Tissue Engineering in Regenerative Medicine (Stem Cell Biology and Regenerative Medicine)

Stem Cell Biology and Regenerative Medicine

Series Editor Kursad Turksen, Ph.D. [email protected]

For further volumes: http://www.springer.com/series/7896

Harold S. Bernstein Editor

Tissue Engineering in Regenerative Medicine

Editor Harold S. Bernstein Cardiovascular Research Institute Eli and Edythe Broad Center of Regeneration Medicine and Stem Cell Research, Department of Pediatrics University of California San Francisco San Francisco, CA, USA [email protected]

ISBN 978-1-61779-321-9 e-ISBN 978-1-61779-322-6 DOI 10.1007/978-1-61779-322-6 Springer New York Dordrecht Heidelberg London Library of Congress Control Number: 2011934681 © Springer Science+Business Media, LLC 2011 All rights reserved. This work may not be translated or copied in whole or in part without the written permission of the publisher (Humana Press, c/o Springer Science+Business Media, LLC, 233 Spring Street, New York, NY 10013, USA), except for brief excerpts in connection with reviews or scholarly analysis. Use in connection with any form of information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed is forbidden. The use in this publication of trade names, trademarks, service marks, and similar terms, even if they are not identified as such, is not to be taken as an expression of opinion as to whether or not they are subject to proprietary rights. Printed on acid-free paper Humana Press is part of Springer Science+Business Media (www.springer.com)

This book is dedicated to my father, Wallace Carl Bernstein (1923–2010), who taught me to ask questions.

Preface

Over the past decade, significant advances in the fields of stem cell biology, bioengineering, and animal models have converged on the discipline of regenerative medicine. Significant progress has been made leading from preclinical studies through phase 3 clinical trials for some therapies. This volume provides a state-of-the-art report on tissue engineering toward the goals of tissue and organ restoration and regeneration. Examples from different organ systems illustrate progress with growth factors to assist in tissue remodeling; the capacity of stem cells for restoring damaged tissues; novel synthetic biomaterials to facilitate cell therapy; transplantable tissue patches that preserve three-dimensional structure; synthetic organs generated in culture; aspects of the immune response to transplanted cells and materials; and suitable animal models for nonhuman clinical trials. Tissue regeneration, and even stem cell therapy, is not a new concept. As discussed in the cautionary first chapter, efforts toward bone and marrow transplantation have been underway for almost half a century. Steady progress has been made in understanding the criteria for successful cell transplantation, and developing a robust structure for clinical oversight. More recently, pluripotent stem cells, with their capacity for self-renewal and tissue-specific differentiation, have become a prime candidate for tissue engineering and regenerative therapies. More than 100 clinical trials have examined the use of mesenchymal stem/stromal cells. Biochemical and mechanical interactions between the extracellular matrix and cell surface receptors, as well as physical interactions between cells, are now recognized as essential for stem cell self-renewal and differentiation. New technologies for scaffold engineering and fabrication have taken advantage of these observations, and hold promise for repairing tissues requiring a highly specialized niche, such as skeletal muscle. These discoveries have led to clinical trials with bioengineered vascular conduits in children with congenital heart disease, complete hollow organs, and complex organs such as bioartificial livers. An evolving understanding of innate and adaptive immune responses, including the foreign body response, has led to novel approaches to modulating the immune system that facilitate tissue repair. Finally, the development of small animal models for discovery, and large animal models for studies of safety and efficacy, has propelled the field of tissue engineering toward the clinic. vii

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The chapters of this book are organized into six sections: Stem Cells, Biomaterials and the Extracellular Environment, Engineered Tissue, Synthetic Organs, Immune Response, and Animal Models. Each section is intended to build upon information presented in the previous chapters, and set the stage for subsequent sections. Throughout the chapters, the reader will observe a common theme of basic discovery informing clinical translation, and clinical studies in animals and humans guiding subsequent experiments at the bench. I thank the members of my laboratory for their helpful discussion, and my colleagues in Pediatric Cardiology for their support – we all strive to improve the lives of our patients. I appreciate always the encouragement I receive from Tricia Foster, Nathaniel Bernstein, and Katharine Bernstein. I am grateful to the 54 colleagues who have contributed their expertise to this project. We hope that this first edition of Tissue Engineering and Regenerative Medicine will serve as an introduction and guide for students of the field at all levels. San Francisco, CA

Harold S. Bernstein

Contents

Part I  Stem Cells   1 Hematopoietic Stem Cell Transplantation: Reflections on Yesterday and Thoughts for Tomorrow....................... Andrew D. Leavitt

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  2 Human Embryonic Stem Cells in Regenerative Medicine.................. Odessa Yabut and Harold S. Bernstein

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  3 Current Status of Induced Pluripotent Stem Cells.............................. Thach-Vu Ho, Grace Asuelime, Wendong Li, and Yanhong Shi

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  4 Mesenchymal Stromal Cells: Latest Advances.................................... Sowmya Viswanathan and Armand Keating

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Part II  Biomaterials and the Extracellular Environment   5 The Role of Mechanical Forces in Guiding Tissue Differentiation............................................................................. Sean P. Sheehy and Kevin Kit Parker   6 Synthetic Multi-level Matrices for Bone Regeneration....................... Nicholas R. Boyd, Richard L. Boyd, George P. Simon, and David R. Nisbet

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  7 Directing Cell Fate Through Biomaterial Microenvironments........... 123 Kelly Clause, Jonathan Lam, Tatiana Segura, and Thomas H. Barker Part III  Engineered Tissue   8 Basic Considerations with Cell Sheets.................................................. 143 Masayuki Yamato and Sebastian Sjöqvist

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  9 Myocardial Repair and Restoration...................................................... 161 Sharon S.Y. Wong and Harold S. Bernstein 10 Skeletal Muscle Engineering: The Need for a Suitable Niche............ 197 Frédéric Trensz, Anthony Scimè, and Guillaume Grenier 11 Restoring Blood Vessels.......................................................................... 211 Narutoshi Hibino, Christopher Breuer, and Toshiharu Shinoka 12 Engineering Functional Bone Grafts.................................................... 221 Sarindr Bhumiratana and Gordana Vunjak-Novakovic 13 Engineering Functional Cartilage Grafts............................................. 237 Andrea R. Tan and Clark T. Hung 14 Adult Stem Cells and Regeneration of Adipose Tissue....................... 251 Daniel A. Hägg, Bhranti Shah, and Jeremy J. Mao Part IV  Synthetic Organs 15 Hollow Organ Engineering.................................................................... 273 Anthony Atala 16 Engineering Complex Synthetic Organs............................................... 297 Joan E. Nichols, Jean A. Niles, and Joaquin Cortiella 17 Liver Regeneration and Tissue Engineering........................................ 315 Ji Bao, James Fisher, and Scott L. Nyberg Part V  Immune Response 18 Immune Modulation for Stem Cell Therapy........................................ 335 Gaetano Faleo and Qizhi Tang 19 Regenerative Medicine and the Foreign Body Response..................... 353 Kerry A. Daly, Bryan N. Brown, and Stephen F. Badylak Part VI  Animal Models 20 Small Animal Models of Tissue Regeneration...................................... 379 Fernando A. Fierro, J. Tomas Egana, Chrisoula A. Toupadakis, Claire Yellowley, Hans-Günther Machens, and Jan A. Nolta 21 Use of Large Animal and Nonhuman Primate Models for Cell Therapy and Tissue Engineering............................................. 393 Alice F. Tarantal and Karina H. Nakayama About the Editor............................................................................................. 415 Index................................................................................................................. 417

Contributors

Grace Asuelime  Department of Neurosciences, Center for Gene Expression and Drug Discovery, Beckman Research Institute of City of Hope, Duarte, CA, USA Anthony Atala  Wake Forest Institute for Regenerative Medicine, Wake Forest University School of Medicine, 5th Floor, Watlington Hall, Medical Center Boulevard, Winston-Salem, NC 27157, USA Stephen F. Badylak  McGowan Institute for Regenerative Medicine, University of Pittsburgh, 450 Technology Drive, Suite 300, Pittsburgh, PA, USA Ji Bao  Mayo Clinic, Rochester, MN, USA Thomas H. Barker  The Wallace H. Coulter Department of Biomedical Engineering at Georgia Tech, Emory University, 313 Ferst Drive, Atlanta, GA 30332, USA Harold S. Bernstein  Cardiovascular Research Institute, University of California San Francisco, Eli and Edythe Broad Center of Regeneration Medicine and Stem Cell Research, Department of Pediatrics, University of California San Francisco, San Francisco, CA 94143-1346, USA Sarindr Bhumiratana  Department of Biomedical Engineering, Columbia University, New York, NY, USA Nicholas R. Boyd  Department of Materials Engineering, Monash University, Clayton, VIC, Australia Richard L. Boyd  Monash Immunology and Stem Cell Laboratories, Monash University, Clayton, VIC, Australia Christopher Breuer  Department of Cardiac Surgery, Interdepartmental Program in Vascular Biology and Therapeutics, Yale University School of Medicine, New Haven, CT, USA

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Bryan N. Brown  Department of Clinical Sciences, Cornell University, Ithaca, NY, USA Kelly Clause  The Wallace H. Coulter Department of Biomedical Engineering at Georgia Tech, Emory University, Atlanta, GA, USA Joaquin Cortiella  Laboratory of Regenerative and Nano-Medicine, Department of Anesthesiology, University of Texas Medical Branch, Galveston, TX, USA Kerry A. Daly  McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, USA J. Tomas Egana  Department of Plastic and Hand Surgery, University Hospital Rechts der Isar, Technical University of Munich, Munich, Germany Gaetano Faleo  Department of Surgery, University of California, San Francisco, CA, USA Fernando A. Fierro  Department of Internal Medicine, Stem Cell Program and Institute for Regenerative Cures, University of California, Davis, CA, USA James Fisher  Mayo Clinic, Rochester, MN, USA Guillaume Grenier  Étienne-Lebel Clinical Research Center, Department of Orthopedic Surgery, Université de Sherbrooke, 3001–12th Avenue North, J1H 5N4, Sherbrooke, QC, Canada Daniel A. Hägg  Tissue Engineering and Regenerative Medicine Laboratory, Columbia University Medical Center, New York, NY, USA Narutoshi Hibino  Department of Cardiac Surgery, Interdepartmental Program in Vascular Biology and Therapeutics, Yale University School of Medicine, New Haven, CT, USA Thach-Vu Ho  Department of Neurosciences, Center for Gene Expression and Drug Discovery, Beckman Research Institute of City of Hope, Duarte, CA, USA Clark T. Hung  Department of Biomedical Engineering, Columbia University, 1210 Amsterdam Avenue, Engineering Terrace 351, New York, NY 10027, USA Armand Keating  Cell Therapy Program, Princess Margaret Hospital, University Health Network, 610 University Avenue, Suite 5-303, Toronto, ON M5G 2M9, Canada Department of Medicine, University of Toronto, Toronto, ON, Canada Jonathan Lam  Department of Biomedical Engineering, University of California, Los Angeles, CA, USA Andrew D. Leavitt  Departments of Laboratory Medicine and Medicine, UCSF Adult Blood and Marrow Transplant Laboratory, University of California, 513 Parnassus Avenue, Box 0100, San Francisco, CA 94143-0100, USA

Contributors

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Wendong Li  Department of Neurosciences, Center for Gene Expression and Drug Discovery, Beckman Research Institute of City of Hope, Duarte, CA, USA Hans-Günther Machens  Department of Plastic and Hand Surgery, University Hospital Rechts der Isar, Technical University of Munich, Munich, Germany Jeremy J. Mao  Tissue Engineering and Regenerative Medicine Laboratory, Columbia University Medical Center, 630 W. 168 St. – PH7 East, New York, NY 10032, USA Karina H. Nakayama  Departments of Pediatrics, Cell Biology and Human Anatomy, School of Medicine, UC Davis Clinical and Translational Science Center, California ­National Primate Research Center, University of California, Davis, CA, USA Joan E. Nichols  Laboratory of Regenerative and Nano-Medicine, Departments of Internal Medicine and Infectious Diseases, University of Texas Medical Branch, Galveston, TX, USA Jean A. Niles  Laboratory of Regenerative and Nano-Medicine, University of Texas Medical Branch, Galveston, TX, USA David R. Nisbet  Research School of Engineering, ANU College of Engineering and Computer Science, The Australian National University, Ian Ross Building 31, North Road, Acton, Canberra, ACT 0200, Australia Jan A. Nolta  Department of Internal Medicine, Stem Cell Program and Institute for Regenerative Cures, University of California, Davis, 2921 Stockton Blvd., Room 1300 , Sacramento, CA 95817, USA Scott L. Nyberg  Mayo Clinic, 200 First Street SW, Rochester, MN 55905, USA Kevin Kit Parker  Disease Biophysics Group, Wyss Institute for Biologically Inspired Engineering, School of Engineering and Applied Sciences, Harvard University, Pierce Hall, Room 321, 29 Oxford St., Cambridge, MA 02138, USA Anthony Scimè  Muscle Health Research Centre, York University, Toronto, ON, Canada Tatiana Segura  Department of Chemical and Biomolecular Engineering, University of California, Los Angeles, CA, USA Bhranti Shah  Tissue Engineering and Regenerative Medicine Laboratory, Columbia University Medical Center, New York, NY, USA Sean P. Sheehy  Disease Biophysics Group, Wyss Institute for Biologically Inspired Engineering, School of Engineering and Applied Sciences, Harvard University, Cambridge, MA, USA

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Contributors

Yanhong Shi  Department of Neurosciences, Center for Gene Expression and Drug Discovery, Beckman Research Institute of City of Hope, 1500 E. Duarte Rd, Duarte, CA 91010, USA Toshiharu Shinoka  Department of Cardiac Surgery, Interdepartmental Program in Vascular Biology and Therapeutics, Yale University School of Medicine, 333 Cedar Street, Boardman 204, PO Box 208039, New Haven, CT 06520-8039, USA George P. Simon  Department of Materials Engineering, Monash University, Clayton, VIC, Australia Sebastian Sjöqvist  Institute of Advanced Biomedical Engineering and Science, Tokyo Women’s Medical University, Tokyo, Japan Andrea R. Tan  Department of Biomedical Engineering, Columbia University, New York, NY, USA Qizhi Tang  Department of Surgery, University of California, Box 0780, 513 Parnassus Avenue, San Francisco, CA 94143-0780, USA Alice F. Tarantal  Departments of Pediatrics, Cell Biology and Human Anatomy, School of Medicine, UC Davis Clinical and Translational Science Center, California National Primate Research Center, University of California, Pedrick and Hutchison Roads, Davis, CA 95616-8542, USA Chrisoula A. Toupadakis  Department of Anatomy, Physiology and Cell Biology, School of Veterinary Medicine, University of California, Davis, CA, USA Frédéric Trensz  Étienne-Lebel Clinical Research Center, Université de Sherbrooke, Sherbrooke, QC, Canada Sowmya Viswanathan  Cell Therapy Program, Princess Margaret Hospital, University Health Network, Toronto, ON, Canada Gordana Vunjak-Novakovic  Department of Biomedical Engineering, Columbia University, 1210 Amsterdam Avenue, Engineering Terrace 351, New York, NY 10027, USA Sharon S.Y. Wong  Cardiovascular Research Institute, University of California San Francisco, San Francisco, CA, USA Odessa Yabut  Cardiovascular Research Institute, University of California, San Francisco, CA, USA Masayuki Yamato  Institute of Advanced Biomedical Engineering and Science, Tokyo Women’s Medical University, 8-1 Kawada-cho, Shinjuku-ku, Tokyo 162-8666, Japan Claire Yellowley  Department of Anatomy, Physiology and Cell Biology, School of Veterinary Medicine, University of California, Davis, CA, USA

Part I

Stem Cells

Chapter 1

Hematopoietic Stem Cell Transplantation: Reflections on Yesterday and Thoughts for Tomorrow Andrew D. Leavitt

Abstract  Biomedical science is entering a new era with exciting prospects for using cellular therapy to treat a wide spectrum of human diseases from nerve injury to diabetes, myocardial infarction, and more. Hematopoietic stem cell (HSC) transplantation has been used to treat patients for nearly half a century. The experiences and lessons learned over those 50 years are both informative and encouraging. This chapter distills the history of HSC transplantation to provide an orientation to the past that can be used to more wisely navigate the future of cell therapy. The details presented help the reader appreciate that developing novel cell therapy can be a struggle and that chance will likely continue to play a role in future success. However, it also becomes apparent that attention to fundamental details, such as choice of cell type or types, where to obtain the cells, how to handle and process the cells, how to prepare and select patients, how to evaluate success and failure, and how to organize the biomedical community to serve the good of patients, are all critical for new cell therapy to become a reality.

Abbreviations BMT GVHD GVL HLA HSCs

Bone marrow transplantation Graft-versus-host disease Graft-versus-leukemia Human lymphocyte antigen Hematopoietic stem cells

A.D. Leavitt (*) Departments of Laboratory Medicine and Medicine, UCSF Adult Blood and Marrow Transplant Laboratory, University of California, 513 Parnassus Avenue, Box 0100, San Francisco, CA 94143-0100, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_1, © Springer Science+Business Media, LLC 2011

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PBSCs Peripheral blood stem cells UCB Umbilical cord blood

1.1 Introduction HSCs are the most studied and well-understood of all adult stem cells, and they provide a model system and paradigm for the more global understanding of stem cell biology [1]. HSCs have also been used clinically for nearly 50  years, with over 55,000 HSC transplants performed around the world in 2009 alone [2]. HSCs and their clinical application, therefore, provide an excellent reference point for discussing the future of stem cell therapy, be it the use of embryonic stem cells and their derivatives or the direct use of tissue-specific adult stem cells. This chapter presents a brief history of HSC transplantation to give perspective and to help inform and orient the reader to issues that will likely be faced as biomedical scientists begin developing tomorrow’s stem cell therapies. Accounts of the history of HSC transplantation have been summarized by others, including a personal account by E. Donnell Thomas who shared the 1990 Nobel Prize in Medicine for his pioneering role in the development of BMT [3, 4].

1.2 Radiation: A Double-Edged Sword Marie Curie (born Maria Sklodowska) shared the 1903 Nobel Prize in physics with Pierre Curie and Henri Becquerel “in recognition of the extraordinary services they have rendered by their joint researches on the radiation phenomena discovered by Professor Henri Becquerel.” She also won the 1911 Nobel Prize in Chemistry “in recognition of her services to the advancement of chemistry by the discovery of the elements radium and polonium, by the isolation of radium and the study of the nature and compounds of this remarkable element.” Tragically, she died on July 4, 1934, from marrow toxicity, reported in various sources as aplastic anemia and/or leukemia, but almost certainly secondary to the chronic radiation exposure she received during her early pioneering studies related to naturally radioactive substances. The bone marrow toxicity of ionizing radiation was appreciated only after much of her initial exposure, and interestingly the field of clinical marrow transplantation relied for decades on the use of ionizing radiation as a preparative regimen to both eradicate underlying malignant disease and to immunosuppress the recipient to facilitate marrow engraftment and HSC repopulation. The highly deleterious effects of radiation on bone marrow were appreciated well before World War II [5], but development and use of the atomic bomb in the 1940s highlighted the marrow toxicity of radium, uranium, and other sources of ionizing radiation. Classified government research to develop treatments for bone marrow toxicity due to atomic bomb radiation exposure was performed in the 1940s under the auspices of the Atomic Energy Commission, but it was not published until

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1950 [6]. Those studies sought “to determine what benefits, if any, may be derived from the transplantation of normal bone marrow in animals that have suffered damage to their bone marrow as a result of single dose roentgen irradiation.” The studies failed to achieve their goal of allogeneic engraftment or to demonstrate any clinically useful effect of marrow transplantation. However, failure was most likely secondary to inadequate radioablation of the recipient animal’s immune system required to achieve engraftment. So, even though the studies failed to achieve their goal [6], they highlighted one of the critical aspects of HSC transplantation – the host is not naturally receptive to foreign cells and the host’s immune system needs to be suppressed to overcome this barrier to cellular therapy. This critical fact is important to consider when developing any future form of allogeneic stem cell therapy.

1.3 Bone Marrow Transplantation: It Is the Cells In 1949, independent investigators reported that lead shielding of the spleen protected mice from the mortality of total body irradiation [7]. Interestingly, it was thought that the beneficial effect was humorally mediated. Even after a 1951 report demonstrated that intravenous or intraperitoneal injections of bone marrow cells protect mice and guinea pigs from the mortality of total body radiation [8], the humoral theory remained the prevailing theory to explain radioprotection. It required an innovative experiment reported in 1955 to begin to convince the research community that radioprotection stemmed from the bone marrow cells themselves engrafting into the recipient [9]. In brief, the investigators knew that skin grafts would not survive if performed between H2-incompatible mice, but the authors showed that skin grafts could survive across H2-incompatible strains if the recipient was first transplanted with marrow from the skin donor [9]. Moreover, skin graft survival required that the irradiated recipient mouse receive marrow from the same mouse strain that provided the skin graft. These findings, as the authors concluded, “are consistent with the cellular repopulation theory of radiation protection.” In 1956, using the then novel technique of genetically traceable donor marrow cells, it was convincingly shown that the radioprotective effect of BMT correlated with engraftment of donor marrow cells in the recipient [10]. The essential role of the marrow cells in radioprotection had finally been established, as had the fact that allogeneic marrow transplantation could work.

1.4 A Rough Clinical Start: Patients Are Always a Bigger Challenge than Mice With the animal transplant data in hand and knowing that radiation could kill leukemic cells, it was only natural for investigators to try to connect these two observations for therapeutic benefit. A 1956 report demonstrated that radiation could be used to

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eradicate leukemia in mice and that bone marrow transplant could rescue the host from the marrow-damaging effects of the radiation treatment [11]. One year later, in 1957, Thomas et al. published the first report of infusing allogeneic marrow into humans when he described his experience with six patients – three with hematologic malignancies (chronic myelogenous leukemia, multiple myeloma, and chronic lymphocytic leukemia), one with ovarian carcinoma, one with metastatic cancer of uncertain origin, and one who had suffered a massive central nervous system bleed [12]. The five patients with malignancies had each received chemotherapy and/or radiation therapy shortly before the marrow infusions. This initial report clearly focused on evaluating the safety and toxicity of the marrow cell infusions and not their therapeutic benefit [12]. There were no deaths attributed to the infused cells, and great effort was taken to assess for pulmonary emboli, which were not found to be a problem clinically or when evaluated at postmortem exam. One case suggested transient engraftment based on circulating blood cell analysis, but no long-term engraftment was demonstrated. The major conclusion was that anticoagulated suspensions of allogeneic marrow cells, strained through fine mesh to remove particulate matter, can be safely given to human recipients, as had been previously demonstrated in animals [13]. In addition to demonstrating relative safety (i.e., no major obvious untoward effects) in a very small number of patients, the authors raised important fundamental issues that are important to consider when developing any type of cell therapy in the future. They discussed the need to establish a clinically relevant cell dose, to develop a preparative regimen to treat the recipient so that their immune system does not reject the allograft, and to define a detailed monitoring system that allows for accurate assessment of toxicity and benefit. The same group reported in 1959 the successful, albeit temporary, eradication of acute lymphocytic leukemia in a patient treated with total body irradiation (Co60) followed by allogeneic BMT from an identical twin [14]. While the patient relapsed 12 weeks later, the case demonstrated that lethal radiation followed by BMT could achieve a remission, even in advanced disease, and it highlighted the importance of immunologically matched donors for efficient engraftment [14]. The authors concluded that transplants of syngeneic marrow are readily achieved in humans, that 1,000 rad of whole body radiation administered properly does not produce troublesome acute radiation sickness in humans, and that whole body irradiation at the 1,000 rad level produces a remission but not a cure of leukemia when followed by infusion of syngeneic marrow. Chemotherapy (cyclophosphamide) was soon added to total body irradiation to help eradicate the underlying disease when employing allogeneic BMT to treat patients with acute leukemia, a preparative regimen that remained in use for several decades. Reports of allogeneic BMT rose steadily over the next few years, with over 60 such transplants reported in 1962. However, enthusiasm rapidly declined as toxicity was clear and success was hard to find; only a few transplants were reported annually through the late 1960s [15]. A 1970 review of all 203 reported allogeneic transplants through 1968 highlighted the dismal state of the field, with few if any true successes. In fact, 125 of the 203 recipients did not even demonstrate evidence of

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engraftment, including 66 of 73 patients with aplastic anemia [15]. Interestingly, the other seven aplastic anemia patients received allogeneic marrow from a syngeneic twin, five of whom had clinical recovery from their disease. This subset of patients provided hope for BMT as a clinical intervention, and the outcome with the identical twins reemphasized the critical importance of immunologic match for a successful engraftment of donor bone marrow. It also highlighted the difference between treating a disease that has a dominant phenotype that is likely to recur, such as leukemia, versus one with a recessive phenotype, such as aplastic anemia. While the late 1950s through the early 1970s was not a good time for clinical success within the BMT field, significant headway was made in critical areas of transplant immunology through the use of animal studies. The advances grew out of studies in the early 1950s that actively developed immune tolerance in young mice [16]. By the mid-1960s, runt disease in mice [17, 18], which is essentially what we call GVHD in the human transplant setting, was becoming well-understood, at least from the perspective of factors related to its development [19, 20]. For example, it was not associated with the injection of syngeneic cells but required antigenic differences between donor and host, and the more pronounced the differences, the more severe the disease. Moreover, persistence of the allogeneic cells was required for persistent disease, and injection of presensitized cells could worsen the problem. Also, one could tolerize the animal prior to transplant and avoid runt disease. These findings continue to influence the field of HSC transplantation today as investigators seek to control GVHD while maintaining therapeutic success, in particular when treating malignant disease. However, as discussed below, the relationship between GVHD and therapeutic success differs with the disease being treated. In parallel with the work in mice, others were using dogs to better understand issues of engraftment, rejection, and GVHD [21, 22]. Dogs, while having a clear disadvantage due to their size and cost of housing, had a distinct advantage in being outbred and large enough for the types of surgical procedures needed to be performed at the time. Dog models demonstrated graft rejection and GVHD, but some became long-term engrafters, true HSC transplant successes, and the search was on to understand why. Ultimately, dog models were used to develop immune serum to allow for the identification of matched allogeneic donors, and it was in this setting that the use of methotrexate to reduce GVHD was developed. By the end of the 1960s, the dog model system had been used to develop a nearly 90% success rate from immunomatched allogeneic outbred donors identified using the serum reagents developed by the investigators [23–25]. They had shown quite clearly in a large animal model that lymphocyte immunophenotyping was critical for the success of allogeneic transplants, something that was proven to be true in human transplants and that continues to be of central clinical importance to this day. GVHD remains a great cause of morbidity and mortality following allogeneic HSC transplantation. Improved antileukemic preparative regimens have made disease recurrence less problematic. However, it is now appreciated that GVHD is a double-edged sword when treating leukemia with allogeneic HSC transplantation. GVHD is itself deleterious, but allogeneic HSC transplant also provides a GVL

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effect that is beneficial and contributes to overall survival. Attempts to separate these two immunologic phenomena are under intense study.

1.5 Finally Some Encouraging Results The disappointing results summarized in 1970 [15] saw many investigators leave the field, but some persevered. They believed that success was possible if they could answer a few key questions – cell dose, patient preparation that can both treat disease and prevent graft rejection, and how to reduce the problem with GVHD. A 1972 publication described four patients with aplastic anemia treated with HLA-A matched sibling donors, giving BMT a much-needed boost. All were opposite sex transplants, so standard karyotyping could determine if blood count return posttreatment was due to endogenous marrow recovery or allogeneic marrow engraftment. One patient died from GVHD with a cellular marrow at 45 days posttransplant, another rejected the transplant and died 67 days after transplant, but two were alive with a robust functioning allogeneic marrow at the time of the report, 138 and 215 days out from transplant. In 1975, the BMT team in Seattle published a two-part review [26, 27] that extensively outlined the scientific rationale for performing BMT and the requirements for successful BMT, including details on the care of the patient, the importance of immunosuppression to allow for engraftment and prevent rejection, the need to eradicate underlying malignancy, and the need for HLA matching. It also established a marrow-nucleated cell count dose that should be met for successful transplant and defined many clinical aspects of GVHD. The review also presented the authors’ results treating 37 patients with aplastic anemia and 73 with end-stage leukemia. While the survivorship was low for the patients with end-stage leukemia, the fact that any were alive 2  years posttreatment was a remarkable success that energized the BMT field. Patients were alive that would otherwise have died if it were not for their BMT. However, the field really took off following a 1977 report describing the outcomes of 100 consecutive patients treated with chemotherapy, total body radiation, and sibling-matched allogeneic transplants for end-stage recurrent leukemia. Thirteen of the patients were apparent “cures” as defined by no recurrence of disease at 2 or more years (some over 4 years) posttransplant [28]. The authors and others realized that success might be much higher if leukemia patients were treated before they relapsed and reached end-stage status of their disease. In 1979, two groups reported on matched, related, allogeneic transplantation for leukemia, demonstrating a nearly 50% survival at 2 years [29, 30]. Bone marrow transplant had worked. Patients were benefiting, and over the next 15 years such transplants became part of mainstream medical care. It is estimated that roughly 60,000 transplants were performed around the world in 2010. The Center for International Blood and Marrow Transplant Research maintains a worldwide database of HSC transplants, including source of cells, underlying disease, and outcome (http://www.cibmtr.org).

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1.6 Not All GVHD Is Bad GVHD was rapidly appreciated to be a major complication of allogeneic transplants, and detailed clinical information on how to define this disorder was included in the 1975 two-part report [26, 27]. However, even as far back as the 1950s, it was speculated that the allogeneic donor cells might also provide a beneficial effect when treating malignant diseases such as leukemia [11]. That is, maybe the same immunologic attack of the normal host tissue could also play a role in destroying the diseased cells. This has turned out to be true, with higher cure rates associated with moderate GVHD. This idea was further supported by findings from identical twin (syngeneic) transplants [31]. It was originally thought that an identical twin would be the ideal donor because of the lack of or minimal GVHD. However, patients with acute myelogenous leukemia who received an allogeneic donation from an identical twin had a significantly higher relapse rate than those who received marrow from an HLA-matched sibling [31]. The twin data highlighted that HLA (-A, -B, -DR, and DQ) matching does not match all immunologic differences, and the ones that remain are sufficient to allow for clinically important GVL effect. This immunological therapeutic value of the allogeneic HSC transplant, GVL, remains a critically important contributor to the cure rate for allogeneic transplants for malignant hematologic diseases. However, it is important to remember that there is no beneficial role for graft-versus-disease when using allogeneic transplantation to treat nonmalignant diseases, such as sickle cell anemia [32] and thalassemia [33]. Innovative approaches to reduce GVHD will be essential if we are to bring this valuable treatment to more patients with nonmalignant hematologic disorders [34]. Congenital immunodeficiencies represent yet another group of disease that can be treated with allogeneic transplantation. As with other nonmalignant diseases, GVHD needs to be minimized at all costs. However, these patients allow for greater HLA mismatch in “the other” direction because the recipient immune system is often unable to mount a host-versus-graft response to reject the marrow. Consequently, more gentle conditioning regimens can often be employed, which translates to less therapy-related toxicity. The immunocompetence of the recipient could have a large impact on trial design and clinical outcomes when identifying initial candidates for novel cell therapies developed in the future.

1.7 Source of Hematopoietic Stem Cells While increasing numbers of people now use the name “hematopoietic stem cell transplantation,” from the start and for many years it was called BMT for obvious reasons. In the original 1957 report entitled “Intravenous infusion of bone marrow in patients receiving radiation and chemotherapy,” the cells infused into the six patients were obtained from fetal (n = 1) or adult (n = 1) cadavers, ribs removed at surgery (n = 1), or the anterior or posterior iliac crest aspiration of a living donor (n = 3).

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By the 1970s, iliac crest marrow aspiration was the standard method for obtaining bone marrow cells for transplantation, a procedure that requires general anesthesia. While the HSCs are required for long-term, sustained engraftment, it is wellappreciated that the transplanted marrow includes many more hematopoietic cell types than just HSCs. The importance of the non-HSC cells for assisting with engraftment remains uncertain, but it is quite clear that the non-HSC progenitor cells play a critical role in providing a more rapid production of circulating allogeneic blood cells following infusion. This aspect of progenitor cells helps protect the patient from infection and bleeding, complications of neutropenia and thrombocytopenia, respectively [35]. Given that transplant morbidity and mortality are directly related to the duration of posttransplant cytopenia, the non-HSC cells in the transplanted material clearly play an important and favorable clinical role. Consequently, as cellular therapy moves to other tissues, it is important to consider the value of cells beyond the stem cells proper. It could be that an overly reductionist or “pure” cell population has less benefit than one that contains critical accessory cells. It became clear in the late 1980s that adequate numbers of HSCs could be obtained from the peripheral blood of patients following administration of newly available human cytokines, such as G-CSF or GM-CSF [36]. Interest grew rapidly in the clinical use of such PBSCs as source material for HSC transplantation, and reports of their use became common in the mid-1990s [35, 37–42]. Clinical trials confirmed their safety and efficacy, and PBSCs rapidly expanded as an HSC source for allogeneic and autologous transplants. G-CSF rapidly became the mobilizing agent of choice [43, 44]. More recently, a CXCR4 inhibitor has been approved as an alternate method for mobilizing PBSCs in a subset of patients. Curiously, the use of PBSCs posed a nomenclature problem for the field. How could PBSC transplants be called bone marrow transplants when the cells were not collected from the bone marrow? Fortunately, BMT is also the acronym for “blood and marrow transplantation,” which is how it is commonly used today. While there is not a simple clinical method to quantify the true HSC content of a PBSC product, standard of care is to use CD34 surface expression as a surrogate marker for HSCs and to dose PBSC transplants based on a desired number of CD34+ cells/kg that ensures engraftment. This contrasts with marrow samples, where the clinical adequacy of the collection is based simply on a nucleated cell count/kg. In either case, it is important to realize that no clear enumeration of HSCs is applied to determine the adequacy of an HSC collection, yet the use of surrogate markers has proven productive and safe for many decades. The limited availability of related, matched allogeneic donors became a problem as the sophistication of HLA matching and the use of transplants grew. While in principle one has a one-in-four chance of finding a sibling match, success is even less in real life. Therefore, the majority of patients who can benefit from a BMT do not have an acceptable sibling donor. The first report of a successful, unrelated HLA-matched (HLA-A, -B, -C, and -DR; four loci, which means eight total alleles) allogeneic transplant for leukemia was reported in 1980 [45]. Finding a match was made possible through the advent of more sophisticated HLA phenotyping, but the

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success in finding this particular donor was the result of pure luck and circumstance. The matched donor was a technician at the Seattle transplant center, where everyone had been HLA typed as part of the center’s studies in HLA typing. The first matched, unrelated allogeneic transplant highlighted the potential value of developing a robust mechanism for identifying unrelated HLA-matched donors. As a direct outgrowth of this particular experience and productive lobbying of the US government by concerned and involved individuals, federal funding was eventually allocated for the development of the National Marrow Donor Program (http://www. marrow.org) in the USA. The program has grown dramatically over the ensuing 25 years, is now linked to other similar programs in Europe and elsewhere, and unrelated donors are identified for thousands of patients each year through the sophisticated international systems. It is a great example of how national boundaries and differences can become invisible when health care and humanity are placed above politics. As a testament to the importance and the success of these programs, more unrelated than related allogeneic transplants were performed in the USA in 2009. UCB HSCs [46] were first demonstrated as a clinically useful option for HSC transplants in 1989 [47] when they were used to treat a patient with Fanconi’s anemia, a nonmalignant, congenital blood disorder. UCB has a number of advantages over other HSC sources, including the lack of risk or discomfort to the donor and the ability to store the product in large banks. The latter point means that one can avoid the need to isolate the HSC product from a donor in a timed fashion relative to the patient’s treatments. It also means that intercurrent health issues do not delay or prevent a donation as they can with a living donor. There is also an apparent advantage related to greater tolerance of HLA mismatching [48]. On the other hand, the limited number of cells in most UCB units precludes their use in older adolescents and adults, a fact that has led to the use of multiple UCB units to treat an adult [49]. Regardless, UCB now occupies a legitimate seat at the table of HSC sources for patients of all ages in need of allogeneic HSC transplantation, and the future establishment of organized public UCB banks will be a big step forward in making UCB cells available to more patients in need [50]. While many efforts have been undertaken, human HSCs have not yet been convincingly generated from human embryonic stem cells, so the clinical application of hESC-derived HSCs remains theoretical.

