094 bioreactor technology, overview for stem cell culture, 1098–1099

cell source, for liver support, 1091–1093 hepatic cells, prospects for extracorporeal liver support, 1087–1091 human liver cells, in bioreactors for liver support, 1093 liver progenitor cells and bioreactor cultures for extracorporeal liver support, 1096–1097 methods of liver support, 1086–1087 modular liver support, concept of, 1094–1096 need for, 1086 regenerative medicine bioreactors, 1086 stem cell research, challenges, 1097–1098 Liver progenitor cells for extracorporeal liver support, 1096–1097 Liver regeneration, 80, 361–362 after partial hepatectomy, 361 adhesion and migration, 80–81 apoptosis, 82–83 and BMSCs, 274–275 bone marrow stem cells and, 274 differentiation, 81–82 following toxic injuries, 361 implications, 362 proliferation, 81 Liver stem cells, 103, 1098 Liverbeads™, 359 Living donor islet transplantation, 801–803 Long-term culture initiating cells (LT-CIC), 290 Lower critical solution temperature (LCST), 608, 609 LTB4, 726 Lung tissue engineering, 971 Lymphocyte/monocyte predominant chronic inflammatory response, 713 Lymphoid cells, 288–289 Lymphoid-primed multipotent progenitors (LMPP), 288 Lysozyme, 596 M-cadherin, 387, 890 Macaque eye, 427 Macrophage interactions, 710–712 Macrophage mannose receptor (MMR), 713 Macroporous biodegradable scaffolds bioactive molecule surface immobilization in, 585–587 sustained release of, 588–590 fabrication of, 582–585 Macroporous hydroxyapatite scaffolds, 323 Macular degeneration, 418, 442

Index

Madin–Darby canine kidney (MDCK), 51, 57 Magnetic-activated cell sorting (MACS), 892 Magnetic resonance imaging (MRI), 483, 556, 557, 837 of bioartificial livers, 938–939 Major histocompatibility complex (MHC), 5, 9, 202–203, 329 MHC-I molecules, 461 Malaria surface antigen merozoite surface protein-1 (MSP-1), 181 Mammalian stem cells, 5 Mammography Quality Standards Act (MQSA), 1369 Mandibular distraction osteogenesis, 680, 905 Mannitol, 242–243, 664 Maple, 935 Marrow-isolated adult multilineage inducible cells (MIAMI cells), 260 MAS5, 566, 567 Mash1, 423 Material-induced wound healing, 1326 Matrigel, 191, 1026, 1028, 1031, 1051, 1277 Matrix metalloproteinases (MMPs), 52, 68, 624, 667, 676, 693, 728–729, 900–901, 1025 MMP-2, 974 MMP-9, 439, 974 MMP-20, 1291 Maturation promoting factor (MPF), 149, 178 Maturational lineage biology, 348–350, 941 Maxillofacial BM, 322 McKinney, Gerry, 1348 Mean arterial pressure, 756 Mechanical determinant, of tissue development bioreactors, overview of modeling, 491 types, 490–491 cell as signal receiver and processor, 483 fates, 489–490 mechanochemical transduction, 487–489 receptors and sensors, 483–484 tensegrity model of, 486–487 ex vivo measurements, 483 forces, in biological tissue, 482 in vivo measurement, 483 mechanical properties, of tissues, 482 mechanics of materials, 481 force and stress, 481 material properties, 481–482 practical examples of bone, 492–494

cartilage, 494–495 vasculature, 491–492 Mechanical forces, in tissue development biological tissues, forces, 482 ex vivo measurements, 483 in vivo measurements, 483 mechanical properties, of tissues, 482 mechanics of materials, 481–482 Mechano-anatomical units (MAUs), 1164 Mechanochemical transduction hard-wired nucleus, 489 intracellular mechanics, alteration, 488–489 ligands and binding sites, 487 mechanosensitive ion channels, 488, 522 Mechanosensitive ion channels, 488, 522 Mechanosensory neurons, 488 Mechanotransduction, 678, 1181 mechanism, 521–525 cellular response, 524–525 force transmission, 521–522 signal propagation, 523–524 see also Mechanochemical transduction, 487 Medial collateral ligament (MCL), 1206 of knee, 1212 ACL reconstruction, 1213–1214 combined ligament injuries, 1214 SIS treatment, effect of, 1218 Medical Device Amendments, 1367 Membrane attack complex (MAC), 714 Membrane permeability, 958 MEMS devices to measure cellular forces, 542–543 Meniscus, 901–902 Mer tyrosine kinase protooncogen (MERTK), 853, 856 Merosin, 783 Merozoite surface protein-1 (MSP-1), 181–182 Mesangioblasts stem cells, 390, 391, 903 Mesenchymal cells, 50, 67, 889 Mesenchymal stem cells (MSCs), 19, 33, 105, 110, 227, 270–271, 318, 390, 438, 444, 448, 536, 746, 768, 770, 825, 891, 1178 based EC, 331 for cartilage in pre-clinical animal studies, 773–774 cellular-based cartilage tissue engineering in vitro, role in, 1181–1183 and collagen gels, 1217 culture expansion, 772, 876, 881 culture procedures, 772–774

1427

1428

INDEX

Mesenchymal stem cells (MSCs), (continued) definition of, 319–320 and ECM interactions, 894 gene therapy using, 904 identification, 444, 893–894 immunomodulatory effects, 328–330 immunoregulation, 895–896 immunosuppressive properties, 22 implantation of, 899 in vitro behavior, 893 in vitro expansion, 444–445 isolation, 444, 892–893 markers, 893 multidifferential potential of, 891, 892 nature of, 320–321 niche, 894 nonskeletal tissue regeneration, 330–332 phenotype, characterization of, 774 regulation, 894–895 safety, in animal models, 772 schematic representation of, 871 self-renewal capacity and multipotentiality of, 897 skeletal tissue regeneration bone, 322–325 cartilage, 325–326 intervertebral disk, 327–328 tendon, 326–327 source and isolation techniques, 321–322, 772 therapeutic applications, 445–446, 448 tissues containing MSCs, 322 umbilical cord blood-derived, 893 whole marrow, 770–771 see also Bone marrow stromal cells Mesenchymal-to-epithelial transition (MET), 50, 51, 53, 104 Mesendoderm, 403 Mesoangioblasts, 127 Mesoderm cells, 30, 52, 57, 100, 260, 1073 Metabolic and secretory applications, in regenerative medicine, 41 Metabolic liver disease clinical transplants for, 919 hepatocyte transplantation for, 918–921 Metachromasia, 776 Metachromatic leukodystrophy (MLD), 328 Metastais-associated protein 3 (MTA3), 58 Micro-computed tomography (micro-CT), 325, 528, 1032 Microcontact printing, 541, 664, 675, 1046 Microencapsulation, 955–956, 1152

Microengineered tools, adhesive environment micropatterned screening arrays, 539–541 spatial patterning, 541–542 Microenvironment, cues in, 35–38 3D environment, 36–37 bioreactors, 37–38 Microfabricated post-array detector (mPAD), 543 Microfluidics, 537–539, 544 soluble cues control, 537–538 Microinjection (MI), 168, 170 pronuclear MI, 169, 184 Micropatterned screening arrays, 539–541, 545, 546, 548 Microphthalmia, 419 Microthalamia-associated transcription factor (MITF), 58, 852 Military needs, and solutions in regenerative medicine, 1324 cardiothoracic and vascular engineering, 1323 in-hospital care, 1326–1328 long-term therapy, 1328–1329 musculoskeletal tissue engineering, 1322–1326 pre-hospital care, 1324–1326 wound healing, 1323–1324 Mill, John Stuart, 1336 Minimal manipulation, 1311, 1313 Minocyline, 1260, 1401 Minor histocompatibility antigen (miHA), 462 Mitochondrial DNA (mtDNA), 462 Mitogen arterial pressure, 895 Mitotic cell cycle, 569 Mixed lymphocyte reaction (MLR), 22 Model based expression index (MBEI), 566 Modular liver support (MLS), 1094–1096 CellbioreactorModule, 1094–1095, 1096 concept, 1095 DetoxModule, 1094, 1095–1096 DialysisModule, 1094, 1095, 1096 Molecular engineering techniques, 1024 Molecular organization, of cells, 50 cellular mechanisms, of EMT, 50–53 ECM signaling, 54 signaling molecule, 54 molecular control, of EMT, 53–59 basal lamina invasion, 53 cell–cell adhesion changes, 51–52 cell–ECM adhesion changes, 52–53 cell motility simulation, 53 transcriptional program, of EMT, 57–59

Index

Molecular weight cutoff (MWCO), 958, 959 Mononuclear cells (MNCs), 246, 321 Mononuclear phagocytic system (MPS), 711 Monster tumors, 345 Morphogen, 127, 498, 499, 1292 Morphogenesis, 498 biomimetic biomaterials, scaffolds of, 505–506 bone marphogenetic proteins, 499–503 cartilage-derived morphogenetic proteins, 506 effects of stress on, 519 regenerative medicine and surgery, 506–508 stem cells, 503–505 steps in, 500 Morphometric analysis, 77 Morula/blastocyst transition, 158 Mosaic, 783 Mosaicplasty, 1184 versus ACI, 777 Mouse embryonic stem cells (mESCs), 128, 160, 190, 211, 258, 428, 971 BMPs effects on, 129 culture of, 212–213 differentiation of, 215 versus hESCs, 192 pluripotency, molecular control of, 216 Wnt signaling in, 130 Mouse nuclear transfer ESCs (mnt-ESCs), 458, 459 mRNA expression patterns, 155–156 Multicellular constructs, organization of heterotypic interactions, 544–546 homotypic interaction, 544–546 three-dimensional patterning, 546–548 two-dimensional patterning, 544–546 Multicoaxial bioreactor, 933 Multidrug resistance protein2 (MRP2), 916 Multi-lineage repopulation, 288–289 Multipotent adult male germline stem cells, 126 Multipotent adult progenitor cells (MAPCs), 8, 126, 271, 331, 402 contribution, to chimeras, 262 differentiation ability of, 261 ESCs, 258 engraftment of, 261–262 greater potency, 262–263 isolation of, 260–261 postnatal tissue-specific stem cells, 258–260 Multipotent cell, 30, 127, 268 Multipotent somatic stem cell, 301

Multipotent stem cells, 258, 346–347, 972 adipocyte-derived, 346–347 bone marrow, 346 umbilical cord, 346–347 Multi-transgenic pigs, 160 Multi-wall carbon nanotube (MWNT), 557 Murky atmosphere, 461 Musashi-1, 1073 Muschler’s selective retention technology, 876, 879, 883 Muscle-derived stem cells (MDSCs), 972, 1328 Muscle interstitial cells, 391 Muscle regeneration, 274, 388, 389, 390, 1232, 1238, 1327 Muscle stem cell niche, 392 Muscular dystrophy, 888, 902–903 Musculoskeletal repair articular cartilage, 900–901 bone, 897–899 cell biology, 888 chondrocytes, 889 muscle cells, 889–890 osteoblasts, 888–889 osteocytes, 888–889 tendons and ligament cells, 890 cell-based therapies for, 888, 896–905 craniofacial tissue, 890–891 strategy of, 896 connective tissue cell, types, 958 craniofacial tissue, 904–905 ESCs, 890–894 gene therapy, 897 in vivo cell-based, reports, 898 invertebral disk, 904 meniscus, 901–902 MSCs, 891 identification, 893–894 immunoregulation, 895–896 in vitro behavior, 893 isolation, 892–893 niche, 894 regulation, 894–895 osteochondral tissue, 902 skeletal muscle, 902–903 tendon and ligament, 903–904 tissue engineering, 896–897 Musculoskeletal tissue engineering, 1322–1323, 1327 Myeloid cells, 288–291

1429

1430

INDEX

Myoblasts, 140, 783–784, 785, 821–822, 903, 1233 culture, 669, 1033–1034 skeletal, 386, 388, 391, 393, 1039, 1232 transplantation, 784, 785, 788, 789 see also Adult myoblasts Myocardial infarction, 448, 1038, 1039, 1042, 1053 neurological regeneration, 245–246 Myocardial repair, 39, 824, 835, 1039, 1040 MyoD, 58, 140, 233, 1041 Myofiber, 108, 783, 784–785, 786, 820, 1232, 1233, 1234, 1236, 1240 Myofibroblast, 688, 710, 730, 1004 apoptosis, 77, 80, 82 differentiation, 73, 76–77, 79–80, 81–82, 83 Myogenic-cell in skeletal muscles, implantation cell implantation, 785 cell injection density, 786–787 risks, 787–788 steps, to improve efficiency, 788 cell survival, in recipient early survival, 788–789 ensuring, 789 long-term survival, 789 properties of donor-derived satellite cells, formation, 785 gene complementation, 783–784 myofibers formation, 784–785 Myogenic-cell transplantation, in skeletal muscle, 782 cell implantation, 782 cell injections, density of, 786–787 risks, in procedure, 787–788 cell survival, in recipient, 788–789 early survival, 788–789 ensuring, 789 long-term survival, 789 cells to graft, 782 properties of, 783–785 Myogenic phenotype induction of pluripotent stem cells, 233 Myostatin (Mstn), 1239 Müller Glia, 422, 427–428, 431 N-acetylglucosamine, 595, 596 N-cadherin, 51, 291, 292 N-vinyl pyrrolidone (NVP), 622 Na glucose cotransporter 1 (SGLT1), 1079 Naive cells, 238 Nanocarrier components, 554, 555, 557

Nanodelivery vehicles, 557 Nanofibrous biodegradable scaffolds, 584 Nano-fibrous PLLA (NF-PLLA) scaffolds, 751, 752 Nanog, 100, 131, 132, 140, 196, 201, 216, 258, 263 Nanoparticulate imaging probes, 556 Nanotechnology, applications of, 554 as multifunctional tool for biomaterial control, 558 for cell-based therapies, 556–558 impact, on regenerative medicine, 555–556 National Bioethics Advisory Commission, 373 National Institutes of Health (NIH), 372, 1337, 1356, 1376 National Research Council, 757 Natural polymers, 637, 1289 hybrid scaffold, 642, 644–646 Natural scaffolds alignates, 1185 chitosan, 595–597, 1186 collagen, 594–595, 1185 ECM, 597–598 hyaluronic acid, 1186 Natural signals, molecular modification of, 1024, 1025 NCBI GEO, 563 Necrosis, 693, 761, 783, 786, 815, 1335 Neo-dermis, 17 Neointestinal cysts, 1074, 1078, 1079, 1080 Neo-muscles, 785 Neomycin selection, 172 Neonatal donor, for determined stem cells, 354 Neovascularization, 271, 710, 728–729, 812, 1020–1021, 1023, 1024, 1027, 1030, 1031, 1201 abnormal retinal, 442 cell-based repair for cardiovascular regeneration and, 812 EPC, role of, 441–442, 443 process of, 1023 screening of, 1031 transplanted ECs, 1027 Nerve cell regeneration and BMSCs, 275 Nerve growth factor (NGF), 960, 1238, 1260, 1261, 1276 Nerve guidance conduits/channels (NGCs), 1270, 1272–1275 Nestin-positive cells, 406, 973 Neural applications, in regenerative medicine, 39 Neural cell adhesion molecule (NCAM), 356, 571, 890 Neural cell surface markers, 244–245

Index

Neural Ceroid Lipofuscinosis, 9 Neural marker, 242, 859 Neural stem cells (NSCs), 7, 104, 115, 138, 139, 1256–1257 cross-differentiation and cell fusion, 302 definition, 301–302 history of, 300 in vitro neurogenesis, 303–305 epigenetic, 303–305 genetic, 305 in vivo neurogenesis, 302–303 therapeutics and clinical approaches in neurodegenerative diseases, 309–310 spinal cord injury, 305–309 stroke, 310–311 Neural tissue regeneration, 970 NeuroD1, 244, 423 Neurodegenerative diseases, 309–310, 1262 Neurofilament (NF), 245 proteins, 947 Neurogenic phenotype induction of pluripotent stem cells, 233 Neuroglial progenitor cell markers, 245 Neurological regeneration aldehyde dehydrogenase expressing cells, 251 Alzheimer’s disease, 243 amyotrophic lateral scelerosis, 243–244 cardiac disorders, clinical trials for, 246–247 cardiac treatment, 245 chondrocytes, 250 diabetes, treatment of, 247 endothelial progenitors, 249–250 ex vivo expansion, 250–251 hepatocyte-like cells, 248–249 Huntington’s disease, 243 myocardial infarction, 245–246 neural cell surface markers, 244–245 Parkinson’s disease, 243 spinal cord and CNS injuries, 244 stroke, 241–243 Neuronal cells, 946 preparation of, 948 production of, 948 arm test, 951 basis of, 946–947 clinical trial design, 948–949 cognitive testing, 951 functional outcomes, 950 in future, 952

imaging, 951 potential mechanisms, 947–948 safety and feasibility, in patients, 951–952 technique, 949 transplantation, 946 Neurospheres, 37, 423, 1263 Neurotrophins, 1260, 1276 New Drug Applications (NDA), 1370 NGF fusion proteins, 960 Ngn2, 423, 425 Nitric oxide (NO), 726, 982 Nitrogen ion beam implementation of, 660 Nodal point, 936 Nodal protein, 142 Nodal signaling, 129, 142, 403, 404 Non-autologous human cells, 1120 Non-canonical pathways, 421 Non-classical type, of molecular regulation, 132 Noncomplicity defenses, 1337 Non-covalent coating method, 661 Non-degradable synthetic polymers polymers with –C–C– backbone hydrolytically stable polyurethanes, 611–612 poly(2-hydroxyethyl methacrylate), 608 poly(ethylene terepthalate), 611 poly(ethylene), 606 poly(meth)acrylates poly(N-isopropylacrylamide), 608–609 poly(propylene), 606 poly(styrene), 606 poly(tetrafluoroethylene), 606–607 polyacrylamides, 607–609 polyethers, 609–610 polysiloxanes, 610–611 Non-human primate models, 461 Nonhuman sialic acid, 214 Nonindividuation argument, 1335–1336 Non-invasive monitoring method, 483, 556, 1031 Non-keratinized epithelium, 1060 Non-linear viscoelastic models, 1210 Non-organ transplant candidates hepatocyte transplantation for, 921–922 Nonprocreative cloning, 1338, 1340 Non-seeded acellular matrices, 1127 Non-seeded allogeneic acellular bladder matrices, 1129 Nonskeletal tissue regeneration, by MSCs, 330–332 Non-steroidal anti-inflammatory drugs (NSAIDs), 1236