1.8 Autologous HSC Transplantation Allogeneic HSC transplantation was for many years the primary focus for HSC transplantation, and the most common application was to treat hematologic malignancies. However, it was clear from the beginning that autologous transplants may prove useful if antileukemia regimens could eradicate the disease, thereby making unnecessary the GVL effect achieved with allogeneic transplants. The use of combined chemotherapy and total body irradiation preparative regimens provided such an opportunity, as did subsequent use of all chemotherapy preparative regimens,

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and autologous HSC transplant was found to be curative in a number of patients with acute leukemia [51–53]. While autologous transplants are performed in the setting of clinical remission, there was great concern that relapse could be due to reinfusion of leukemia clones with the transplanted cells. This question was addressed with some of the very first gene therapy trials in which viral vectors were used to mark harvested cells prior to their reinfusion. If the viral vector marked relapsed disease, investigators would know that it came from the harvested and reinfused cell product. Such studies showed that a fraction of relapsed disease does in fact come from reinfusion of malignant cells [54–56]. The risks of autologous and allogeneic transplants differ, with the former having a much higher risk of relapse and no risk of GVHD-related morbidity and mortality. In contrast, allogeneic transplants have a much lower risk of relapse but a significant risk of GVHD-related morbidity and mortality. As risk stratification has evolved, different subsets of patients are preferentially treated with one or the other approach.

1.9 Regulatory Agencies BMT grew up in an era quite different from today when it comes to regulation and oversight. In fact, one might wonder if HSC transplantation could have ever gotten off the ground in today’s regulatory environment. For many decades, procedural decisions and standards were established by individual transplant programs without outside scrutiny. However, as programs grew and more centers opened, it became important for professional organizations to establish rules to guide the field. From this appreciation was born the Federation for Accreditation for Cellular Therapy, the major professional organization that now accredits BMT programs, and accreditation has become an important goal for all centers in the USA. The Federation for Accreditation for Cellular Therapy, originally called the Foundation for the Accreditation of Hematopoietic Cell Therapy, was established in 1996 to develop and implement the inspection and accreditation program of the parent organizations, the International Society for Hematotherapy and Graft Engineering and the American Society of Blood and Marrow Transplant. Training of inspectors began in September 1996 and the first on-site inspections began in September 1997. The Foundation for the Accreditation of Hematopoietic Cell Therapy changed its name to the Federation for Accreditation for Cellular Therapy in December 2001 when it became clear that cellular therapy was growing beyond traditional hematopoietic stem and progenitor cells. The Federation for Accreditation for Cellular Therapy inspects an entire program, including collection, laboratory, and clinical care. The Joint Accreditation Committee of the International Society for Cellular Therapy and the European Group for Blood and Marrow Transplantation launched their first official inspection programs in January 2004, providing Europe a similar accreditation program. The American Association of Blood Banks also inspects and accredits BMT laboratories.

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The Federation for Accreditation for Cellular Therapy and like organizations have done much to make BMT programs safer and more responsive to patient needs. In addition to professional accreditation agencies, such as the Federation for Accreditation for Cellular Therapy and the American Association of Blood Banks, BMT programs in the USA must have all or part of the program licensed with state health care agencies and be registered with the Food and Drug Administration. The governmental organizations work to ensure good practices and to provide an avenue to disseminate information relevant to maintaining a safe operation. They make on-site inspections on a regular basis to ensure that procedures are in place and followed and that clinical outcomes and support are consistent with high-quality care. It behooves the cellular therapy community to put energy into professional organizations that provide oversight of any new cellular therapies that develop. Selfpolicing by informed and interested professionals is the best way to ensure safety and reproducibility and to avoid unwanted and unproductive regulations from outside agencies. For BMT programs in the USA, the Federation for Accreditation for Cellular Therapy and the American Association of Blood Banks provide excellent avenues for working with the states and with the Food and Drug Administration to ensure rational and productive systems.

1.10 Conclusions The history of HSC transplantation offers an informative glimpse into the past, providing a number of experiences that can help guide the future of stem cell therapy. First and foremost is the appreciation that HSC transplantation did not “work” right away. In fact, it took decades before people could speak of meaningful clinical success. However, unlike today’s stem cell activities, the field of HSC transplantation grew up in relative anonymity, a truth that made its initial struggles less likely to derail its efforts. Therefore, the first issue for the stem cell field is to not oversell its product or its timeline for success and to articulate clear and simple goals. While the field of HSC transplantation took a while to gather momentum, there were observations even in the early years that proved informative. For example, the relatively early successful transplant of patients with immunodeficiency syndromes highlighted the fact that some patients provide a more receptive environment for transplant engraftment than do others. Such experiences demonstrate the significant impact of highly selected patient populations on successful outcomes. People developing new cellular therapies need to keep this in mind because nothing breeds success and maintains public support like success. Unlike the development and application of HSC transplantation, most novel cellular therapies being considered today are for nonmalignant diseases. This is an advantage because it typically means not having to eradicate a phenotypically dominant disease and replace it with a normal (phenotypically recessive) new stem cell population. For example, replacing injured nerves or destroyed pancreatic islet cells does not require therapy to remove the diseased cells. However, it could be that the

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environment, i.e., cellular niche where the new cells need to engraft, is damaged or altered in the diseased state leaving it less receptive to new cells, such as in myocardial infarction or diabetes. Consequently, understanding the health and makeup of the engraftment location might be critical for success. The field of cellular therapy, both stem cell and other, also needs to keep in mind that the fundamentals are the key. Just like for HSC transplant, one needs to determine the (minimum) number of cells needed to achieve one’s goal and how to best prepare the patient to receive and accept the transplanted cells. It is envisioned that some cellular therapies will ultimately be developed through modification of autologous cells, but that will not happen tomorrow, so selective immunomodulation will be just as important as it is for current day tissue and organ transplantation. Moreover, consideration should be given to the possible use and benefit of accessory cells, much as the non-HSC progenitor cells help with the clinical success of HSC transplants. Of course, well-designed systems to monitor for toxicity and efficacy are essential to keep the field developing productively. Modern stem cell therapy is growing up under an intense public spotlight. The better the cell therapy community polices itself, the more care it takes to learn from the accreditation and inspection organizations that have developed within the HSC transplant community, the more trust it will be given by the public. Involved members of the scientific community must actively engage regulatory agencies and develop professional oversight groups, much like the HSC transplant community has done. This has resulted in better and safer HSC transplant programs, better data monitoring, and it affords the involved community an efficient mechanism for communication and engagement with government organizations. The future for cellular therapy is promising and exciting, and lessons learned along the way must be carefully and actively used to everyone’s advantage.

References 1. Orkin SH, Zon LI (2008) Hematopoiesis: an evolving paradigm for stem cell biology. Cell 132(4):631–644 2. Pasquini MC, Wang Z (2010) Current use and outcome of hematopoietic stem cell transplantation: CIBMTR Summary Slides. http://www.cibmtr.org. Accessed 9 Jul 2011 3. Perry AR, Linch DC (1996) The history of bone-marrow transplantation. Blood Rev 10(4): 215–219 4. Thomas ED (2005) Bone marrow transplantation from the personal viewpoint. Int J Hematol 81(2):89–93 5. Shouse SS, Warren SL, Whipple GH (1931) II. Aplasia of marrow and fatal intoxication in dogs produced by roentgen radiation of all bones. J Exp Med 53(3):421–435 6. Rekers PE, Coulter MP, Warren SL (1950) Effect of transplantation of bone marrow into irradiated animals. Arch Surg 60(4):635–667 7. Jacobson LO, Marks EK, Robson MJ, Gaston E, Zirkle RE (1949) The effect of spleen protetction on mortality following x-irradiation. J Lab Clin Med 34:1538–1543 8. Lorenz E, Uphoff D, Reid TR, Shelton E (1951) Modification of irradiation injury in mice and guinea pigs by bone marrow injections. J Natl Cancer Inst 12(1):197–201 9. Main JM, Prehn RT (1955) Successful skin homografts after the administration of high dosage X radiation and homologous bone marrow. J Natl Cancer Inst 15(4):1023–1029

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10. Ford CE, Hamerton JL, Barnes DW, Loutit JF (1956) Cytological identification of radiationchimaeras. Nature 177(4506):452–454 11. Barnes DW, Corp MJ, Loutit JF, Neal FE (1956) Treatment of murine leukaemia with X rays and homologous bone marrow: preliminary communication. Br Med J 2(4993):626–627 12. Thomas ED, Lochte HL Jr, Lu WC, Ferrebee JW (1957) Intravenous infusion of bone marrow in patients receiving radiation and chemotherapy. N Engl J Med 257(11):491–496 13. Congdon CC, Uphoff D, Lorenz E (1952) Modification of acute irradiation injury in mice and guinea pigs by injection of bone marrow: a histopathologic study. J Natl Cancer Inst 13(1):73–107 14. Thomas ED, Lochte HL Jr, Cannon JH, Sahler OD, Ferrebee JW (1959) Supralethal whole body irradiation and isologous marrow transplantation in man. J Clin Invest 38:1709–1716 15. Bortin MM (1970) A compendium of reported human bone marrow transplants. Transplantation 9(6):571–587 16. Billingham RE, Brent L, Medawar PB (1953) Actively acquired tolerance of foreign cells. Nature 172(4379):603–606 17. Nisbet NW, Heslop BF (1962) Runt disease-II. Br Med J 1(5273):206–213 18. Nisbet NW, Heslop BF (1962) Runt disease. Br Med J 1(5272):129–135,contd 19. Billingham RE (1966) The biology of graft-versus-host reactions. Harvey Lect 62:21–78 20. Billingham RE, Silvers WK (1959) The induction of tolerance of skin homografts in rats with pooled cells from multiple donors. J Immunol 83:667–679 21. Cavins JA, Kasakura S, Thomas ED, Ferrebee JW (1962) Recovery of lethally irradiated dogs following infusion of autologous marrow stored at low temperature in dimethylsulphoxide. Blood 20:730–734 22. Thomas ED, Collins JA, Herman EC Jr, Ferrebee JW (1962) Marrow transplants in lethally irradiated dogs given methotrexate. Blood 19:217–228 23. Epstein RB, Storb R, Ragde H, Thomas ED (1968) Cytotoxic typing antisera for marrow grafting in littermate dogs. Transplantation 6(1):45–58 24. Storb R, Epstein RB, Bryant J, Ragde H, Thomas ED (1968) Marrow grafts by combined marrow and leukocyte infusions in unrelated dogs selected by histocompatibility typing. Transplantation 6(4):587–593 25. Storb R, Rudolph RH, Thomas ED (1971) Marrow grafts between canine siblings matched by serotyping and mixed leukocyte culture. J Clin Invest 50(6):1272–1275 26. Thomas E et  al (1975) Bone-marrow transplantation (first of two parts). N Engl J Med 292(16):832–843 27. Thomas ED et al (1975) Bone-marrow transplantation (second of two parts). N Engl J Med 292(17):895–902 28. Thomas ED et al (1977) One hundred patients with acute leukemia treated by chemotherapy, total body irradiation, and allogeneic marrow transplantation. Blood 49(4):511–533 29. Blume KG, Beutler E (1979) Allogeneic bone marrow transplantation for acute leukemia. JAMA 241(16):1686 30. Thomas ED et al (1979) Marrow transplantation for acute nonlymphoblastic leukemia in first remission. N Engl J Med 301(11):597–599 31. Gale RP et al (1994) Identical-twin bone marrow transplants for leukemia. Ann Intern Med 120(8):646–652 32. Johnson FL, Look AT, Gockerman J, Ruggiero MR, Dalla-Pozza L, Billings FT III (1984) Bone-marrow transplantation in a patient with sickle-cell anemia. N Engl J Med 311(12): 780–783 33. Thomas ED et al (1982) Marrow transplantation for thalassaemia. Lancet 2(8292):227–229 34. Hsieh MM et al (2009) Allogeneic hematopoietic stem-cell transplantation for sickle cell disease. N Engl J Med 361(24):2309–2317 35. Korbling M et al (1995) Allogeneic blood stem cell transplantation for refractory leukemia and lymphoma: potential advantage of blood over marrow allografts. Blood 85(6):1659–1665 36. Socinski MA, Cannistra SA, Elias A, Antman KH, Schnipper L, Griffin JD (1988) Granulocytemacrophage colony stimulating factor expands the circulating haemopoietic progenitor cell compartment in man. Lancet 1(8596):1194–1198

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37. Korbling M et al (1995) Allogeneic blood stem cell transplantation: peripheralization and yield of donor-derived primitive hematopoietic progenitor cells (CD34+ Thy-1dim) and lymphoid subsets, and possible predictors of engraftment and graft-versus-host disease. Blood 86(7):2842–2848 38. Schmitz N et al (1995) Primary transplantation of allogeneic peripheral blood progenitor cells mobilized by filgrastim (granulocyte colony-stimulating factor). Blood 85(6):1666–1672 39. Azevedo WM et al (1995) Allogeneic transplantation with blood stem cells mobilized by rhGCSF for hematological malignancies. Bone Marrow Transplant 16(5):647–653 40. Russell JA et  al (1995) Collection of progenitor cells for allogeneic transplantation from peripheral blood of normal donors. Bone Marrow Transplant 15(1):111–115 41. Bensinger WI et al (1995) Transplantation of allogeneic peripheral blood stem cells mobilized by recombinant human granulocyte colony-stimulating factor. Blood 85(6):1655–1658 42. Dreger P, Suttorp M, Haferlach T, Loffler H, Schmitz N, Schroyens W (1993) Allogeneic granulocyte colony-stimulating factor-mobilized peripheral blood progenitor cells for tretment of engrftment failure after bone marrow transplantaion. Blood 81:1404–1407 43. Bensinger WI et  al (1993) The effects of daily recombinant human granulocyte colonystimulating factor administration on normal granulocyte donors undergoing leukapheresis. Blood 81(7):1883–1888 44. Caspar CB, Seger RA, Burger J, Gmur J (1993) Effective stimulation of donors for granulocyte transfusions with recombinant methionyl granulocyte colony-stimulating factor. Blood 81(11):2866–2871 45. Hansen JA, Clift RA, Thomas ED, Buckner CD, Storb R, Giblett ER (1980) Transplantation of marrow from an unrelated donor to a patient with acute leukemia. N Engl J Med 303(10):565–567 46. Broxmeyer HE et al (1989) Human umbilical cord blood as a potential source of transplantable hematopoietic stem/progenitor cells. Proc Natl Acad Sci USA 86(10):3828–3832 47. Gluckman E et al (1989) Hematopoietic reconstitution in a patient with Fanconi’s anemia by means of umbilical-cord blood from an HLA-identical sibling. N Engl J Med 321(17):1174–1178 48. Wagner JE, Gluckman E (2010) Umbilical cord blood transplantation: the first 20 years. Semin Hematol 47(1):3–12 49. Brunstein CG, Laughlin MJ (2010) Extending cord blood transplant to adults: dealing with problems and results overall. Semin Hematol 47(1):86–96 50. Anonymous (2005) Cord blood: establishing a national hematopoietic stem cell bank program, a 2005 report from The Institue of Medicine of The National Academy of Sciences. http://iom. edu/Reports/2005/Cord-Blood-Establishing-a-National-Hematopoietic-Stem-Cell-BankProgram.aspx. Accessed 9 Jul 2011 51. Linker CA (2003) Autologous stem cell transplantation for acute myeloid leukemia. Bone Marrow Transplant 31(9):731–738 52. Linker CA, Damon LE, Ries CA, Navarro WA, Case D, Wolf JL (2002) Autologous stem cell transplantation for advanced acute myeloid leukemia. Bone Marrow Transplant 29(4):297–301 53. Linker CA, Ries CA, Damon LE, Rugo HS, Wolf JL (1993) Autologous bone marrow transplantation for acute myeloid leukemia using busulfan plus etoposide as a preparative regimen. Blood 81(2):311–318 54. Brenner MK et  al (1993) Gene marking to determine whether autologous marrow infusion restores long-term haemopoiesis in cancer patients. Lancet 342(8880):1134–1137 55. Brenner MK et  al (1993) Gene-marking to trace origin of relapse after autologous bonemarrow transplantation. Lancet 341(8837):85–86 56. Deisseroth AB et al (1994) Genetic marking shows that Ph+ cells present in autologous transplants of chronic myelogenous leukemia (CML) contribute to relapse after autologous bone marrow in CML. Blood 83(10):3068–3076

Chapter 2

Human Embryonic Stem Cells in Regenerative Medicine Odessa Yabut and Harold S. Bernstein

Abstract  Human embryonic stem cells have the capacity for self-renewal and pluripotency, making them a primary candidate for tissue engineering and regenerative therapies. To date, numerous human embryonic stem cell (hESC) lines have been developed and characterized. In this chapter, we discuss how hESC lines are derived, the means by which pluripotency is monitored, and how their ability to differentiate into all three embryonic germ layers is determined. We also outline the methods currently employed to direct their differentiation into populations of tissuespecific, functional cells. Finally, we highlight the general challenges that must be overcome and the strategies being developed in order to generate highly purified hESC-derived cell populations that can safely be used for clinical applications.

Abbreviations bFGF DKK1 hEB hESC HLA miR RPE

Basic fibroblast growth factor Dickkopf homolog-1 Human embryoid body Human embryonic stem cells Human lymphocyte antigen MicroRNA Retinal pigment epithelium

H.S. Bernstein (*) Cardiovascular Research Institute, University of California San Francisco, San Francisco, CA, USA Eli and Edythe Broad Center of Regeneration Medicine and Stem Cell Research, University of California San Francisco, San Francisco, CA, USA Department of Pediatrics, University of California San Francisco, San Francisco, CA, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_2, © Springer Science+Business Media, LLC 2011

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TGFb Transforming growth factor-b VEGF Vascular endothelial growth factor

2.1 Introduction Stem cells have the ability to maintain long-term proliferation and self-renewal. Under specific conditions, stem cells can differentiate into a diverse population of mature and functionally specialized cell types. There are two main types of human stem cells classified according to their source and developmental potential: embryonic and adult, or tissue-specific, stem cells. Human embryonic stem cells are pluripotent cells that can differentiate into all types of somatic and in some cases, extraembryonic tissues. Human adult stem cells are derived from nonembryonic tissues and are capable of generating specific cells from its organ or tissue of origin. Because of the unrestricted potential of human embryonic stem cells (hESCs), these cells have become a highly desirable experimental tool for understanding human development, and are especially attractive for therapeutic applications.

2.2 Sources and Derivation of Human Embryonic Stem Cells hESCs were first derived from the inner cell mass of the blastocyst-stage preimplantation embryo (Fig. 2.1). The inner cell mass is composed of pluripotent cells that are capable of differentiating into the extraembryonic endoderm and the three germ layers that will eventually generate all tissues of the embryo: ectoderm, mesoderm, and endoderm. To generate a hESC line, the cells encompassing the inner cell mass are microsurgically removed and cultured in vitro under specific conditions designed to select cell populations with the capacity to proliferate in the undifferentiated state. Thomson and colleagues reported the first derivation of pluripotent hESCs using this method [1] and were quickly followed by a number of other groups [2–4]. To date, there are 82 hESC lines that adhere to US federal guidelines, many of which are widely used in basic and clinical research. A current list may be found at http://www.grants.nih.gov/stem_cells/registry/current.htm. hESC lines have also been derived from earlier stages of embryonic development, including single blastomeres of 4- or 8-cell stage embryos [5–8] and 16-cell morulae [9, 10] (Fig. 2.1). A single blastomere is considered totipotent and can produce an entire embryo. Thus, blastomere-derived hESCs could circumvent ethical issues surrounding the use of hESCs in biomedical research, since the removal of a single blastomere from an early-stage embryo will, theoretically, not impede the ability of the remaining blastomeres to develop into a normal embryo. hESC lines can also be obtained from parthenogenetic embryos, which are generated when a single egg is fertilized in the absence of male sperm (Fig. 2.1). hESC lines derived using this method can circumvent ethical concerns about the use of embryonic cells since viable embryos are neither created nor destroyed. Parthenote-derived hESC lines have been generated through artificial fertilization of

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Fig. 2.1  Generation of hESC lines from various embryonic sources. Generation of hESC lines undergo three stages. First, donor embryos are obtained after in vitro fertilization or by egg activation (parthenogenetic embryos) and allowed to develop in vitro. Second, pluripotent cells are isolated either from the inner cell mass (ICM) of pre-implantation blastocysts or from 4, 8, or 16-cell stage morulae. Finally, isolated cells are plated in defined hESC medium with or without feeder cell layers to propagate and select for pluripotent cell populations

donor oocytes [11–14]. The ability to derive hESC lines from parthenote-blastocysts is especially attractive not only because of their normal karyotype and their pluripotent properties, but also because these lines contain homozygous major HLA alleles, which could circumvent immunological rejection involved in transplantation therapies (discussed below).

2.3 Characteristics of Pluripotent Human Embryonic Stem Cells hESC lines have been derived from different sources using different methods which can introduce variability between lines. Thus, defining the specific properties and identifying the features of hESCs are critical to their use. In this section, we discuss the guidelines currently used when characterizing new hESC lines.

2.3.1 Cell Morphology and Density Pluripotent hESCs maintain a specific cell morphology and density. hESCs have a high cell nucleus-to-cytoplasm ratio due to an enlarged nucleus and distinct nucleoli.

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Fig. 2.2  Phase contrast images of undifferentiated and differentiating hESCs in culture. (a) A compact colony of proliferating pluripotent hESCs can be seen when cultured in defined medium on mouse embryonic fibroblasts. (b) Floating hEBs are observed at 2 days after induction of differentiation. (c) Differentiating cardiomyocytes appear within adherent cultures at 48 h after plating hEBs onto a gelatin-coated culture dish. Bar, 25 mm

Proliferating pluripotent hESCs form compact and spherical cell colonies when grown on mouse embryonic fibroblast cell layers (Fig. 2.2a). Differentiating hESCs are easily distinguished by the loss of compact morphology and the appearance of flattened cells that form at the edges of the colony. This can be controlled with regular supplementation of fresh growth medium [15].

2.3.2 Expression Profiling A systematic study has been conducted by the International Stem Cell Initiative, a consortium of stem cell researchers from more than 15 countries, on 59 independently derived and commonly used hESC lines in order to identify a panel of molecular markers that are consistently and strongly expressed in pluripotent hESCs [16]. These included developmentally regulated genes such as NANOG, POU domain class 5 homeobox 1 protein (POU5F/OCT4), teratocarcinoma-derived growth factor 1, DNA (cytosine-5-)-methyltransferase 3b, g-aminobutyric acid A receptor b3, and growth differentiation factor 4. The study also established that the collective expression of Stage-Specific Embryonic Antigens 3 and 4, along with keratin sulfate (TRA-1-60, TRA-1-81, GDTM2, and GCT343) and protein (CD9 and Thy1) antigens, are reliable cell surface markers of pluripotent hESCs. Other characteristics of hESCs include the expression of the enzyme alkaline phosphatase, Stem cell factor (or c-Kit ligand), and class 1 HLA. The expression profile of small, noncoding RNAs known as microRNAs, which regulate translational efficiency of target mRNAs [17], has been evaluated in hESCs by several groups [18–22]. These studies identified a number of miR family clusters specifically expressed in pluripotent hESCs. Among these are miR-92b, the miR302 cluster, miR-200c, the miR-368 and miR-154* clusters, miR-371, miR-372, miR-373*, miR-373, and the miR-515 cluster [18, 22]. Functional studies of some

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of these miRs, such as miR-302 and miR-92b, have established roles in pathways that control self-renewal and maintenance of the pluripotent stem cell state [23, 24]. To date, studies that compare the miR expression profiles of all available hESC lines are still lacking. A comprehensive analysis of miR expression profiles is warranted, as this will identify miRs that are expressed across hESC lines, and could be used to select for pluripotent populations, evaluate newly derived hESC lines, and understand mechanisms that regulate basic hESC biology.

2.3.3 Epigenetic Properties Epigenetic mechanisms influence gene expression through heritable modifications in chromosomal or DNA structure, such as DNA methylation, histone modification, and X-chromosome inactivation. Similar to expression patterns of coding genes discussed above, the epigenetic properties of pluripotent hESCs can be used as molecular signatures to distinguish them from other cell types. The chromatin structure of hESCs is generally in an open conformation, making it readily accessible to the transcriptional machinery necessary for the maintenance of pluripotency [25]. It has also been observed that hESC lines display DNA methylation profiles distinct from most other cell types [26]. One study of 14 different lines revealed markedly reduced methylation patterns of CpG dinucleotides when compared to somatic cells. Further analysis revealed that the observed differential methylation of hESCs was specific to promoter regions of pluripotency genes such as OCT4 and Nanog [27]. Thus, the unique epigenetic properties of hESCs likely promote maintenance of the pluripotent state and can be used as a hallmark of undifferentiated hESCs. To date, almost all established female hESC lines analyzed exhibit partial or complete X-chromosome inactivation, a process that occurs as early as the blastocyst stage and leads to methylation of promoter regions. The states and levels of X-inactivation appear to differ between hESC lines, and also between subcultures of each hESC line that are propagated by different laboratories [28–31]. These observations point out that in addition to genetic heterogeneity, the environment and culture conditions can lead to variable and unstable epigenetic states. The variability in X-chromosome inactivation could result in inconsistencies as hESCs are developed for therapeutic applications. Thus, generating an epigenetically naïve hESC line, in which X-chromosome inactivation or other epigenetic modifications have not yet occurred, is an important goal.

2.3.4 Pluripotency of hESCs hESCs are defined in part by their capacity to differentiate, which can be tested using in vivo and in vitro methods. A test of pluripotency in vitro involves determining

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Fig. 2.3  In vivo differentiation of hESCs by teratoma formation. Proliferating cultures of hESCs were used to form teratomas by renal capsule grafting using established methods [35]. (a) An explanted teratoma is shown. (b–f) Teratomas were sectioned and stained with hematoxylin and eosin to identify embryonic tissues. Representative tissues from all three embryonic germ layers can be seen, including endoderm (b), mesoderm (c, d), and ectoderm (e, f). (b) Glandular intestinal structure. (c) Nascent renal tubules and glomeruli within a bed of primitive renal epithelium. (d) Cartilage surrounded by capsule of condensed mesenchyme. (e) Nascent neural tube. (f) Primitive squamous epithelium. Bar, 100 mm

the ability of hESCs to form hEBs when cultured in a nonadherent cell suspension in the absence of feeder cell layers (Fig. 2.2b). hEBs are spherical colonies of differentiating hESCs that contain cell types representative of all three embryonic germ layers [32]. hEBs can be differentiated into specific tissues under suitable culture conditions (Fig. 2.2c). The most commonly used in vivo method to test pluripotency involves the transplantation of undifferentiated hESCs into immunodeficient mice to induce the formation of teratomas [33–36]. Teratomas are benign tumors comprised of disorganized tissue structures characteristic of the three embryonic germ layers. Analysis of embryonic tissues found in teratomas from engrafted hESCs can be used to test their differentiation potential (Fig. 2.3). The ability of hESCs and hEBs to mimic in vitro and in vivo the events occurring during human embryonic development makes them valuable tools for understanding the mechanisms involved in developmental processes, and steppingstones toward the generation of desired cell types suitable for cell therapies.

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2.4 Stem Cell Derivatives and Their Uses for Cell-Based Therapies Cellular insufficiency or deficiency, due to dysfunction or degeneration, respectively, is the root of diseases such as heart failure, neurodegenerative disorders, diabetes, bone marrow failure, and spinal cord injury. For centuries, therapeutic approaches have been limited to the surgical removal of damaged tissues or treatment with pharmacological therapies to ameliorate symptoms and fight infection. Thus, the prospect of replacing damaged or missing cells with new functional cells has shifted the therapeutic paradigm toward restoring tissue function. Deriving specific cell populations from hESCs that could either replace damaged cells or coax neighboring cells to function normally provides a promising strategy for cell-based therapy. With hESCs, it is possible to generate lineage-restricted progenitors that are capable of differentiating into specialized postmitotic cell types such as cardiomyocytes, pancreatic islet cells, chondrocytes, hematopoietic cells, endothelial cells, or neurons. Furthermore, the ability of hESCs to divide indefinitely makes these cells an inexhaustible large-scale source of specific progenitors. Current research studies are focused on identifying and refining ways for directing the differentiation of hESCs that will enrich for pure, homogenous populations of specific cell types. In the following sections, we provide some examples of how differentiation of hESCs can be directed toward specific cell/tissue types, and the potential use of these cell types for clinical applications (Table 2.1).

2.4.1 Endodermal Cell Derivatives of hESCs Endodermal derivatives include cells that populate the lung, liver, and pancreas. Directing the differentiation of hESCs toward definitive endoderm would help generate specific cell types, such as islet cells or hepatocytes, which could be used toward treatment of diseases such as diabetes or liver disease, respectively. D’Amour et al. [37] showed that selective induction of endoderm could be achieved through the addition of high concentrations of activin A, under low serum conditions, and in a stage-specific manner. Activin A mimics the action of Nodal, a ligand that activates TGFb signaling, which in turn leads to the induction of endoderm differentiation. The effect of activin A in inducing definitive endoderm is enhanced when additional factors such as Wnt3a [38] and Noggin [39] are present, or when coupled with the suppression of the phosphoinositide 3-kinase pathway [40]. Induction of definitive endoderm can lead to the generation of specific progenitor populations following the addition of other factors. Among the most successful examples to date is the generation of pancreatic islet progenitors devised by Kroon et al. [41], accomplished through the sequential exposure of hESCs to activin A and Wnt3A, followed by the addition of keratinocyte growth factor or fibroblast growth factor-7 to induce the formation of the primitive gut tube. Subsequently, retinoic

Table 2.1  Examples of methods used to differentiate hESCs into specific cell types Methods to induce differentiation Differentiation factors and/or culture conditions FGF, BMP4, hepatocyte growth factor, Derivation of endodermal Differentiation of hESC into oncostatin M, dexamethasone cells from hESCs definitive endoderm, followed by sequential exposure to Activin A, Wnt3A, keratinocyte growth differentiation factors factor/FGF7, retinoic acid, cyclopamine, noggin Recombinant keratinocyte growth factor Genetic modification of hESCs followed by spontaneous differentiation Human embryoid body formation Serum-free conditions; BMP4 Derivation of mesodermal cells Micromass of dissociated embryoid bodies; BMP2 from hESCs High-density culture of dissociated embryoid bodies; ascorbic acid, dexamethasone Serum-free conditions; bFGF Spin embryoid body formation Serum-free conditions Co-culture with stromal cells Co-culture with stromal cell line M210-B4 for enhanced proliferation of CD34+/CD45+ hematopoietic progenitor cells Dense monolayer of hESCs; activin A, BMP4 Directed differentiation from hESCs by sequential exposure BMP4, BMP4/bFGF/activin A, VEGF/DKK1, to differentiation factors VEGF/DKK1/bFGF Co-culture with primary chondrocytes; poly-d, Directed differentiation from l-lactide scaffold hESCs by the addition of differentiation factors on 3D polymeric scaffolds Cardiac-specific reporter Genetic modification of hESCs followed by spontaneous differentiation Cardiomyocytes [53]

Chondrocytes [72]

Cardiomyocytes [55] Cardiomyocytes [56]

Cardiomyocytes [50] Blood cells [45] T and NK cells [47]

Dendritic cells [46] Chondrocytes [70] Chondrocytes [73]

Lung alveolar cells [68, 69]

Pancreatic islet progenitors [41]

Example of differentiated cells Hepatocytes [42, 43]

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Derivation of ectodermal cells from hESCs

Directed differentiation from hESCs with sequential exposure to differentiation factors Direct differentiation with sequential exposure to differentiation factors on 3D culture with extracellular matrix components

Co-culture with stromal cells and addition of differentiation factors Formation of neural rosettes and addition of differentiation factors

Methods to induce differentiation

Dopaminergic neurons [62] Schwann cells [64]

FGF8, SHh Ciliary neurotrophic factor, neuregulin 1b, dbcAMP Retinoic acid, SHh Withdrawal of FGF2, BDNF; addition of GDNF, NGF, dibutyryl cyclic AMP Serum-free conditions; activin A, nicotinamide B27, thyroid hormone, retinoic acid, FGF2, EGF, insulin BMP4, ascorbic acid

Dopaminergic neurons [59]

FGF8, SHh

Basal keratinocytes [67]

Motor neurons [63] Peripheral sympathetic and sensory neurons [64] Retinal pigment epithelium [66] Oligodendrocytes [65]

Example of differentiated cells

Differentiation factors and/or culture conditions

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acid, cyclopamine, and Noggin are added to inhibit hedgehog and TGFb signaling, and thus induce the differentiation of posterior foregut cells, the source of pancreatic cell progenitors. These are cultured further to generate pancreatic endoderm cells. When these cells are engrafted in immunodeficient mice, they display the histological and structural characteristics of pancreatic islet cells, and are able to sustain insulin production for at least 100 days [41]. In a similar manner, hepatocytes can be obtained after differentiation of hESCs into definitive endoderm [42, 43]. A robust population of functional hepatocytes was generated with the sequential addition of low serum medium, collagen I matrix, and hepatic differentiation factors that include FGF, BMP4, hepatocyte growth factor, oncostatin M, and dexamethasone [43]. These cells expressed known markers of mature hepatic cells, exhibited appropriate function, and were able to integrate and differentiate into mature liver cells when injected into mice with liver injury [43].

2.4.2 Mesodermal Derivatives of hESCs Directing the differentiation of hESCs into mesoderm requires the activation of the TGFb signaling pathway and can be accomplished through the stepwise and dosage-dependent addition of activin A, BMP4, and growth factors, VEGF and bFGF [44]. Mesodermal derivatives have also been successfully obtained by spontaneous differentiation of hESCs through hEB formation without first directing them toward mesoderm. Robust differentiation of hESCs into hematopoietic lineage cells, which give rise to all blood cell types and components of the immune system, has been achieved under serum-free conditions through spin hEB formation [45]. Specific hematopoietic cells, such as functional dendritic cells, have been successfully differentiated from hESCs through spontaneous hEB formation under serum-free conditions with the addition of BMP4 at specific time points [46]. Hematopoietic progenitor cells that give rise to functional T and natural killer cells capable of targeting human tumor cells both in vitro and in vivo have also been derived from hESCs co-cultured with stromal cells [47]. Thus, the ability to differentiate hESCs into hematopoietic lineage cells promises to be useful in improving existing therapies that require blood cell transplantation, and in immune therapies that require induction of the immune response in an antigenspecific manner [48]. Cardiomyocytes, which represent another therapeutically important derivative of mesoderm, have been successfully generated from hESCs using several methods [49]. Through hEB formation, hESCs can spontaneously differentiate into cardiomyocytes under appropriate culture conditions. These cardiomyocytes exhibit morphological, molecular, and electrophysiological properties similar to adult cardiomyocytes [50], and display quantifiable responses to physiological stimuli reminiscent of atrial, ventricular, and pacemaker/conduction tissue [51–54]. Cardiomyocytes have also been generated by directed differentiation with activin A and BMP4 on a dense monolayer of hESCs; these cells successfully form specific

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cardiac lineages when transplanted in  vivo [55]. Another study used additional medium supplements that included VEGF, and the Wnt inhibitor, DKK1, followed by the addition of bFGF to promote cardiomyocyte differentiation from hEBs [56]. Success of these studies was measured by the expression of proteins specific for mature cardiac cells such as cardiac troponin T, atrial myosin light chain 2, and the cardiac transcription factors, Tbx5 and Tbx20.

2.4.3 Ectodermal Derivatives of hESCs The dominant differentiation pathway in hESC cultures leads to the formation of ectoderm, which makes up cells of the nervous system and the epidermis. hESCderived neural progenitor cells are characterized by rosette-like neural structures that form in the presence of growth factors, FGF2 or EGF, through either spontaneous differentiation from an overgrowth of hESCs or after hEBs are plated onto adherent substrates [57, 58]. These neural rosettes have become the signature of hESCderived neural progenitors, capable of differentiation into a broad range of neural cells in response to appropriate developmental signals. Thus, many studies are exploring ways to enhance the formation of neural rosettes in order to generate an enriched population of specific neural cell types. One example is the use of specific stromal cell lines [59]. With this method, stromal cells provide ectodermal signaling factors required for neural induction, as determined in animal model studies, and therefore promote the formation of neural rosettes [60, 61]. The withdrawal of FGF2 and EGF, and the addition of specific compounds can lead to the differentiation of neural rosettes into specific neural subtypes. For example, hESC-derived neural progenitors treated with FGF8 and sonic hedgehog give rise to dopaminergic neurons [62], while treatment with sonic hedgehog and retinoic acid induce motor neuron differentiation [63]. Neural crest stem cells derived from neural rosettes can differentiate into peripheral sympathetic and sensory neurons by withdrawing FGF2/EGF and adding BDNF, GDNF, NGF, and dbcAMP, or into Schwann cells in the presence of CNTF, neuregulin 1b, and dbcAMP [64]. Neuroglial cells, such as oligodendrocytes, are generated with B27, thyroid hormone, retinoic acid, FGF2, epidermal growth factor, and insulin [65]. In 2010, the biotechnology firm, Geron, initiated the first clinical trial with hESCs in the USA using hESC-derived oligodendrocytes to treat acute spinal cord injuries (http://www.clinicaltrials.gov/ct2/archive/NCT01217008). Oligodendrocytes are rapidly lost following acute spinal cord injury leading to demyelination and neuronal loss. In these trials, purified oligodendrocyte progenitor cells derived from hESCs will be injected into the spinal cord of paralyzed patients within 2 weeks of the acute injury. While this first trial is a safety study, the expectation is that these progenitor cells will terminally differentiate into oligodendrocytes and produce myelin, which insulates neuronal cell membranes and is critical for efficient conduction of nerve impulses. Thus, the transplantation of newly differentiated oligodendrocytes is expected to restore myelination of damaged neurons preventing further neuronal death and restoring function.

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Retinal pigment epithelium (RPE) cells are another specific cell type derived from neuroectoderm. These support the neural retina by phagocytosing and renewing the photoreceptor outer segments of rhodopsin. Recent reports have shown that RPE can be induced from hESCs in the presence of nicotinamide and activin A under serum-free conditions [66]. hESC-derived pigmented cells exhibit the morphological and functional properties of RPE cells after transplantation in an animal model of macular degeneration, a disease caused by dysfunction and loss of RPE. These data have led to the second and third clinical trials using hESCs by the biotechnology company, Advanced Cell Technology. For these trials, hESC-derived RPEs will be transplanted directly into the degenerating retinae of patients with Stargardt’s Macular Dystrophy, a juvenile form of macular degeneration, or Dry Age-Related Macular Degeneration, to rescue visual acuity. The launch of these clinical trials heralds the translation of hESC research into therapy for neurodegenerative disease.

2.5 The Promise of hESCs in Tissue Engineering Tissue engineering and regeneration utilize biological substitutes to restore or maintain tissue function. As with cell transplantation, a successfully engineered tissue depends on the generation of the appropriate cell type that is able to provide normal cellular function. Thus, cells suitable for tissue engineering should have the ability to enter a desired differentiation program to produce a specific cell type, and be expandable in vitro to meet the needs of cell transplantation. hESCs provide much promise in tissue engineering and regeneration since hESCs can act as an inexhaustible in  vitro source of differentiated cell types. The potential use of hESCs in tissue engineering include, but are not limited to, organ substitutes, vascularization, and ex vivo cartilage/bone construction. While these applications are discussed in detail in subsequent chapters, brief examples are provided below. Basal keratinocytes, the cells that make up the pluristratified epidermal layer of the skin, have been successfully differentiated from hESCs. Guenou et al. [67] have shown that long-term culture of hESCs in defined medium supplemented with BMP4 and ascorbic acid leads to the directed differentiation of hESCs into basal keratinocytes. These cells express keratins 14 and 5, a6- and b4-integrins, collagen VII, and laminin 5 at levels comparable to postnatal keratinocytes. More importantly, these hESC-derived keratinocytes form a cohesive pluristratified epidermis when placed in 3D culture or when engrafted into immunodeficient mice. These findings prove the feasibility of using hESC-derived keratinocytes as a source of allograft for patients requiring skin restoration. The use of hESCs to treat lung injury has also been an area of active investigation. A significant step toward directed differentiation of lung-specific cells was reported by Wang et al. [68, 69], in which genetically modified hESCs carrying lung-specific reporters under the control of promoters from tissue-specific genes such as surfactant

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protein C, aquaporin 5, and T1a, resulted in the purification of type I and type II alveolar epithelial cells. When engrafted into mice suffering from acute lung injury, these cells terminally differentiated in vivo into type I and type II alveolar epithelial cells and exhibited functional properties that include the capacity for gas exchange and histological amelioration of lung injury. hESCs readily form connective tissue, such as bone or cartilage, as can be appreciated from teratoma formation assays (Fig. 2.3). Thus, hESCs are a valuable source of cells suitable for connective tissue replacement therapy for a number of bone and joint diseases, such as osteoarthritis, which is characterized by the breakdown of cartilage within joints. Most successful and efficient protocols for directing chondrocyte differentiation from hESCs utilize 3D culture systems created by seeding hESCs at high density leading to the formation of a pellet, or by introducing the cells into a synthetic 3D scaffold. Such systems enable cell–cell signaling between the undifferentiated hESCs and mature chondrocytes to stimulate homogeneous and sustained chondrogenic differentiation. For example, single-cell suspension of dissociated hEBs cultured as high-density micromass with BMP2 leads to efficient chondrocyte formation [70]. hESCs co-cultured with primary chondrocytes, or in the presence of osteogenic supplements and polymeric scaffolds, yield cartilaginous- or osteogenic-like cells [71, 72]. More recently, feeder-free 3D culture systems have successfully derived multipotent connective tissue progenitors from hESCs yielding tendon-like structures [73]. The engraftment of these in vitro differentiated tendon structures in injured immunosuppressed mice restored ankle joint movements that rely on an intact Achilles tendon [73]. Furthermore, there is evidence that transplanted chondrogenic cells may exert a stimulatory effect through paracrine mechanisms that promote growth and repair of endogenous cells [74].

2.6 Current Challenges As discussed above, cell therapy with hESCs has begun to enter clinical trials. The International Stem Cell Banking Initiative has been created by the International Stem Cell Forum, a group of national and international stem cell research funding bodies, to develop a set of best practices and principles when banking, testing, and distributing hESCs for clinical application [75]. In the USA, the Food and Drug Administration also monitors these guidelines and has issued recommendations for reviewers of proposals for clinical trials of stem cell therapy (http://www.fda.gov/ BiologicsBloodVaccines/GuidanceComplianceRegulatoryInformation/Guidances/ Xenotransplantation/ucm074131.htm). It is important to note that these recommendations do not ensure the quality or efficacy of hESC-derived cells used for clinical application. Rather, these guidelines warrant that the cells used for therapy are reproducible and meet specific criteria to ensure patient safety (Table  2.2). The major safety concerns for the use of hESCs are discussed in the following sections.

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Table 2.2  Requirements for standardization and optimization of hESCs for clinical use Important factors Examples of test methods Cell line identity: must match Short tandem repeat (STR) testing or human leukocyte all alleles of parent cell line antigen (HLA) testing Sterility and pathogen screening Bacteria/fungi/mycoplasma testing by microbiological culture; qPCR analysis for murine viral short interspersed elements (SINE) Genetic/chromosomal stability Analysis of multiple single nucleotide polymorphisms (SNP); karyotype by G-band analysis of 20 metaphase spreads or fluorescent in situ hybridization Epigenetic stability MicroRNA profiling, methylation analysis, X-inactivation Pluripotency Formation of teratomas in immunodeficient mice; flow cytometry to determine hESC-specific antigens such as SSEA-3/4, TRA-1-60, TRA-1-81 Quality and differentiation Gene expression profiling by DNA microarray or qPCR ability analysis to analyze expression of markers of pluripotency or differentiated cell types; ability to form embryoid bodies Functional assays Report on potency, efficacy, and lot-to-lot variability

2.6.1 Xenobiotic-Free Conditions Many of the hESC lines currently in use have been exposed to animal products during isolation of the inner cell mass and propagation of hESCs in vitro. Under these conditions, hESCs could possess animal viruses and other unknown substances capable of eliciting a detrimental immune response in transplanted hosts. Currently, hESC lines under development for clinical use undergo extensive microbiological testing as strictly recommended by the International Stem Cell Banking Initiative. In the USA, the Food and Drug Administration legally requires documentation of the source, potential genetically modified components, and pathogenic agents in any hESCderived cell intended for therapeutic use. Thus, avoiding exposure to xenobiotics is emphasized by law. Recently, replacement media have been developed that would allow maintenance of hESCs in xenobiotic-free conditions. These include xenobioticfree serum replacements such as knockout serum replacer (KSR; Invitrogen) or xenobiotic-free culture media such as HESGRO (Millipore) or TeSR (STEMCELL). Feeder-free culture systems are now being developed to reduce the risk of contamination with foreign agents when hESCs are cultured on animal feeder cell layers. Feeder-free and xenobiotic-free, defined culture media that consist of a combination of recombinant growth factors known to inhibit differentiation and maintain hESCs in the pluripotent state are now commercially available. However, some reports have associated feeder-free culture conditions with greater chromosomal instability and an increased risk of propagating genetically altered hESCs [76]. For this reason, most hESC laboratories practice a surveillance program for genomic instability in cultured lines [36, 53].

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hESC lines derived using human feeder cells have been reported. For example, hESC lines have been successfully derived on human fibroblasts generated from neonatal foreskin [77, 78] and adult skin fibroblasts [79]. Some laboratories deriving new lines have moved exclusively to xenobiotic-free conditions [80]. The ability to derive and maintain new hESC lines using human fibroblast feeder cells represents a significant step toward generating clinical-grade hESCs.

2.6.2 Genetic Abnormalities in hESC Lines The best characterized hESC lines to date are among the earliest lines derived. However, they may not be the best lines for therapeutic applications as many of these lines were derived using animal products. Chromosomal and genomic instability has been detected among several hESC lines, including loss of heterozygosity or copy-number variation in cancer-related genes [81, 82]. Many of these mutations appeared to be induced by prolonged culture, since these changes were not observed in low passage cells. It has been proposed that such karyotypic aberrations occurred with adaptation to the original culture conditions used when the first few lines were being derived and expanded [83]. These observations emphasize the need for complete characterization of hESC lines, particularly the effects of long-term culture, and the design of guidelines for designating therapeutic-grade hESCs.

2.6.3 Enrichment, Directed Differentiation, and Purification Protocols for hESCs A primary safety concern when using pluripotent hESCs is their potential to form germ layer tumors. As discussed above, in vivo transplantation of undifferentiated hESCs in immunodeficient mice results in teratoma formation. Evidence of tumorlike growths has also been observed in differentiated hESC derivatives transplanted in vivo [84, 85]. Thus, it is essential that candidate hESC derivatives intended for use in cell transplantation are free of tumorigenic cells. Another concern is the differentiation of hESC-derived cells into unwanted cell types. For example, the engraftment of inappropriate muscle cells into the myocardium could alter the electrical activity of recipient tissue, provoking arrhythmias [86]. Thus, developing and further optimizing differentiation and purification protocols are necessary to minimize the generation of unwanted cell types for preclinical transplantation experiments and clinical therapy. As discussed earlier in this chapter, enrichment of specific cell types can be achieved using molecules introduced at specific time points during culture. However, many of these methods yield only moderate enrichment that is not yet scalable for clinical application. It may be desirable to enrich first for partially differentiated, proliferative hESC intermediates with specific cell fates. These could then be

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expanded before further differentiation into cells for therapy. For example, the expression of the cell surface antigen, CD133, on proliferating hESCs identifies cells predestined toward a neuroectodermal fate [34]. CD133-positive cells have been selected from cultures of undifferentiated hESCs, and have been observed to differentiate primarily into neuroectodermal cells in vitro and in vivo [34]. In the absence of specific cell surface antigens such as CD133 to identify tissuespecific precursors, molecular beacons have been used to select for specific subpopulations of hESCs. King et al. [33] first demonstrated the utility of this system in isolating live Oct4-expressing pluripotent hESCs in a specific and high-throughput manner. Molecular beacons are single-stranded oligonucleotides that generate fluorescent signals when bound to their target mRNAs, making these cells detectable and selectable by fluorescence-activated cell sorting. More importantly, molecular beacons have a short lifespan within cells and do not alter the function or genomic structure of hESCs. Thus, this method can be used to enrich for desired hESCderived cell populations or used to select against unwanted cell types, such as undifferentiated hESCs that could form tumors.

2.6.4 Circumventing Immune Rejection Using Transplanted hESC-Derived Cells Transplanted hESCs encounter immune rejection [87] because proliferating and differentiated hESCs express class I and II HLA as well as minor histocompatibility antigens at levels sufficient to activate the immune system [87, 88]. Another potential barrier to hESC engraftment can occur through mismatch between donor hESC and recipient ABO blood group antigens. While studies to determine the effects of ABO incompatibility on hESC transplantation are still lacking, this has long been a criterion for successful organ transplantation and thus, it is likely that ABO incompatibility between hESC-donor cells and the recipient would also trigger immune rejection. Ideally, having genetically identical donor and patient cells is the best way to circumvent immune rejection. Thus, there is high interest in developing and using somatic cell nuclear transfer to generate patient-specific hESC lines. Using this technique, the DNA obtained from either a patient’s skin or muscle cell would be transferred into an unfertilized egg that has had its DNA removed. Subsequently, the egg is artificially fertilized and allowed to develop until it reaches the blastocyst stage to derive hESCs. The resultant hESC line would have an immunologic profile matching the patient and could be used for cell therapy. This technique has been conducted successfully in animals using species-specific ESCs, but the bona fide derivation of hESCs through somatic cell nuclear transfer has not yet been reported. Another strategy is to generate hESC lines with the closest match to potential transplant patients. Suggestions have included engineering “universal donor hESCs,” a blood antigen O cell in which the expression of HLA is suppressed, or chimeric hematopoietic cells derived from hESCs capable of inhibiting the immune response when co-transplanted with the desired hESC-derived cells [89]. Alternatively, creating

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hESC banks that store lines representing HLA/ABO combinations that match the majority of potential patients has been proposed. Studies have provided estimates on how many hESC lines would be needed in order to support the needs of a specific population. Taylor et al. [90] estimated that approximately 150 hESC lines could provide an HLA match for most of the population in the United Kingdom. Alternatively, approximately ten parthenote-derived hESC lines that are homozygous for HLA types could be sufficient for a majority of the population. Studies by Nakajima et al. [91] estimated that approximately 170 hESC lines, or 55 hESC lines with homozygous HLA types, would be sufficient for 80% of patients in the Japanese population. These findings demonstrate the feasibility of creating and maintaining a hESC bank with sufficient representation to support a large number of patients. However, in countries such as the USA, many more hESC lines would need to be established to serve its ethnically and genetically diverse population. Given the ethical issues and restrictions on hESC research, and the small number of approved hESC lines currently available, the creation of a hESC bank with a highly diverse collection of cell lines will undoubtedly face enormous challenges.

2.7 Conclusions Research on hESCs has progressed significantly since their first derivation in 1998. The international scientific community has discovered the enormous potential of hESCs as newly derived lines continue to be developed, and differentiation methods into various types of cells are optimized for scientific investigation and clinical use. It is clear that there are still major scientific challenges as well as ethical and legislative issues that must be addressed, especially in the USA. Certainly more questions will emerge as more is understood in the coming years. However, it is encouraging to see that clinical trials involving the use of hESCs in spinal cord injury and macular degeneration have begun. These studies will pave the way toward determining the therapeutic benefit of hESCs in regenerative medicine. Acknowledgments  The authors thank members of the Bernstein Laboratory for helpful discussion. H.S.B. is supported by grants from the National Institutes of Health, the California Institute for Regenerative Medicine, and the Muscular Dystrophy Association. O.Y. is supported by a fellowship from the National Institutes of Health.

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57. Reubinoff BE, Itsykson P, Turetsky T, Pera MF, Reinhartz E, Itzik A, Ben-Hur T (2001) Neural progenitors from human embryonic stem cells. Nat Biotechnol 19(12):1134–1140 58. Zhang SC, Wernig M, Duncan ID, Brustle O, Thomson JA (2001) In vitro differentiation of transplantable neural precursors from human embryonic stem cells. Nat Biotechnol 19(12):1129–1133 59. Perrier AL, Tabar V, Barberi T, Rubio ME, Bruses J, Topf N, Harrison NL, Studer L (2004) Derivation of midbrain dopamine neurons from human embryonic stem cells. Proc Natl Acad Sci USA 101(34):12543–12548 60. Kawasaki H, Mizuseki K, Nishikawa S, Kaneko S, Kuwana Y, Nakanishi S, Nishikawa SI, Sasai Y (2000) Induction of midbrain dopaminergic neurons from ES cells by stromal cellderived inducing activity. Neuron 28(1):31–40 61. Kawasaki H, Suemori H, Mizuseki K, Watanabe K, Urano F, Ichinose H, Haruta M, Takahashi M, Yoshikawa K, Nishikawa S, Nakatsuji N, Sasai Y (2002) Generation of dopaminergic neurons and pigmented epithelia from primate ES cells by stromal cell-derived inducing activity. Proc Natl Acad Sci USA 99(3):1580–1585 62. Yan Y, Yang D, Zarnowska ED, Du Z, Werbel B, Valliere C, Pearce RA, Thomson JA, Zhang SC (2005) Directed differentiation of dopaminergic neuronal subtypes from human embryonic stem cells. Stem Cells 23(6):781–790 63. Li XJ, Du ZW, Zarnowska ED, Pankratz M, Hansen LO, Pearce RA, Zhang SC (2005) Specification of motoneurons from human embryonic stem cells. Nat Biotechnol 23(2):215–221 64. Lee G, Kim H, Elkabetz Y, Al Shamy G, Panagiotakos G, Barberi T, Tabar V, Studer L (2007) Isolation and directed differentiation of neural crest stem cells derived from human embryonic stem cells. Nat Biotechnol 25(12):1468–1475 65. Nistor GI, Totoiu MO, Haque N, Carpenter MK, Keirstead HS (2005) Human embryonic stem cells differentiate into oligodendrocytes in high purity and myelinate after spinal cord transplantation. Glia 49(3):385–396 66. Idelson M, Alper R, Obolensky A, Ben-Shushan E, Hemo I, Yachimovich-Cohen N, Khaner H, Smith Y, Wiser O, Gropp M, Cohen MA, Even-Ram S, Berman-Zaken Y, Matzrafi L, Rechavi G, Banin E, Reubinoff B (2009) Directed differentiation of human embryonic stem cells into functional retinal pigment epithelium cells. Cell Stem Cell 5(4):396–408 67. Guenou H, Nissan X, Larcher F, Feteira J, Lemaitre G, Saidani M, Del Rio M, Barrault CC, Bernard FX, Peschanski M, Baldeschi C, Waksman G (2009) Human embryonic stem-cell derivatives for full reconstruction of the pluristratified epidermis: a preclinical study. Lancet 374(9703):1745–1753 68. Wang D, Haviland DL, Burns AR, Zsigmond E, Wetsel RA (2007) A pure population of lung alveolar epithelial type II cells derived from human embryonic stem cells. Proc Natl Acad Sci USA 104(11):4449–4454 69. Wang D, Morales JE, Calame DG, Alcorn JL, Wetsel RA (2010) Transplantation of human embryonic stem cell-derived alveolar epithelial type II cells abrogates acute lung injury in mice. Mol Ther 18(3):625–634 70. Toh WS, Yang Z, Liu H, Heng BC, Lee EH, Cao T (2007) Effects of culture conditions and bone morphogenetic protein 2 on extent of chondrogenesis from human embryonic stem cells. Stem Cells 25(4):950–960 71. Bielby RC, Boccaccini AR, Polak JM, Buttery LD (2004) In vitro differentiation and in vivo mineralization of osteogenic cells derived from human embryonic stem cells. Tissue Eng 10(9–10):1518–1525 72. Vats A, Bielby RC, Tolley N, Dickinson SC, Boccaccini AR, Hollander AP, Bishop AE, Polak JM (2006) Chondrogenic differentiation of human embryonic stem cells: the effect of the micro-environment. Tissue Eng 12(6):1687–1697 73. Cohen S, Leshansky L, Zussman E, Burman M, Srouji S, Livne E, Abramov N, Itskovitz-Eldor J (2010) Repair of full-thickness tendon injury using connective tissue progenitors efficiently derived from human embryonic stem cells and fetal tissues. Tissue Eng Part A 16(10):3119–3137 74. Toh WS, Lee EH, Guo XM, Chan JK, Yeow CH, Choo AB, Cao T (2010) Cartilage repair using hyaluronan hydrogel-encapsulated human embryonic stem cell-derived chondrogenic cells. Biomaterials 31(27):6968–6980

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Chapter 3

Current Status of Induced Pluripotent Stem Cells Thach-Vu Ho, Grace Asuelime, Wendong Li, and Yanhong Shi

Abstract  The discovery of induced pluripotent stem cells (iPSCs) has “spiced up” the stem cell research field in the last few years. It has made tremendous progress in a very short time by demonstrating that adult fibroblasts could be reprogrammed into iPSCs using pluripotency factors. This suggested that cell fates are not as permanent as initially thought, but rather possess a degree of plasticity. Unsurprisingly, induced pluripotent stem cell technology still faces many technical obstacles before safe and high-quality human iPSCs can be generated for therapeutic applications. This chapter examines the current status of iPSC technology and new methods for ­inducing pluripotency and its use in modeling human disease.

Abbreviations iPSCs ESCs ICM SCNT MEFs OSKM

Induced pluripotent stem cells Embryonic stem cells Inner cell mass Somatic cell nuclear transfer Mouse embryonic fibroblasts Oct4, Sox2, Klf4, and c-Myc

Y. Shi (*) Department of Neurosciences, Center for Gene Expression and Drug Discovery, Beckman Research Institute of City of Hope, Duarte, CA, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_3, © Springer Science+Business Media, LLC 2011

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3.1 Introduction Stem cells are a unique population of cells that possess the ability to self-renew and differentiate into multiple lineages. ESCs are derived from the ICM of the blastocyst [1]. Cells isolated from ICM have the ability to differentiate into any cell type derived from the three germ layers (Fig. 3.1a), but are unable to generate extraembryonic tissues and thus lack the capacity to become a complete organism [2]. This restriction in cell potency suggests that ESCs are pluripotent but not totipotent. Nonetheless, ESCs can be differentiated into many tissue-specific cells, which hold promise for tissue and stem cell replacement therapies. Derivation of stem cells via SCNT or therapeutic cloning presents a potential advantage over ESCs for clinical use, since donor-derived stem cells could theoretically be used to treat human disease without fear of immune rejection. SCNT is the introduction of a nucleus from a somatic cell into an enucleated oocyte [3]. After inoculation, the oocyte is coaxed into becoming a fertilized embryo (Fig.  3.1b). ICM is removed at the blastocyst stage and is allowed to differentiate into tissue-specific cells in  vitro. Evidence indicates that SCNT stem cells are indistinguishable from ESCs, supporting the therapeutic potential of SCNT stem cells [4]. Unfortunately, ESCs and SCNT stem cells also carry limitations in therapeutic application. Evidence shows that ESC transplantation can evoke an immune response due to the lack of patient specificity [5]. While SCNT stem cells may be donor-specific, and therefore bypass immune rejection, SCNT stem cells have not yet been achieved in humans despite significant effort. It is also important to point out that the efficiency of SCNT is low due to epigenetic changes during the cloning process [4, 6]. Moreover, the use of SCNT to make ESCs requires the generation of possibly viable embryos that raises ethical concerns for many people. For these reasons, stem cell researchers have been exploring alternative approaches that would eliminate the use of human embryos. A major breakthrough in the stem cell field came when Takahashi and Yamanaka demonstrated the conversion of somatic cells to pluripotent stem cells using four transcription factors, OSKM [7] (Fig.  3.1c). Thereafter, several studies have ­verified that iPSCs could be derived using various combinations of transcription factors and have demonstrated the use of only one or two factors to induce ­pluripotency [8–13]. The ability of iPSCs to differentiate into multiple lineages and give rise to chimeric mice suggested that iPSCs are functionally similar to ESCs [14]. Moreover, iPSCs could lead to patient-specific cell lines, which would overcome concerns of immune rejection. The aim of this chapter is to review the current status of iPSC technology and discuss the challenges that iPSCs face before their application in the clinical setting.

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Fig. 3.1  A schematic overview of the generation of pluripotent stem cells. (a) Embryonic stem cells are derived from the inner cell mass of a fertilized egg in  vitro. (b) In somatic cell nuclear transfer, the nucleus of the somatic cell is inserted into an enucleated egg. The nucleus is stimulated to develop into a blastocyst. (c) Adult somatic cells are reprogrammed by overexpression of defined factors into induced pluripotent stem cells. Pluripotent stem cells can differentiate into tissue-specific cells of the three germ layers

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3.2 Molecular Profiling of iPSCs Since the advent of iPSCs, many studies have examined the properties of iPSCs, hoping to identify the key factors that regulate pluripotency and improve the reprogramming process [8, 15, 16]. Several studies have compared iPSCs and ESCs at the genomic level to determine whether iPSCs and ESCs are distinct at the molecular level [17–20]. While no study has yet proven that iPSCs are functionally equivalent to ESCs, many studies have shown that iPSCs may be very similar to ESCs [14]. To be ESC-like, iPSCs have to demonstrate self-renewal and several pluripotency criteria. The most stringent criterion of pluripotency is the yield of live organisms through tetraploid complementation [21]. Recent work demonstrated that iPSCs have passed the pluripotency criteria by generating live mice through tetraploid complementation assay [22–24]. While iPSCs exhibit ESC-like characteristics (Fig. 3.2), ESCs and iPSCs have subtle differences in gene expression profiles, epigenetic modification, and mitochondrial regulation [16–20]. Recent gene expression data demonstrated that iPSCs and ESCs may be distinguishable at the gene expression level [17, 20]. Data from Chin et al. [17] suggested that the gene expression profile of late-passage human iPSCs (beyond 37 passages) was highly correlated to human ESCs while the gene expression profile of earlypassage human iPSCs (less than 5 passages) was significantly different from that of human ESCs [17]. This observation suggests that reprogramming is a gradual process and that passage numbers must be taken into consideration when evaluating iPSCs.

Fig. 3.2  A general comparison between ESCs and iPSCs. Many studies have demonstrated that ESCs and iPSCs have common as well as distinct characteristics. The significance of these similarities and differences to tissue regeneration remain to be elucidated

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Human iPSCs are also distinguished from human ESCs at the epigenetic level [25, 26]. Experimental data suggested that the epigenetic variations observed between iPSCs and ESCs may have been caused by incomplete reprogramming of human somatic cells to human iPSCs [17, 19, 25]. In fact, the methylation pattern of CpG islands in human iPSCs can be significantly different from the parent somatic cells and ESCs [25]. This suggests that certain loci are incompletely reprogrammed. Ghosh et al. [19] proposed that human iPSCs have the ability to retain “transcriptional memory” of donor cells. Based on evidence from mice, Kim et al. [16] similarly suggested that iPSCs retain “epigenetic memory” of their tissue donors. DNA methylation is an important criterion for comparison of iPSCs and ESCs, and the state of histone H3 lysine 27 trimethylation and histone H3 lysine 4 trimethylation is commonly used to analyze chromatin modification in pluripotent stem cells. Through pairwise comparison of genes occupied by trimethylated histone H3 lysine 27 and lysine 4 in human iPSCs and human ESCs, Guenther et al. [20] found that human iPSCs were not significantly different from human ESCs for genes studied by this method. Data from Chin et al. [17] provided similar observation on histone H3 lysine 27 trimethylation within promoter regions and concluded that human ESCs and human iPSCs were nearly identical in their histone methylation pattern. Thus far, the question as to whether iPSCs and ESCs are equivalent is clear. They are not equivalent, but rather quite similar.

3.3 Progress in Reprogramming the Pluripotent State iPSC technology has exciting potential for disease modeling, drug development, and toxicity screening as previously mentioned. Nonetheless, the molecular mechanism of reprogramming remains elusive and it is only with further understanding of the cellular pathways involved in reprogramming that iPSC technology can possibly move from the bench to the clinic. In recent years, much effort has been focused on refining reprogramming methodologies. For example, iPSC technology needs to overcome the low reprogramming efficiency observed using Yamanaka’s original method, along with finding suitable viral-free methods for iPSCs’ derivation before moving toward clinical application. Viral-mediated transduction is the most used method so far for delivering the reprogramming transcription factors into somatic cells. One concern regarding the use of viral vector-mediated transduction is the integration of viral DNA into the host genome. Permanent integration of viral DNA can result in the development of cancer and also holds the possibility of being passed through the germ line [27]. As such, the expression of retroviral transgenes for OSKM may hold serious consequences as recently reported [13, 28]. Several groups, however, have successfully generated iPSCs using plasmids, piggyBac transposons, adenoviruses, recombinant proteins, and synthetic mRNAs [29–33]. Despite the fact that virus-free iPSCs can be generated, the efficiency remains extremely low (0.0001–0.1%). Reprogramming efficiency is influenced by many variables, some of which may be unknown. Several groups have attempted to remove or substitute the four transcription

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factors with small molecules [8, 10, 13]. However, it is important to note that certain factors such as Oct4 cannot be excluded, confirming that Oct4 is a key regulator of pluripotency. For example, Shi et al. [10] demonstrated that mouse embryonic fibroblasts could be induced to iPSCs by Oct4, Klf4, and small molecules, BIX and BayK. BIX and BayK are specific inhibitors of histone methyltransferase G9a and L-type calcium channel, respectively [10]. The authors concluded that BIX/BayK improved the reprogramming efficiency. Similarly, Huangfu et  al. [8] increased reprogramming efficiency (~1%) of human fibroblasts with a combination of Oct4, Sox2, and valproic acid, a histone deacetylase inhibitor. Recently, iPSCs were generated from MEFs with only Oct4 and small molecules (tranylcypromine, valproic acid, CHIR99021, and 616452) [13]. The Oct4-iPSCs were capable of germ line transmission in chimera formation, which was indicative of pluripotency. The authors argued that tranylcypromine (H3K4 demethylation inhibitor), valproic acid, CHIR99021 (GSK3-b inhibitor), and 616452 (TGF-b inhibitor) improved reprogramming efficiency by reducing epigenetic barriers [13]. Meanwhile, reprogramming efficiency was reported to improve as much as 50-fold with sodium butyrate treatment leading to a reprogramming efficiency of 15–20% [34]. In these studies, Mali and colleagues virally transduced IMR90 fibroblasts with OSKM followed by small-molecule treatment (RG108, BIX01294, valproic acid, sodium butyrate) 2 days after viral transduction. They observed that sodium butyrate facilitated epigenetic changes and stimulated pluripotency-associated genes in iPSCs. Sodium butyrate also efficiently stimulated reprogramming using a PiggyBac transposon delivery system [34]. More recently, human iPSCs were generated using Oct4 and chemical compounds, including sodium butyrate [35]. Small molecules may provide an important new path to reprogramming since these molecules can sufficiently replace the expression of one or more of the originally described reprogramming transcription factors.

3.4 iPSCs as Models of Disease The fundamental goal of regenerative medicine is to replace or restore normal function to damaged or aged and/or congenitally defective human cells, tissues, or organs. Animal models have been extensively used despite the fundamental differences and limitations compared to humans [36–38]. This raises concerns as to whether animal models accurately predict the effectiveness of the proposed therapies [37]. iPSCs have several advantages, including an abundant source of cells (self-renewal), ability to generate tissue-specific cell types (differentiation), and ability to generate autologous cell lines. Nonetheless, Saha and Jaenisch [39] reiterated the difficulty of using iPSCs to study human diseases in culture. The disease progression in the patient is much more dynamic than any model can possibly mimic. Therefore, Saha and Jaenisch suggested that the progression of the disease can be accelerated by exposing human iPSCs to environmental stimulus, such as oxidative stress. Moreover, modulating culture conditions to mimic the microenvironment similar in the patient may be essential when using iPSCs to model human diseases.