1431

1432

INDEX

Non-traumatic peripheral nerve injuries, 1270 Non-viral vectors, 747, 874 Normal stress, in tissue development, 481, 517 Normosol®, 759 Notch intracellular domain (NICD), 110 Notch ligand, 250, 739 Notch signaling pathway, 56, 107, 140, 404 NS-398, 1236 Ntera 2/c1.D1 (NT2), 947 Nuclear export sequence (NES), 59 Nuclear localization, 58, 59 Nuclear magnetic resonance spectroscopy (NMRS), 938, 939 Nuclear reprogramming process, 131 Nuclear transfer (NT), 34, 148, 184, 457 and embryonic gene expression pattern, 155–156 ntESCs, 459–460, 461, 462, 463 parameters, effects of calcium oscillation, 179 cell cycle, 176 cell isolation, 177 cell passage number, 179 cell type, 175–176 cycloheximide, use of, 178 embryo culture, 179–180 fusion/cleavage, screening tool, 179 number of passages, 177 ultraviolet versus polarized light enucleation, 177–178 somatic NT, 149–151, 158–160 Nuclear transplantation, 1109, 1338 see also Somatic cell nuclear transfer (SCNT) Nucleofection, 194, 325 Nucleus pulposus (NP) cells, 327, 904 Oasis™, 114 Occludin, 57 Octamer-binding protein 4 (Oct-4), 192, 201, 229, 230, 571, 855, 856 expression of, 258, 260, 261, 262, 263 knock-out embryos and stem cells, 131 locus, 132 Ocular citracial pemphigoid, 1060 Ocular trauma, 1060 Ocusert®, 606 Odontoblasts, 1287 Office of Combination Product (OCP), 1370 Oil-O-Red, 232

Olfactory ensheathing glia (OEG), 1256 Olfactory nerve neurons, 104 Oligo(poly(ethylene glycol) fumarate) (OPF), 623, 624 Oligodendrocytes, 244, 303, 310, 1248 Oligodeoxynucleotides (ODN), 1216, 1217 Oligomaltose, 664 Oligopotent stem cells, 288 Ontogeny, of satellite cells, 394 Oocytes, 1340, 1341 donation, 970, 1343 enucleation, 149–150, 173, 175 in vitro activation, 571 metaphase II stage, 178 parthenogenesis, 575 quality, 462 Operational tolerance, 805 Opsonins, 709 Optic cup, 419, 421, 422 Optic vesicle cells, 418 Optipress II, 239 Optx2 (Six6), 418, 421 Orapharma, 1400, 1401 Orcel, 1378 Orkin®, 1394 Ornithine transcarbamylase (OTC), 913, 919, 920 Orphan Drug Act, 1386 Orthopedic applications, in regenerative medicine, 40–41 Orthotopic liver transplantation (OLT), 912, 1086 Osteoarthritic TMJ, 904 Osteoarthritis (OA), 512, 767, 888, 900, 902, 1176 Osteoblast (OB), 105, 272, 291, 493, 639, 750, 752, 875, 888–889, 894 niche, 272 Osteocalcin, 232, 328 Osteochondral autografting, 1184 see also Mosaicplasty Osteochondral tissue, 902 Osteochondral transplantation, 769 Osteoclast, 493, 815 Osteoconduction, 897, 1165, 1166–1167 Osteoconductive materials, 871 beta-tricalcium phosphate, 1167 ceramics, 1166 collagen, 1166–1167 Osteoconductive scaffolds, 871 Osteocytes, 493, 888–889 Osteogenesis imperfecta, 40, 874

Index

Osteogenetic proteins (OPs), 747 OP-1, 501, 1168 Osteogenic cells, 746, 870, 871 delivery of, 870–871 sources for cell therapy allogeneic bone marrow, 873 autologous bone marrow, 872–873 gene therapy, 874–875 novel tissue sources, 873–874 Osteogenic phenotype induction of pluripotent stem cells, 232 Osteogenin, 501 Osteoinduction, 897, 1165 Osteoinductive graft materials, 872 Osteoproduction, 897 Osteoprogenitor cells, 745, 875, 876, 889, 1166 Osteopromotive graft materials, 871 Osterix, 889 “Oval cell” response, 361 Ovarian follicles in vitro culture of, 1157–1159 Ovary anatomy, 1157 in vitro culture, of ovarian follicles, 1157–1159 Overcoating technologies, of surface modification, 660 covalent coatings, 661–662 non-covalent coatings, 661 schematic representation of, 658 Oxygen transport, for cardiac tissue engineering, 1050–1051 P-selectin, 726 P130CAS, 71, 488 PA6, 215–216, 428, 429 PAI-1, 79 Pancreatic beta cells, 7, 139, 331 Pancreatic development, in islet cell therapy, 403 β-cell differentiation, 404 endocrine specification, 404 endoderm/gut endothelium generation, 403–404 pancreatic differentiation, 404 Pancreatic differentiation, 404 Pancreatic islet cells, 973 encapsulation, 956 Pancreatic stem cells β-cell replacement therapy, 389–392 benefits, 389–390 limitations to, 391–392 in future, 398–399

insulin-producing cells alternative sources of, 392–393 islet cell therapy and, 398 islet neogenesis, 394–397 from adult stem cells, 395–397 from ES cells, 394–395 pancreatic development, 393–394 endocrine specification, 394 endoderm/gut endothelium generation, 393–394 β-cell differentiation, 394 transdifferentiation, 393, 397–398 Paneth cells, 1074 Paracrine effects, of MSCs, 324 Parenchymal progenitors, 357 Parkinson’s disease, 2, 39, 243, 969, 1120, 1262 Parthenogenesis in vitro activation, of oocyte, 575 Parthenogenetic embryonic like stem cells (PGESC), 571, 575 Partial thickness lesion. See Chondral lesion Patch-clamp studies, 488 Patellar tendon (PT) fibroblast, 1216 Patent, 184, 1343, 1387 provisional applications role, 1393 Patent claims, 1389, 1396, 1397 Patent Extension Act, 1394 Patent infringement, 1343, 1389, 1397 claims, 1343 Patent rights, 1386, 1392, 1397 Pax2, 852 Pax6, 244, 420, 421, 428, 430, 852–853 Pax7, 388, 390 Paxillin, 71 Penetrating keratoplasty (PK), 1060 Penile prothesis, for reconstruction, 1148 Penis, tissue reconstruction of, 1143 anatomy, 1144 corpus cavernosum reconstruction acellular collagen matrix preparation, 1145–1146 human endothelial cell culture, 1146–1148 smooth muscle cell culture, 1144–1145 penile prothesis, 1148–1149 cartilage tissue harvest, 1148 Percoll, 892 Percutaneous transluminal coronary angioplasty, 1001 Perfluorocarbons, 758 Perfusion bioreactors, 37, 490, 529 Perfusion scaffold systems, 935

1433

1434

INDEX

Perichondrium, 768, 1184 transplantation, 767 Periodontal ligament (PDL), 1286, 1287, 1291 Periodontium, 1287, 1292 tissue regeneration, 1291–1292 Periosteal grafts, 1184 transplantation, 767 Peripheral blood stem cells therapeutic applications, 445 of EPC, 446–448 gene therapy, 448–449 of MSC, 448 types and source, 438 bone marrow cells, 439 EPC, 439–444 HSC, 445 MSC, 444–445 Peripheral nerve regeneration, 1270 animal models, 1280–1281 anisotropic nerve grafts, 1278 aligned anisotropic scaffolds, 1278–1279 cell-seeded longitudinally aligned NGCs, 1279 ECM molecules, 1279 neurotrophic factors, 1279 current strategies, 1271 ECM molecules, 1276–1277 historical background, 1256 isotrophic nerve grafts, 1272 agarose gels, 1275 collagen gels, 1275–1276 hydrogels, as scaffolds, 1275 natural materials, as scaffolds, 1272 synthetic scaffolds, for nerve repair, 1272–1275 natural nerve grafts, 1279–1280 neuronal support cells for, 1277–1278 neurotrophic factors, 1276 problems and challenges, 1256 Peripheral nervous system (PNS), 1248, 1250, 1270, 1275, 1277 Permacol®, 114 Peroxisome proliferation-activated receptor 2 (ppar(2)), 232 Pfizer, 1400 Phagocytosis, 709 Phalanges and small joints composite tissue structures, creation, 1201–1203 future perspectives, 1203 principles of, 1199–1201

Phenobarbital, 917 Phenotype, cell change, 31, 136, 139–140 cell fusion, 138–139 extrinsic controls, 141–142 gene activity control, 140–141 plasticity, 137–138 stem cells, 136–137 Phorbol 12-myristate 13-acetate (PMA), 329 Phosphacan, 68 Phosphate-buffered saline (PBS), 194, 227, 961 Phosphatidylinositol bis-phosphate (PIP2), 72 Phosphatidylserine, 557 Phospholipase C (PLC), 523 PLCγ, 72 Photografting, 661 Photolithography, 660, 664, 674–675, 698, 1028 Photopatterning method, 546, 547 Photopolymerizable hydrogels, 698 Photopolymerized (meth)acrylated polymer networks, 624, 625 Phronesis, 1349 Physical entrapment method, 663 Physical stress, 512, 517 effects on repair, and remodeling, 519–520 effects on tissue growth, 519 in vitro mechanical conditioning blood vessel bioreactors, 528–530 bone bioreactors, 527–528 cartilage bioreactors, 525–527 mechanotransduction mechanism, 521–525 and regenerative medicine, 520–521 structural hierarchy and continuum concept, 513 Physiochemical surface modification chemical modifications, 657–660 topographical modifications, 660 Phytohemagglutinin (PHA), 329 PI3K, 54, 55 Piezoelectric-driven micromanipulator, 460 Piezoelectric microinjection tool, 150 Pigment epithelium-derived factor (PEDF), 853, 860 Pigmented epithelium, 425–426 Pipecolic acid, 191, 921 Placenta and amniotic fluid differentiated cells, 227 isolation and characterization, 231–234 mesenchymal cells, 227–230 pluripotent stem cells, 230–234 Placental growth factor (PLGF), 104, 439, 1023