3  Current Status of Induced Pluripotent Stem Cells Table 3.1  Disease-specific iPSCs Disease Sickle cell anemia Amyotrophic lateral sclerosis Adenosine deaminase deficiency-related severe combined immunodeficiency Shwachman-Bodian-Diamond syndrome Gaucher disease type III Duchenne and Becker muscular dystrophy Parkinson’s disease Huntington disease Juvenile-onset, type 1 diabetes mellitus Down syndrome (trisomy 21) Lesch-Nyhan syndrome Acute myocardial infarction Spinal muscular atrophy Fanconi anemia Myeloproliferative disorder LEOPARD syndrome Angelman and Prader–Willi syndromes Rett syndrome Fabry disease Globoid cell leukodystrophy Mucopolysaccharidosis VII Long QT syndrome Inherited metabolic disorders of the liver and other liver diseases Lung diseases Fragile X syndrome Bombay blood group Vascular disease

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Stem cell origin Mouse Human Human

Reference [59] [60] [61]

Human Human Human Human, rats Human, monkey Human Human Human Mouse Human Human Human Human Human Human Mouse Mouse Mouse Human Human

[61] [61] [61] [48, 61–63] [61, 64] [61, 65] [61] [61] [66] [67] [68] [69] [41] [47] [70] [71] [71] [71] [72] [73, 74]

Human Human Human Human

[75] [76] [77] [78]

To date, iPSC technology has already been used to generate disease-specific iPSC lines (Table 3.1). For example, LEOPARD syndrome, also known as multiple lentigines syndrome, is an autosomal dominant condition [40] caused by a missense mutation in the nonreceptor protein tyrosine phosphatase type 11 gene [41]. The clinical manifestations in LEOPARD syndrome patients may include hypertrophic cardiomyopathy, multiple lentigines, and deafness [42]. Two common mutations associated with LEOPARD syndrome are Y279C and T468M. Clinical testing for LEOPARD syndrome involves sequence analysis of the nonreceptor protein tyrosine phosphatase type 11 gene after clinical manifestations have been established. Currently, there is no cure for LEOPARD syndrome. Recent studies have established human iPSC lines with a heterozygous T468M mutation from two patients [41]. Carvajal-Vergara and colleagues derived human iPSC lines from the fibroblasts of two LEOPARD syndrome patients using OSKM. They demonstrated that these human iPSC lines were able to differentiate into three germ layers. Hypertrophic cardiomyopathy is manifested in a majority of LEOPARD

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syndrome patients. Data from iPSC studies revealed that human iPSCs were able to differentiate into hematopoietic and cardiac lineages and demonstrated that LEOPARD syndrome iPSC-derived cardiomyocytes showed similar hypertrophic features when compared to wild-type iPSC-derived and human ESC-derived cardiomyocytes [41]. The generation of functional cardiomyocytes from human iPSCs also supports their potential for cardiomyocyte replacement therapy for myocardial infarction, which is characterized by the loss of cardiomyocytes due to the imbalance of blood supply [43]. However, the current inefficient processes for reprogramming patient-specific cells have to be improved before this becomes a therapeutic reality. Neurodegenerative diseases represent another area of research that requires disease-specific models to study the mechanism underlying the disease. Most neurodegenerative diseases in humans, including Angelman and Prader–Willi syndromes, are caused by genetic mutations. Angelman syndrome is characterized by a mutation in the UBE3A gene that is expressed only from maternal chromosomes. In contrast, Prader–Willi syndrome is distinguished by the loss of paternal expression of SNORD116 snoRNAs [44]. Studies demonstrated that both Angelman and Prader– Willi syndromes are caused by the deletion or lack of expression of seven genes on chromosome 15q11–15q13 [45]. Clinical manifestations of Angelman syndrome include developmental delay, movement disorder, speech disorder, and behavior problems [46]. Children with Prader–Willi syndrome, on the other hand, experience obesity, hypogonadism, short stature, and mental retardation [45]. Currently, there is no specific treatment for patients with Angelman and Prader–Willi syndromes and no autologous disease model with which to test potential therapies. Recently, Chamberlain et  al. [47] generated human iPSCs from children with Angelman and Prader–Willi syndromes using retroviral vectors expressing OSKM and LIN28. Additionally, they assessed DNA methylation to screen for epigenetic changes in the resulting iPSC lines. Both Angelman and Prader–Willi syndrome human iPSC lines showed normal DNA methylation patterns compared to control human iPSCs. The group also demonstrated that Angelman syndrome iPSCs differentiated into functional neurons in  vitro. This is an important step toward designing an Angelman syndrome human iPSC model. They also explored the regulatory mechanism of UBE3A in Angelman syndrome human iPSCs. Within the last few years, several iPSC models have been derived from patient fibroblasts, including ones for Parkinson’s disease [48], familial dysautonomia [49], and several others, to study the mechanisms and explore novel compounds for treatment. However, many more human diseases await mechanistic elucidation using iPSCs. Another application for iPSC technology is in drug efficacy and toxicity screening. Current technologies allow screening and evaluating thousands of compounds to target human diseases [50]. Increased knowledge of biological networks allows pharmaceutical companies to narrow the number of targets and relevant compounds for specific human diseases [51, 52]. However, nonhuman models used for current screening provide inefficient reaction mimicry of the compounds in humans [36, 50]. Disease-specific iPSC lines would alleviate this problem by allowing direct analysis of compound efficacy in a human system. For example, Lee et  al. [49]

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derived patient-specific familial dysautonomia iPSCs to study the disease mechanism. Familial dysautonomia is caused by gene mutations in the IkB kinase complexassociated protein, which results in the degeneration of sensory neurons [53]. Lee and colleagues used three parameters, mutant IkB kinase complex-associated protein splicing, neurogenesis, and diseased iPSC-derived neural crest precursor function, to monitor the effects of drug treatment. This study suggested that the potential of disease-specific iPSC model for drug screening is feasible.

3.5 Bypassing the Pluripotency State In recent years, several groups have explored the idea of direct conversion (or induced transdifferentiation) to generate tissue-specific cells. For example, rat exocrine pancreatic cells were treated with leukemia inhibitory factor and epidermal growth factor in vitro to generate insulin-producing beta cells [54]. Shortly after, Zhou and colleagues found that the expression of transcription factors Ngn3, Pdx1, and Mafa could reprogram pancreatic exocrine cells to b-cells in adult mice in  vivo [55]. Recently, mouse tail fibroblasts were converted to neurons capable of generating action potentials and forming functional synapses using the transcription factors Ascl1, Brn2 (also known as Pou3f2), and Myt1l [56]. Most recently, Ieda et al. [57] examined whether the key regulators of cardiac development could directly convert cardiac fibroblasts into cardiomyocytes. They found that three transcription factors, Gata4, Mef2c, and Tbx5, were able to convert mouse cardiac fibroblasts and dermal fibroblasts into cardiomyocytes in vitro and in vivo while maintaining a global gene expression pattern intermediate between the ICM and endogenous cardiomyocytes. They also demonstrated that the transdifferentiation of functional beating cardiomyocytes was more rapid and up to 20% more efficient then iPSC reprogramming [57]. While these groups successfully provided evidence for the potential of transdifferentiated cells in animal models, Szabo et al. [58] were the first to illustrate this concept in human cells. Through the transduction of human fibroblasts with lentivirus expressing Oct4 and cultured with cytokine supplements known to support hematopoietic progenitor development, they were able to derive multipotent hematopoietic progenitors that give rise to the myeloid, erythroid, and megakaryocytic lineages [58]. Concurrently, they verified that the conversion does not require passage through a pluripotent stem cell state [58]. To date, the advantages of using directly converted cells for therapy include, but are not limited to, higher induction efficiency and lower risk of tumorigenesis. It is important to consider that further verification of the similarities between transdifferentiated cells and primary tissue-specific cells is required, as well as an appreciation that unlike stem cells, adult cells cannot easily be expanded and may require larger numbers of primary cells.

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3.6 Conclusion iPSCs hold much promise for therapy, as well as for drug screening and disease modeling, especially when one considers the ethical and immunological complications associated with ESC and SCNT use. Currently, iPSC technology is at an early stage of development, but is making rapid progress. The future of iPSC technology has enormous potential, but as with many approaches using stem cells, more work is needed to bring iPSC technology to therapeutic application. Acknowledgments  We apologize to colleagues whose work could not be cited due to space limitations. T.H. is supported by a stem cell research internship program of the California Institute for Regenerative Medicine and California State University at Long Beach. W.L. is supported by a postdoctoral fellowship from the California Institute for Regenerative Medicine. Y.S. is supported by the National Institutes of Health/NINDS (R01 NS059546 and RC1 NS068370) and the California Institute for Regenerative Medicine (TR2-01832).

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71. Meng XL, Shen JS, Kawagoe S, Ohashi T, Brady RO, Eto Y (2010) Induced pluripotent stem cells derived from mouse models of lysosomal storage disorders. Proc Natl Acad Sci USA 107:7886–7891 72. Moretti A, Bellin M, Welling A, Jung CB, Lam JT, Bott-Flugel L, Dorn T, Goedel A, Hohnke C, Hofmann F, Seyfarth M, Sinnecker D, Schomig A, Laugwitz KL (2010) Patient-specific induced pluripotent stem-cell models for long-QT syndrome. N Engl J Med 363:1397–1409 73. Rashid ST, Corbineau S, Hannan N, Marciniak SJ, Miranda E, Alexander G, Huang-Doran I, Griffin J, Ahrlund-Richter L, Skepper J, Semple R, Weber A, Lomas DA, Vallier L (2010) Modeling inherited metabolic disorders of the liver using human induced pluripotent stem cells. J Clin Invest 120:3127–3136 74. Ghodsizadeh A, Taei A, Totonchi M, Seifinejad A, Gourabi H, Pournasr B, Aghdami N, Malekzadeh R, Almadani N, Salekdeh GH, Baharvand H (2010) Generation of liver disease-specific induced pluripotent stem cells along with efficient differentiation to functional hepatocyte-like cells. Stem Cell Rev 6:622–632 75. Somers A, Jean JC, Sommer CA, Omari A, Ford CC, Mills JA, Ying L, Sommer AG, Jean JM, Smith BW, Lafyatis RA, Demierre MF, Weiss DJ, French DL, Gadue P, Murphy GJ, Mostoslavsky G, Kotton DN (2010) Generation of transgene-free lung disease-specific human induced pluripotent stem cells using a single excisable lentiviral stem cell cassette. Stem cells 28:1728–1740 76. Urbach A, Bar-Nur O, Daley GQ, Benvenisty N (2010) Differential modeling of fragile X syndrome by human embryonic stem cells and induced pluripotent stem cells. Cell Stem Cell 6:407–411 77. Seifinejad A, Taei A, Totonchi M, Vazirinasab H, Hassani SN, Aghdami N, Shahbazi E, Yazdi RS, Salekdeh GH, Baharvand H (2010) Generation of human induced pluripotent stem cells from a Bombay individual: moving towards “universal-donor” red blood cells. Biochem Biophys Res Commun 391:329–334 78. Freund C, Davis RP, Gkatzis K, Ward-van Oostwaard D, Mummery CL (2010) The first reported generation of human induced pluripotent stem cells (iPS cells) and iPS cell-derived cardiomyocytes in the Netherlands. Neth Heart J 18:51–54

Chapter 4

Mesenchymal Stromal Cells: Latest Advances Sowmya Viswanathan and Armand Keating

Abstract  Over the past decade, the study of mesenchymal stromal cells (MSCs) has moved rapidly from in vitro and animal models to randomized clinical trials. Despite the challenges of defining MSCs, a consensus has emerged on culture methodology and their characterization, including the requirement for a minimum immunophenotype. Mechanisms of action in tissue regeneration have matured from the simple notion of transdifferentiation to effects on endogenous progenitors and promotion of an anti-inflammatory environment. Clinical investigation with MSCs now covers a wide variety of diseases, and sources of MSCs other than the bone marrow continue to be identified. Genetically engineered MSCs may provide more effective agents of tissue regeneration but will require careful preclinical study. Nonetheless, challenges remain: the need for appropriate preclinical models, informative clinical trials, good manufacturing practice cell production, and long-term trials follow-up.

Abbreviations ALS AMI AT AT-MSC

Amyotrophic lateral sclerosis Acute myocardial infarction Adipose tissue Adipose tissue-derived mesenchymal stromal cell

A. Keating (*) Cell Therapy Program, Princess Margaret Hospital, University Health Network, ON, Canada Department of Medicine, University of Toronto, Toronto, ON, Canada e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_4, © Springer Science+Business Media, LLC 2011

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bFGF BDNF BM BM-MSC CD CDAI DC Dkk-1 DMSO EAE FBS FGF GvHD GM-CSF GMP HGF HLADR HLA-G5 HUCPVCs IBD IDO IGF-1 LIF MI MHC MSCs NGF NK NO NT-3 PD PDGF PGE2 SDF-1 SSEA-3/4 TGF-b T-reg TSG-6 UC UCB VCAM-1 VEGF

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Basic fibroblast growth factor Brain-derived neurotrophic factor Bone marrow Bone marrow-derived mesenchymal stromal cell Crohn’s disease Crohn’s disease activity index Dendritic cell Dickkopf-1 Dimethyl sulfoxide Experimental autoimmune encephalomyelitis Fetal bovine serum Fibroblast growth factor Graft-versus-host disease Granulocyte macrophage colony-stimulating factor Good manufacturing practice Hepatocyte growth factor Human leukocyte antigen DR Human leukocyte antigen G5 Human umbilical cord perivascular cells Inflammatory bowel disease Indoleamine-pyrrole 2,3-dioxygenase Insulin-like growth factor-1 Leukemia inhibitory factor Myocardial infarction Major histocompatibility complex Mesenchymal stromal cells Nerve growth factor Natural killer Nitric oxide Neurotrophin-3 Parkinson’s disease Platelet-derived growth factor Prostaglandin E2 Stromal cell-derived factor-1 Stage specific embryonic antigen-3/4 Transforming growth factor-beta Regulatory T cell Tumor necrosis factor inducible gene-6 protein Umbilical cord Umbilical cord blood Vascular cell adhesion molecule-1 Vascular endothelial growth factor

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4.1 MSC Definition Since the first published report by Friedenstein and colleagues [1] describing the expansion of an adherent, spindle-shaped population of cells from whole human BM, MSCs have been expanded from numerous sources including skeletal muscle, AT, UC, synovium, dental pulp, amniotic fluid, and other tissues [2]. There has been a tendency in the field to identify this heterogeneous population with different terminology, including multipotent stromal cells, mesenchymal stem cells, marrow stromal cells or MSCs, without rigorous discrimination of their “stemness” properties [3]. For this reason, we prefer the term “mesenchymal stromal cells” (the acronym MSC still applies) as recommended in a white paper from the International Society for Cell Therapy [4]. Here, we define MSCs according to the minimum criteria established by the International Society for Cellular Therapy [5], i.e., greater than 95% cells must express CD105, CD73, and CD90, as measured by flow cytometry, and less than 2% should be positive for CD45, CD34, CD14 or CD11b, CD79a or CD19 and HLA Class II, and must be able to differentiate into osteoblasts, adipocytes, and chondroblasts under standard in  vitro differentiating conditions. It is reassuring that McGonagle et  al. [6] have demonstrated that MSCs are not merely an in  vitro manifestation of an unknown cell in vivo by detecting primary cells in the BM with an immunophenotype indistinguishable from that of cultured MSCs: CD45lo, CD271+, CD105+, CD90+, and CD10+. The origin of MSCs remains unproven. There are several hypotheses, some suggesting that MSCs may be skeletal stem cells [7, 8], others suggesting an embryonic remnant of pluripotent cells [9, 10], and still others suggesting a neural crest origin [11]. It is especially intriguing that MSCs bear a resemblance in immunophenotype, location and function to pericytes [12–14] and thus would be expected to be present in all vascularized tissues, especially during inflammation or injury. This change in paradigm of MSCs from self-renewing multilineage precursors to pleiotropic pericytes is the subject of ongoing investigation.

4.2 Sources of MSCs 4.2.1 Adipose Tissue In the adult, MSCs have been isolated outside the BM in AT [15]. This highly complex tissue contains adipocytes, pre-adipocytes, fibroblasts, vascular smooth cells, endothelial cells, monocytes, macrophages, and lymphocytes [16], stores and provides energy, and also serves as a dynamic hormone-producing organ involved in a number of physiological and pathological processes. AT-MSCs are

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readily available in large amounts (100 mL to 1 L) with relatively little morbidity through liposuction procedures [17]. MSCs are found at a higher frequency in AT than in BM, ranging from 1–10 per 1,000 cells [18] to 25–50 per 1,000 cells [19]. AT-MSCs also exhibit higher proliferation rates than BM-MSCs [20]. Comparative analyses of MSCs from BM and AT show that there is little difference in terms of morphology or immunophenotype [21]. Similar to their BM counterparts, AT-MSCs differentiate into multiple cell types including adipocytes, myocytes, osteocytes, chondrocytes, hepatocytes, neurons, pancreatic cells, endothelial cells, and cardiomyocytes (reviewed in [22]). Similar to BM-MSCs, AT-MSCs are immunosuppressive [23], lack HLA-DR expression and can therefore be used therapeutically in allogeneic transplantation with low risk of immune-mediated rejection. AT-MSCs exhibit similar cell surface antigens to BM-MSCs [20, 24] and secrete growth factors such as VEGF, HGF, and IGF-1 [25].

4.2.2 Umbilical Cord Blood The UC contains two arteries and one vein, which are surrounded by mucoid connective tissue called Wharton’s Jelly. Nonhematopoietic cells have been isolated from the connective tissue of the cord using different isolation and enzymatic processing techniques by many groups [26–28]. UCB-MSCs are a multipotent stromal cell population that are plastic-adherent and share BM-MSC surface markers such as CD73, CD90, and CD105 (reviewed in [29]), but have lower expression of CD106 and HLA-DR. UCB-MSCs also exhibit pluripotent transcription factors such as Oct-4, Nanog, and Sox-2 but several magnitudes lower than that expressed in embryonic stem cells [30]. UCB-MSCs also express Rex-1, SSEA-3, SSEA-4, Tra-1-60, and Tra-1-81 [31]. Additionally, UCBMSCs, like myofibroblasts, express vimentin, desmin, and/or alpha-smooth muscle actin [29, 32, 33]. Isolation of this rare cell (1:200 million) is found in only an average of 29% of cords although this can be improved to 60% by optimizing the isolation process [21]. UCB-MSCs proliferate more rapidly compared with BM-MSCs [34] and maintain expansion and differentiation properties longer in culture [35].

4.2.3 Human Umbilical Cord Perivascular Cells UCB-MSCs are a heterogeneous group, coming from different zones of the connective tissue (subamniotic, intervascular, and perivascular) and exhibiting overlapping but distinctive features [36]. A distinct population of MSCs are obtained by enzymatic digestion of UC perivascular tissue [37]. These cells, HUCPVCs, can be isolated with high efficiency (100%) compared with UCB and have high clonogenic potential with a CFU-F frequency of 1:300 [37]. HUCPVCs exhibit multilineage differentiation potential in vitro and in vivo as recently confirmed at the clonal level [38]. They express cell surface markers similar to BM-MSCs but have higher levels

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of CD146 [39]. HUCPVCs, like UCB-MSCs, exhibit a higher proliferation rate and can be propagated in culture for longer periods than BM-MSCs [37].

4.3 Immune Modulation MSCs have inherently low immunogenicity, lacking MHC Class I and co-­ stimulatory molecules for T-cell recruitment [40], rendering them safe for mismatched allogeneic transplantation. Conflicting evidence [41], however, suggests that allogeneic, but not syngeneic, gene-modified murine MSCs may elicit an immune response in immunocompetent mice [42]. MSCs are generally considered immunosuppressive and numerous mechanisms have been proposed. MSCs secrete soluble factors such as IDO [43], NO [44], TGF-b1, HGF [45], PGE2 [46], HLA-G5 [47], LIF, and IL-10 (reviewed in [48]) to exert their immunosuppressive effects in a systemic manner. Best studied is the suppression of T-cell proliferation, although the exact mechanism of inhibiting this proliferation is not fully understood with some groups suggesting MSC-mediated cell-cycle arrest in G1 phase while others argue for MSC-mediated apoptosis [49–52]. MSCs have also been shown to recruit and support proliferation of regulatory T-cells [53], which in turn inhibit T-cell proliferation and cytokine production. MSCs also mediate the ­suppression of other cell types such as NK cells, although they can also be targets of NK cell killing [54]. MSCs further inhibit DC proliferation [55] and maturation [56] and can decrease production of pro-inflammatory cytokines [46]. MSCs also inhibit monocyte differentiation and modulate macrophage activity [57]. High doses of MSCs appear to inhibit B-cell proliferation and differentiation through paracrine action, while cell–cell contact increases antibody production [58]. MSC-mediated immunosuppression has been demonstrated in vivo in disease animal models, including experimental autoimmune encephalomyelitis [59, 60] and allogeneic skin grafts [61]. In vivo administration leads to inhibition of pathogenic antibodies, due to metalloproteinase processing of CCL2 produced by MSCs [60, 62]. Xenogeneic infusions of human MSCs are not immunologically recognized by immunocompetent rodents and can account for dramatic improvements in models of MS, stroke, colitis, IBD, and MI. The ability of MSCs to inhibit T-cell responses in a non-MHC-restricted manner [45] has led to the successful treatment of steroid-resistant acute GvHD [63]. However, the mechanisms of immune suppression by MSCs in patients are not well understood, in part, because the in vivo fate of MSCs is poorly documented.

4.4 MSC Therapy for Autoimmune Diseases Because of their immunosuppressive properties and demonstration of safety and feasibility in GvHD, there is increased interest in MSCs for treating autoimmune and inflammatory disorders, including Crohn’s disease, diabetes, and MS.

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4.4.1 Crohn’s Disease The immunosuppressive properties of MSCs make them well suited for inflammatory diseases such as IBD and CD. In a murine model of IBD, AT-MSCs reduced systemic and mucosal levels of pro-inflammatory cytokines, increased IL-10 secretion, induced T-regs in mesenteric lymph nodes, and reduced microscopic signs of colitis [64]. In a pilot study, MSCs isolated from patients with refractory CD had intact immunosuppressive properties [65]. MSCs were expanded and infused into nine patients; three showed a clinical response as measured by CDAI but in three others disease progressed requiring surgery. This pilot study showed no apparent benefit for patients with severe refractory luminal disease. AT-derived cells encapsulated in fibrin glue were used to treat complex perianal fistula associated with CD (n = 14) and those of cryptoglandular origin (n = 35) in a Phase II trial [66]. There was a 71% response rate (fistula closure) in the treatment group compared with a 16% response in the group receiving fibrin glue alone. In a Phase II trial sponsored by Osiris Therapeutics, two infusions of allogeneic MSCs were administered at 2 or 8 × 106 cells/kg in nine moderate-to-severe CD patients. This led to a reduction in CDAI in all nine by day 28; three of nine patients had a clinical response determined by a reduction in CDAI by 150 [67]. There was no correlation of dose with outcome. Based on these results, a Phase III trial has been initiated with 207 patients with a CDAI between 250 and 450, and who have previously failed therapy with at least one steroid, an immunosuppressant, and a biological agent. The primary endpoint for this trial is disease remission, defined as a CDAI at or below 150 by day 28.

4.4.2 Diabetes Mellitus Human MSCs can improve diabetes although the mechanism is not well understood [68]. For example, paracrine secretion is implicated in a study of type 1 diabetic mice in which infusion of MSCs resulted in a significant reduction of blood glucose within 1 week that reached near euglycemic values a month later. These animals showed an increase in morphologically normal beta-pancreatic islets and normal glomeruli, suggesting therapeutic potential of these cells [69]. In contrast, other studies report the derivation of insulin-secreting cells by differentiating MSCs isolated from UCB. Blood sugar levels were reduced upon xenotransplantation into NOD mice and the histological presence of insulin-secreting cells with human nuclei and C-peptide was detected in the liver [70]. Another strategy for insulin treatment involves electroporating MSCs with the endogenously active glucoseresponsive promoter, EGR1 [71]. Mice receiving modified human MSCs exhibited dose-responsive corrections of hyperglycemia, improved glucose tolerance, and reduced body weight. In a clinical trial, 25 patients with type II diabetes received

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concomitant intrapancreatic autologous BM stem cell infusions and hyperbaric oxygen treatment and showed improved metabolic control and reduced insulin requirements 1 year later [72].

4.4.3 Multiple Sclerosis Preclinical evaluation in EAE models shows that MSCs can suppress clinical manifestations [59], although the effect was apparent only when cells were injected at disease onset or peak and not during the chronic phase. Rodents exhibiting EAE receiving intravenous and intraventricular infusions of MSCs had almost twice the number of axons as control animals [73]. A case report using intrathecal and intravenous infusions of UCB-MSCs to treat a patient with MS showed improvements in sensory impairment, an expanded disability status scale score, and a reduced T2 lesion load by MRI, with no side effects [74]. A pilot study of autologous MSCs in ten patients similarly demonstrated safety and feasibility [75].

4.5 MSC Therapy for Neurodegenerative Diseases MSCs are highly interactive with their microenvironment, and share protein, RNA, and mitochondria with damaged tissue, which may be particularly relevant in treating neurodegenerative diseases [76]. Increasing numbers of preclinical and clinical studies are being reported investigating the effects of MSCs on diseases such as ALS, PD and on acute and chronic ischemic stroke, and spinal cord injuries. MSCs have been shown to stimulate the proliferation, migration, and differentiation of endogenous neural stem cells [77], promote neuronal survival [78] and neurite outgrowth [79], and protect neurons against oxidative stress [80] through the secretion of soluble factors such as BDNF [81], Wnt antagonist Dkk-1 [82], and NGF. Trophic factors secreted by MSCs promote neuronal growth while modifying the tissue microenvironment [79]. MSCs also promote oligodendrogenesis [83]. In animal studies, transplantation of BM-MSCs into the brain of immunodeficient mice markedly increased the proliferation of endogenous neural stem cells [77]. The common theme of these studies is that the effect of MSCs lies in stimulating endogenous cells to enhance the repair of neural tissue.

4.5.1 Amyotrophic Lateral Sclerosis Several groups have demonstrated that intraparenchymal delivery of human MSCs is safe and can delay loss of motor neurons in rodents [84]. Human MSCs

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t­ ransplanted directly into the spinal cords of transgenic SOD1 mice [85] migrated throughout the spinal cord and delayed loss of motor neurons, prolonging motor performance. A pilot study in ALS was conducted by transplanting autologous, cultureexpanded MSCs into a surgically exposed spinal cord (T7–T9 level). No significant side effects were reported [86]. Magnetic resonance imaging performed 3 and 6 months after transplantation did not show structural changes of the spinal cord or abnormal cell proliferation when compared with baseline scans. Three months after cell implantation, four patients exhibited a significant slowing of muscle strength decline in the proximal muscle groups of lower limbs. A Phase II clinical trial using MSCs is underway in Europe, and the FDA has recently approved a Phase I trial in the USA.

4.5.2 Parkinson’s Disease The goal of cellular therapy of PD is to replace lost neurons in the substantia nigra with healthy dopaminergic neurons or to prevent further neuronal loss. In a rat model of PD, MSC transplantation resulted in behavioral changes that correlated with partial restoration of dopaminergic markers and vesicular striatal pool of dopamine [87]. MSCs engineered to express neurotrophic factors may be superior, since BDNF-modified MSCs transplanted into a 6-hydroxydopamine-induced lesion model of PD showed behavioral improvements and reduced dopamine depletion compared to unmodified MSCs [88]. Recently, Venkataramana et  al. [89] reported the first open-label clinical pilot study with a single dose of autologous MSCs transplanted by stereotaxic surgery into the striatum of seven patients with advanced PD. At 10–36 months follow-up, there were no safety concerns or serious adverse events reported, but no efficacy outcomes could be concluded.

4.5.3 Stroke MSCs have been used in the treatment of experimental stroke [90–93] and exert an effect in rats even when administered 1 month after the stroke [92]. However, very few transplanted cells were detected during the 1-year tracking experiment, suggesting that MSC differentiation and replacement of neurons is an unlikely primary mechanism [92]. Secretion of factors from MSCs including BDNF, NT-3, VEGF, NGF, bFGF, and IGF-1 may play a role and promote functional repair [90, 91]. This was confirmed in a recent study that showed MSC-treated grafts had higher levels of BDNF, NT-3, and VEGF compared with saline control grafts, which accelerated proliferation of neuronal progenitors in the subventricular zone [94].

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A single randomized study of culture-expanded MSCs was conducted in 30 patients with cerebral infarcts within the middle cerebral arterial territory, and severe neurological deficits [95]. Only five received autologous 108 culture-expanded MSCs intravenously while 25 other patients served as controls. Outcomes improved in the MSC-treated patients compared with control patients as measured by the Barthel Index and Rankin Score at 3 and 6 months, but differences were not statistically significant at 12 months. No significant toxicities were identified. Clinical trials of a MSC-like multipotent cellular product, MultiStem®, were recently approved by the FDA to treat stroke in the USA.

4.5.4 Spinal Cord Injuries Many studies have documented successful engraftment of MSCs into the injured spinal cord [96]. Some evidence suggests that MSCs may reduce the acute inflammatory response to spinal cord injury and decrease astrocyte, microglia, and macrophage reactivity [96]. In 2005, Park et al. [97] reported the first trial of MSCs in six patients with spinal cord injury and showed that autologous whole BM injected directly into the site of spinal cord injury, along with intravenous infusions of GM-CSF slightly improved neurological function in five patients. The same researchers later treated 18 patients with MSCs and 13 patients in a control group (decompression and spinal fusion surgery) [98]. Thirty percent of patients who had received cells during the acute injury stage, 33% during the chronic injury stage, and 8% of the control group demonstrated an increase in neurological function. In another study, MSC injection into the cerebrospinal fluid produced improvement in the quality of life score only for patients with acute but not chronic injuries [99].

4.6 MSC Therapy for Cardiovascular Disease MSCs have been shown to mediate functional improvement in a number of ­animal models of cardiac injury, including a porcine AMI model [100], a canine chronic myocardial ischemia model [100, 101], a pig heart failure model [102], and a rodent model of dilated cardiomyopathy [103]. Mechanisms mediating this repair, however, remain unclear; early reports claimed that MSCs transdifferentiate into a cardiomyocyte phenotype [104–107] although differentiation to a mature functioning cardiomyocyte has not been demonstrated. Other groups citing evidence of low engraftment have proposed paracrine mechanisms [108, 109] that include recruitment of resident cardiac stem cells [110], inhibition of fibrosis [111], cardioprotection (likely through secretion of HGF, TGF-b, VEGF, IGF-1, stanniocalcin 1, and GM-CSF), promotion of neoangiogenesis [112], and improvement in cardiac contractility [113]. Another hypothesis is

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that MSCs, which only appear to be transiently present in the infarcted myocardium, modulate the inflammatory microenvironment of ischemic tissue by up-regulation of a number of genes including TSG-6 [114], an anti-inflammatory protein produced by monocytes, macrophages, and DCs [115]. This was confirmed when the beneficial effect of MSCs was abrogated in siRNA knock-down of TSG-6 and rescued by infusion of recombinant TSG-6. Clinical studies involving interventional delivery [intramyocardial (PROMETHEUS trial) and transendocardial (TAC-HFT)] of autologous MSCs are currently ongoing. The only trial involving systemic delivery of allogeneic MSCs is the Provacel trial by Osiris Therapeutics. They reported a randomized Phase I trial of allogeneic MSCs administered intravenously in 53 patients after AMI [116]. There were comparable adverse events between the MSC-treated and placebo groups and no serious events related to cell treatment at 2 years. Unlike preclinical studies [117], the MSC-treated group reported improved pulmonary function. At 6 months however, there was no difference in ejection fraction between MSC-treated versus placebo groups by echocardiography, but analysis of a subgroup using MRI revealed significant persistent improvement at 12 months. A larger Phase II study of 220 patients is underway. MSCs have been evaluated in small numbers of patients with administration via the intracoronary route [118]. In all trials, preliminary data showed improvement in ejection fraction and perfusion defects in MSC-treated groups, although controls were not included in early studies.

4.7 Manufacturing Considerations Culture expansion of MSCs from BM, AT, UC, fetal tissue, and other sources is required for clinical use because the cells are present in very low frequencies in these tissues. GMP is required to produce clinical-grade MSCs to meet multiple morphological, immunophenotypic, functional, karyotypic, safety, and sterility criteria prior to infusion into patients. A common protocol for the expansion of MSCs from the European Group for Blood and Marrow Transplantation has since been adopted by many groups participating in multicenter trials [119]. A tenet of GMP is the use of certified, pathogen-free reagents. Despite remarkable clinical advancements in this field, MSCs are still expanded in traditional culture media containing FBS [119]. This can be concerning as MSCs have been shown to retain FBS proteins and increase the risk of sensitization [120]. Platelet lysates represent an efficient alternative to FBS as demonstrated by many groups [121, 122] although there may be some limiting effects on the functional behavior of differentiated MSCs [123]. MSCs expanded in platelet lysate were administered to GvHD patients; although there were no safety concerns, a lower overall response of 54% was reported at a 28-day time point [124]. It is not clear whether this reduced efficacy is related to the design of the clinical protocol or due to the FBS substitution.

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Human MSCs can also be expanded in serum-free medium supplemented with a cocktail of factors, including recombinant human PDGF-BB, bFGF, and TGF-b1 and appear to retain their immunophenotypic, colony-forming, and differentiation potential [125]. However, exposure to growth factors may induce MHC Class II expression and cause aneuploidy [126]; consequently, follow-up is needed to assess the long-term safety of these defined culture conditions. Production of xenogeneic contaminant-free adhesion proteins (e.g., fibronectin, lamenin, vitronectin) to facilitate attachment of MSCs to culture surfaces also poses a challenge. GMP-grade MSCs used for clinical trials are typically early passage, although there are no significant differences between MSCs passaged up to P7 in terms of immunosuppressive properties [127]. Early-passage MSCs are still preferred for increased safety and efficacy as continuous culture of MSCs over several passages may result in the accumulation of karyotypic abnormalities [128]. A confounding factor is how passage numbers are really measured in different systems with different seeding densities and definitions of confluency. Population doublings are a more instructive term, as we have proposed [129] and MSCs expanded for fewer population doublings (<25–30) should be used clinically. To generate large doses of MSCs for clinical use, approximately 108 cells per patient [130, 131], a larger starting volume of source material (e.g., 100–200 cm3 of BM aspirate) or bioreactors can be used to achieve the requisite clinical dose in fewer population doublings. MSCs are traditionally grown in static cultures in nonscalable plates, bags, or flasks, which limit production to subtherapeutic dose range for most applications. Scalable bioreactor systems, including spinner flasks or rotating-wall bioreactors, can generate eightfold more expansion of MSCs than static cultures [132]. Importantly, it has been demonstrated that MSCs grown in these suspension cultures retained the standard immunophenotypic markers including primitive adhesion proteins, colony formation ability, and tri-lineage differentiation potential [132]. In our hands, using a different cocktail of cytokines in a stirred-suspension bioreactor, we observed that while MSCs retained their immunophenotypic markers, they exhibited lower levels of adhesion molecules such as VCAM-1 and CD44 [133]. Other groups have preserved MSC adhesion properties without compromising scalability by culturing the MSCs on microbeads in suspension cultures. Rat MSCs grown on gelatin-coated beads appeared to retain tri-lineage differentiation potential, although full functionality was not tested in this study [134]. These results have been replicated for human MSCs grown on different microcarriers [135]. Importantly, using microbeads would eliminate the need for constant trypsinization as the MSC-loaded restorable microbeads may be directly transplanted.