Index

Plaque activation, 818 Plasma deposition, 661 Plasma fibronectin, 81 Plasma treatment, 671 Plasmin, 81 Plasminogen activator, 53 Plasticity, 5, 137–138 embryonic development, 137 hematopoietic stem cells, 138 larval development, 138 neural stem cells, 138 trophectoderm, 137 Platelet activating factor (PAF), 726 Platelet-derived growth factor (PDGF), 67, 214, 667, 694, 706, 710, 725, 726, 745, 901, 1001, 1165, 1167, 1171–1172, 1212, 1216, 1239, 1289, 1292 Platelet-derived growth factor-BB (PDGF-BB), 271, 1021, 1023, 1026, 1215 Platelet factor 4, 706, 707 Platelet-rich plasma (PRP), 871, 905 Platelets, 706 Pluripotency, 218 of BMSCs, 268 of hepatic stem cells, 344 of human ES cells, 215 molecular control of, 216 molecular basis, 126–133 BMP/GDF, 128–129 epigenetic and environmental regulation, 131–132 extracellular signaling factors, 127–131 extracellular signaling transduction, 127–131 fibroblast growth factor (FGF), 130 leukemia inhibitory factor (LIF), 127–128 TGF-β/activin/nodal, 129 transcriptional networks, 131 Wnt, 130 Pluripotent cells, 30 in vitro, 127 in vivo, 127 from umbilical cord blood cells, 241 Pluripotent embryonic carcinomas, 132 Pluripotent parenchymal progenitors, 357 Pluripotent stem cells, 130, 230, 330, 571, 1342, 1354–1355 isolation and characterization, 231–234 from placenta and amniotic fluid, 230–234 Pluronics®, 609, 618 PMPs (pancreas-derived multipotent precursors), 259, 262 Podocytes, 1114, 1121

Point-of-care osteogenic cell enrichment, 875 Poisson’s ratio, 514, 518 Poloxamers, 609–610 Poly tetrafluoroethylene (PTFE), 606–607, 979, 1063, 1077 Poly(2-hydroxyethyl methacrylate) (PHEMA), 607, 608, 637, 664, 1063 Poly(2-hydroxyethyl methacrylate-co-methyl methacrylate (P(HEMA/MMA)), 1253 Poly(amino acids), as synthetic polymers, 617 Poly(D, L-lactic acid-co-glycolic acid) (PLGA), 4, 581, 583, 589, 613, 639, 698, 720, 747, 749, 881, 890, 1005–1006, 1027, 1127, 1275, 1289 scaffolds, 627 Poly(diallyldimethylammonium chloride) (PDAC), 752 Poly(dimethyl siloxane) (PDMS), 537, 541, 610, 644, 675, 1043, 1063, 1064 Poly(DTE carbonate), 617 Poly(ethylene glycol) (PEG), 113, 583, 609, 672, 693, 900, 956, 1025, 1064, 1183 and diacrylate, 623, 624, 625 and dimethacrylate, 624, 625 vinylsulfones, 624 Poly(ethylene glycol) diacrylate (PEGDA), 546, 695, 698, 699 Poly(ethylene oxide) (PEO), 609, 615, 641, 670, 1186 Poly(ethylene terephthalate) (PET), 611 Poly(ethylene-co-vinyl acetate) (PEVAc), 606 Poly(ethyleneimine) (PEI), 588, 747 Poly(glycolic acid) (PGA), 582, 613, 614, 618, 637, 639, 668, 693, 696, 749, 775, 900, 984, 990, 1005, 1127, 1138, 1200, 1254, 1275, 1289 fibers, 1074 scaffolds, 1006, 1009, 1026, 1044, 1081, 1131, 1181 Poly(hydroxybutyrate) (PHB), 1254 Poly(iminocarbonates), 617, 641 Poly(L-lactic acid) (PLLA), 581, 614, 749, 1005, 1027, 1074, 1149 Poly(L-lactide), 588, 668 Poly(L-lysine), 1151–1152 Poly(lactic acid) (PLA), 613, 614, 637, 693, 900, 990, 1046, 1127, 1182, 1275 Poly(lactic acid-co-lysine), 582 Poly(lactide-co-glycolide) (PLG), 1026, 1048, 1201 Poly(meth)acrylates, 608–609 Poly(methyl methacrylate) (PMMA), 604, 607–608, 712 Poly(N-isopropyl acrylamide), 608–609, 610, 620, 637, 1066

1435

1436

INDEX

Poly(N-isopropylacrylamide-co-acrylic acid-coacryloxysuccinimide), 1067 Poly(p-dioxanone), 616 Poly(propylene fumarate) (PPF), 621–622, 623, 637 Poly(propylene fumarate-co-ethylene glycol) (P(PF-co-EG)), 622, 623 Poly(propylene oxide) (PPO), 609 Poly(propylene) (PP), 606–607 Poly(SA–CPP), 619 Poly(styrene) (PS), 606–607 Poly(α-hydroxy acids), 613–615, 639–641, 749 Poly(α-hydroxy esters), 615, 637–638 Poly(ε-caprolactone) (PCL), 616, 643, 668, 901, 1138, 1200 Poly-3-hydroxybutyrate, 1275 Poly-4-hydroxybutyrate, 985, 1006 Polyacrylamide, 608–609, 669 Polyactive, 1327 Polyanhydrides, 581, 582, 619, 696 Polycarbonates, 617, 1254 Polycarprolactone, 581 Polycomb genes, 288 Polyelectrolyte multilayers (PEMs), 671, 752 Polyesters, synthetic polymers poly(α-hydroxy acids), 613–615 polycarbonates, 617 polylactones, 616 polyorthoesters, 616–617 Polyethers, 609–610 Polyethylene, 606 derivatives poly(propylene), 606 poly(styrene), 606 poly(tetrafluoroethylene), 606–607 Polylactones, 616 Polymer demixing, 675 Polymerase chain reaction (PCR), 918, 1185, 1201 Polymeric encapsulation systems, 956 Polymorphonuclear neutrophils (PMN), 745 Poly-N-acetyl glucosamine (p-NAG), 1324 Polyorthoesters (POEs), 616–617 Polyphosphazenes, 619–620 Polyploid adult cells, 349, 358 Polypropylene oxide, 1066 Polysiloxanes, 610–611, 658 Polyurethanes (PUs), 605, 618 Polyurethaneurea, 612, 618, 691 Polyvinyl alcohol (PVA), 624 Polyvinylpyrolidone (PVP), 758, 1063

Porcine dermal matrix, 114 Porcine small intestine submucosa, 114, 597, 1218 Porous scaffolds, 596, 691–692 in cardiac tissue engineering, 1044 and high interconnectivity, 748 Portal vein thrombosis, 795, 799 Posterior definitive endoderm, 403 Postnatal tissue-specific stem cells, 258–260 Post-operative rehabilitation protocol, 768 Post-zygotic reprogramming telomere length, 158 X-chromosome inactivation, 157–158 Potency, of cells, 126–127 pp125FAK, 488 Preatheroma, 1000 Precursor cells. See Progenitor cells Pre-implantation mammalian embryo, 131 Pre-market approval (PMA), 610, 1172, 1309 Prenatal mortality, 457 “Pressure relief valve” mechanism, 488 Pre-zygotic reprogramming DNA methylation patterns, cloned embryos and fetuses, 156–157 embryonic gene expression patterns, 155–156 imprinted gene expression, in cloned embryos and fetuses, 154–155 Primary cells, 172, 954–955 Primary mesenchyme cells (PMCs), 51, 52 Primatrix™, 114 Primitive endoderm, 131 Primitive hematopoiesis, 285 Principles, of regenerative medicine current and future perspectives, 2–10 Prion disease, 183 Processed lipoaspirate (PLA) cells, 322 Procreative cloning, 1338 Progenitor cells, 29, 130, 245, 1314 Progestasert®, 606 Progressive familial intrahepatic cholestasis (PFIC), 919, 920 Pro-HGF, 102 ProNectin-L, 1115 Proneural bHLH gene, 423 Prostaglandin (PG), 493, 1151, 1236–1237 PGE, 319, 528, 1167 PGE-2, 319, 528 Prostanoids, 726 Prosthetic tube, 1075