4.8 Genetically Modified MSCs The intrinsic ability of MSCs to home to sites of inflammation, tissue injury, and tumors, to interact with their microenvironment, along with their capacity to be transduced by adenoviruses, retroviruses [136], adeno-associated viruses [137], and

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lentiviruses make them excellent cellular vehicles for gene therapies. A 10-year study has documented the safety and adverse events of genetically engineered MSCs in vivo [138]. Genetically engineered MSCs can be used for the functional repair of tissuespecific diseases, for instance to express antiapoptotic factors (e.g., Bcl-2) to improve cell survival [139] or angiogenic factors and chemokines (VEGF and SDF-1) to improve MSC homing and angiogenesis [140]. Over-expressing growth factors such as BDNF ameliorates disease progression in a neurodegenerative disease model [141]. Genetically modified MSCs need to be considered with caution even though current research is focused on inserting genetic materials into “safe harbor” areas of the MSC genome [142]. There is still the risk of neoplastic transformation, infection, and immune responses, as most gene transfer methodologies use viruses. Refinement of novel, nonviral delivery methods using lipids or nanoparticles may reduce these risks [143, 144], although low transduction efficiencies and high titer toxicity issues must still be addressed.

4.9 MSCs and Safety Concerns Although there have been reports of spontaneous transformation of AT-MSCs [145] and even BM-MSCs [146] in vitro, it now appears that the neoplastic cells may have been due to contamination by an epithelial cancer cell line during the culture process [147]. These results underscore the need for sensitive and relevant in-process testing and monitoring of MSCs during in vitro culture expansion (i.e., checking expression and epigenetic status of critical genes involved in transformation such as p53, p21, p16Ink4a, hTERT, and c-myc). Others have further demonstrated the safety and nontransformative nature of MSCs despite the presence of chromosomal abnormalities [126], likely because MSCs exhibit replicative senescence in culture [21, 148–150] and escape senescence at very low frequencies [129]. Concerns that MSCs may induce tumor formation have arisen from several studies in which sarcomas have been modeled in mice by transducing MSCs with oncogenic fusion genes. This suggests that MSCs or some mesodermal precursor cell might constitute the target cell for transforming mutations in sarcomas including myxoid liposarcoma, rhabdomyosarcoma, and Ewing’s sarcoma [151–154]. The fusion gene, EWS–FLI1, resulting from a translocation of EWS with FLI1, may induce the transformation of MSCs into Ewing’s sarcoma cells [155, 156]. In fact, cancer-initiating cells in Ewing’s sarcoma display MSC-like properties [157]. The fusion gene SYT–SSX1 in MSCs induces a transcriptomic profile similar to that of synovial sarcoma [158]. It has also been reported that some gastric cancers may originate from MSCs, which may have fused with gastric mucosa cells infected with Helicobacter pylori [159]. The mechanism for potential transformation of MSCs is unknown, although accumulation of chromosomal instability is suspected. In a recent study, AT-MSCs

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from p53 and/or retinoblastoma protein-deficient mice were shown to have differential in vitro growth properties and some formed leiomyosarcoma-like tumors in immunodeficient mice [160]. Circumstantial evidence for the involvement of MSCs in transformative events includes a positive correlation between MSC numbers and formation of aggressive fibromatosis tumors in mature mice. The aberrant in vitro growth of fibromatosis-like tumor from MSCs derived from mice predisposed to aggressive fibromatosis tumors also supports the notion linking MSCs to tumor formation [161]. Another safety concern involves the controversial role of MSCs in stimulating tumor growth. There is some evidence that tumor formation entails recruitment of endogenous MSCs, which in turn affects the tumor stroma microenvironment and enhances fibrovascular desmoplasia, tumor formation, and metastasis [162, 163]. The multiple properties of MSCs that make them attractive for cellular therapies may, in fact, play a dual role in supporting the formation of tumors [164, 165]. The immunosuppressive properties, for example, may facilitate tumor progression by disabling the antitumor immune response; melanoma cells formed tumors when co-injected with MSCs but not when injected alone, suggesting a role for MSCs in tumor formation [166]. In this scenario, exogenous MSCs appeared to localize to the stroma of the tumor and may have modulated the tumor niche to support tumor proliferation [167]. MSCs may interact with tumor cells through cell–cell interactions and by secretion of numerous paracrine factors as shown for breast cancer cell lines [168–170]. Importantly, secretion of bioactive molecules by MSCs may not require the local presence of MSCs in the tumor, but may be mediated systemically. Other reported safety concerns include ossification and calcification in a model of AMI [171], potential arrhythmia induced by MSC engraftment [172], and clinical adverse events caused by exposure to DMSO, a common cryopreservative for culture-expanded MSCs, during intrathecal delivery [173]. Despite these concerns, no tumors have been found in human recipients of MSCs to date, and remarkably, even aneuploid MSCs may not give rise to tumors [126]. We estimate that while several thousand patients world-wide have been treated with different types of MSCs, very few serious adverse events have been documented, although long-term follow-up data are not available. Currently, 182 trials are registered with the FDA (clinicaltrials.gov) to treat various conditions, using autologous or allogeneic MSCs, isolated and processed under different conditions, making it difficult to compare effects from different protocols.

4.10 Future Directions The field is replete with conflicting data that can be resolved in part, by the development of more appropriate preclinical animal models. Better models will inform the design of more clinically appropriate prospective trials. Additional consideration needs to be given to the design of more informative clinical trials: they may provide information that could otherwise be very difficult to obtain, even with the best

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p­ reclinical models, for example, in assessing perturbations of the immune system in patients. It will be important to use rigorously manufactured, quality-assured, and animal substance-free MSC products for all trials, but especially for large-scale randomized, controlled studies. Considerable effort should be placed on developing real-time tracking and imaging of administered MSCs, a development that is highly likely to improve scheduling, dose, delivery and ultimately, the efficacy of the cells. Much work needs to be accomplished to develop database registries of the recipients of MSCs to better assess efficacy and safety, especially for long-term adverse effects. Further studies of the influence of MSCs on tumor biology in patients need to be conducted with preclinical models and in patients with malignancies receiving the cells. A greater understanding the mechanisms of action of MSCs will lead to a rational application of genetic engineering to produce more efficacious treatment for the enormous variety of disorders that are the target of this extraordinary population of cells.

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146. Rosland GV, Svendsen A, Torsvik A et al (2009) Long-term cultures of bone marrow-derived human mesenchymal stem cells frequently undergo spontaneous malignant transformation. Cancer Res 69:5331–5339 147. Garcia S, Bernad A, Martin MC et al (2010) Pitfalls in spontaneous in vitro transformation of human mesenchymal stem cells. Exp Cell Res 316:1648–1650 148. Fehrer C, Brunauer R, Laschober G et al (2007) Reduced oxygen tension attenuates differentiation capacity of human mesenchymal stem cells and prolongs their lifespan. Aging Cell 6:745–757 149. Stenderup K, Justesen J, Clausen C et al (2003) Aging is associated with decreased maximal life span and accelerated senescence of bone marrow stromal cells. Bone 33:919–926 150. Stolzing A, Jones E, McGonagle D et al (2008) Age-related changes in human bone marrowderived mesenchymal stem cells: consequences for cell therapies. Mech Ageing Dev 129:163–173 151. Castillero-Trejo Y, Eliazer S, Xiang L et al (2005) Expression of the EWS/FLI-1 oncogene in murine primary bone-derived cells results in EWS/FLI-1-dependent, ewing sarcoma-like tumors. Cancer Res 65:8698–8705 152. Charytonowicz E, Cordon-Cardo C, Matushansky I et al (2009) Alveolar rhabdomyosarcoma: is the cell of origin a mesenchymal stem cell? Cancer Lett 279:126–136 153. Riggi N, Cironi L, Provero P et al (2005) Development of Ewing’s sarcoma from primary bone marrow-derived mesenchymal progenitor cells. Cancer Res 65:11459–11468 154. Riggi N, Cironi L, Provero P et al (2006) Expression of the FUS-CHOP fusion protein in primary mesenchymal progenitor cells gives rise to a model of myxoid liposarcoma. Cancer Res 66:7016–7023 155. Tirode F, Laud-Duval K, Prieur A et  al (2007) Mesenchymal stem cell features of Ewing tumors. Cancer Cell 11:421–429 156. Burns JS, Abdallah BM, Schroder HD et al (2008) The histopathology of a human mesenchymal stem cell experimental tumor model: support for an hMSC origin for Ewing’s sarcoma? Histol Histopathol 23:1229–1240 157. Suva ML, Riggi N, Stehle JC et  al (2009) Identification of cancer stem cells in Ewing’s sarcoma. Cancer Res 69:1776–1781 158. Cironi L, Provero P, Riggi N et al (2009) Epigenetic features of human mesenchymal stem cells determine their permissiveness for induction of relevant transcriptional changes by SYT-SSX1. PLoS One 4:e7904 159. Houghton J, Stoicov C, Nomura S et al (2004) Gastric cancer originating from bone marrowderived cells. Science 306:1568–1571 160. Rubio R, Garcia-Castro J, Gutierrez-Aranda I et al (2010) Deficiency in p53 but not retinoblastoma induces the transformation of mesenchymal stem cells in vitro and initiates leiomyosarcoma in vivo. Cancer Res 70:4185–4194 161. Wu C, Nik-Amini S, Nadesan P et  al (2010) Aggressive fibromatosis (desmoid tumor) is derived from mesenchymal progenitor cells. Cancer Res 70:7690–7698 162. Karnoub AE, Dash AB, Vo AP et al (2007) Mesenchymal stem cells within tumour stroma promote breast cancer metastasis. Nature 449:557–563 163. Kidd S, Spaeth E, Dembinski JL et  al (2009) Direct evidence of mesenchymal stem cell tropism for tumor and wounding microenvironments using in vivo bioluminescent imaging. Stem Cells 27:2614–2623 164. Bexell D, Scheding S, Bengzon J (2010) Toward brain tumor gene therapy using multipotent mesenchymal stromal cell vectors. Mol Ther 18:1067–1075 165. Loebinger MR, Janes SM (2010) Stem cells as vectors for antitumour therapy. Thorax 65:362–369 166. Djouad F, Plence P, Bony C et al (2003) Immunosuppressive effect of mesenchymal stem cells favors tumor growth in allogeneic animals. Blood 102:3837–3844 167. Spaeth EL, Dembinski JL, Sasser AK et  al (2009) Mesenchymal stem cell transition to tumor-associated fibroblasts contributes to fibrovascular network expansion and tumor progression. PLoS One 4:e4992

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168. Rhodes LV, Muir SE, Elliott S et al (2010) Adult human mesenchymal stem cells enhance breast tumorigenesis and promote hormone independence. Breast Cancer Res Treat 121:293–300 169. Martin FT, Dwyer RM, Kelly J et al (2010) Potential role of mesenchymal stem cells (MSCs) in the breast tumour microenvironment: stimulation of epithelial to mesenchymal transition (EMT). Breast Cancer Res Treat 124:317–326 170. Klopp AH, Lacerda L, Gupta A et  al (2010) Mesenchymal stem cells promote mammosphere formation and decrease E-cadherin in normal and malignant breast cells. PLoS One 5:e12180 171. Breitbach M, Bostani T, Roell W et al (2007) Potential risks of bone marrow cell transplantation into infarcted hearts. Blood 110:1362–1369 172. Price MJ, Chou CC, Frantzen M et al (2006) Intravenous mesenchymal stem cell therapy early after reperfused acute myocardial infarction improves left ventricular function and alters electrophysiologic properties. Int J Cardiol 111:231–239 173. Karussis D, Kassis I, Kurkalli BG et al (2008) Immunomodulation and neuroprotection with mesenchymal bone marrow stem cells (MSCs): a proposed treatment for multiple sclerosis and other neuroimmunological/neurodegenerative diseases. J Neurol Sci 265:131–135

Part II

Biomaterials and the Extracellular Environment

Chapter 5

The Role of Mechanical Forces in Guiding Tissue Differentiation Sean P. Sheehy and Kevin Kit Parker

Abstract  Stem cell differentiation is regulated by a diverse array of extracellular cues. Recent evidence suggests that mechanical interactions between extracellular matrix (ECM) and cell surface receptors as well as physical interactions between neighboring cells play important roles in stem cell self-renewal and differentiation. It is also becoming clear that the ECM effects cellular behavior through many physical mechanisms, such as ECM geometry, elasticity, and the propagation of mechanical signals to intracellular compartments. Considerable effort is being targeted at developing biomaterials that exploit cellular microenvironments in guiding cells to desired phenotypes and organizing these into functional tissues. Improved understanding of the interactions between stem cells and their physical environment should yield new insight into the mechanisms governing their activity and allow the fabrication of artificial ECM to promote tissue development.

Abbreviations CAD ECM LINC MRTFs MSCs SRF STARS

Computer-aided design Extracellular matrix Linker of nucleoskeleton and cytoskeleton Myocardin-related transcription factors Mesenchymal stem cells Serum response factor Striated muscle activator of Rho signaling

K.K. Parker (*) Disease Biophysics Group, Wyss Institute for Biologically Inspired Engineering, School of Engineering and Applied Sciences, Harvard University, Cambridge, MA, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_5, © Springer Science+Business Media, LLC 2011

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5.1 Mechanotransduction Embryonic development is marked by dynamic, adaptive self-assembly, and self-organizational processes over the course of gastrulation and subsequently, during the formation of nascent organs. While chemical gradients and genetic regulatory networks certainly play important roles in morphogenesis, it is clear that the expression of genetic markers is necessary, but not sufficient, to explain differentiation. Microenvironmental chemistry and genetic synchrony are choreographed with mechanical signaling cues to drive development [1]. Increasing evidence suggests that epigenetic factors include mechanical and structural cues that play essential roles in embryogenesis and organogenesis [1, 2]. For example, mechanical tension in the cytoskeleton arising from physical interactions between neighboring cells and adhesion of cells to the ECM has been shown experimentally to contribute to epithelial branching and angiogenesis during lung development [3]. Moreover, branching morphogenesis during angio- and vasculogenesis arises from a complex interplay between tension exerted by epithelial cells on the ECM and regional differentials in ECM turnover by matrix metalloproteinases that creates localized fluctuations in ECM rigidity [4]. Direct physical interactions between cells play a vital role in development as well. Regulation of transcriptional programs via the Wnt/b-catenin signaling pathway mediated by cadherins junctions has been found to play a key role in the epithelial budding that gives rise to structures such as hair follicles [5]. This process of converting physical forces into intracellular biochemical responses is referred to as mechanotransduction [2]. As research in the field of stem cell biology has progressed, an increasing interest has arisen in evaluating the role of mechanotransduction in stem cell lineage commitment and its potential for exploitation in the development of regenerative therapies [6, 7]. Coordinated interactions with soluble factors, other cells, and extracellular matrices define a local biochemical and mechanical niche that stem cells occupy in vivo [8]. The ECM in this niche influences stem cell behavior both by providing mechanical signals and by physically trapping growth factors, limiting their diffusion, and regulating the temporal dynamics of paracrine signaling within the niche. A better understanding of the mechanisms of mechanical interaction between stem cells and the niche microenvironment will be important for directing the development of synthetic niches for therapeutic stem cell delivery [9].

5.1.1 The Role of Cell–Extracellular Matrix Interactions in Differentiation Cellular interactions with the ECM play an essential role in tissue formation, as shown in the heart where coordinated expression of specific ECM and integrin isoforms direct the proliferation and differentiation of early myocytes [10]. During fetal

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development, the ECM undergoes rapid changes in its composition and this change is associated with alterations in the expression of a-integrin isoforms that specifically recognize various ECM components [11]. Stem cells play an important role in tissue homeostasis and injury repair throughout the lifetime of an individual and thus must reside in an environment that maintains a balance between self-renewal, quiescence, and cell fate commitment. The mechanisms through which the stem cell niche maintains a population of self-renewing undifferentiated cells while simultaneously expelling differentiating daughter cells have been studied extensively, such as in bone marrow and intestinal crypts, where stem cell niches have been found to reside and participate in tissue development [12]. Niche localization and asymmetric division of stem cells is widely regarded to be a product of the specific intercellular and cell–ECM interactions that are characteristic of the stem cell compartment (Fig. 5.1a) [12, 13]. Uncommitted stem cells have been observed to express high levels of b1-integrins in the niches of a number of tissue types [13]. Thus, transmission of mechanical signals from the ECM to intracellular signaling pathways via transmembrane integrin receptors may play a prominent role in regulating cell cycle entry and stem cell fate decisions. 5.1.1.1 Signaling Through the Integrin–ECM Interphase Magnetic twisting cytometry experiments have shown that the transmembrane integrin receptors form a direct mechanical linkage between the ECM and the cytoskeleton [14]. Since this report, integrins have been demonstrated to serve as the primary conduit of bi-directional signaling between cells and the ECM, despite the fact that they do not possess intrinsic kinase activity [15]. Rather, integrins transmit information, encoded as mechanical forces, to the cytoskeleton that in turn activate mechanosensitive signal transducers, such as focal adhesion kinase that are able to translate the mechanical signal into a biochemical response. Integrin-mediated mechanotransduction has been shown to activate a myriad of chemical signaling pathways, including the Rho kinase, PI3K, ILK, Src, ERK, and MAP kinase pathways that modulate gene expression and direct important cellular activities, such as cell cycle progression and the induction of apoptosis (Fig. 5.1b) [2, 15, 16]. Many of these signaling molecules, along with biochemical mediators of transcription and protein synthesis, do not freely diffuse throughout the cytoplasm. Rather, they are immobilized on the cytoskeleton, and are thus subject to mechanical perturbations of the cytoskeleton, modulating their activity and translocation to cellular compartments, such as the nucleus [13]. 5.1.1.2 Mechanical Force Balance and ECM Stiffness in Mechanotransduction Mechanotransduction may be mediated simultaneously at multiple locations inside the cell through force-induced rearrangements within a tensionally integrated

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Fig. 5.1  Mechanotransduction in the stem cell niche. (a) Mechanical interactions between neighboring cells and with the ECM govern the response of stem cells to physical signals, such as tensile, compressive, and fluid stresses present within their local microenvironment. (b) Magnetic twisting cytometry experiments show that transmembrane integrin receptors form a direct mechanical linkage between the ECM and the cytoskeleton that can activate a number of intracellular signaling pathways. (c) The force balance between the ECM and the cytoskeleton allow naïve mesenchymal stem cells to adopt different fates depending on physical properties of the ECM, such as elastic modulus. (d) In addition to mechanotransduction through the integrin–ECM interface, stem cells also respond to mechanical signals from neighboring cells through intercellular junctions and direct transmembrane ligand–receptor interactions

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cytoskeleton [14]. This force balance between the cytoskeleton and the ECM allows cells to respond to variations in matrix compliance in a distinctive manner [17, 18]. Physical properties, such as elastic modulus, can vary considerably between, and within, organs. The elastic modulus of brain tissue has been measured to be on the order of 1  kPa, while those of muscle and bone are approximately 10 and 100 kPa, respectively [19]. These variations in stiffness are as a result of variety of factors, including cell demographics, extracellular heterogeneities such as ECM, sinuses, and the extent of the interstitial space. Cells have developed a variety of intra- and intercellular mechanisms to optimize these material properties for physiological function. For example, myosin-II motors play an essential role in forcefeedback response of stem cells to matrix elasticity [19, 20]. Phenomena observed in vitro, such as durotaxis where cells crawl up stiffness gradients, have lead many researchers to postulate that the mechanical microenvironment can influence tissue morphogenesis and stem cell fate choices [17]. Studies on embryonic cardiomyocytes reveal that changes in matrix rigidity associated with heart morphogenesis and fibrotic ECM remodeling caused by myocardial infarction dramatically affect rhythmic contraction of the cells [21]. In the case of marrow-derived MSCs, studies have shown that their lineage commitment is influenced by the elastic modulus of the substrate they are grown on. Culturing naïve MSCs on elastic substrates with a modulus of around 1 kPa promoted neurogenic differentiation, whereas growth on stiffer substrates, 10  kPa modulus, induced myogenic differentiation, and 100  kPa modulus substrates resulted in osteogenic lineage commitment (Fig. 5.1c). Further, experiments with the myosin II ATPase inhibitor blebbistatin showed that the elasticity-dependence of stem cell fate specification could be ameliorated by the inhibition of nonmuscle myosin II activity [19].

5.1.2 Intercellular Contact-Based (Juxtacrine) Mechanotransduction In addition to force transmission across the integrin–ECM interface, cells also receive mechanical signals from their neighbors through intercellular junctions and through direct transmembrane ligand–receptor interactions (Fig. 5.1d). The specification and proper arrangements of new cell types during tissue differentiation require the coordinated regulation of gene expression and precise interactions between neighboring cells, interactions that target transmembrane Notch receptors and the Wnt signaling intermediates localized to adherens junctions [22]. Cytoskeletal tension plays a key role in the formation and maintenance of adherens junctions during embryogenesis. Studies quantifying force transmission between endothelial cells across adherens junctions showed that this “intercellular tugging force” was associated with increases in the size and strength of adherens junctions, and in turn, regulated tissue architecture [23].

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The synthesis of gap junction channels is closely tied to the formation of adherens junctions [24]. Gap junctions are intercellular channels that allow the direct exchange of ions and biomolecules smaller than 1  kDa between the cytoplasm of adjacent cells. In cardiomyocytes, it has been found that N-cadherin and connexin 43 share a temporal relationship in their expression and spatial co-localization during adherens junction formation [25]. Further, mechanical forces acting on myocytes during contraction in  vivo and pulsatile stretch in  vitro were found to cause a dramatic increase in the expression of connexin 43 and a concomitant increase in conduction velocity due to increased electrical coupling between myocytes [24]. Mechanical loads placed on myocytes by contraction and pulsatile stretch were found to induce mechanotransductive signaling events through the b1–integrin–ECM interface that were responsible for the upregulation of N-cadherin and connexin 43 observed [26]. Studies of human embryonic stem cells reveal that they express both connexin 43 and connexin 45 that are assembled into functional gap junction channels [27]. Altogether, these results reveal that mechanical interactions between cells can influence chemical signaling by providing alternative pathways for signal transmission that potentially act on faster time scales than paracrine signaling through extracellular diffusion gradients. 5.1.2.1 Notch Signaling Pathway The Notch family of transmembrane receptors participates in an evolutionarily conserved signal transduction pathway that has been found to affect stem cell differentiation in a time- and context-dependent manner. Notch receptors mediate cell fate decision in multiple organs, including the skin, brain, and heart [8, 28]. Neighboring lineage committed cells present a transmembrane ligand known as Delta that interacts with and activates the extracellular domain of Notch receptors presented by an uncommitted stem cell when the cells come into physical contact with one another. Thus, the Notch receptor acts as a “touch sensor” for cells sharing the same tissue compartment, allowing them to sense and respond to the developmental activity of their neighbors. The spatial localization of Notch receptors and ligands in the cell membrane has also been found to affect the signaling response initiated upon Notch activation, although the mechanisms of Notch trafficking are still largely unknown [29]. Upon activation, an intracellular fragment of the Notch receptor is proteolytically cleaved and subsequently translocates to the nucleus where it initiates transcription to promote either proliferation or lineage commitment in a context-dependent manner [29]. Activation of Notch in neural stem cells has been associated with expansion of the uncommitted cell population both during development and in response to ischemic injury [30]. Notch1 activation in cardiac progenitor cells gives rise to a population of Nkx2-5 expressing transit amplifying myocytes that mediate postnatal growth of the myocardium [28]. Taken together, the results of these studies suggest that Notch serves as a mechanical signaling relay between cells within the stem cell niche that provides greater spatial and temporal precision than soluble cytokine gradients.

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5.1.2.2 Wnt/b-Catenin Signaling Pathway Wnts are secreted lipid-modified signaling peptides that play a ubiquitous role in development. Canonical Wnt signaling involves translocation of b-catenin from cadherins junctions to the nucleus where it interacts with a number of transcription factors to mediate transcription [22]. During embryonic development, Wnt signaling is necessary for the establishment and maintenance of cell polarity during gastrulation by modulation of actin cytoskeletal organization and contraction via its activation of the Rho signaling pathway. It is speculated that mechanical regulation of Wnt activity during embryonic development could serve as “mechanical checkpoints” that ensure certain structural criteria are met before the next stage of development proceeds [31]. Maintenance of the hematopoietic stem cell niche in bone marrow has been shown to depend on N-cadherin intercellular junctions with osteoblasts cells that serve to regulate b-catenin activation by Wnt [22]. The activity of the Notch and Wnt/b-catenin signaling pathways has been found to have reciprocal effects in cardiac progenitor cells. Notch1 signaling promotes differentiation of cardiac progenitor cells and negatively regulates the activity of b-catenin. Activation of b-catenin by the canonical Wnt pathway inhibits differentiation by negatively regulating cardiac transcription factors and instead promotes proliferation of cardiac progenitor cells [32]. Altogether, these studies reveal that cells possess signaling modalities beyond just the traditional chemical signaling pathways associated with development and that these mechanical signaling intermediates play important roles in tissue formation.

5.2 Role of Cell Geometry and Cytoskeletal Dynamics in Differentiation The ECM provides a number of contextual signaling cues during tissue formation that act by exerting tension on the cytoskeleton [13]. Cells respond to these signals from the ECM to “tune” their mechanical properties through cytoskeletal remodeling. Human MSCs cultured on micropost arrays adopted either an adipogenic or osteogenic phenotype depending on the stiffness of the microposts, with stiffer microposts promoting the osteogenic lineage and softer microposts promoting adipogenesis. It was postulated that the observed dependence of lineage commitment was due to changes in Rho-mediated cytoskeletal contractility in response to matrix elasticity and that the cytoskeletal architecture of naïve MSCs could be used to predict the fate they will ultimately adopt [7, 33]. Several studies have examined the influence of specific physical stimuli, such as tension, compression, and fluid shear stress on stem cell behavior to characterize the biophysical mechanisms that govern lineage commitment [34]. Experiments on Drosophila melanogaster embryos showed that acto-myosin-mediated tensional forces promoted proliferation, while compression suppressed it. These opposing physical forces are transmitted throughout the developing tissue and continually feed back to regulate tissue shape and organization [31].

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5.2.1 Effects of Mechanical Microenvironment on Cellular Organization Local variations in ECM mechanics act as motility cues that direct local cell growth differentials critical in organogenesis and wound healing [35]. The orientation of the mitotic spindle in dividing cells, and thus the division plane and spatial arrangement of daughter cells is affected by the spatial distribution of ECM proteins [36]. When grown on planar substrates in vitro, mammalian cells exhibit random walk motion. However, observations of pattern formation in epithelial and endothelial tissues revealed that cells migrated in a coordinated fashion [37]. Haptotaxis is widely regarded to be responsible for cohort migration at the macroscopic tissue scale, but at the scale of the cell’s local microenviroment, boundary conditions imposed by ECM topology, adjacent cells, and heterogeneities in the interstitial space provide the symmetry breaking cues that initiate the formation of specialized tissue patterns (Fig. 5.2a) [37]. Evaluation of the motility of cells grown on isolated ECM islands reveals that the direction of cell motility is defined by the topological organization of the cytoskeleton with respect to geometric cues in the ECM, the resulting tractional forces exerted by cells on the substrate, and the subsequent, spatially segregated activation of Rac, Rho, and cdc42 [35]. Cells grown on polygonal ECM islands exerted the greatest tension forces at the corners of the islands and this localization of mechanical force was associated with the localization of lamellipodia and filopodia to the corners as well. Taken together, the results of these studies indicate that the spatial organization of cells during tissue morphogenesis is the product of a complex interplay between mechanical guidance cues imposed by the ECM and tractional forces exerted on the ECM by cells mediated by dynamic rearrangement of the cytoskeleton and focal adhesions that serve to “steer” the direction of cell movement in response to the force-balance between cells and the ECM (Fig.  5.2b). Examination of these mechanisms in differentiating stem cells could prove useful in linking multicellular organization to spatial differentials of cell differentiation within a tissue [38, 39].

5.2.2 Effects of Mechanical Microenvironment on Cellular Shape and Function During embryonic development, changes in the mechanical microenvironment exert tensile and compressive forces that alter cell shape. Alterations to cell shape have been associated with stem cell fate decisions, as in the differentiation of embryonic stem cells into vascular endothelial cells [1]. Studies of human MSCs in vitro showed that they adopt an osteogenic phenotype when they were allowed to flatten and spread out, whereas they became adipocytes when they were restricted from spreading and maintained a rounded morphology [40]. In the heart, interactions between myocytes and the ECM give rise to changes in cell shape that direct actin filament orientation, sarcomere organization, and

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Fig. 5.2  Effects of the mechanical environment on tissue morphogenesis. (a) Localized boundary conditions imposed on cells by the ECM and the degree of mechanical coupling between neighboring cells provide cues for the coordinated migration of vascular smooth muscle cells on micropatterned square fibronectin islands. (b) Studies of cells grown on square ECM islands have shown that tractional forces imparted on the cell at focal adhesions cause dynamic rearrangement of the cytoskeleton in response to geometric constraints

myofibrillogenesis [41]. Changes in cell shape are the product of Rho-mediated rearrangement of the cytoskeleton. Examination of capillary network formation by human microvascular endothelial cells in vitro and retinal angiogenesis in vivo using the Rho inhibitor p190RhoGAP revealed that Rho-induced changes in

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cytoskeletal architecture regulated angiogenesis by modulating the activities of two antagonistic transcription factors, TFII-1, and GATA2, that govern the expression of the VEGF receptor. Moreover, the activity of p190RhoGAP was found to be sensitive to ECM elasticity [42]. Dynamic assembly and disassembly of cytoskeletal elements generate directed forces that perturb cell shape and guide the organization of cellular components. This mechanical force-balance influences cellular behavior by modulating gene expression activity and could serve as an important factor in cell fate decisions made by stem cells during tissue morphogenesis.

5.2.3 Actin Cytoskeletal Remodeling and Transcriptional Regulation The mechanical stiffness of the local microenvironment and the contractile activity of cells influence gene expression during embryogenesis [31]. In particular, genes encoding proteins involved in tissue remodeling processes have been found to be susceptible to changes in cellular morphology induced as a consequence of direct perturbation of cytoskeletal structure with actin and microtubule disrupting agents, such as cytochalasin D and colchicine [43]. Indeed, coordination between protein synthesis and cell motility is necessary for the timely generation of the structural components that support remodeling of the cytoskeleton. Examination of the link between cytoskeletal dynamics, motility, and gene expression revealed that MRTFs are physically bound to globular actin monomers until they are incorporated into actin filaments. Upon release from actin monomers, the MRTFs are free to translocate to the nucleus where they interact with the transcription factor SRF to promote the expression of genes under its control (Fig. 5.3a) [44]. This actin–MRTF–SRF mechanotransduction pathway may be particularly important in striated muscle development, as studies have identified a muscle-specific actin binding protein known as STARS that activates SRF through a Rho-dependent mechanism [45]. It is postulated that the upre­ gulation of STARS during myogenesis provides a feed-forward mechanism for driving the expression of genes regulated by MRTF and SRF and reinforcing the differentiation process during the formation of skeletal and cardiac muscle tissues [45]. RhoA-dependent regulation of the actin cytoskeleton also plays a central role in regulating transcription during smooth muscle differentiation as well. Most smooth muscle-specific differentiation marker genes code for proteins associated with contractility, suggesting that Rho-dependent changes in smooth muscle contractility may be coupled to long-term regulation of smooth muscle-specific gene expression [46].

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Fig. 5.3  Mechanical regulation of gene expression. (a) Myocardin-related transcription factors associates with globular actin in the cytoplasm translocate to the nucleus and alter gene expression as globular actin is incorporated into actin filaments during cytoskeletal remodeling. (b) Mechanical continuity between integrins, the cytoskeleton, and nuclear scaffolds could provide a path for mechanical signal transfer between the ECM and the nucleus

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5.3 Nuclear Mechanics and Regulation of Gene Expression The induction of gene expression by mechanotransduction has traditionally been assumed to occur via activation of established transcriptional regulatory pathways through biochemical signaling molecules localized to the surface of the plasma membrane. Experimental data suggest that individual filaments of the cytoskeleton bear tensile and compressive loads and give rise to a mechanical network under isometric tension that propagate physical signals throughout the cell at a velocity far exceeding the limits of chemical diffusion [14, 47]. An intriguing alternative signaling paradigm is the transduction of mechanical signals through the ECM–cytoskeletal network to structures deep within the cytoplasm, such as the nucleus, where they can alter enzymatic activity or gene expression by altering nuclear shape or physically deforming genomic structures within the nuclear compartment.

5.3.1 Mechanical Continuity Between ECM and Nucleus It is widely recognized that focal adhesions serve as a mechanical conduit between the ECM and the cytoskeleton. However, much speculation remains about the physical continuity between the cytoskeleton and the nucleus and whether this mechanical linkage serves as an epigenetic regulator of gene expression [13]. Molecular connections between integrins, cytoskeletal filaments, and nuclear scaffolds may therefore provide a discrete path for mechanical signal transfer through cells as well as a mechanism for producing integrated changes in cell and nuclear structure in response to changes in the ECM. Studies involving the application of force to focal adhesions using micropipettes and RGD-coated microbeads provided evidence of mechanical continuity between membrane-localized integrin receptors and the nucleus via the actin cytoskeleton [48]. Interactions between nesprins, SUN, and lamins form a specialized nuclear anchoring structure for cytoskeletal filaments referred to as the LINC complex [47]. Emerin proteins within the nucleus provide a physical connection between the LINC complex and many proteins involved in chromatin modification. Chromosomes are traditionally regarded as discreet, physically separate entities, but microsurgery experiments revealed that isolation of one chromosome from living cells under isotonic conditions resulted in the removal of all of the chromosomes within the nucleus. Analysis of chromosome positioning and movement suggested that different chromosomes often behave as if they were physically connected during interphase and this mechanical coupling may coordinate dynamic alterations in chromatin structure [49]. Taken together, the results of these studies provide strong evidence that a direct physical linkage between the ECM and genome exists, and raises the question of whether this mechanical continuity provides a mechanotransduction pathway for modulating gene expression by directly altering chromatin architecture (Fig. 5.3b).