Index

Protein kinase C (PKC), 72, 524 Protein manufacturing, 346 hepatic stem cell, use of, 366 Protein–protein cross-linking agents, 598 Proteins, 1024–1025 control orientation antibodies, control orientation, 739 collagen and ECM, orientation, 739–740 control of packing density, 737 G-protein, 739 histidine (HIS6) tags, 737 hydroxyapatite for orientation control, 739 immobilization in lipid layers, 739 ionic charge (pH) orientation, 738 preservation agents, 737 streptavidin and avidin orientation, 738–739 templating methods, 740 tethered lipid bilayers, 739 implications, 740 matrices, 214 neovascularization, 1023 non-specific protein absorption, 735–736 inhibition of, 736 precision control for tissue engineering, 734 Proteoglycan, 66, 81, 350, 351, 1176, 1181, 1208, 1280 Provisional matrix formation, 705, 706, 713–714 Proximal tubule cells isolation and culture, 1115, 1116, 1117 Public Expression Profiling Resource, 563 Public Health Service (PHS) Act, 1355, 1367, 1397 Publication, and patentability, 1389–1390 Pulmonary radiotracer uptake, 920 Pulsatile flow bioreactor, 1009 Quantile normalization, 566 Quantitative reverse transcription-polymerase chain reaction (qRT-PCR), 1008 Quantum dots (QD), 556 Quasi-linear viscoelastic theory, 1210 QuickClot, 1325 Radiation grafting, 661 Radio-frequency glow discharge (RFGD), 658, 661 Rana pipiens, 148 RANKL (receptor activator of NFkB ligand), 474 Rap1, 52–53 Ras, 52, 524

Rat amniotic epithelial (RAE) cells, 229 Rat pancreatic extract (RPE), 331, 567 Reactive oxygen species (ROS), 53, 726 Receptor–ligand interactions, 78–79, 292, 1261 Receptor tyrosine kinase (RTK) ligands, 55–56 Recombinant human albumin (rhA), 183 Recombinant human bone morphogenic protein-2 (rhBMP-2), 624, 1165, 1169–1170 Recombinant human growth hormone (rhGH), 720 Recombinant human OP-1 (rhOP-1), 1168 Recombinant human platelet-derived growth factor (rhPDGF), 1166, 1167, 1170 Recombination-activating gene 2 (Rag-2), 459 Recoverin, 429, 430 Recurrent autoimmunity, 805 Re-epithelialization, 74, 76, 79, 729 Regeneration index (RI), 107 Regeneration mechanisms, 101 adult stem cells, activation of, 102–106 aging effects, 108 compensatory hyperplasia, 101–102 dedifferentiation, 106–108 Regeneration phase, in skeletal muscle healing, 1234–1235 exercise and muscle regeneration, 1240 growth factors, 1238–1239 hyperbaric oxygen, 1240 myostatin, neutralization of, 1239 therapeutic ultrasound, 1239–1240 Regeneration templates, 114 Regional anatomical domains, 1164 see also Bone graft Regional gene therapy, 747 Relaxin, 1241 Renal cell culture, 1106 Renal cell proliferation, 1108 Renal proximal tubule cells, 1115 Renal structures ex vivo, 1132 in vivo, 1132 creation, 1108–1109 Renal transplantation, 1111 Renal tubule assist device preclinical characterization, 1116–1117 Replicative senescence cell culture modeling, of T cell, 470–471 gene expression and function, 471–472 retarding/preventing, in aging and HIV disease, 474–475

1437

1438

INDEX

“Reproductive cloning”, 458 of transgenic animals, 158, 160 Reproductive system, tissue engineering of, 1138 in female ovary, 1157–1159 uterus, 1155–1157 vagina, 1152–1154 in male penis, 1143–1149 testes, 1149–1152 urethra, 1139–1143 Request for Designation (RFD), 1370 Retinal diseases ciliary marginal zone (CMZ), 424–425 eye-field transcription factors, 418–421 retinal progenitors, 421–423 retinal regeneration and transdifferentiation ciliary epithelium, 427 intrinsic stem cells, 427–428 Müller Glia, 427–428 pigmented epithelium, 425–426 retinal neurons, from ES cells, 428–430 rod precursors, 427–428 Retinal pigment epithelium (RPE), 420 characteristics, 853 development, 852–853 from embryonic stem cells, 855–859 HES-derived, culture and properties of, 859–861 transdifferentiation, in culture, 425–426, 853–854 transplantation, 854–855 Retinal progenitors, 421–423 Retinal regeneration ciliary epithelium, 427 intrinsic stem cells, 427–428 Müller Glia, 427–428 pigmented epithelium, 425–426 retinal neurons, ES cells from, 428–430 rod precursors, 427–428 Retinal stem cells, 424, 427 Retinitis pigmentosa, 418, 852 Retinoic acid, 117, 141, 142, 242, 428, 890, 947 Retroviral vector, in viral transduction, 195–196 Reverse transcriptase polymerase chain reaction (RT-PCR), 155, 234, 347, 506, 562, 773, 853, 1109, 1217 Rhabdomyosarcomas, 393, 785 RHAMM (receptor for hyaluronate-mediated motility), 70 Rheumatoid arthritis, 888 Rhodopsin, 429, 430

RLBP1 gene, 853 RNA interference (RNAi), 190 RNA-induced silencing complex (RISC), 193 RNAs, 565 biological roles of double-stranded RNA (dsRNA), 193 RNA interference (RNAi), 192–193 short hairpin RNA (shRNA), 193–194 small interfering RNA (siRNA), 193 Robust differentiation protocols, 460 Robust Multichip Analysis (RMA), 566 Rod-cone dystrophy, 853 Rod precursors, 427–428 Rosuvastatin, 819 Rotating-wall bioreactor, 490, 492 RPE65, 853, 856, 860 RPT file, 565 Runx2, 105, 889, 899 Rx, 418, 419–420 Rx/Rax, 419 Salt-leaching technique macroporous scaffolds fabrication, 583 Sarcomeres, 386, 1051 Sartorelli profiled C2C12, 572 Satellite cells, 106, 387, 388–389, 783, 785, 820, 890, 1234, 1240 Sca-1, 260, 262, 292, 444, 831, 1040 Scaffolds, 1178, 1184, 1218–1220 as biomimetic biomaterials, 505–506 bone tissue engineering, 748 composite, 749 design criteria, 748 interconnectivity, 748 nano-fibrous, 751–752 porosity, 748 for cardiac tissue engineering cell-based cardiac patches, 1043–1044 cell-free cardiac patches, 1043 combination approach, 1046 fibrous scaffold approach, 1044–1046 hydrogel approach, 1043–1044 porous scaffold approach, 1044 thin films, 1046 for cartilage tissue engineering autologous scaffolds, 1184–1185 natural, 1185–1186 synthetics, 1186–1187

Index

for nerve repair, 1272, 1278–1279 hydrogels, 1275–1276 natural materials, 1272 synthetic, 1272–1275 for small diameter tissue-engineered vessels biologic scaffolds, 582–590, 1006–1008 cell seeding techniques, 1008–1009 scaffold-free approach, 1009 synthetic scaffolds, 648–651, 1005–1006 histogenesis, 686 hybrid scaffold biomaterial, 639–648 ceramic, 641–644 miscellaneous, 646–648 natural polymers, 644–646 poly(α-hydroxy acid) family, 639–640 tissue engineering biomaterials, 636–637 bioceramic scaffolds, 638 importance, 636–637 natural polymers, 637 poly(α-hydroxy esters), 637–638 synthetic polymers, 637 Scanning electron microscopy (SEM), 325, 543, 584, 585, 589, 648, 649, 749, 750, 751, 855, 1110, 1189, 1278 Scar tissue formation, 1301–1302 Schwann cells, 106, 115, 442, 970, 1254, 1256, 1272 Screening arrays, 539–541 SeaPrep® agarose hydrogel, 1275 Sebacic acid (SA), 582, 619 Secrete human leukocyte antigen-I (sHLA-I), 918 Seeded tubularized collagen matrices, 1127 Self-assembled monolayers (SAMs), 549, 612, 661, 737 “Self ” tissue, 595 Sendai virus, 138 Sepax, 239 Septic shock, 1117 Serine/threonine kinases, 54, 72, 504, 524 Serum replacement (SR), 191, 213, 856, 858 Serum starvation, 151, 176 Service mark, 1394 Severe combined immunodeficiency (SCID), 295, 325, 366, 459, 571, 1053, 1376 “Shake-off ” method, 177 Shear strain in tissue development, 481 Shear stress, 490, 491, 512, 517, 528, 679–680 in tissue development, 481 Short bowel syndrome, 1076 Short hairpin RNA (shRNA), 193

Sickle cell anemia phenotype, 460 Side population (SP) cells, 33, 269, 293, 872–873, 973 Signal transduction events, during cell–ECM interactions adhesion and migration, 70–72 apoptosis, 73 differentiation, 72–73 proliferation and survival, 72 Signal transduction pathways, 73, 676 Signaling molecules, 54, 484, 485 Notch pathway, 56 RTK ligands, 55–56 TGF-β pathway, 54 Wnt pathway, 55 Silicones, 605, 610 Simultaneous islet kidney (SIK), 400 Singapore, 373 stem cell research policy, 1358 Single cell suspension, 17 Single channel. See Affymetrix Single-pass albumin dialysis (SPAD), 1095 Single-wall carbon nanotube (SWNT), 557 Sirolimus, 799, 800 Six3, 418, 419, 421, 852 Skeletal muscle, 106, 902–903 myogenic cells, implantation, 782–790 of stem cells, 386 after injury or disease, 387 challenges facing the potential therapeutic uses, 392–393 characteristics, 390–391 criterion, 387–388 historical perspective in, 388–390 muscle stem cell niche, 392 in natural environments in vivo, 391–392 regeneration and BMSCs, 274 regenerative medicine implications, 1232 degeneration phase, 1233–1234 fibrosis phase, 1235, 1240–1241 inflammation phase, 1234, 1235–1238 regeneration phase, 1234–1235, 1238–1240 Skeletal muscle cells, 889–890 Skeletal muscle healing, inflammation role in, 1235–1238 cyclooxygenase pathway, 1236–1237 inflammation and muscular dystrophy, 1237–1238 inflammatory cells, 1235–1236 prostaglandins, 1236–1237 Skeletal myoblasts, 1039