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5.3.2 Modulation of Nuclear Shape and Its Effect on Gene Expression Nuclear shape, structure, and stiffness strongly correlate to cellular function and phenotype in physiological and pathological situations where force is involved. The nucleus of most cells is roughly ellipsoid, or spheroidal, in shape and is regarded as the stiffest of the organelles. Studies of differentiating human embryonic stem cells noted that uncommitted cells possessed large, round nuclei with little lamin A and highly mobile chromatin [50–52]. As the cells adopted a particular lineage, researchers found that the nuclei demonstrated concomitant changes in nuclear shape and structure, revealing a strong correlation between nuclear shape change and changes in cellular phenotype [53]. Forces applied directly to the surface of cells, such as shear forces during fluid flow, can increase the load on the cytoskeleton and subsequently deform the physically connected nucleus. Examination of neonatal cardiomyocytes in vitro showed that the spatial organization of myofibrils in response to geometric cues provided by the ECM caused the aspect ratio of the nucleus to increase as the aspect ratio of the myocytes increased [54]. Measurements of gene and protein expression in primary osteogenic cells cultured on micropatterned islands of ECM protein revealed that changes in nuclear shape affected the activity of transcription factors that govern the expression of collagen I and osteocalcin, markers for the osteogenic phenotype [55]. Together, the results of these studies provide strong evidence for a possible role for mechanotransductive regulation of gene expression through alterations in the transfer of mechanical forces from the cytoskeleton to the nucleus. Recent experiments confirm that gross epigenetic modifications that occur during stem cell differentiation can be detected as changes in the shape and stiffness of the nucleus, clearly demonstrating a relationship between nuclear architecture, chromatin organization, and transcription [52]. The intriguing, recently proposed concept of cytoskeletal epigenetics raises the question of whether the continued reorganization of long-lived cytoskeletal structures in a cell can serve as an epigenetic mechanism to record the “mechanical history” of a cell and influence the behavior of its daughter cells. The implications of this hypothesis are that stable cytoskeletal structures could potentiate variability in cell behavior and guide cell fate decisions toward certain phenotypes across generations of cells [36]. Nuclear shape is emerging as an important indicator of mechanical continuity between the nucleus, cytoskeleton, and ECM that has been implicated in providing an alternative pathway for regulating gene expression in response to the mechanical microenvironment of the cell.

5.4 Utilization of Mechanical Cues to Guide Engineered Tissue Formation Advances in the field of cellular biomechanics are beginning to explain how physical forces and mechanical structures impact information processing and cellular decisionmaking [9]. Increased understanding of the relationship between cellular behavior

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and the mechanical characteristics of their environment is motivating the development of new biomaterials that take advantage of this phenomenon to drive stem cell differentiation and tissue morphogenesis with greater precision [56]. One example of this is efforts to fabricate functional myocardial tissue grafts to repair damaged areas of the heart after myocardial infarction. Current efforts aim to derive cardiac progenitor cells that can be expanded in vitro and then selectively differentiated into the muscular, vascular, and conduction system cells that comprise the myocardium. Of equal importance is the development of ECM scaffolds that provide appropriate mechanical cues to guide the differentiation and organization of cardiac progenitor cells into a functional tissue structure that can be incorporated into highly complex structure of the native myocardium [57]. In addition to the structural guidance cues provided by the ECM, the behavior of cells during embryonic development is also influenced by tractional forces created by contracting cells and propagated through the ECM to neighboring cells. The application of a 10% static stretch to mouse embryonic stem cells was found to increase the number of contracting cells, whereas application of 10% cyclic stretch to human embryonic stem cells was found to decrease differentiation and maintain them in a pluripotent state [58]. The results of these studies clearly indicate that a better understanding of the influence of the mechanical environment on stem cell activity and the development of novel biomaterials that take advantage of this knowledge is required to advance the field of regenerative medicine.

5.4.1 Computational Modeling of Mechanotransductive Effects With refinements in our understanding of mechanobiology and the in vitro experimental platforms used to study mechanotransduction, mathematical models of force distribution in tissues and the parameters that dictate mechanosensing are starting to emerge. The development of in vitro techniques to regulate ECM composition and geometry has made it possible to explore the effects of cell–ECM interactions on specific parameters of cell behavior [59]. Such techniques have been used to develop a computational model of the relationship between cell shape and calcium dynamics in developing cardiomyocytes [60]. It has also been used to develop a computation model to forecast the fate specification of human MSCs based on the early cytoskeletal arrangement imposed on the cells by the geometry of the ECM [7]. As techniques to fabricate free-form engineered tissues emerge, mathematical models are being developed that attempt to describe their behavior and predict their performance characteristics given some change to tissue architecture. For example, a finite element model was recently developed to simulate the performance characteristics of engineered myocardial constructs and provide predictions about the effects of changing myofibrillar orientation on their contractile function [61]. In addition to mathematical descriptions of in vitro model systems, researchers have also begun to develop computation simulations of the injured in  vivo tissue environments for which engineered tissues are being developed to repair. A multiscale mathematical

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model of strain-driven eccentric growth and stress-driven concentric growth of the myocardium during ischemic injury has recently been developed that allows researchers to explore the effects of local changes in cardiomyocyte morphology due to alterations in stress/strain distribution caused by fibrosis on cardiomyocyte function from the multicellular tissue scale down to the molecular scale of the sarcomere [62]. These reports are the first effort to develop CAD tools for engineered tissues. While CAD tools are routinely used in many engineering disciplines, in tissue engineering these tools, combined with medical imaging data, will require understanding of biotic–abiotic interface physics and a hierarchal understanding of self-organizing biological systems.

5.4.2 Fabrication of ECM Substrates That Promote Functional Maturation Artificial tissues suitable for regenerative applications will require scaffolds that can promote controlled differentiation of a stem cell population and impose precise cellular organization. A number of synthetic polymer compounds have been evaluated for their ability to support the efficient clonal expansion and differentiation of stem cells based on structure–function relationships between cell behavior and substrate material properties [63]. By mimicking the physicochemical properties and self-assembly fabrication of natural materials, artificial scaffolds are beginning to be developed that incorporate peptide motifs that support the engagement of specific pairs of integrins and allow remodeling of the synthetic matrix by proteases secreted by cells [56, 63, 64]. As our understanding of the influence of mechanical cues on stem cell fate decisions matures, this information can be used to direct the development of cell substrates that utilize these mechanical cues to create stem-cell derived tissue constructs with desirable functional properties. 5.4.2.1 Control of Cell Shape and Organization The intercellular and cell–ECM interactions within a tissue govern the shape and organization that the cells comprising that tissue will ultimately adopt, and these interactions clearly play an important role in regulating the survival and functionality of those cells [16]. Microcontact printing is a well-established technique for fabricating planar cell growth substrates with precisely defined ECM geometry (Fig. 5.4). An elastomeric stamp with micrometer-scale features can be “inked” with an ECM protein of choice, such as fibronectin, laminin, or collagen, and transferred to a flat substrate that promotes protein adsorption. Cells seeded onto these substrates preferentially bind to the portions of the substrate coated with the patterned ECM protein, giving rise to a large population of cells with shapes defined by the ECM pattern [65]. This technique has been used extensively to study the relationship between shape and behavior and a number of cell types, including differentiating

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Fig. 5.4  Utilization of mechanical cues to guide differentiation. Microcontact printing allows the fabrication of ECM substrates with defined microscale geometry using photolithographic templates. This technique has been used extensively to study the contribution of cell geometry and tissue organization on stem behavior in vitro

stem cells. The results of these studies indicate that cell geometry is indeed and important factor in directing the lineage commitment of stem cells and continues to influence their behavior throughout their lifetime. A critical limitation of this technique is the fact that it can only be used with rigid planar substrates that do not mimic the mechanical properties of natural tissues and give rise to monolayers of cells. Three-dimensional scaffolds with natural tissue-like mechanical properties need to be developed that incorporate precise ECM cues for controlling cell shape in a nonplanar substrate. 5.4.2.2 Evolution of Biomaterials for Regenerative Medicine Biomaterials made today are routinely information rich and incorporate biologically active components inspired by natural analogs [66]. Researchers have begun to design materials that combine synthetic polymer compounds with peptide motifs that can be proteolytically cleaved by matrix metalloproteinases secreted by cells to create scaffolds that can be sculpted by cells during tissue formation [64]. Advances in the construction of three-dimensional polymeric scaffolds are also starting to make the fabrication of therapeutically relevant artificial tissue constructs a reality [56, 67, 68]. A recently developed method derived from the microcontact printing approach to fabricating two-dimensional tissues in vitro allows the fabrication of free-standing protein nanofabrics. These protein nanofabrics are constructed by microcontact printing successive layers of ECM protein onto a rigid substrate coated with a thermosensitive polymer. These nanofabrics can be comprised of a heterogeneous

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composition of ECM proteins and the microcontact printing technique provides control over the shape, size, and orientation of the protein “threads” with respect to one another. Further, cells will readily adhere to these ECM fabrics and stacking of these nanofabrics may allow the construction of ECM scaffolds with precise organizational cues throughout the volume of the scaffold [69]. Another promising approach for fabricating three-dimensional ECM tissue scaffolds with precise geometry is the recently developed rotary jet spinning technique for generating fibrous tissue scaffolds [70]. This technique overcomes the limitations of the traditional electrospinning technique to produce highly aligned nanoscale fibers using a nozzle rotating at high speed to produce a jet of polymer solution that undergoes extensive stretching before polymerization. The primary advantage of this technique over other methods of three-dimensional scaffold production is its ability to quickly produce large quantities of tissue scaffolds of arbitrary size composed of precisely aligned protein nanofibers. The focus of future biomaterials design will likely be focused on the development of “smart” materials that integrate multiple inputs from both chemical and mechanical stimuli to direct their behavior [56]. Such materials could simplify and optimize engineered tissue fabrication by more closely reproducing the dynamic microenviroment presented to differentiating cells during development, allowing researchers to take advantage of the natural interactions between cells and their environment during tissue morphogenesis to reproducibly drive the fate commitment of cells without the need for complex experimental manipulations.

5.4.3 Measurement of Maturation and Tissue Function An important final consideration in the fabrication of engineered tissues from uncommitted stem cells is the evaluation of functional performance characteristics of the artificial tissue. Traditionally, differentiation has been assessed by measuring the expression of specific marker genes. However, this metric requires destruction of the tissue to isolate mRNA for measurement and is not informative for cells and tissues that require the precise assembly and organization of macromolecular structures, such as the sarcomeres of striated muscle for their functionality. Biomimetic microfluidic devices are emerging as a promising platform for measuring the performance characteristics of engineered tissues in  vitro. A recent study provided the first proof of principle demonstration of this approach to model the structural, functional, and mechanical properties of the alveolar–capillary interface of the human lung. This microfluidic device was not only able to reproduce the functionality of an alveoli, but it also allowed the identification of novel mechanosensitive responses of the lungs to nanoparticulates [71]. Application of these organ-on-chip devices to the fabrication of tissues using stem cells could provide a powerful tool for the quantitative analysis of stem cell-derived artificial tissues. Evaluation of the functional characteristics of muscle tissue is especially challenging, as traditional assays are not able to provide direct measurements of their contractile performance.

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A novel muscular thin film assay was recently developed that allows direct measurement of the contractile force of engineered muscle tissues [72]. This assay has been successfully used to demonstrate the myogenic potential of mouse cardiac progenitor cells isolated from the primary and secondary heart fields during various stages of cardiogenesis [73]. Subsequent modifications to the muscular thin film assay have made it amenable to the evaluation of smooth muscle cell contractility, in addition to striated muscle contractility, and allow the simultaneous measurement of multiple engineered muscle constructs in the same dish [74]. As the field of regenerative medicine advances, and the complexity of engineered tissues increases, new approaches will be needed to evaluate the utility of these tissues for therapeutic applications. Cell-based biochips represent an attractive test system that negate the need for costly animal models and allow quantitative analyses of tissue function that are not possible in traditional cell culture systems.

5.5 Opportunities and Challenges for Utilizing Mechanical Cues to Guide Tissue Formation It is now commonly accepted that mechanotransduction plays an important role in stem cell differentiation and tissue morphogenesis. However, much remains to be discovered about the cellular mechanisms that provide the interface between mechanosensation and activation of biochemical processes, such as gene expression. Much evidence points to the cytoskeleton as this nexus, since it provides the mechanical continuity between the ECM and intracellular structures, and dictates the shape and spatial organization of a cell. As the biomechanics of mechanotransduction are elucidated, these findings must be incorporated into next-generation, multiscale biomaterials to provide stem cells with a mechanical microenvironment that directs their behavior in a predictable and reproducible manner.

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54. Bray M-AP, Adams WJ, Geisse NA, Feinberg AW, Sheehy SP, Parker KK (2009) Nuclear morphology and deformation in engineered cardiac myocytes and tissues. Biomaterials 31(19):5143–5150 55. Thomas CH, Collier JH, Sfeir CS, Healy KE (2002) Engineering gene expression and protein synthesis by modulation of nuclear shape. Proc Natl Acad Sci USA 99(4):1972–1977 56. Huebsch N, Mooney DJ (2009) Inspiration and application in the evolution of biomaterials. Nature 462(7272):426–432. doi:10.1038/nature08601 57. Chien KR, Domian IJ, Parker KK (2008) Cardiogenesis and the complex biology of regenerative cardiovascular medicine. Science 322(5907):1494–1497 5 8. Jacot JG, Martin JC, Hunt DL (2010) Mechanobiology of cardiomyocyte development. J Biomech 43(1):93–98 59. Mrksich M, Dike LE, Tien J, Ingber DE, Whitesides GM (1997) Using microcontact printing to pattern the attachment of mammalian cells to self-assembled monolayers of alkanethiolates on transparent films of gold and silver. Exp Cell Res 235(2):305–313 60. Adams WJ, Pong T, Geisse NA, Sheehy SP, Parker KK (2007) Engineering design of a cardiac myocyte. J Computer-Aided Mater Des 14:19–29 61. Bol M, Reese S, Parker K, Kuhl K (2008) Computational modeling of muscular thin films for cardiac repair. Comp Mech 43:535–544 62. Latimer DC, Roth BJ, Parker KK (2003) Analytical model for predicting mechanotransduction effects in engineered cardiac tissue. Tissue Eng 9(2):283–289 63. Mei Y, Saha K, Bogatyrev SR, Yang J, Hook AL, Kalcioglu ZI, Cho SW, Mitalipova M, Pyzocha N, Rojas F, Van Vliet KJ, Davies MC, Alexander MR, Langer R, Jaenisch R, Anderson DG (2010) Combinatorial development of biomaterials for clonal growth of human pluripotent stem cells. Nat Mater 9(9):768–778. doi:10.1038/nmat2812 64. Chau Y, Luo Y, Cheung ACY, Nagai Y, Zhang SG, Kobler JB, Zeitels SM, Langer R (2008) Incorporation of a matrix metalloproteinase-sensitive substrate into self-assembling peptides – a model for biofunctional scaffolds. Biomaterials 29(11):1713–1719. doi:10.1016/j. biomaterials.2007.11.046 65. Singhvi R, Kumar A, Lopez GP, Stephanopoulos GN, Wang DI, Whitesides GM, Ingber DE (1994) Engineering cell shape and function. Science 264(5159):696–698 66. Lutolf MP, Hubbell JA (2005) Synthetic biomaterials as instructive extracellular microenvironments for morphogenesis in tissue engineering. Nat Biotechnol 23(1):47–55 67. Kraehenbuehl TP, Zammaretti P, Van der Vlies AJ, Schoenmakers RG, Lutolf MP, Jaconi ME, Hubbell JA (2008) Three-dimensional extracellular matrix-directed cardioprogenitor differentiation: systematic modulation of a synthetic cell-responsive PEG-hydrogel. Biomaterials 29(18):2757–2766 68. Eschenhagen T, Zimmermann WH (2005) Engineering myocardial tissue. Circ Res 97(12):1220–1231 69. Feinberg AW, Parker KK (2010) Surface-initiated assembly of protein nanofabrics. Nano Lett 10(6):2184–2191 70. Badrossamay MR, McIlwee HA, Goss JA, Parker KK (2010) Nanofiber assembly by rotary jet-spinning. Nano Lett 10(6):2257–2261 71. Huh D, Matthews BD, Mammoto A, Montoya-Zavala M, Hsin HY, Ingber DE (2010) Reconstituting organ-level lung functions on a chip. Science 328(5986):1662–1668 72. Feinberg AW, Feigel A, Shevkoplyas SS, Sheehy S, Whitesides GM, Parker KK (2007) Muscular thin films for building actuators and powering devices. Science 317(5843):1366–1370 73. Domian IJ, Chiravuri M, van der Meer P, Feinberg AW, Shi X, Shao Y, Wu SM, Parker KK, Chien KR (2009) Generation of functional ventricular heart muscle from mouse ventricular progenitor cells. Science 326(5951):426–429 74. Alford PW, Feinberg AW, Sheehy SP, Parker KK (2010) Biohybrid thin films for measuring contractility in engineered cardiovascular muscle. Biomaterials 31(13):3613–3621

Chapter 6

Synthetic Multi-level Matrices for Bone Regeneration Nicholas R. Boyd, Richard L. Boyd, George P. Simon, and David R. Nisbet

Abstract  Current bone replacement strategies are clinically inadequate, yet there is great promise in the use of synthetic adjuvant matrices. Electrospinning provides a three-dimensional platform in which matrices can be designed to mimic features of the extracellular matrix and improve bone regeneration. Composite nanofibers can be functionalised with therapeutic molecules, and/or may permit the delivery of growth factor combinations as required to stimulate bone healing. Collectively, these should more precisely direct repair by exogenous and endogenous stem and progenitor cells. The real novelty will be in combining multiple levels of scaffold-based tissue engineering developments in an “off the shelf” clinic-ready product. Until then, application of bioactive nanofiber analogues, with dual-scale three-dimensional porosity that can be co-interfaced within effective stem cell treatment regimes, will be crucial in develo­ ping smart matrices for skeletal repair. This review presents holistic concepts for more effective bone regeneration and the methods in which they can be incorporated into nanotechnology-based scaffolds from a materials engineering perspective.

Abbreviations ALP BMP-2 BMSCs BSP ECM FAp

Alkaline phosphate Bone morphogenetic protein-2 Bone-marrow-derived stromal cells Bone sialoprotein-2 Extracellular matrix Fluoroapatite

D.R. Nisbet (*) Research School of Engineering, ANU College of Engineering and Computer Science, The Australian National University, Canberra, ACT, Australia e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_6, © Springer Science+Business Media, LLC 2011

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FDA FGF-2 HAp Hep Hep-S LbL MMT MSCs MWCNT MWNT OC OP OPG PCL PDGF PEO PDLLA PGA PLA PLLA RANKL RGD rhBMP-2 SBF SEM SIS SPARC TCP TEM TGF-b TSP VEGF

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Food and Drug Administration Fibroblast growth factor-2 Hydroxyapatite Heparin Heparan sulphate Layer-by-layer Montmorillonite Mesenchymal stem cells Multi-walled carbon nanotube Multi-walled nanotube Osteocalcin Osteopontin Osteoprotegerin Poly(e-caprolactone) Platelet-derived growth factor Polyethylene oxide Poly(DLlactide) Polyglycolic acid Polylactic acid Poly-l-lactide acid Receptor activator of NF-kB ligand Arg-Gly-Asp Recombinant human BMP-2 Simulated body fluid Scanning electron microscopy Small intestinal submucosa Osteonectin Tricalcium phosphate Transmission electron microscopy Transforming growth factor b Thrombospondin Vascular endothelial growth factor

6.1 Introduction One of the ironies of improvements to general health and living conditions that lead to an increase in average life span is a higher incidence of skeletal disease. While some treatments do exist, they are usually problematic due to a lack of long-term success and insufficient clinical efficacy. More sophisticated approaches, such as artificial bone substitutes, are thus required on a global scale. The rapidly evolving interface between nanotechnology and biological sciences provides an optimistic outlook. Bone-related diseases, skeletal abnormalities and physical trauma can often lead to defective skeletal support, such that bone may no longer be able to regenerate and

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repair naturally [1]. Osteoporosis, osteonecrosis, bone cancer, osteoarthritis, rickets, Paget’s disease and osteogenesis imperfecta [2–8] cause an enormous strain economically and on the patients quality of life. In 1997, it was estimated that osteoporosis affected 75 million people in Europe, USA and Japan alone [9], with alarming projections estimating that the incidence in hip fracture as a result of osteoporosis will double by 2025 [10]. Throughout orthopaedics, neurosurgery and dentistry, some 2.2 million bone-graft surgeries are already performed annually [11]. Two major issues are rapidly becoming clear: (1) the need for bone grafts is increasing and (2) the longevity of existing bone-graft therapies needs to improve in order to avoid revision surgery and to accommodate the expected increase in patient lifespan post-surgery [12, 13]. Consequently, new strategies and more sophisticated materials are required for restoration of bone/joint functionality. Current bone grafts are burdened by limited availability, infection and morbidity, poor mechanical properties and over-elevated bone resorption [1, 14]. While the traditional “gold standard” autograft has shown a degree of clinical success, only small skeletal defects can be treated and they are very invasive [14]. Alternatively, problems with allografts are compounded by immune rejection, and immunosuppressive drugs carry the added risk of susceptibility to infection [1, 15]. Synthetic biomaterials have provided an alternative solution with advantages in availability, versatility, precision, reduced immune rejection and the potential for smarter, biologically instructive and active bone-graft templates. Most common hard orthopaedic biomaterials, e.g., titanium hip implants, are largely bio-inert which leads to limited integration with endogenous tissue and poorly sustained functional restoration [16]. Numerous attempts to improve the integration of the implant via surface treatment of the titanium [17] have included mechanical roughening, chemical treatment, sol-gel coating, ion implantation and thermal spraying of HAp. However, major limitations still persist and the ability to achieve appropriate mechanical and biological properties has yet to be achieved simultaneously. Potentially, the most promising fabrication technique for the next generation of synthetic bone grafts is electrospinning. This provides a simple, yet highly versatile bottom-up approach [18], in which a vast array of porous nanofiber matrices can be manufactured by accelerating a jet of charged polymeric liquid under the presence of an electric field [19–21]. A wide range of polymeric, ceramic and composite nanofibers can be produced via electrospinning [18]. In turn, the diversity of electrospun membrane properties and functionality can be manipulated relatively easily. Electrospun fibers are amenable to post-processing surface treatments including deposition of bioactive nanoparticles, attachment of growth factors and bioactive proteins. However, while the approach is simple, the physical science behind electrospinning is quite complex [20, 21] and difficult to control. The challenge is clear: can innovative new strategies be developed to optimise the utility and function of three-dimensional tissue analogues? Currently, the mechanical properties of electrospun scaffolding for bone tissue engineering are generally inadequate for therapeutic applications as load bearing bone-graft substitutes. While their in vivo implementation is in its infancy, direct avenues for their clinical application exist and have advanced as guided-boneregeneration membranes in areas such as dental restoration. Moreover, augmentation in non-union/delayed fracture healing and in replacement of segmental gap defects

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may support revolutionary therapeutic advances if improvements can be made to tailor their mechanical properties. This review firstly introduces basic bone biology and healing redevelopment processes as a platform for scaffold design. A hierarchy of physical and biochemical inputs have been utilised for stimulating bone regeneration, yet have not been successfully co-incorporated to date. Hence, the focus of this review discusses the most recent, novel and promising advances in scaffold engineering designed primarily for bone tissue reconstruction. Effective three-dimensional synthetic biomatrices for bone regeneration are thought to require surface and structural biomimicry, in conjunction with effective stem cell treatment strategies.

6.2 The Platform for Design: Biology The successful growth of cells on synthetic matrices is dependent on synergies between topographical, biochemical and mechanical stimuli that best reflect the in vivo microenvironment. Understanding these biological functions, structures and healing processes is therefore crucial for creating relevant inputs into the adjuvant bone graft. Material engineering strategies alone have not yet progressed to the point where they can recreate bone analogues accurately, due to the dynamic nature and intricate hierarchy of bone.

6.2.1 Bone Structure Bone is a highly organised connective tissue that principally provides structural support for skeletal motion, protection of internal organs and a mineral reservoir for homeostasis of chiefly calcium and phosphate [22–24]. Furthermore, bone supports the primary production of blood cells via haematopoiesis [16], has unique selfregenerative capabilities [25] and an exceptional light-weight combination of strength, stiffness, fatigue resistance, fracture toughness and porosity [26, 27]. The specialised physical and biological capabilities of bone are founded on the complex ECM hierarchy (reviewed in refs. [26, 28]) and subsequent dynamic relationship with bone-associated cells. At the nanoscale, intricately organised mineralised collagen fibrils (individual fibril (1.5–3.5 nm), fibers (50–70 nm), and bundles of fibers (150–250 nm) [26] form the building blocks of the ECM. A complex, interpenetrative and cooperative network of organic and inorganic components underpins the highly advanced mechanical properties of bone at all scale lengths [26, 29–31]. Bone-apatite nanoplatelets based around the chemical structure of HAp, Ca5(PO4)3OH, occupy approximately 70  wt% of bone [32] and are dispersed throughout a highly organised and cross-linked [33, 34] periodic stacking [26, 30, 34] of triple helical collagen type I molecules [31, 35], referred to as tropocollagen. HAp platelets act

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as stiffening components and mineral reservoirs, and have been reported to be located on the fiber surface, as well as within the fiber [36, 37]. The organic component makes up about 25 wt% of bone [32], with more than 90% being collagen type I [28]. A small fraction of type V collagen coexists, which facilitates the assembly of fibrillar structures [38, 39], as do several non-collagenous proteins [40]. While only accounting for a minor contribution (approximately 10% [28]) [41], OC [42], SPARC [43, 44], fibronectin [44], BSP [45] and TSP [46–48] play crucial biochemical roles in the ECM.

6.2.2 Bone Maintenance and Remodelling Bone, like most tissues, undergoes continual self-renewal and requires repair and remodelling following damage. Homeostatic maintenance of bone structure essentially involves a continuing flux between bone resorption by osteoclasts and reformation by osteoblasts, which is initiated through mechanical stimuli and ECM-cellular communication. Osteoclasts are bone-resorbing macrophage-cells, which locally secrete hydrochloric acid that facilitates localised mineral dissolution, followed by up-regulation of cysteine-proteinases (importantly cathepsin K) that degrade the organic matrix [49, 50]. Osteoblasts are bone-forming cells derived from MSCs which are present predominantly in the endosteal niche and surrounding marrow vasculature [51]. New bone is developed through osteoblastic synthesis, deposition, mineralisation and organisation of the ECM. This includes production and secretion of type I collagen, bone apatite and a vast variety of non-collagenous proteins [40], cytokines and growth factors [24, 52]. Virtually all tissues derive growth and differentiation occurs through dynamic interaction with the surrounding ECM. Bone, however, displays uniquely adaptive properties, as the cells and bone remodelling is distinctly sensitive to mechanical stress [53]. Physical forces imposed on bone are translated and interpreted via three-dimensional osteocyte networks through mechanotransduction (ECM-cell cross-talk) [54], which correspondingly feeds instructive signals to regulate osteoblastic and osteoclastic behaviour [53]. On a molecular basis, the structure of bone converges around two key molecules: OPG and OPG-ligand, also known as RANKL [55]. With structural damage to the ECM, including tropocollagen uncoiling, fibrillar sliding and tearing, generation of microcracks, lamellar sliding [34] and macroscopic fracture, new remodelled bone is required.

6.2.3 Bone Healing Bone healing is complex. A highly regulated sequence of multi-tissue activated events, referred to as the regeneration cascade, proceeds at the initial point of healing. Overlapping of each developmental stage occurs, with multiple cell types and signal

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cascades acting in harmony [56, 57]. Unlike soft-tissue healing, bone healing does not leave an obvious scar, the process being regenerative, rather than solely reparative [32, 58, 59]. There are direct links between the mechanisms involved in initial skeletal tissue formation and the fracture healing cascade [60]. Fracture size and mechanical stability are crucial in defining the success of the healing mechanism [61]. The intricacies and multiple-tissue development involved in secondary bone healing have yet to be effectively incorporated in synthetic bone tissue engineering scaffolds. It is thought that the endogenous healing mechanisms activated in response to bone-graft implantation follow similar progressive regeneration as are involved in fracture repair. Indeed understanding the involvements that the various components play in fracture healing is a key feature for designing relevant scaffolds for bone grafts. Bone fracture is followed by an inflammatory response that involves haematoma formation, recruitment of MSCs and activation of inflammatory signals and osteogenic stimulators. A temporary cartilaginous matrix replaces the initial haematoma in order to stabilise the fracture site, which is crucial for facilitating subsequent angiogenesis and primary bone formation. Finally, the woven primary bone fracture callus is reorganised into functional bone. The biological processes, cellular involvement and predominating growth factors involved in each stage of bone regeneration have been previously reviewed in detail [56, 57, 62–64]. Many of these growth factors have successfully been incorporated into synthetic adjuvant scaffolding with therapeutic stimulation of bone repair.

6.3 Levels of Matrix Sophistication The evolving forefront of nanotechnologies, cellular and molecular microbiology and biomaterial science provides locus for developing highly effective tissue engineering materials on a range of size scales. Each level of sophistication provides an essential biomimetic design input to assist in instructing and stimulating the in vivo cellular niche.

6.3.1 Electrospun Fibers 6.3.1.1 Natural and Synthetic-Based Polymeric Nanofibers Electrospun fibers alone are by no means ideal for biomatrices, despite some preliminary success in bone tissue engineering [65]. They are, however, of crucial importance as they form the base-framework in the design of advanced bone tissue constructs since they are morphologically and dimensionally similar to collagen in the endogenous ECM. Initial studies were based on this premise and made important contributions to the mimetic prototypes. Investigation of cellular response on different fiber diameters have lead to a variety of results [66–70], few of which have been well substantiated [71].

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Cellular preference to fiber size is intrinsically difficult to isolate, as with changes in electrospun fiber size, there is subsequent change in the inter-fiber spacing [72, 73], mechanical properties [74] and biodegradability; all of which have been shown to mediate cellular behaviour and underpin the entire performance of the synthetic bone graft. Furthermore, depending on the cell phenotype, preferential fiber size and porosity for maximum cell adhesion may peak at multiple points [71]. It has been well shown that osteoblastic cells respond more favourably to nano-topographic cues demonstrating superior cellular adhesion, growth, proliferation and differentiation [75], similarly seen with co-incorporating nanofibers in microporous scaffolds [73, 76]. Interestingly, cellular behaviour on nanofibers was improved with enhanced levels of bio-stimulative factors in the culture medium [68, 69], thus suggesting synergistic interactions and the necessity for multiple microenvironmental inputs to satisfy the cellular niche. More recent developments have focussed on advanced electrospun systems and the incorporation of bone-specific bioactivity. As bone has the inherent capacity to regenerate, biodegradable materials have been predominately used. Synthetic polymers, most commonly polyesters (PLA, PGA, PCL, etc.) and their various copolymers, offer the advantage of substantial flexibility in electrospinning processability, mechanical modification, biodegradability and are of relatively low expense. Furthermore, of clinical translational importance, many have FDA-approved biocompatibility and have been widely investigated in other areas of tissue engineering and medical applications. Polymers, such as PCL, have the capacity to be synchronised with the redevelopment rate of endogenous bone in vivo, and the degradation by-products do not overly increase the cellular microenvironment acidity [77], making it a more viable solution for sustained cellular mobilisation and angiogenesis [77]. However, most synthetic polymers are hydrophobic (hence showing poor cell attachment) with limited bioactivity and functional groups for post-tethering of growth factors and proteins [78]. Electrospun natural polymers, which include collagen [79], gelatin [80], silk fibrin [81, 82] and chitosan [83], have the key advantages of exhibiting enhanced biofunctional motifs, are much more hydrophilic and in turn support superior cell adhesion [78]. Initial cell adhesion and survival is eminent in the consequential success and effectiveness of the bone graft. However, natural polymers have greater structural, chemical and molecular weight variability and a higher possibility of immune rejection. Furthermore, there is much less versatility in regards to electrospinning control and ease, in that only few solvents do not denature natural polymers, narrowing the window of solution and operating conditions to permit electrospinnability. There has been substantial investigation into combining the properties of both natural and synthetic polymers via electrospinning various polymer blends such as PCL with Hep-S [84], gelatin [85] and silk fibrin [78, 86]. The benefits of coincorporating natural and synthetic polymers have been demonstrated in a comparative study by Lee et  al. [87], highlighting that the addition of collagen into PCL significantly increases hydrophilicity, degradation rate, mechanical elongation, cell adhesion and growth in vitro, cellular infiltration in vivo, and up-regulation of OP, ALP activity and collagen type I in osteoprogenitor cells. Furthermore, the blending of SIS with PCL in an layered micro/nano dual fiber size construct provided

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approximately a fourfold increase in proliferation of BMSCs, possibly linked to an increase in hydrophilicity and thus cell adhesion [86]. Importantly, incorporation of layers of electrospun fibers without SIS had no significant impact on the cellular response. Blending PCL with Hep-S [84], a carrier for human MSCs, improved the in vitro growth of these osteoprogenitor cells. Heparin alone showed some improvement, but only half that of Hep-S. Indeed Hep-S has the intrinsic ability to regulate numerous functional growth factors involved in post-fracture bone [88]. When implanted, there was no protective carrier for the cells, so any impact of the Hep-S on bone formation could not be accurately evaluate given the likely intervention of host inflammatory responses [84]. 6.3.1.2 Inorganic Nanofibers A novel approach to overcoming the polymer nanofibers shortcomings (which include limited integration with the host tissue and insufficient mechanical properties) is to electrospin inorganic (ceramic) nanofibers, shown in Fig. 6.1. While the material processing of such nanofibers has been well reported [89], their involvement in the bio-industry has been relatively untapped, in particular their effects on cellular responses in vitro and in vivo. Inorganic electrospun fibers have predominately been obtained via electrospinning sol-gels of polymer and ceramic components, followed by heat treatment up to 700 °C [90] to burn the binding polymer off, thus leaving ceramic fibers. Nanofibrous matrices of silicate [91], bioactive glass [90, 92, 93], FAp [94] and HAp [95] are potential integrative materials for bone tissue regeneration. Bioactive glass has intrinsic anti-inflammatory and anti-microbial capabilities [96], the capacity to up-regulate the secretion of growth factors [97], and can stimulate mineralisation [92]. Cell viability and ALP activity of bone-marrow-derived human MSCs on bioactive glass nanofibers significantly superseded that of bioactive glass discs and PCL nanofiber controls [90]. Despite superior bioactivity and dramatic increases in modulus and mechanical strength, inorganic nanofibers have the limitations of being too brittle as bone-graft candidates [95, 98]. 6.3.1.3 Composite Nanofibers The hybrid material properties obtained via biodegradable polymers and bioactive particles within electrospun nanocomposites are intuitively more suitable for functional bone substitutes. Pre-blending HAp nanoplates [99, 100], carbon nanotubes [101], bioactive glass [102] and various calcium phosphates within electrospun biodegradable polymers has been a logical, exciting and successful approach for promoting cell anchorage and bone-specific bioactivity with improvements in fiber mechanical properties, nano-roughness and hydrophilicity. Advances in osteoconductivity (the recruitment and stimulated in-growth of osteoprogenitor cells and neo-blood vessels from the host tissue into the scaffold [59, 103]) and osteoinductivity (the capacity to induce bone formation directly [59, 103]) have been achieved with the integrated complexity of such a composite system.