1439

1440

INDEX

Skeletal tissue regeneration, by MSCs bone, 322–325 cartilage, 325–326 intervertebral disk, 327–328 tendon, 326–327 Skin cell-based therapy, for burn victims, 1298 clinical studies, 1303–1305 conventional therapy and current commercial products, 1298–1301 regulatory issues, 1309–1310 21 CFR 1271.10(a), 1311 21 CFR parts 1270 and 1271, 1311 “351” HCT/P, 1313–1314 “361” HCT/P, 1313–1314 biologics, licensing of, 1310–1311 HCT/Ps, 1311 skin progenitor cells, 1307–1309 spray transplantation method, 1302–1303, 1305–1307 wound healing, problems of, 1301–1302 Skin-derived progenitors (SKPs), 259 Skin repair, topical agents for, 113–114 Skin tissue engineering, 971 Skin wound healing, cell–ECM interactions adhesion and migration, 73–75 apoptosis, 77–78 differentiation, 76–77 proliferation, 76 Slug, 57, 58 Smad, 54, 129, 502–503, 1182 Smad interacting protein (SIP), 503 Sip1, 57 Small bowel, 1076–1080 Small diameter tissue-engineered grafts, 1000 bioreactor cultures, 1011–1012 cell sourcing, 1003–1005 cellular interactions, 1010–1011 in vivo construct fabrication, 1012 scaffolds, 1005–1009 Small interfering RNA (siRNA), 193 Small intestinal submucosa (SIS), 597, 637, 719, 984, 1008, 1075, 1077, 1078, 1127, 1143, 1200, 1218 and ECM, 598 Small molecules, for chemical induction, 117 Smooth muscle alpha actin (SMA), 1026 Smooth muscle cells (SMCs), 442, 491, 818, 978, 980, 981, 982, 992, 1000–1001, 1004, 1007, 1010, 1021 Snail, 57 Sodium alginate, 1151

Soft lithography, 541, 664, 675 Soluble ephrin-B2, 1023 Soluble factors, 114–115 Soluble signals, 350, 358, 361 Solvent casting/salt-leaching technique, 583, 649 Somatic cell nuclear transfer (SCNT), 111, 169, 174, 202, 212 blastocyst, 458 cell line development, 171 copy number, 172 FISH analysis, integration site, 172–173 neomycin selection, 172 time line, 171–172 transfection, 172 derived ESCs, in regenerative medicine, 456 development in, 459–461 DNA construct development, 170–171 history, 457–458 prospects and challenges, 461–463 Somatic cells, 20, 468 growth and expansion of T lymphocytes, 468 T cells cell culture modeling, of replicative senescence, 470–471 growth and expansion potential, 469–470 HIV disease, retarding/preventing, 474–475 immune system, 469 replicative senescence, gene expression and function, 471–472 senescent cells, 472–474 Somatic cloning, in animals applications reproductive cloning, 158–160 therapeutic cloning, 160 history, 148–149 success rates and normality of cloned offspring, 151–152 technical aspects, 149–151 see also Epigenetic reprogramming Sonic hedgehog (Shh), 108, 141, 404, 419, 426, 1042, 1181 Sourcing, of human cells and tissues, 367 Southern blotting analysis, 172 Sox-2, 6, 100, 112, 131, 140, 201, 216, 244, 258 L-Sox-5, 889 Sox-6, 889 Sox-9, 105, 573, 889, 904, 1183 SPARC (secreted protein acidic and rich in cysteine), 67, 76

Index

Spatial patterning, of cellular adhesive environment, 541–542, 543 Spemann, Hans, 418, 420, 457, 463, 502 Spheroid encapsulation, 934 of HF bioartificial liver, 934 Spinal cord injury (SCI), 1248 animal models acute versus chronic injury, 1252 compression models, 1251–1252 contusion models, 1251 transaction models, 1250 cell therapy, 1262–1263 and CNS injuries, 244 drug delivery, 1257 methods, 1258–1260 neuroprotective agents, 1260–1261 neuroregenerative agents, 1261–1262 routes, 1257–1258 entubulation, 1252–1255 chemotactic cues, 1256 haptotactic cues, 1255–1256 nerve guidance scaffolds, cell seeding in, 1256–1257 NSC therapy, 305–309 psychological events, 1248–1250 Spinal fusion, 323 Spinner flasks, 494, 899, 1008, 1012, 1044, 1046 Split-skin transplantation method, 1298, 1306 Sponsors, 1371 Spontaneous differentiation model, 31 Spray transplantation method, 1305–1307 Sprouting models, 1030, 1031 Stabilizing pressure input orthosis (SPIO), 557 Stage-specific embryonic antigens (SSEAs), 100, 855 SSEA-I, 444 STAT 3, 59, 100, 128, 213 Static flask, bioreactor, 490 Static versus dynamic, nutrient inputs, 938 Statins, 837 Stellate cells, 82 Stem cell-derived hepatocytes, 913, 922 Stem cell factor (SCF), 191, 244, 272, 277, 292 Stem cell research policy in Israel, 1358 Korea, 1358 Stem cells, 20–21, 29, 136–137, 268, 275, 320, 347, 386, 438, 503–505, 955, 1073, 1096–1097 adult stem cells, 32–34 autologous approach, limitations of, 768

based gene therapy versus BMP-2 protein delivery, 323, 324 behavioral model, 129 from bone marrow, 407 in bone tissue engineering, 745–746 and cancer, 351–352 cell-based therapies, issues in, 34–35 clinical, commercial, and research applications future uses, 364–367 properties, 362–364 cloning, 231 compartment, of human livers, 356–358 from cord blood, 238–251 derived from placenta and amniotic fluid, 226–235 differentiation, 569 AFSC, 571–572 hESC, 569–571 hierarchical model of, 29–30 meta-analysis, 572–573 PGESC, 571 EGCs, 29, 32 in embryos committed progenitors, 347 determined stem cells, 347 embryonic stem cells, 345–346 multipotent stem cells, 346–347 totipotent stem cells, 345 ESCs, 29, 30–32 ethics and legal issues in usage general issues, 367–368 survey by philosophy and religion, 368–369 survey opinions by countries, 369–373 United States federal and state law, 373–374 from fetal tissue, 969 of fetal origin, 402 in human liver, 354–362 insulin-producing cells, source of, 398–399 isolation and purification, 347–348 microenvironment, cues in, 35–38 niches, 271–272, 1288 in peripheral blood therapeutic applications, 445–449 types and source, 438–445 plasticity, 30, 302 and progenitor cells, 344–345 in regenerative medicine, 38 cardiovascular applications, 39–40 hematopoietic and autoimmune applications, 41

1441

1442

INDEX

Stem cells, (continued) in regenerative medicine, (continued) metabolic and secretory applications, 41 neural applications, 39 orthopedic applications, 40–41 research challenges, 1097–1098 revolutionary therapeutics for regenerative medicine, 562 of skeletal and cardiac muscle, 106 stroke, therapy, 310 transplantation, 803 US research policy, 1354 Stemness genes, 127, 132, 305, 392, 562, 569, 575 Stethoscopes, 1349 Stevens–Johnson syndrome (SJS), 1060 Stirred-flask bioreactor, 490–491 Stomach, 1080–1082 Strain, 513–517, 1209 in tissue development, 481 Strategies, of regenerative medicine bioartificial tissues, 112–113 cell transplants, 109–112 chemical/physical induction, of repair and regeneration, 113–117 Strength development, of cardiac tissue, 1047 Streptavidin and avidin orientation, 738–739 Stress, 512, 517 effects on morphogenesis, 519 effects on repair and remodeling, 519–520 effects on tissue growth, 519 in tissue development, 481 Stress fibers, 485, 667, 672 Stress urinary incontinence (SUI), 1134 Strictures, 1074, 1127, 1142 Stro-1, 271, 274, 319, 321, 746, 893 Stroke, 233–235 neuronal transplantation action research arm test, 951 basis, 946–947 clinical trial design, 948–949 cognitive testing, 951 functional outcomes, 950 imaging, 951 neuronal cells, 948 potential mechanisms, 947–948 safety and feasibility, 951–952 technique, 949 NSC therapy, 300–301