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Fig. 6.1  Electrospun polymer, inorganic and composite fibers. (a) SEM image of electrospun PLLA (860 ± 110 nm) (adapted with permission from [100]). (b) TEM image of electrospun PVA/ collagen nanofibers (~245 nm) (adapted with permission from [104]). (c) SEM image of aligned HAp nanofibers post-removal of poly(vinyl pyrolidone) via heat treatment at 600 °C for 6 h. (d) High magnification SEM image of HAp nanofibers displaying fusing of adjacent HAp nanorods (diameter of ~7 nm and length of ~27 nm); (Both (c) and (d) adapted with permission from [95]). (e) SEM image of blended PLLA/HAp (wt ratio 4:1) electrospun composite fibers: 845 ± 140 nm (adapted with permission from [100]) (f) TEM image of composite PVA/collagen/HAp nanorod blended nanofibers (~320 nm) (adapted with permission from [104])

Apatite/Polymer Composite Fibers Biomimetic electrospun structures consisting of nano-HAp with various biodegradable polymers have been well explored [99, 104] (refer to Fig. 6.1). The effectiveness of nanoparticulate, reinforced polymeric matrices is critically dependent on the dispersion uniformity and binding affinity at the polymer–particle interface. Surfactant adsorption of hydroxysteric acid [105] and Hep [106], grafting of PCL [107], electrostatic dispersion via pH changes [108] and HAp silica coating [109] feature as some of the possible techniques to reduce the inherent tendency for nanoHAp particles to aggregate, thereby improving their interaction with the matrix. Recently, more sophisticated additions including incorporation of chemical variations of apatite have shown potent effects in stimulating the cellular response. TCP [110] and carbonated-apatite [111] dissolve faster than nano-HAp and are more readily taken up by osteoclasts, whereby an initially rich calcium and phosphorous microenvironment facilitates more rapid bone redevelopment. Stable bone-specific inorganics, such as FAp [94], sustain slower calcium dissolution rates, where combining it with HAp (40% FAp:60% HAp) enhanced osteoblastic cell survival and expression of OPG, OC and BSP-1 in vitro on flame-sprayed orthopaedic coatings [112]. Similarly, Arinzeh et al. reported a combination of HAp with b-TCP (20:80) promoted far higher MSC bone deposition than each individual component in vitro and in vivo [113]. Intuitively, a combination of inorganic crystals that can lead to rapid osteoconductivity and osteoinductivity for primary fixation stability and integration with the host tissue, while maintaining osteogenesis until the defect has healed, are best suited. However, dissolution is only one parameter that determines the choice of inorganic bioactive ceramics; structural integrity during degradation, biocompatibility, particle size, shape, crystallinity, surface chemistry, and binding and dispersive interaction with the polymer co-component are all vital for controlling the endogenous response.

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The necessity to provide a combination of bio-stimulative phases has been supported by statistically higher cell number and ALP activity of adult human BMSCs on combinations of HAp (50:50) with two different bioactive glasses than on each bioactive glass and HAp nano-rough ceramic surface alone [114]. The combinatorial benefits of HAp with bioactive glass has also recently been applied to Ti-orthopaedic implants and shown to supersede the initial cell adhesion, proliferation and sustained osteogenic differentiation of commercially available HAp coatings [96]. Carbon Nanotubes in Electrospun Fibers As a result of their highly specialised combination of aspect ratio (surface area), strength, stiffness, light weight and versatile surface chemical functionalisation [115], the incorporation of carbon nanotubes within electrospun polymers has provided a unique platform for bone tissue engineering. Dramatic increases in stiffness, strength and elongation have also been reported [116]. The addition of 0.02 wt% MWCNTs into silk fibrin nanofibers increased the modulus from 231.2 ± 13.7 MPa to 342.7 ± 16.2 MPa; tensile strength 6.7 ± 0.7 MPa to 10.0 ± 1.0 MPa, but elongation decreased from 16.8% to 11.7% [117]. However, the properties were dependent on dispersion, functionalisation and nanotube concentrations. Reduction in mechanical properties, slower biodegradation rate and increases in fiber diameter were observed with increases in non-functionalised MWCNT concentration beyond 0.5 wt% [116]. Similar to blending apatite nano-crystals, surface functionalisation has been critical for dispersion to enable the significant addition of nanotubes without aggregation and decline in material properties. These properties will also improve adhesion with the polymeric matrix [118] and biocompatibility [119]. PCL has been grafted onto MWNTs to create a hard-core/soft-shell structure, again targeted for “artificial bones” [120]. However, the addition of carbon nanotubes into electrospun fibers has had varied effects on scaffold mechanical properties. While covalently-functionalised carbon nanotubes may provide better dispersion throughout the fiber, their effect can be mechanically detrimental. Functional groups in the graphene lattice also act as defect points that can weaken the nanotube and corresponding electrospun fiber [121]. Biologically, the presence of nanotubes has also enhanced bone-specific bioactivity: fibrous blends of biocompatible MWCNTs and PLGA (2.15 mm), stimulate apatite mineralisation from SBF without the presence of any cells [101]. Furthermore, coblending of both HAp nanoparticles and MWNTs within PLGA [122] has been proposed as attractive nanocomposites for bone tissue engineering, exhibiting both attachment and proliferation of bone-marrow-delivered MSCs. Core-Shell Electrospinning Despite the advantages in pre-blending particles with biodegradable polymers to yield electrospun nanocomposite fibers, the spatial organisation within the fiber is

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still random. As a result, it is not yet possible to mimic the physical and chemical organisation of mineralised fibrils in native bone. However, improvements may be realised through core-shell electrospinning (also referred to as co-axial electrospinning). This is conceptually similar to single jet electrospinning, however, it involves a dual-orifice spinneret whereby two separate solutions are simultaneously fed though an outer and inner needle to create dual-layered nanofibers [123, 124]. One of the crucial advantages in core-shell electrospinning is that there are fewer restrictions on the solution properties, with the shell acting as a guide to form appropriately layered structures [124]. While not yet extensively explored, core solutions including low-molecular weight polymers, molecules with limited solubility, metallic and ceramic particles [124] and even cells (which have been demonstrated to maintain full functionality post-electrospinning) [125] are potential candidates for unique advances in bone tissue engineering.

6.3.2 Dual Porosity, Cellular Infiltration and Three-Dimensional Structures Scaffold three-dimensionality with interconnected microporosity for cellular infiltration, vascularisation and integration with the patient’s tissue is paramount for the clinical viability of the adjuvant synthetic matrix. Host tissue in-growth and consequential mechanical interlocking is of key importance for effective osseointegration and implant stability. Not only does the stability of the implant influence healing time, pain and physical support, the micromechanics imposed in the endogenous niche have a critical influence over the formation of neo-blood vessels, osteoprogenitor differentiation, proliferation, migration, inflammatory response [59, 126], and ultimately the quantity and quality of newly formed bone. This is one of the major inherent limitations with electrospun matrices: the pores are too small. The increasingly prevalent need to provide micropores while maintaining nanomimetic features and porosity for which cell nutrients and metabolic waste are transported [127] has lead to combining a range of fabrication technologies with electrospinning. Dual-scale porosity in electrospun nanofiber matrices has been improved through co-incorporating electrospun water-soluble sacrificial fibers [128–130], ice crystals [131–133], micron-sized fibers produced by heated melt deposition [76, 86, 134] or co-electrospinning [135], salt leaching [136, 137], salt leaching/gas foaming techniques [127, 138], drawing a metallic comb through the fibers to expand the mesh mechanically [139], using a metallic spherical dish as a collector of the electrospun fibers [140] and cutting pores out using laser ablation [141], some of which are shown in Fig. 6.2. While many of these approaches increase cellular infiltration and/or survival, a major pitfall that accompanies increased porosity is the subsequent reduction in mechanical strength. In an attempt to overcome this issue, Lee et al. [127] co-blended MMT nano-sized platelets as reinforcements in electrospun PLLA and used salt leaching/gas foaming to create a reinforced three-dimensional nanofibrous mesh

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Fig. 6.2  Methods for generating microporosity within electrospun nanofibrous scaffolds. (a) Schematic simultaneous electrospinning of PLLA with PEO onto a common rotating drum (concepts derived from [130]). (b) SEM image of electrospun PLLA/PEO (1:3), post-removal of PEO and fiber mineralisation in SBF (adapted with permission from [130]). (c) Electrospun PCL collected in an electrically grounded reservoir of water. Nanofibers drained into a lower reservoir where the vortex fluid motion confined PCL fibers into yarned bundles that were collected in reservoir 2. Slurry of PCL fibrous yarns and water from reservoir 2 was moulded into cylindrical three-dimensional structures, immersed in liquid nitrogen (−196 °C) for 15 min and freeze dried; (d) corresponding SEM image; ((c) and (d) adapted with permission from [133]). (e) Simultaneous deposition of salt particles (100–200 mm) during electrospinning of collagen/sodium hyaluronate

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with microsized pores. Opportunistic bone-specific strengthening systems as discussed in sect. 6.3.1.3 may be important in combination with dual-scale porous scaffolds to create fully functional three-dimensional bone-graft implants. Open microporosity may also be crucial in effective tissue regeneration, yet without active surface chemistry as discussed below, host infiltration is likely to be ineffective. It has been suggested that modified surface chemistry on its own can encourage tissue infiltration, and is likely to be a prerequisite. This has been supported through in vitro and in vivo studies of Hep-coated PLLA nanofibers, where in-growth of bovine aortic endothelial cells into three-dimensional scaffolds was improved with Hep coating, and interestingly through using aligned fibers [142]. More potently, as neural cell migration can be driven by neurotrophin-2 and nerve growth factor gradients [143], there may be potential for encouraging bone-specific infiltration via bone-specific gradients.

6.3.3 Surface Functionalisation The ability to modify the surface chemistry on the nanofibers facilitates manipulation and tuning of biological responses for bone regeneration. Greater surface exposure of bone-specific inorganics, cell adhesive molecules and growth factors attachment has been achieved through various novel surface modification techniques discussed below. The ability to expose multiple drugs and growth factors from electrospun systems is an exciting prospect for sophisticated bio-regenerative signalling, yet currently in its developmental infancy. 6.3.3.1 Surface Activation Techniques and Hydrophilicity An alternative technique for increasing cell affinity and biocompatibility of synthetic electrospun polymers involves modification of the surface of electrospun fibers via the attachment of hydrophilic chemical groups such as amides, carboxyl and hydroxyl groups [144]. In situ grafting of hydrophilic acrylic acid on to PGA, PLLA and PLGA electrospun fibers for increased presentation of carboxyl groups facilitated significant improvements in fibroblast attachment, spreading and proliferation in vitro [145]. Importantly, these reactive groups are prerequisites for providing docking sites for adsorption and/or the immobilisation of an extensive range of growth factors, proteins and peptides. Often spacer molecules are also attached onto the fiber via polymerisation grafting, to change the surface biochemistry and provide docking sites for immobilisation of more powerful biological molecules and to maintain bioactivity [144]. Fig. 6.2 (continued)  (80/20); (f) High magnification SEM image of dried scaffold post-collagen/ sodium hyaluronate cross-linking and salt leaching; ((e) and (f) adapted with permission from [136]). (g) Schematic stacking of PCL/collagen type I blended electrospun nanofibers with microfibers via direct polymer melt deposition; (h) SEM image of electrospun nanofibers on top of microfibers; ((g) and (h) adapted with permission from [76])

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6.3.3.2 Macromolecules, Natural Polymers and Signalling Factors Involved in Bone Healing Following surface pre-activation, collagen [146], gelatin [147] and fibronectin [29] have been tethered onto electrospun synthetic polymers to increase biofunctionality. Cell adhesive ligands, including RGD-containing peptides, have also been attached to electrospun synthetic biodegradable polymers, leading to enhanced control over attachment, spreading [148], proliferation and differentiation of a variety of osteoprogenitor stem cells [149, 150]. Owing to its high negative charge, Hep attachment onto electrospun fibers has been important for electrostatic adsorption of a variety of growth factors involved in bone morphogenesis, such as FGF-2 [151], VEGF, TGF-b [152] and PDGF [153]. Furthermore, dual-immobilisation of laminin and FGF onto Hep functionalised PLLA nanofibers showed enhanced bioactivity [154]. Conversely, Hep has been shown to suppress BMP-2 osteogenic bioactivity [155] and thus control over Hep location, concentration and timed exposure should therefore be taken into careful consideration in designing multiple growth factor delivery systems for bone tissue engineering. Not only does the selection of bio-signalling molecules have a great influence over regulating cellular response, the immobilisation mechanism is crucial. The two major surface attachment techniques include covalent conjugation and adsorption via secondary forces. Typically, conjugation has been linked with sustained signal exposure yet greater chances of protein bio-deactivation, while adsorption only supports burst delivery. Importantly, a comparison between these two techniques has been shown for attachment of BMP-2 [156]. BMPs can stimulate osteoblastic differentiation of MSCs in  vitro [157], initiate bone osteoinduction and osteoconduction in endochondral and intramembranous regenerative cascades [158], and even play a concurrent role in mediating angiogenesis [159]. Conjugation of BMP-2 onto cast PCL films showed greater concentrations of BMP attached with a slower accumulated release, in contrast to adsorption of BMP [156]. Greater ALP and OC expression from BMSCs (15 days in culture) was observed with covalent attachment of BMP-2, suggesting that this conjugation maybe a more favourable attachment mechanism [156]. 6.3.3.3 Calcium Phosphate Coatings One convenient technique utilised in developing bone-specificity on electrospun fibers has been coating biodegradable fibers with HAp and various calcium phosphates via immersion in SBF [160]. Surface presentation of selective chemical groups on biodegradable fibers is a prerequisite for stimulating surface nucleation and growth of HAp when incubated in calcium- and phosphate-rich SBF solutions [161]. Various techniques have been used to create surface active groups which can support the nucleation of calcium phosphate phase, some of which include plasma treatment [145, 162], alkaline erosion [163] and LbL coating [164]. An elegant study by Cui

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et al. utilised covalent grafting of chitosan on amino-functionalised PDLLA fibers to facilitate controlled HAp deposition with uniform surface coverage. The presence of chitosan and mineralisation showed improved proliferation and differentiation of osteoprogenitor cells in comparison to controls [165]. However, a distinct disadvantage in surface deposition of HAp, rather than blending within the fiber, is that there is virtually no improvement in mechanical properties. 6.3.3.4 Electrospin/Electrospray Systems Utilisation of electrospraying has been recently proposed as an alternative technique to enable strategic engineering control over the spatial organisation of each constituent phase in electrospun nanocomposites [166, 167]. Francis et  al. [168] recently utilised simultaneously electrospraying of HAp and electrospinning of gelatin fibers onto a common rotating drum collector. Gelatin fibers were crosslinked, which increased the mechanical strength and elongation. As HAp particles were electrosprayed onto the surface of the gelatin fibers, this resulted in greater HAp exposure and enhanced surface “nano-roughness,” alleviating problems such as the loss of HAp platelet bioactivity that may result if blended within the fiber. This spin-spray system provided appreciable enhancement in human foetal osteoblast proliferation, mineralisation and ALP after 15 days of culture, in comparison to blended gelatin/HAp electrospun fibers [168]. Spraying technologies also offer potentially more effective incorporation of mechanically mismatched materials, which may otherwise be detrimental if physically mixed together. However, despite various improvements in mechanical properties, material technologies are yet to be able to produce electrospun bioactive composites that can be utilised as stand-alone, temporary bone substitutes. 6.3.3.5 Layer-by-Layer Surface Modification of Nanofibers Alternate electrostatic assembly of anionic and cationic electrolyte layers is a highly attractive means of altering nanofiber surface chemistry and coating a nano-thin biological reservoir of growth factors, drugs and proteins onto the electrospun fibers. Within the cellular environment, growth factors are released by degradation of interlayer bonding as a result of aqueous, enzymatic or cellular activity. A wide variety of lipid vesicles, nucleic acids, DNA, proteins and more recently growth factors have been immobilised in LbL coatings with broad biological benefits [169–171]. Higher concentrations of molecules can be accumulated in a robust and controlled manner by simply increasing the number of assembled layers [144, 170]. Loading efficiency currently limits the clinical application of many growth factors such as rhBMP-2 [172]. Encapsulation of rhBMP-2 in cross-linked poly-l-lysine/hyaluronan LbL films, which could be coated onto nickel-titanium implant surfaces, maintained near zero-order release kinetics and protein bioactivity over 10 days, and promoted in  vitro differentiation of myoblasts into osteoblasts [173]. A variety of bonding

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forces, beyond the most common electrostatic interactions, including hydrogen bonding, hydrophobic interaction and covalent bonding, can also be used to control the biomolecular retention and release kinetics [174].

6.3.4 Cellular-Based Matrices The design of artificial microenvironments for repair needs to accommodate the specific cell types and their interface with region-specific physical and pharmacological parameters. With bone, these should be specific to those activated during endogenous regeneration. As cells ultimately define the repair process, their involvement with the scaffold is an eminent step in therapeutic tissue engineering. Various cell-encapsulating devices have been implemented to increase delivery efficiency, longer-term viability and function as opposed to direct cellular re-injection. MSCs can proliferate and differentiate on two-dimensional culture surfaces with the addition of growth factors, proteins, cytokines and hormones active in the bone healing cascade along with various anti-inflammatories and antimicrobials added into the culture medium [175]. Once dense differentiated populations of MSCs are achieved, they can be seeded on three-dimensional electrospun scaffolds as discussed throughout this review and correspondingly implanted. Alternatively, proliferation could be encouraged directly on the scaffolds in  vitro. There have been few innovative approaches to cell-encapsulation within electrospun scaffolds. Recently, this has began to change with research involving electrospraying of cells [176], encapsulating cells inside individual electrospun fibers [125, 177] and three-dimensional stacking of cells between electrospun layers [178].

6.4 Summary and Future Perspectives There is an ever-increasing need to improve the clinical management of skeletal disorders. While not necessarily life threatening, these conditions severely impact quality of life and present a major economic burden on society. Recently, the process of bone repair has received great impetus from progress made in engineering (nanotechnology) and biomedicine (bone stem and progenitor cells). Innovative strategies to combine each field have provided a rational and strategic framework to transform the clinical management of skeletal disease repair. Establishing a nano-framework using natural and synthetic polymers, in combination with apatite crystals, functionalised nanotubes and/or bioactive glass underpins the mechanical properties of scaffolds that are critical for functional bone development. It is envisaged that combining techniques will yield more biomimetic and structurally sound synthetic ECM. Importantly, improved design of smart biomatrices that target bone regeneration is of great importance. Surface engineering via adsorption immobilisation, covalent conjugation and LbL coatings provide unique platforms

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for releasing/presenting growth factors that stimulate bone regeneration from electrospun scaffolds. It is visioned that cooperative interaction of tissue engineering at multiple levels of matrix sophistication with stem cell-based therapies will ultimately expedite bone healing. The challenge is how they can be strategically combined.

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Chapter 7

Directing Cell Fate Through Biomaterial Microenvironments Kelly Clause, Jonathan Lam, Tatiana Segura, and Thomas H. Barker

Abstract  Biomaterials offer discrete advantages over standard ECM systems, like matrigel, in the context of both fundamental stem cell biology and control of stem cell fate/phenotype. In particular, one can specifically design features into the 3D microenvironment with high levels of control. The fundamental limitation to date is that we currently lack the design rules for eliciting specific cellular and/or multicellular behaviors that may lead to true regenerative medicine. As the matrix biology and stem cell biology fields mature, biomaterial scientists continually look to these fields for inspiration and understanding of what features should be considered in the design of the “optimal” material for their specific application. In this chapter, we outline the basic emerging biological concepts leading toward a first set of design rules; those include physical or mechanical signals, chemical and/or biochemical signals, spatial orientation and positioning of signals, and finally, time-resolved display of signals. We highlight current biological findings that support these design rules (physical, chemical, and x, y, z, and t) and outline current efforts to develop biomaterial systems that enable both the decoupling or integration of these design criteria for tissue engineering and regenerative medicine applications.

Abbreviations ECM EDC EGF MSCs

Extracellular matrix N-(3-Dimethylaminopropyl)-N¢-ethylcarbodiimide hydrochloride Epidermal growth factor Mesenchymal stem cells

T.H. Barker (*) The Wallace H. Coulter Department of Biomedical Engineering at Georgia Tech, Emory University, 313 Ferst Drive, Atlanta, GA 30332, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_7, © Springer Science+Business Media, LLC 2011

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PDGF PEG PLA PLGA RGD Shh sulfo-NHS VEGF

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Platelet-derived growth factor Polyethylene glycol Polylactide Poly(lactic-co-glycolic acid) Arginine-glycine-aspartic acid Sonic hedgehog N-Hydroxysulfosuccinimide Vascular endothelial growth factor

7.1 Introduction The use of biomaterials as a platform for the presentation of bioactive molecules to direct cellular behaviors, such as stem cell differentiation, leading to regenerative medicine applications has lead to only relative success. The fundamental basis for the majority of biomaterial technologies is the homogeneous presentation of simple ECM mimetics and growth factors and while these systems have established a strong foundation, they have not displayed the level of success envisioned. Part of this lack of success is due to inadequate knowledge on the biological “design rules” that one needs to follow in order to direct cell behaviors. As the field of biomaterials slowly converges with the fields of matrix biology, mechanobiology, and stem cell biology, new information becomes available that may provide key insights into these design rules. In this chapter, we outline the core concepts emerging in the design of biomaterials toward the specific regulation of stem cell phenotype. Current biological advances clearly outline the need for biomaterial systems that are designed to display specific mechanical properties and/or physical cues, i.e., stiffness and elasticity, specific chemical and/or biochemical cues, i.e., adhesion ligands and growth factors, and the integration of these cues in all four dimensions, specifically the three spatial coordinate axes x, y, z, and time.

7.2 Mechanical Signals Mechanical forces have long been implicated in regulating many physiologic processes and pathologic processes. A cell responds to mechanical signals depending upon its substrate/environmental material properties and the surroundings it interacts with. Individual molecules, cells, and tissues are exposed to several kinds of mechanical forces and stresses. At an organism level, morphogenic events in which changes in tissue shape or position via mechanical forces occur. On a tissue scale, muscles are tensed over bones in load-bearing events and shear stress is exerted by blood flow against the blood vessel wall. On a cellular level, cell generated forces and externally applied forces from the ECM act upon each other. These mechanical

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cues are then converted into biochemical signals at the molecular level to provide and ensure the correct cellular response, i.e., the cells and their environment reach a “mechano-homeostasis.” Given recent technological advances, significant insights into mechanotransduction (how cells receive these forces and transduce them into biochemical and functional responses) have been and continue to be made and are becoming an increasingly important component in biology. As a consequence of this increase in our understanding of how mechanics regulates cell behavior, it becomes essential to consider mechanical properties in the design of biomaterials.

7.2.1 Mechanotransduction Many molecules, cellular components, and extracellular structures have been shown to contribute to mechanotransduction including focal adhesions (specifically integrins). Integrins, because of their role in both transmitting mechanical forces (both from within the cell due to actomyosin contractility and from the extracellular environment) and regulating a number of intracellular signaling pathways, have emerged as a critical component in mechanotransduction that can be engaged specifically through biomaterials design. Integrins, a large family of heterodimeric transmembrane glycoproteins composed of many different combinations of a and b subunit pairs, function to attach cells to ECM proteins and/or ligands on other cells. Both global and local effects are observed when integrins are manipulated by exerting different types of forces on them. Integins respond to external application of force and environmental stress fields locally by enhancing their association with cytoskeletal proteins via integrin clustering and activation and strengthening and enlarging small focal complexes into mature focal adhesions. Integrins respond to strain globally by influencing the cytoskeletal tension of the cell and by initiating cell signaling cascades that can control cell responses [1, 2]. As the cells become stiffer and resist mechanical deformation at higher levels of applied stress, force transfer correlated with the cell stress levels and actin cytoskeletal rearrangements takes place [3, 4]. These studies suggest that local forces are transmitted via integrins onto the actin cytoskeleton and are then propagated throughout the entire cell [5, 6]. The maturation of focal adhesions via integrin clustering and activation is an actin polymerization-dependent process and leads to increased stiffness of the integrin-cytoskeletal connections [2, 7]. For example, integrin binding to the ECM protein fibronectin locally stiffens the link between integrins and the cytoskeleton in proportion to the stiffness of the substrate and in an integrin-fibronectin binding site occupancy-dependent manner [1]. While initial responses to stress include integrin-cytoskeletal attachment strengthening or resistance against displacement, later responses involve biochemical changes and activation of signaling cascades [1, 2]. Focal adhesions contain several tyrosine kinases and phosphatases, as well as scaffolding proteins and molecules that link integrins to the actin cytoskeleton (for in-depth reviews, see refs. [8–10]).

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Thus, integrins transduce forces coming from both outside and inside the cell and respond by stiffening their link with the cytoskeleton as well as forming mature focal adhesions. This ensures cell shape and tensional balance is maintained when local force is applied as well as the induction of a number of signaling cascades that control the cell response. This subset of studies suggests that there is an inherent interplay between integrin engagement and mechanotransduction. Both physical and chemical (discussed later) material attributes are critical to driving cell adhesion, focal adhesion formation, and force transduction. Thus, altering the biophysical attributes of the material could elicit altered integrin clustering and downstream signaling cascades, and altering the biochemical attributes of the material (e.g., the immobilization of celladhesion motifs) could alter the integrin specificity of cellular interaction, which is also likely to have significant affects on how the cell senses and responds to the mechanical environment.

7.2.2 Mechanics Cues from Developmental Biology While mechanical forces via integrins or other cellular components are capable of modulating almost all aspects of cell function, they can also directly affect the form and function of tissues, especially during embryogenesis and as a result may provide important design considerations for biomaterials in stem cell therapy and tissue regeneration. The mechanical contraction process (compaction of the morula) of embryogenesis causes the first polarized cell structures to form at the 8-cell stage of development. The cells actively generate tensional forces within their cytoskeleton (actomyosin contraction) and apply these forces to cell–cell contacts with neighboring cells. At the next stage, the inner and outer cells then rearrange to form a hollow ball, the blastula, which invaginates in on itself to form a two-layered structure in a process known as gastrulation. In Xenopus laevis, dorsal involuting and noninvoluting are controlled by tissue mechanical signals – not the external forces that are generated in a different place in the embryo [11]. Similarly, laser-ablation studies have suggested that global cellular movements and dorsal closure in a number of species are also regulated mechanically [12]. Flow also plays a role in the mechanics of embryogenesis. The leftward movement of fluid at the ventral node, called nodal flow, is the central process in left– right asymmetry formation. Nodal flow is autonomously generated by the rotation of cilia that are tilted toward the posterior on cells of the ventral node [13]. Interstitial flow, the movement of fluid through the ECM of tissues, is driven by physical stresses and can affect the distribution of factors that are necessary for correct development and pattern formation [14]. These important tissue-level mechanics are important considerations when designing biomaterials. The development of biomaterials systems that enable these larger-scale gross movements and mechanical events have yet to be realized and provide opportunities for further advancement.

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7.2.3 Translating Mechanical Signals to Biomaterials Given the growing body of evidence supporting the role of mechanical forces in development and maintenance of homeostasis of tissues in precise spatio-temporal patterns, there is both a need and an opportunity for translating those mechanical signals to biomaterials. Chemical and biological modifications to biomaterials can directly influence cell behavior by altering substrate properties, surface interactions, microenvironment architecture, etc., and ultimately manipulating signal transduction pathways [15] and will be covered in the following subsection. With regard to altering mechanics, a number of new materials offer a wide range of properties and/or controlled responses that may be used for directing specific cell behaviors. Hydrogels, crosslinked polymer networks that are insoluble but swellable in aqueous medium, provide an environment that more closely resembles the native hydrated state of the ECM. Through the modulation of crosslinking, one can alter the mechanical properties at the macroscale, e.g., elasticity, of the hydrogel. Several types of crosslinking strategies have emerged, with more under development, that allow not only control of cell-level mechanics via crosslinking but also, through the incorporation of structural pores, enable modification of gross biomaterial mechanical properties, including interstitial fluid flow patterns [16, 17]. Smart biomaterials, materials that can respond to their surroundings, are an emerging class of materials that may provide dynamic mechanical features. This is attractive for controlling drug release/availability, cell adhesiveness, mechanical properties, or permeability in response to specific stimuli such as pH, temperature, and light [16]. Coupling these emerging technologies with biomaterial design enables one to precisely control dynamic events in response to cellular needs and to potentially realize mechanical events on the scale of developmental biology, i.e., tissue scale movements. Biologically inspired materials, such as self-assembling materials or biomimetic materials, are designed to mimic properties or processes that occur in the native organism, specifically at the level of ECM and may provide a means to recapitulate microarchitectural features such as surface topology that are known to influence cell behavior. Self-assembly is based on the formation of weak noncovalent bonds or hydrophobic interactions and can form distinct 3D structures such as micelles, vesicles, and tubules in the presence of an external stimulus (i.e., pH, temperature, or salt) [16, 18]. Biomimetic materials tend to have inherent patterning that can emulate the ECM and tissue architecture. The primary focus has been on incorporating short oligopeptides, with the most studied sequence being RGD, others have been reviewed elsewhere [19]. Similarly modifying material micro- and nanoscale physical parameters can also be an important component in translating mechanical signals to biomaterials. Microscale and nanoscale modifications to the material, density, porosity, and component size or alignment can also impart a specific local geometry to the material that can be as important as the material chemistry [20, 21]. The specific architecture of the native ECM provides geometric cues to cells in the form of fiber diameter,

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length, and crosslinking patterns, as well as surface irregularities. Surface specificity can be obtained by incorporating short oligopeptides that can exhibit specific binding domains or whole proteins on the biomaterial surface which are important in modulating cell adhesion. Surface microarchitecture depends on fiber diameter and alignment, which can also play a role in porosity. Thus, surface modifications, surface chemistry, and topography of a biomaterial are important parameters that influence cell morphology, adhesion, and migration [16, 20].

7.2.4 Mechanics in Stem Cell Differentiation We have illustrated that biomaterial mechanical modifications as well as mechanotransduction via integrins can have an effect on cell shape (i.e., spreading). Cell shape is a key regulator of many aspects of development, including differentiation. Stem cells can also undergo differentiation depending on deformational forces (i.e., tension or compression), the ECM (i.e., its geometry or elasticity), as well as a number of other factors. Changes in cell shape, via mechanical cues, and binding of specific growth factors and ECM proteins to their respective cell surface receptors can switch cells between discrete fates of growth, differentiation, apoptosis, and migration [22]. Changing cell shape from a round to flattened morphology can profoundly alter the organization of the actin cytoskeleton and the assembly of focal adhesions [23, 24] and reliably switch MSCs between different lineages (i.e., osteoblastic versus adipogenic) [25]. The importance of the ECM on stem cell fate has been shown with particular emphasis on the interactions of ECM ligands with cell surface receptors and ECM geometry or elasticity. Engler et al. recently showed that culturing MSCs on matrices of increasing stiffness, which resemble the compliance of various tissues in  vivo, affected lineage specification: soft substrates induced a neurogenic commitment, intermediate compliance resulted in differentiation into myoblasts, while cells grown on the stiffest matrices gained an osteoblast fate [26]. In addition to the macroscale changes to matrix stiffness, topographical patterns, either micro- or nanoscale, of the ECM could also be potent regulators of stem cell differentiation [27–29]. Stem cell differentiation can thus be determined by matrix properties at the macro-, micro-, and nanoscale. Thus, controlling the mechanical environment of biomaterials at the macro-, micro-, and nanoscale can further improve our ability regulate stem cell differentiation. However, stem cell fate decisions can be complicated not only by the biophysical and biochemical aspects of the material (reviewed below) but also by the cell biophysical aspects (i.e., tissue-specific patterns of ligand and receptor expression) or sequential autocrine and paracrine inductive loops [30]. This indicates that more precise spatial and temporal control of material factors, either mechanical or chemical, is necessary (reviewed below) to create an environment that is similar to the native stem cell niche.

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7.3 Incorporating Bioactive Signals into Biomaterials As the field of regenerative medicine and tissue engineering continues to expand and evolve, researchers become increasingly better at creating biomaterials that mimic the natural ECM. The native ECM is a heterogeneous mixture that consists of different protein fibers, proteins, hormones, and glycosaminoglycan chains. This network not only provides structural support and mechanical cues as described above, but also a myriad of physically coupled and soluble signals that drive cell fate. As shown in Fig. 7.1, the molecular/biochemical signals can be split into three main categories: (1) insoluble molecules (fibronectin, vitronectin, laminin, etc.), (2) soluble molecules (growth factors, cytokines, chemokines, etc.), and (3) proteins via cell–cell contact (cadherins, CAMs, etc.). These signals work in concert to determine the fate of the cell: replication, differentiation, migration, or apoptosis. A number of different growth factor delivery approaches have been used to incorporate them into 3D hydrogel matrices. Key points to consider when incorporating signals into biomaterials include: loading capacity, distribution of growth factor, binding affinity, release kinetics, long-term stability of protein, and economical viability.