Stroma-supportive system, 321 Stromal cell derived factor 1 (SDF-1), 271, 277 Stromal cell-derived inducing activity (SDIA), 428, 856 Stromal cells, 199, 290, 330, 504, 770, 869, 1063 Stromal layer, of cornea, 1060 Stromal stem cells, 321 Stromal vascular fraction, 322 Structural cartilage, 1189 Structural molecules, of cytoskeleton, 486–487 Successful in vitro cartilage tissue engineering, 1187 bioreactors, 1188–1189 mechanical stimulation, 1188 Sulfated glycosaminoglycan (sGAG), 901 Superarrays, 564 Suprathel®, 1306 Suramin, 1241 Surface analyses techniques, 657 Surface charge, cell-substrate interactions, 670 Surface immobilization, of biomaterials on macroporous biodegradable scaffolds, 585–587 Surface mechanisms signal receiver and processor, 484, 492 Surface micromachining, 1119 Surface modification, of biomaterials interfaces, in regenerative medicine, 656 overcoating technologies, 658, 660 covalent coatings, 661–662 non-covalent coatings, 661 physiochemical modification chemical modifications, 657–660 topographical modifications, 660 roughness and topography, 660 strategies, overview of, 656–657 surface chemical patterning, 664 surfaces, biological modification, 662–664 Surface modification, of scaffolds, 585, 586, 752 Surface stability, 657 Surgery, regenerative medicine, 506–508 Surgical peripheral nerve injuries, 1270 Surgisis®, 114, 1078 SV40 large T (tumor) antigen-transformed h1RPE7, 854 SwissProt keyword, 569 Syndecan, 68, 356 Synovial cells, 774 Synthetic grafts, properties of, 446, 868, 1142 Synthetic materials, 604, 695, 735, 1108, 1200 Synthetic peptides, 617, 1041, 1187 Synthetic polymeric biomaterials, 637

Index

Synthetic polymers, 604, 1138, 1186, 1257 applications, 625 biodegradable aliphatic polyesters, 581–582 amino acid-derived polymers, 617–618 block copolymers of polyesters with PEG, 618–619 cross-linked polymer networks, 620–625 peptides, 618 poly(amino acids), 617 polyanhydrides, 582, 619 polyesters, 613–617 polyphosphazenes, 619–620 polyurethanes, 618 non-degradable, 606 Hydrolytically Stable Pus, 611–612 poly(2-hydroxyethyl methacrylate), 608 poly(ethylene), 606 poly(ethylene terephthalate), 611 poly(meth)acrylates, 607 poly(methyl methacrylate), 607–608 poly(N-isopropylacrylamide), 608–609 poly(propylene), 606 poly(styrene), 606 poly(tetrafluoroethylene), 606–607 polyacrylamides, 607 polyethers, 609–610 polysiloxanes, 610–611 Synthetic scaffolds, 594, 1005–1006 for cartilage engineering, 1186–1187 for ligament and tendon healing, 1218 for nerve repair, 1272 T cells, 9, 22, 575, 715, 717, 806, 1074 cell culture modeling, of replicative senescence, 470–471 growth and expansion potential, 469–470 immune system, 469 negative health outcomes, variety, 473–474 of replicative senescence gene expression and function, 471–472 retarding or preventing, HIV disease, 474–475 senescent cells, 472–473 see also Somatic cells T lymphocytes. See T cells T7 olig(dT), 565 Tacrolimus, 797, 798, 800, 804, 805 Tadpole stage Xenopus embryos, 457 Taurine, 664

TDGF1, 569, 571, 573 Teeth bioengineering. See Dental tissue engineering Teflon®, 604, 605, 606, 1131 Telogen, 103 Telomerase, 100, 158, 471, 475, 569, 571, 575 Telomerase repeat amplification protocol (TRAP) assay, 232 Telomerase reverse transcriptase (TERT), 158, 1005, 1027 Telomere length, 444 and somatic cloning, 158 Temporomandibular joints (TMJ), 904–905 Tenascin C, 67–68, 79, 1221 Tendon, 326–327 and ligament, 680, 890, 903–904 injuries, 1206 Tensegrity model, of cell, 486–487 Tensile testing, 1208–1210 Tensin, 71 Tensional integrity. See Tensegrity model, of cell Teratocarcinomas, 210–211 Teratomas, 21, 39, 100, 575, 955, 968, 1041 Testes anatomy, 1149 testosterone delivery systems, 1150–1151 testosterone therapy, cell encapsulation for, 1151–1152 transplantation, 1149–1150 Testicular hormone replacement, genitourinary system, 1134 Testoderm, 1151 Tet-off system, 324 Tether arm, 664 Tethered model for voltage-gated ion channels, 523 Tetraethylenepentamine (TEPA), 250 TGF-β/Activin/Nodal pathway, 127, 129 Th1 and Th2 lymphocyte, 720 Th2 helper lymphocytes, 713 Therapeutic cloning, 5, 160, 458, 1199 Therapeutic ultrasound, 1239–1240 Thermally induced phase separation method (TIPS), 588–589 Thiery–Vella loop, 1077 Thin polyurethane films, 1046 Three-dimensional biomaterial scaffolds. See Biological scaffolds Three-dimensional cell culture methods, 1028 Three-dimensional microfabrication, 698–699

1443

1444

INDEX

Thrombogenicity, 1003 of cardiac tissue, 1047–1048 Thrombopoietin (TPO), 244, 272 Thrombospondin, 69, 76, 727 Thrombus formation, 446, 705 Thy1, 260, 569 Tie-1, 249 Tie-2, 249, 292, 440, 443 Timeline for cell line development, 171–172 Tisseel®, 800, 1078 Tissue culture polystyrene (TCPS), 668 Tissue development, mechanical determinants, 480 bioreactors, 490–491 cell, as signal receiver and processor, 483–489 cell fates, 489–490 mechanical forces, experienced by tissues, 481–483 practical examples, 491–495 Tissue donor source, for determined stem cells, 354 Tissue-engineered blood vessels (TEBV) functions, 978 process, 984, 985–995 studies, 984–985 vascular physiology, 980–981 Tissue-engineered construct, 539, 709 Tissue engineered heart valve (TEHV), 1380–1381 Tissue engineered medical products (TEMPS), 1379 Tissue-engineered vascular grafts cell seeding and preconditioning, 992–995 cell source, 992 large diameter TEBV, 985 scaffolds, 984, 985–992 small diameter TEBV, 984–985 Tissue engineering, 38, 536, 580, 734, 888–889, 1127 biomaterials for, 636 of bone, 745, 897, 972 BMPs, 747 combination of, growth/differentiation factors, 747 delivery of, growth/differentiation factors, 747–748 ESCs, 746 MSCs, 746 principles, 745 regional gene therapy, 747 scaffolds, 748–752 cartilage production, 900–901 for cartilage repair, 1176

for casualty care, 1322 cardiothoracic and vascular tissue engineering, 1323 military needs, 1324 musculoskeletal tissue engineering, 1322–1323 wound healing, 1323–1324 and cell-based therapies, 888 corneal advancement in, 1068 EPC participation, 434–435 of fetus, 1132–1134 of functional corneal equivalents, 1066 of heart, 970–971 in humans, ethical issues, 1346 interfaces, controlling proteins, with precision, 737–740 of ligament and tendon injuries, 1206 of liver, 358–359 biological issues, for, 359–361 of lung, 971 of muscle, 972 non-specific protein absorption, 735–736 of peripheral blood stem cells, 446–447 of reproductive system, 1138 in female, 1152–1159 in male, 1139–1152 of skin, 971 strategies, 896 of teeth, 1286 of urologic structures, 1127–1132 Tissue growth, 519 effects of stress on, 512 Tissue healing regulation, 724–725 Tissue inhibitor of metalloproteinases (TIMPs), 79 Tissue perfusion, 756, 761 Tissue regeneration, 78, 322 and BMSCs, 272 EPC participation, 447–448 hematopoietic reconstitution, 240–241 of kidney, 1106–1108 Tissue repair effects of stress on, 519–520 factors, for fracture repair and bone regeneration, 1167–1168 Tissue rules, 1373 Tissue template, 1066 Tissue therapy in central nervous system, 1248 for skeletal muscle injury, 1232

Index

Titanium, 658–660, 1291 Title 21 Code of Federal Regulations (CFR), 1368, 1371 1270 and 1271, 1311, 1313, 1373 3.2(e), 1370 800–1200, 1309 820, 1310 Topographical modifications, of surface modification, 660 Totipotent stem cells, 30, 126, 301, 345, 349 Trade name, 1394 Trade secrets, 1395 Trademark, 1394 Trans Cyte™, 1299 Transaction models, of SCI, 1250 Transcription factor Otx1, 852 Otx2, 419, 852 Transcriptional networks, 131 Transcriptional signature, 562 Transcriptome, 294, 567 analysis, 260 Transdetermination, 137–138 see also Plasticity; Trandsdifferentiation Transdifferentiation, 30, 137, 138, 346, 407–408 in hES-derived RPE cultures, 859 of retina, 425–430 of RPE in culture, 855 see also Plasticity; Transdetermination Transfection, of transgenes, 172 chemical transfection, 194 physical transfection, 194 Transferase proteins, 132 Transforming growth factor beta (TGF-β), 51, 58, 129, 501, 589, 667, 725, 872, 889, 1180, 1231, 1292 and fibrosis, 1241 pathway, 54 TGF-β1, 76, 81, 83, 104, 113–114, 191, 624, 646, 679, 726, 890, 1022, 1215, 1234, 1235 TGF-β2, 81, 1182, 1216 TGF-β3, 81, 116, 525, 730, 895 Transgene copy number, 172, 183, 184 Transgenic production, of therapeutic proteins, 168 caprine SCNT, 173–175 nuclear transfer parameters, 175–180 cloned cows, expressing human albumin, 182–184 cloned goats, management of, 180–182 somatic cell nuclear transfer, 169, 170

DNA construct development, 170–171 cell line development, 171–173 transgenic animals, generation of pronuclear MI, 169 Transglutaminases, 77, 727 Transmission electron microscopy, 77, 1218 Traumatic care and ensuing shock, 756 Traumatic peripheral nerve injuries, 1270 Triblock copolymers, 609 Trichostatin, 572, 574 Tri-fluoroacetic anhydrides, 658 Tripeptide RGD, 68, 549, 586, 671, 694 Trolley-mounted liver support system, 1094 Trophectoderm, 131, 137, 211, 258 Tropomyosin, 233, 1044, 1046 Trypsin, 618, 857, 858, 1115 treatment, 150 Trypsinization, 177 Tubulin, 488 Tubulin βIII, 857, 858, 859 Tumor necrosis factor-alpha (TNFα), 77, 272, 471, 708, 1008, 1023, 1238, 1260 Turkeyís Biweight algorithm, 566, 567 Twist, 57, 58 Type 1 diabetes mellitus (T1DM), 247, 398, 408, 794, 795, 805 Type 2 diabetes mellitus (T2DM), 398, 794 Type I collagen, 351, 645, 727, 1166–1167, 1207 Type V collagen, 1207, 1212 L-Tyrosine, 617 Tyrosine-based polycarbonates, 617 Tyrosine-derived polyacrylates, 617 Tyrosine-derived polyethers, 617 Tyrosine kinase receptors, 70, 272, 292 Tyrosinemia Type I, 919, 920 Ulex europaeus I (UAE-I)-coated Dynabeads, 1145 Ultrafiltration membrane development, of kidney, 1121–1123 Ultrasonic pulser, 483 Ultraviolet versus polarized light enucleation, 177–178 Umbilical cord blood (UCB) stem cells, 8, 34, 228, 241, 246, 322, 346, 746, 762, 826, 969, 973 see also Cord blood Umbilical cord matrix stem (UCMS) cells, 8 Unipotent cell, 127 Unipotent parenchymal progenitors, 357

1445

1446

INDEX

United Kingdom, 369 stem cell research policy, 1357–1358 United States, 2, 28, 372, 373–374, 812, 852, 868, 888, 946, 978, 979, 1000, 1038, 1086, 1114, 1176, 1198, 1248, 1270, 1337, 1355, 1387, 1393, 1397 patent law, current issues in intellectual property rights copyrights, 1395 importance, 1396–1398 patents, 1386–1394 service marks, 1394 trade marks, 1394 trade names, 1394 trade secrets, 1395 valuation, 1398–1401 stem cell research policy commercialization and access, 1362–1363 current law and policy, 1355–1357 egg donors compensation, 1361–1362 international comparisons, 1357–1358 national academies guidelines, 1359–1361 sources, stem cells, 1354–1355 state and private funding, 1358–1359 United States Codes (USC), 1368, 1386, 1391 United States Food and Drug Administration, 16, 239, 345, 401, 456, 506, 581, 637, 696, 837, 872, 900, 948, 1118 Unrestricted somatic stem cells (USSCs), 8, 230, 241, 260 Ureter, 1128, 1131 Urethra, 1127 anatomy, 1139–1140 cell growth, 1140 acellular collagen matrix preparation, 1143 bladder smooth muscle cell culture, 1141–1143 urothelial cell culture, 1140–1141 Urinary bladder submucosa (UBS), 1075 Urokinase plasminogen activator (uPA), 79, 102 Urologic structures, in tissue engineering bladder cell transplantation, 1129 matrices, 1129 seromuscular grafts, 1128–1129 tissue expansion, 1128 genital tissues corporal smooth muscle, 1130 engineered penile prostheses, 1130–1131 female genital tissues, 1131 renal structures ex vivo, 1132 in vivo, 1132

ureter, 1131 urethra, 1127 Urothelial cell culture, 1140–1141 Uruguay Round Act, 1393 US Defense Advanced Research Project Agency, 1328 US Patent and Trade mark Office (US PTO), 1386, 1387, 1388, 1394 US patent law, current issues, 1386 intellectual property rights, 1386 copyrights, 1395 importance of, 1396–1398 novelty, 1389–1390 obviousness, 1391 patent rights, 1392 patent term extension, 1392–1394 patents, 1386–1389 service marks, 1394 trade names, 1394 trade secrets, 1395 trademarks, 1394 written description, 1391–1392 patentability, 1389 patentable subject matter, 1387, 1388, 1389 PTO guidelines, 1387–1388 valuation of, 1398–1401 “real world” context of use, 1388 US Renal Data System (USRDS), 1114 US state law, on stem cell research, 374 Uterus anatomy, 1155–1156 uterine tissue reconstruction, 1156–1157 Utilitarianism, 1336 Vaccine production, 366–367 Vagina, 1152–1153 anatomy, 1153 tissue engineering, 1153–1154 Valvular replacements, 39, 40 Variant Creutzfeldt–Jakob Disease (vCJD), 1373 Vascular assembly analysis and assessment animal models, 1031–1032 in vitro techniques, 1030–1031 quantitative techniques, 1032–1033 biomaterials genes and proteins, presentation, 1024–1026 cell therapies genetically modified, 1027

Index

mature, 1026–1027 progenitor and stem, 1027 engineered and natural tissues, 1020 genes, 1023–1024 mechanisms angiogenesis, 1020–1021 arteriogenesis, 1022–1023 patterning/structure, 1022 remodeling/stabilization, 1021–1022 vasculogenesis, 1020–1021 natural signals, molecular modification, 1024 proteins, 1023 vascularization in vitro photolithographic techniques, 1028–1029 three-dimensional cell culture methods, 1028 vascularization in vivo, 1029–1030 Vascular cell adhesion molecules (VCAMs), 701 Vascular differentiation. See Endothelial cells Vascular disease, 446–447 dialysis vascular access, need for, 979–980 prevalence and impact, 978–979 Vascular endothelial-cadherin, 249 Vascular endothelial growth factor (VEGF), 67, 249, 271, 275–276, 285, 361, 439, 441, 691, 728, 748, 804, 814, 961, 974, 1022, 1042, 1165, 1200, 1239 VEGF-A, 104, 361 VEGFR-1, 249 Vascular integrity injury versus repair, 815–816 Vascular niche, 272 Vascular permeability factor (VPF), 915 Vascular smooth muscle cell (VSMC), 679, 978, 1048, 1156 Vascular therapy, 447 Vascular tissue engineering, 1323 Vascularization of cardiac tissue, 1049 genitourinary system, 1127 in vitro techniques photolithographic techniques, 1028–1029 three-dimensional cell culture methods, 1028 in vivo techniques, 1029–1030 endothelial cell, 491–492 smooth muscle cell, 491–492 Vasculogenesis, 441, 728, 1020–1021 VE-cadherin, 198, 199, 441, 1004 Ventral diencephalons, 418, 419 Ventricular remodeling, 22, 272, 1039

Vertically integrated regenerative therapies, 1329 Very small embryonic-like cells (VSELs), 260, 262 Vesicoureteral reflux (VUR), 1133 Vesicular stomatitis virus glycoprotein (VSV-G), 195 Vietnam War, 1324 Vimentin, 57, 993, 1045 Vinculin, 667 Viral transduction, 195–198 lentiviral vector, 196–198 retroviral vector, 195–196 Viral vectors, 198, 747, 899 Visceral endoderm, 403, 405 Visceral mesoderm, 331 Vitamin D3-upregulated protein 1 (VDUP1), 971 Vitelliform macular dystrophy gene (VMD2), 853 Vitronectin, 484, 666, 671 Voltage-gated ion channels, 522 von Kossa, 232, 564 von Willebrand factor, 725–726 Wavy walled bioreactor, 490, 494–495 Waxman-Hatch Act, 1386, 1401 Western blot analysis, 234, 564, 1131, 1202 Wettability, 670, 752 Wilson’s disease, 919 Wiskott–Aldrich syndrome, 192 Wnt, 130 β-catenin, 421 catenin signaling pathway, 1073 inhibitor, 130, 142, 325 pathway, 55, 127, 214, 288 signaling, 55, 58, 130, 214, 1182 signaling pathway, 288, 894, 895 Wnt2b, 429 Wnt3a, 106, 130 Wnt4, 55, 421 Wnt11, 421 Wolff ’s law, 492, 512 Women’s ischemia syndrome evaluation (WISE) study, 817 Woodchuck hepatitis virus post-transcriptional regulatory element (WPRE), 197 World Intellectual Property Organization (WIPO), 1394 Wound contraction, 729–730 Wound healing, 20, 77, 724, 815, 1301, 1323–1324, 1326 blood–material interactions, 705

1447

1448

INDEX

Wound healing, (continued) essential elements fetal healing, 78–80, 730 fibroplasia, 726–728 inflammation, 725–726 neovascularization, 728–729 re-epithelialization, 729 regulation, 724–725 wound contraction, 729–730 inflammatory response, initiation of, 705–706 problems of, 1302–1303 provisional matrix formation, 706–707 spraying cells, into wound, 1303–1304 and temporal sequence of inflammation, 707 X-chromosome inactivation, after somatic cloning, 157–158

Xenogeneic collagen, 595 Xenogeneic components, 31, 191, 719 Xenopus eye formation, 421 Xenotransplantation, 39, 803, 1375–1376 XIAP (X-chromosome linked inhibitor of apoptosis), 157, 804 Xist (X-chromosome inactive specific transcript), 154 YIGSR synthetic peptide, 691, 692, 697 ZD6474, 444 Zeolite, 1325 Zinc-finger transcription factors, 5 Zyderm I®, 595 Zyderm II®, 595 Zygote, 126, 345 Zyplast®, 595