Fig. 7.1  A myriad of signals play a role in directing and controlling cell behavior. The dynamic and intricate nature of the ECM affects cell response and its subsequent fate. These signals include: soluble/bound signals (i.e., growth factors, cytokines, etc.), physical signals (i.e., fibronectin, laminin, etc.), and cell–cell interactions (i.e., cadherins, cell-adhesion molecules, etc.). In engineering synthetic matrices, some of the strategies researchers have utilized to control and direct cell fate include: creating short peptide fragments, controlling distribution of bioactive signals through a clustering effect, and spatially/temporally patterning signals into the matrix

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The most straightforward method of growth factor delivery involves the direct loading of unmodified proteins into the biomaterial. Typical release profiles of unmodified proteins consist of an early “burst-release” which occurs during the initial swelling of the matrix. It is difficult to completely change the release kinetics; however, modifications to the matrix can be made to help limit the initial release of the growth factor. Modification to the hydrogel, such as altering the crosslinking density, can help alter the release profile of the signal. In addition, changes can be made to the water content during gelation, amount of crosslinker, introducing additional double bonds, decreasing the molecular weight of monomer, and/or adding additional reactive groups to modify signal release [31]. Alternative alterations to the hydrogel matrix have been studied to prevent the quick release of bioactive signals. Introducing positively charged molecules to the matrix induces secondary binding interactions between negatively charged DNA and proteins which can change the release profile [32]. Furthermore, the inclusion of hydrophobic segments to reversibly adsorb proteins can be used to control the signal diffusivity [33]. However, these alterations can change the biomaterial’s swelling properties and should be considered during matrix design. Whereas these previously described systems rely on the passive diffusion of bioactive signals in static conditions, other biomaterials have been designed to release bioactive signals based on mechanical stimuli. To take advantage of the dynamic, ever-changing state of native re-modeling tissue, hydrogels have been developed to release unbound molecules upon mechanical stress. Compression of the biomaterial leads to increased pressure within the biomaterial and the subsequent release of the bioactive signal. Once the hydrogel relaxes, additional growth factors are freed from the matrix and become free to release out of the hydrogel when placed under compression again. Bioactive signals can be immobilized to the biomaterial to allow for a longer and more controlled release profile. Some signals, like adhesion peptides, must be bound to the biomaterial as a rigid mechanical support for the signal to carry out its function. Both desires can be fulfilled through covalent bonds between the biomaterial and signal. Side chain functional groups of amino acids have been used to covalently link proteins to the polymer backbone of a biomaterial. When choosing the amino acid to link, it is essential to avoid amino acids that affect the protein’s functionality and activity; special care must be taken in choosing the type of chemistry for linkage. Michael addition chemistry using the thiol group on a cysteine and unsaturated groups on the substrate is becoming an increasingly popular method for signal incorporation. The effect of neighboring charged amino acids on the Michael-type reactivity of these thiol groups has been characterized and provides researchers with further control over reaction kinetics and selectivity [34]. This chemistry has been utilized to incorporate an RGD adhesion peptide (Ac-GCGYGRGDSPG-NH2) to vinylsulfone-functionalized PEG (PEG-VS) in a study that uses synthetic PEG hydrogels as matrices for in situ bone regeneration [35]. Similar hydrogel studies have been performed linking cysteine containing RGD peptides to acrylated hyaluronic acid hydrogels as well. Furthermore, VEGF

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has been modified to contain unpaired cysteine residues for covalent linkage to PEG-VS [36]. An alternative method of covalently incorporating signals into a hydrogel is to modify the N-terminus of the peptide. It has been shown that N-hydroxysuccinimidylactivated ester can be used to link the N-terminal a-amine of an RGD sequence containing peptide to an acrylate moiety [37]. This modified peptide was copolymerized with PEG diacrylates to form a peptide containing gel solution. In a similar vein, a 2D study conducted by Fan et al. showed that EGF covalently bound to poly(ethylene oxide) via the N-terminus promoted MSC survival when compared with soluble EGF. Additional studies have shown covalent binding of other growth factors (i.e., TGFb1, bFGF, TGFb2) to PEG chains [38]. Carbodiimide chemistry has also been used to bind growth factors to a biomaterial. Carboxylic acids in a porous collagen scaffold were activated using sulfo-NHS and EDC and reacted with amine groups on VEGF for immobilization [39]. EDC chemistry has also been used to couple proteins like rhBMP-2 onto biomaterials [40]. A newer strategy employing ultraviolet irradiation has been shown to successfully immobilize growth factors and adhesion peptide sequences onto a substrate. This technique requires the presence of a photoreactive compound to drive the formation of a new bond. One study was able to introduce a photoreactive compound to EGF by coupling it with N-(4-azidobenzoyloxy)succinimide [41]. The azidobenzoyl-derivatized EGF was coupled to polystyrene through ultraviolet exposure for 10 s. This immobilization of EGF was shown to enhance the growth of anchorage-dependent cells seeded onto the plates. An alternative to coupling the signal itself with the photoreactive compound is to add a photoreactive compound that causes bond formation. Studies by Jennifer West’s group have used 2,2-dimethyl2-phenyl-acetophenone in N-vinylpyrrolidone as a photoinitiator to cause polymerization between PEG diacrylate monomers and acrylated RGD peptides and VEGF [42, 43]. These chemistries not only continue to expand the tools available to researchers in designing matrices for tissue regeneration but can be used to spatially pattern biomolecules within the biomaterial. The methods previously described in linking biomolecules to tissue engineering matrices often require the alteration or addition of a functional group. An alternative approach is to bind the growth factors to ECM proteins and/or heparin/heparin sulfate through electrostatic interactions. Because these interactions have been shown to have a natural regulatory role in growth factor activity in native tissue, it may prove to be a useful tool in designing matrices for tissue regeneration. The most prominent ECM/growth factor interaction involves heparin/heparin sulfate [44]. Heparin is a linear polysaccharide mostly composed of repeating uronic acid and glucosamine units. It has a highly negative charge that enables binding to several different families of growth factors. This interaction helps stabilize the growth factor against proteolysis and thermal denaturation. To take advantage of this, researchers have found ways to bind heparin to their matrix. The presence of heparin allows for the incorporation of growth factors into the biomaterial without having to modify the protein itself.

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One type of chemistry begins by functionalizing heparin with maleimide groups. These molecules can be reacted with thiol-containing PEG polymers to form a PEG–heparin complex. Other PEG molecules can be crosslinked to form a heparin decorated hydrogel. Results show that this hydrogel was able to bind bFGF with the release dependent on hydrogel erosion [45]. In a similar study, heparin was modified with methacrylate groups and then polymerized with methacrylated PEG to form hydrogels capable of delivering bFGF. Results showed that human MSCs cultured inside these gels differentiated into osteogenic cells [46]. Other ECM components like collagen and gelatin have been shown promise in delivering growth factors as well. For systems that require a longer time-release profile, signals can be introduced through microparticle carrier systems that provide protection for the molecule. The signal is encapsulated inside the microparticles and offer another level of control over release kinetics as the carriers can be made with various degradable and nondegradable materials that can be tuned to specific needs. As the microparticle degrades, the signal will be released to act on cells inside and neighboring the hydrogel. Some of the more popular polymers used to synthesize the microparticle shells are PLA and PLGA. These polymers are ideal for biological studies because they exhibit good biocompatibility. Furthermore, the polymer degradation rate can be controlled by adjusting the molecular weight of polymer, crystallinity, and ratio of subunits. Different release profiles of fluorescently labeled albumin were observed when altering the degradation properties of the microparticles [47]. PLGA has been used to encapsulate bFGF via a “water-in-oil-in-water” double emulsion process. These particles were added to an alginate matrix to study their release in vitro. The growth factor remained active as the PLGA shell degraded and induced the proliferation of cardiac fibroblasts. In vivo experiments showed that the bFGF, once released from the scaffold, was able to promote blood vessel formation in rat peritoneum [48]. Microspheres constructed of PLGA have also been used to deliver VEGF. These microparticles, loaded into a dextran-based hydrogel, resulted in a steady-state release of the growth factor [49]. A landmark study performed by Dr. Mooney’s group showcased a system that was able to release two different growth factors with different release kinetics. They incorporated one growth factor directly into the polymer (VEGF), and encapsulated another (PDGF) into polymer microspheres [50]. The scaffold was implanted into subcutaneous pockets in rats. The VEGF exhibited a “burst-release” and the PDGF, whose release was determined by microsphere degradation, exhibited a much steadier release. Histological analysis showed that the scaffold and its dual-growth factor release approach helped form a mature vascular network [50]. An alternative to degradable polymer microparticles are solid lipids. One group loaded the lipid microparticles with insulin and showed that it was remained biologically active after release. Chondrocytes grown on a fibrin gel with the microparticles showed higher amounts of collagen formed per cell when compared to controls. However, the collagen formed was not as high as a positive control consisting of fresh insulin supplements added directly [51].

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7.4 Engineering Heterogeneous (x, y, z) and Dynamic (t) Biomaterials It is clear that mechanical and (bio)chemical modifications to biomaterials can directly influence cell behavior, but despite this, controlling the desired cell response is still difficult. Natural tissue morphogenesis depends on complex interactions within a dynamic four-dimensional environment (x, y, z, and t). Physical and biochemical signals are required at specific locations and times to result in the complex structures found in nature. The goal is to engineer microenvironments that display control in the four dimensions to guide and study tissue morphogenesis.

7.4.1 Coordinate Control of x, y (2D) and x, y, z (3D) in Biomaterials We have already described a number of methods for altering the properties of biomaterials, however, the spatial control of these signals can impart further control. A number of strategies have been employed to introduce 2D and 3D control into biomaterials. Modifications to either the material properties: material integrity, crystallinity, crosslinking density, overall micro- and macroporosity, etc., or the addition of chemical or biological stimuli can be done in both 2D and 3D settings. 2D (x, y) approaches allow well-controlled analysis of the impact of individual components on the cells, whereas 3D (x, y, z) approaches allow reconstruction and realization of the complexity of the natural tissue. In 2D materials, individual signal molecules can be displayed on the substrate, or combinatorial mixtures of signals (i.e., generated by protein spotting) can be displayed. Similarly, protein spot sizes can also be controlled, which when using cell adhesion or cell-regulatory proteins can allow control of cell shape [52]. In addition to the above mentioned, the desired substrate stiffness can be controlled by differential crosslinking of the material. Lutolf et  al. used microcontact printing, a technique that crosslinks chemically altered proteins to a hydrogel, to structure hydrogel arrays topographically to contain thousands of spatially segregated micropatterns (i.e., round microwells with proteins printed specifically at the bottom of each well) to observe specific alterations in cell division (asymmetrical vs. symmetrical self-renewal) [53]. Three-dimensional patterning approaches can be designed further to allow the spatial arrangement of cues all the way down to the level of an individual cell. Current microfabrication methods for 3D hydrogel matrices with controlled intrinsic structures mainly include photolithographic patterning, microfluidic patterning, electrochemical deposition, and 3D printing [54]. The arrangement of cells (either hetero- or homotypic) can also be spatially controlled (via electropatterning) to affect the cell response. Electropatterning localizes live cells within hydrogels by using dielectrophoretic forces. Using these technologies, cell–cell interactions have been modulated by varying cell cluster size in one recent study which then resulted

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in altered synthesis of glycosaminoclycans of chondrocytes (large clusters produced smaller amounts per cell) [55]. Morphogen gradients of individual signals on a small-scale with tethered signals or on a large scale with nontethered signals can be included to provide further regulatory cues to govern cell responses to 3D biomaterials [15, 22, 52].

7.4.2 x, y and x, y, z Biomaterial Control Directing Stem Cell Fate In the case of stem cells, their fate is dependent on, among other things, the local biochemical niche defined by growth factors, cell–cell contacts, and cell–matrix adhesion sites. Since tissues have a number of different cell types working in concert, the ability to locally control specific regions inside the same biomaterial is paramount in designing effective matrices for tissue regeneration. A number of the above approaches have been used in controlling stem cell fate. Khademhosseini et al. used micropatterned composites of hyaluronic acid hydrogels with photoreactive methacrylates and docked or encapsulated mESCs to promote cell viability. These cells could also be manipulated to colonize specific patterns and shapes based on the material microarchitecture [56]. A laser-based layer-by-layer (LbL) stereolithography technique has been used to incorporate precise, pre-designed spatial patterning of molecules into 3D scaffolds to create complex microenvironments more indicative of native tissues. Mapili et al. used this technique to photocrosslink RGD and an ECM component capable of sequestering growth factors (heparin) to an acrylated polyethylene glycol to form a 3D scaffold with specific pore/channel dimensions. They also demonstrated that the spatial pattern could be maintained during the culture period [57]. Bioprinting uses custom-designed inkjet printers to deposit onto solid substrates, in a controlled layer-by-layer fashion, combinatorial mixtures of liquid precursors of hydrogel networks and/or cell in minute volumes (picoliter sized drops) and at a high rate (tens of thousands per second) and density. In one of the first examples of this strategy, 3D PEG-hydrogel arrays were produced to screen for the individual and combinatorial effects of gel degradability, cell-adhesion-ligand type, and cell-adhesion-ligand density on the viability of human MSCs. Increased PEG-network degradability and greater cell-adhesion-ligand density were both found to increase the viability of the stem cells in a dose-dependent manner [58]. Using photopolymerized hydrogels (which can be cleaved by a controlled light source to modulate biophysical and biochemical properties locally), MSCs responded to local changes in stiffness and cell-adhesion properties [22]. In a densely crosslinked gel, the decrease in crosslinking density obtained through cleavage of the backbone of the photolabile chain induced a significant morphological change in encapsulated stem cells (initially round, they became more spread). Moreover, the manipulation of cell-adhesive peptide ligands (RGD) led to inducible change in chondrocyte differentiation [22]. This is by no means an exhaustive list on the fast expanding approaches of using spatially controllable biomaterials to control stem

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cell fate, however, there is yet another factor, time, that must be considered when attempting to regulate physiologic cell responses.

7.4.3 Temporal Control in Biomaterials It is clear that the precise coordination of and interpreting spatial and temporal cues from the microenvironment is essential for stem cells to create complex, functional tissues. Temporal control can be introduced to biomaterials by using diffusion, polymer degradation, and environmental stimuli-driven release. There are two basic modes of controlled release via diffusion: (1) the molecule of interest is uniformly dispersed or dissolved, or (2) the molecule of interest is contained in a core, which is surrounded by a material, and it is released by diffusion through this rate-controlling material. In these modes, the drug is released either by passing through the pores or between polymer chains, and these are the processes that control the release rate [59]. Another way of controlling release via diffusion is to use water as the main agent controlling the release of the drug. In these materials, the molecules cannot physically diffuse out of the device without water molecules diffusing in. There are generally two types of water penetration-controlled systems [60]. Swelling-controlled materials incorporate molecules into a hydrophilic polymer that swells in an aqueous environment; however, it is difficult to control the release rates. Osmotically controlled delivery materials have a semipermeable membrane that allows water to move in, but prevents salt and the molecule from exiting. Instead, the molecule exits through an opening in the material caused by an increase in pressure due to the volumetric increase. Degradable materials can also be used to incorporate temporal control over the release of signaling molecules. These materials erode (with or without changes to the chemical structure) or degrade (breakdown of the main chain bonds) as a result of exposure to chemicals or enzymes and thus release the molecules. Delivery of growth factors and chemicals can also be mediated using degradable particles encapsulated within the biomaterial for timed release of signaling molecules [61]. Osteogenic studies of rat MSCs using recombinant human TGFb1 encapsulated in polymer blends of PEG-PLGA particles has been reported by Peter et al. showed that growth factor delivery enhanced MSC proliferation and transformation into osteoblasts [62]. Responsive biomaterials, described in detail earlier, can be externally regulated systems (like mechanical pumps) via magnetism or ultrasound or they can be regulated by environmental stimuli. In these materials, the release is in direct response to the conditions detected, be it bulk-triggered (i.e., in response to temperature, pH, or concentration), local-triggered (i.e., in response to light), or materials with a hybrid of the two. Light responsive materials can be triggered by conformational changes of dye molecules in the material backbone or by heat generation of dye molecules linked to a thermally responsive material [63]. Combining thermal responsive materials with metallic nanoparticles has been used to generate hybrid materials. The metallic nanoparticles interact with light at wavelengths through

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scattering and absorption that raises the particles surrounding’s temperature when illuminated. They can be embedded into the material using the thermal responsive properties [64]. Cell-triggered release where release is mediated by cell-released proteases has been explored to deliver nanoparticles with temporal control [65]. Enhanced understanding of the mechanics and signals within the microenvironment that regulate cell fate has led to the development of increasingly sophisticated biomaterials. Biomaterials with precisely controlled scaffold architecture that regulate the spatio-temporal release of growth factors and morphogens, and respond dynamically to microenvironmental cues are necessary to recapitulate the microenvironmental stem cell niche.

7.4.4 Temporal Control Directing Stem Cell Fate Temporal control of stem cell differentiation in the form of sequential factor addition has been shown to be effective in generation of functional motor neurons and pancreatic cells. Motor neurons differentiate through two stages, each requiring a distinct set of soluble factors. First, signaling by BMP, FGF, and Wnt proteins, next by retinoic acid-mediated differentiation followed by the progression to terminally differentiated motor neurons in the presence of Shh. Jessell et al. were able to guide mouse ESCs down this pathway by applying these soluble factors in a step-wise manner that emulates natural neural development, resulting in the in vitro generation of motorneurons that can survive and engraft in vivo [66]. The application of soluble signals in sequence (Shh-antagonists then FGF10 then Notch inhibition) to undifferentiated embryonic stem cells in vitro resulted in the generation of functional insulin-producing cells [67]. These and other studies highlight the importance of temporal control in stem cell differentiation in which biomaterial control has been added. Dexamethasone released from PLGA microspheres combined with a hyaluronic acid hydrogel affected the chondrogenic differentiation potential of MSCs in a release-dependent manner [68]. Similarly, proliferation and chondrogenic differentiation of MSCs cultured with gelatin hydrogel microspheres releasing TGFb1 was enhanced [69]. Thus, major advances in the understanding of the spatial and temporal cues in the microenvironment necessary for directing lineage commitment and differentiation of stem cells are necessary to incorporate these factors into biomaterials to better control the cell responses.

7.5 Conclusions Current biology has established the critical importance of physical and chemical signals in directing both individual cell phenotypes and coordinated, multicellular tissue morphology, as well as the spatial and time dependence of these environmental cues. As the biomaterials field progresses, it would be sage advice to look to

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biology, specifically matrix biology, mechanobiology, and stem cell biology for the design rules operative in the development of functional biomaterials that facilitate the delivery of stem cells and other cellular therapies toward the realization of regenerative medicine. The biomaterials field has come a long way in the development of bio-inspired design and it is only through the continued pursuit of biomaterials that display dynamic responses in the presence of both environmental and cellular cues that we will realize the full potential of stem cells in regenerative medicine.

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Part III

Engineered Tissue

Chapter 8

Basic Considerations with Cell Sheets Masayuki Yamato and Sebastian Sjöqvist

Abstract  The foundation of cell sheet engineering is to harvest cells in a minimally destructive way. This way, the cells that retain their cell-to-cell contact show a higher expression of surface proteins, and retain their more highly developed functions than cells harvested enzymatically. In contrast to using biodegradable scaffolds for tissue engineering, cell sheets triggers a significantly lower immune response and do not cause an excess of extracellular matrix to be secreted around the transplanted cells. Cell sheets also tend to integrate more easily with host tissue. In this chapter, the basic concept and history of cell sheet engineering are discussed, including some of our own studies. This chapter focuses on using cell sheets for constructing corneas, minimizing postoperative complications after esophageal surgery, and transplantation of hormone-producing endocrine tissues.

Abbreviations ATP ECM EDTA EMR ESD fT3 fT4 LCST LSCD

Adenosine triphosphate Extracellular matrix Ethylenediaminetetraacetic acid Endoscopic mucosal resection Endoscopic submucosal dissection Free triiodothyronine Free thyroxine Lower critical solution temperature Limbal stem cell deficiency

M. Yamato (*) Institute of Advanced Biomedical Engineering and Science, Tokyo Women’s Medical University, Tokyo, Japan e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_8, © Springer Science+Business Media, LLC 2011

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PEG Poly(ethylene glycol) PIPAAm Poly(N-isopropylacrylamide)

8.1 History of Cell Sheet Engineering The most commonly used method for regenerative medicine is single cell suspension injection. There are, however, limitations to this approach such as low integration to host tissue, difficulties in controlling size, shape, and location of injected cells as well as a risk for embolization [1]. Another method using biodegradable scaffolds has shown some success in creating tissues with low cell concentrations, such as bone and cartilage. After degradation, the space previously occupied by the scaffold is filled with large amounts of ECM that makes it less suitable for cell-rich tissue engineering [2]. Furthermore, due to limitations of passive diffusion, the interior of larger scaffolds tend to become necrotic [3], and inflammation is commonly seen during degradation [4]. Harvesting cells with the help of temperature-responsive dishes was developed in the 1990s with the aim of minimizing cell damage normally seen with conventional enzymatic cell harvesting techniques (Fig.  8.1) [5]. The key to this

Fig. 8.1  Two types of cell harvest. Confluent cells are subjected to two types of harvest methods. Conventional trypsinization digests almost all the membrane proteins as well as deposited ECM (left). Therefore, all cells are harvested as single cells. To the contrary, cell sheet harvest by temperature reduction does not need proteolytic enzymes, so that all cultured cells are harvested as a single contiguous cell sheet together with deposited ECM (right). Green materials depict deposited ECM (adapted with permission from Elsevier [1])

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technique, use of the thermoresponsive polymer PIPAAm, a material with reversible hydrophobic–hydrophilic properties, is discussed below. Endothelial cells and hepatocytes harvested this way were shown to maintain highly developed cell functions compared to cells harvested enzymatically [6]. In contrast to enzymatic treatment, cell-to-cell connections remained intact after harvesting, which would be essential for some tissues, for example, in cardiac tissue engineering [7]. Cell sheets harvested this way also maintain cell–ECM interactions [6], presumably the reason why they easily attach to other surfaces [8] without the need for suturing [9–11] or tissue glue. In addition, such cell sheets exhibit no [12] to low amounts [13] of inflammation in the recipient tissue, as well as sustained display of cell surface proteins [14] and cellular functions [6].

8.2 Temperature-Responsive Polymer PIPAAm has the unique property of shifting from hydrophobic to hydrophilic properties at temperatures below LCST of 32°C in water (Fig. 8.2) [5]. Below LCST, the polymer becomes hydrophilic and therefore is extended and hydrated, while above LCST, the polymer is hydrophobic, has a globular shape, and is dehydrated. The mechanism of cell release is still not yet fully understood. There is support for a two-step mechanism; a passive step where the polymer is hydrated and an active step mediated by cell metabolic activity. Studies with ATP-generation inhibitor, and the fact that detachment as well as metabolic activity are reduced in lower temperatures, support this theory [15]. The amount of deposited ECM from the cultured cells is an important factor for how long detachment takes. The amount of ECM varies greatly with different cell types and in some cases detachment is not achievable. To combat this, porous membranes and co-grafting with PEG have been used. The hydration of PIPAAm is important for detachment. Normally, this occurs only from the periphery of the cell sheet, but with the help of porous membranes, water can access the polymer from beneath the cells, hydration can be greatly accelerated, and

Fig. 8.2  Thermoresponsive polymer. Above the LCST, the polymer has a globular shape, is hydrophobic, and cells attach to the surface. When temperature is lowered below LCST, the polymer is extended, hydrophilic, and the cells detach (adapted with permission from Wiley-VCH Verlag GmbH & Co. KGaA [16])

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Fig. 8.3  Cell detachment time. Average detached areas for cell sheets recovered from PIPAAm-tissue culture polysterene (TCPS) (open square), PIPAAm-porous membrane (PM) (filled square), PIPAAm(PEG0.1%)-PM (open circle), and PIPAAm(PEG0.5%)-PM (filled circle) surfaces with culture area of 4.2 cm2 as a function of incubation time in culture medium at 20°C (adapted with permission from Elsevier [1])

detachment times can be reduced (Fig.  8.3). Polymer thickness is also of great importance. Layers thinner than 15 nm fail to detach cells while layers thicker than 30 nm show cell repellent properties even above LCTS [16].

8.3 Clinical Applications 8.3.1 Corneal Regeneration Corneal epithelial cells are constantly shed and removed in tears. To maintain corneal homeostasis, cells need to proliferate and migrate from stem cells located in the vascularized limbus area [17]. Severe trauma such as chemical or thermal burns, and eye diseases such as Stevens-Johnson syndrome or ocular pemphigoid may result in total LSCD [9]. Such absence often results in defective or abnormal corneal re-epithelization that predisposes the cornea to opacification and neovascularization [18]. Unilateral LSCD can be treated by autologous transplantation of limbal stem cells from the healthy eye, however, this requires a large graft and there is risk of causing bilateral LSCD. Allograft is another possibility, but requires long-term

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immunosuppression, with accompanying risks such as infection. For these reasons, we developed a method for culturing cell sheets for corneal restoration from auto­ logous oral mucosal cells. Oral mucosal cells were selected because this epithelium resembles the target tissue [19] and there had been earlier reports using a similar approach in animal models [20, 21]. In our studies [19], a 3 × 3 mm biopsy of oral mucosal tissue was excised under local anesthesia (Fig.  8.4). The cells were separated from the epithelial layer by dispase treatment after which trypsin and EDTA were used to create single cell suspensions. Mitotically inactivated NIH 3T3 fibroblasts were used as a feeder layer in tissue culture wells, but were separated from the mucosal cells by temperatureresponsive cell-culture inserts. After 14  days of culture, the cell sheets were harvested by simple temperature reduction. The patients’ conjuctival and subconjunctival tissues were removed with a 3 mm margin outside the limbus. The cell sheet was moved from the dish to the transplantation site with the help of a doughnutshaped supporter and adhered to host tissue without the need of sutures. Histologically, the cultured cell tissue, with 3–5 cell layers, resembled corneal epithelium. The optical transparency of the cell sheets was equivalent to cell sheets cultured from limbal cells [19]. Ultrastructural features such as microvilli, tight junctions, desmosomes, hemidesmosomes, and basement membrane resembled those of corneal epithelium in vivo [19]. Immediately after surgery, the cornea was clear (Fig. 8.5). Maximum corneal transparency was achieved by 2 weeks after transplantation and was maintained during the 14-month follow-up period [19].

8.3.2 Wound Healing After Esophageal Surgery Endoscopic surgery of early esophageal cancer has been recommended since the 1980s. Modern ESD shows less incidence of recurrence than conventional EMR [22]. A common complication of both ESD and EMR, however, is stricture of esophagus due to postoperative inflammation [10]. In 2006, we reported successful trials using cell sheets to improve healing in animal models [10], and clinical trials are now in progress. For animal studies, 10 × 10  mm2 biopsies were taken from oral mucosa and treated enzymatically to create single cell suspensions. Cells were seeded on culture dishes with a temperature-responsive central area. After 2 weeks, cells were harvested as intact sheets by temperature reduction. ESD was performed on test subjects leaving a 180° ulceration, 5 cm in length (Fig. 8.6). Histological analysis of the engineered cell sheet indicated that it resembled esophageal epithelium in terms of stratification and histochemical markers (i.e., cytokeratins) [10]. Four weeks after ESD, the transplanted animals showed complete wound healing while controls were still in intermediate healing with a visible fibrin mesh (Fig. 8.7). The results of this study imply that cell sheet transplantation could prevent postoperative constriction in humans.

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8.3.3 Pleural Defect Closure Intraoperative air leak is an inevitable complication of thoracic surgery [23]. The incidence of prolonged air leaks (over 7 days) is reported to be 18.2% after pulmonary resection, leading to a 17% rate of readmission and 29% rate of prolonged hospital stay [24]. Different methods for treatments of air leak include suturing, stapling, and electrocautery, as well as fibrin glue and other biological sealants, however, these approaches have met with limited success. An ideal air leak sealant must be easily applied, adherent to visceral pleura, and also able to withstand pressure from breathing and coughing, while not restricting lung expansion and contraction during the respiratory cycle [25]. In 2007, we reported successful trials using cell sheets derived from skin cells in small animals (rat and rabbit) [23]. These early preclinical trials showed encouraging results, and led to subsequent experiments in a pig model. For these, skin biopsies were taken from the abdomen of pigs. Excised specimens were cut and treated enzymatically to create single cell suspensions, which were seeded on temperature-responsive dishes and expanded in culture. Three weeks after biopsy, cell sheets were harvested by temperature reduction. Histological analysis revealed 1–3 stratified cell-dense layers with relatively little ECM (Fig. 8.8). A 15 × 10 mm lung puncture was made in recipient animals, and injury to visceral pleura was confirmed by the presence of air bubbles in the thoracic cavity. The cell sheets were transplanted using a square-shaped support membrane and attached to host tissue without the use of glue or suture (Fig. 8.9). To reinforce the seal, a second sheet was transplanted on top of the first one. Closure of air leak was confirmed by the absence of air bubbles. The seal was strong enough to withstand air leaks at up to 25 cm-water airway pressure. Pressure over 30 cm-water caused air leak, but once the pressure was lowered, the seal once again achieved integrity [23]. Four weeks after transplantation the sheet maintained a tight seal, while also allowing for normal lung expansion and contraction [23]. Histologically, the cells were tightly

Fig. 8.4  Creating artificial cornea. Transplantation of autologous tissue-engineered epithelial-cell sheets fabricated from oral mucosal epithelium. Panel a shows the removal of oral mucosal tissue (3 by 3 mm) from patient’s cheek. Isolated epithelial cells are seeded onto temperature-responsive cell-culture inserts. After 2 weeks at 37°C, these cells grow to form multilayered sheets of epithelial cells. The viable cell sheet is harvested with intact cell-to-cell junctions and extracellular matrix in a transplantable form simply by reducing the temperature of the culture to 20°C for 30 min. The cell sheet is then transplanted directly to the diseased eye without sutures. In panel b (the scale bar represents 50 mm), harvested cell sheets have 3–5 cell layers and do not resemble the original oral mucosa as shown in panel c (the scale bar represents 100  mm) as closely as they resemble normal corneal epithelium as shown in panel d (the scale bar represents 100 mm). Panel e shows a transmission electron micrograph of developed microvilli on the apical surface of the cell sheet. Specimens of human tissue-engineered epithelial-cell sheets harvested by reducing the temperature of the culture are immunostained green with antikeratin 3 antibodies (panel f), antib1-integrin antibodies (panel g), and anti-p63 antibodies (panel h). The nuclei in panels f, g, and h are shown in red. The scale bars represent 50 mm in panels f, g, and h. The specimens in panels b, c, and d are stained with hematoxylin and eosin (adapted with permission from The New England Journal of Medicine [19])

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Fig. 8.5  Results of transplantation. Eyes of patients before and after transplantation of sheets of tissueengineered autologous epithelial cells. These photographs were taken just before transplantation of the cell sheets and postoperatively at 13, 14, or 15 months (adapted with permission from The New England Journal of Medicine [19])

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Fig. 8.6  Transplantation to esophagus. Endoscopic transplantation of autologous oral mucosal epithelial-cell sheets. After endoscopic submucosal dissection, a flat esophageal ulcer is created (a). The cultured oral mucosal epithelial cell sheet, attached to a white PVDF support membrane, is then grasped by endoscopic forceps and transferred to the dissection site (b) and gently placed on the ulcer wound bed (c). After carefully withdrawing the endoscopic forceps (d), the endoscopic mucosal resection tube is used to apply gentle pressure to the PVDF support membrane and the underlying cell sheet (e). The cell sheet along with the support membrane is then left undisturbed for 10 min to allow for direct attachment to the host tissues (f). The support membrane is then easily removed (g), leaving the autologous cell sheet on the ulcer wound bed (h) (adapted with permission from BMJ Publishing Group Ltd. [10])

attached to pulmonary parenchyma without any air spaces present between transplant and host tissue [23]. The cell sheets were thicker than the original tissue, but had several histological similarities such as an abundance of ECM and expression of fibronectin and elastin [23]. ECM has an important role in providing a strong and expandable framework for the thin alveolar walls. Positive vimentin staining revealed presence of ECM-producing fibroblasts. In addition, no keratin expression, indicative of epithelial tissue, was seen in the area [23].

8.3.4 Endocrine Uses Maintaining the correct hormone level after surgery or in the presence of autoimmune syndromes, such as thyroid disease or diabetes, remains a great challenge. Using tissue engineering to restore endocrine tissue and physiological hormone homeostasis has been the focus of many research studies. However, studies involving the systemic delivery of cells, such as the injection of insulin-producing islets cells into the hepatic portal system, spleen, abdominal cavity, and other organs have shown limited success at achieving a functional benefit, in this case, the ability to regulate blood glucose levels in patients with diabetes. Complications such as low integration and inflammatory reactions have been reported [26]. Using cell sheet engineering may prove advantageous for these conditions, as discussed below.

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Fig. 8.7  Results of cell sheet transplantation. Transplantation of oral mucosal epithelial cell sheets promotes wound healing and reduces postoperative inflammation. Left and right panels represent transplant and control groups, respectively. (a, c) Endoscopic photographs taken 4  weeks after operation. (b, d) Macroscopic images of the esophageal sites receiving endoscopic submucosal dissection, after 4 weeks. (e, f) Hematoxylin and eosin (H&E) staining of the central portions of the ulcer sites. (g) H&E staining of the border region between the transplanted cell sheet and the outer portions of the ulcer site. (h) Comparison of the number of inflammatory cells present in the surgical sites between transplant and control groups (adapted with permission from BMJ Publishing Group Ltd. [10])

8.3.4.1 Thyroid Tissue Primary hypothyroidism is a common disease with an annual incidence of 3.5 per 1,000 women and 0.6 per 1,000 men in the UK. About 3% of the population receives long-term thyroid hormone replacement therapy, but studies indicate that 40–48% of patients are either under- or overtreated [27]. In 2009, we attempted to transplant

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Fig. 8.8  Cell harvesting and sheet histology. Autologous dermal fibroblast sheets can be harvested by low-temperature treatment. Dermal fibroblasts are harvested as intact sheets from temperatureresponsive culture surfaces with the use of a square-shaped supporter (a). Hematoxylin and eosin (b) and Azan (c) staining show that the fibroblast sheets are composed of 1–3 cell layers with relatively little extracellular matrix. Scale bars indicate 1 cm in (a), 50 mm in (b), and 50 mm in (c) (adapted with permission from Elsevier [23])

Fig. 8.9  Repairing lung puncture. Transplantation of autologous dermal fibroblast sheets to seal lung punctures. For transplantation procedures, the thoracic cavity is surgically opened (a) and a 15 mm diameter, 10 mm deep incision is made in the right lung (b). Air leaks are confirmed by the presence of air bubbles on the lung surface (c). The harvested cell sheet along with the supporter (d) is then placed directly on the pleural surface over the defect site (e). After allowing 5 min for the cell sheet to attach to the lung surface, the supporter is carefully removed (f). A second autologous fibroblast sheet is then transplanted directly over the first (g) and the support membrane is removed (h). The transplanted cell sheets immediately act to seal the air leak sites (i) (adapted with permission from Elsevier [23])

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Fig. 8.10  Thyroid hormone-producing cell sheets. (a) Functional analysis by free triiodothyronine measurement in serum. Significant differences (*p < 0.05) were observed in the 1/4 cell sheet transplantation group at 3 and 4 weeks versus the total thyroidectomized group. (b) Functional analysis by free thyroxine measurement in serum. In the cases of all cell sheet-transplanted groups, the levels of free thyroxine were increasing significantly. Also, significant differences (*p < 0.05) versus the total thyroidectomized group were seen in the 1/4 cell sheet transplantation group at 3, 4, and 5 weeks after total thyroidectomy and in the 1/8 cell sheet transplantation group at 4 and 5 weeks after total thyroidectomy (adapted with permission from Mary Ann Liebert, Inc. publishers [28])

allogenic thyroid cell sheets to thyroidectomized rats [28]. For these studies, thyroid glands were dissected from 4-week-old rats. Cells were treated enzymatically, cultured for 1 week and seeded on temperature-responsive dishes until confluent, then harvested through temperature reduction. The resulting cell sheets were transplanted to thyroidectomized recipient rats. Functional analysis of the grafts was accomplished by weekly blood analysis of fT3 and fT4. Both hormone levels increased in a time-dependent manner (Fig. 8.10), and the size of the transplanted sheet tended to be of importance [28]. However, none of the thyroidectomized groups reached the levels of nonthyroidectomized control animals. Histological examination of cultured cell sheets showed typical follicle morphology and antithyroid transcription factor-1 staining revealed a follicle epithelial cell lining on the follicle inner surface (Fig. 8.11) [28]. Analysis of engrafted cell sheets showed that these had become thicker, had organized into honeycomb-like structures, and contained colloid, microvessels with red blood cells and parafollicular cells, indicative of tissue very similar to native thyroid gland [28]. 8.3.4.2 Islet Cell Tissue Islet cell transplantation by infusion into the portal vein has been used clinically for more than 10 years. Forty-four percent of patients show restoration of insulin production for 1 year; however, this number decreases to 14% after 2 years [29]. We wanted to examine if cells harvested as sheets, with sustained cell-to-cell

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Fig. 8.11  Histological analysis of thyroid cell sheet 4 weeks after transplantation. (a) Hematoxylin– eosin staining demonstrates the presence of reconstructed thyroid tissue. (b) Microvessels including red blood cells can be recognized (arrows). (c) Immunostaining for the antithyroid transcription factor-1 antibody demonstrates the presence of follicle epithelial cells surrounding colloid. (d) The microvessels can be recognized inside the thyroid cell sheet (arrows) (adapted with permission from Mary Ann Liebert, Inc. publishers [28])

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Fig. 8.12  Insulin-producing cell sheets. In vitro insulin secretion assay using islet cells following glucose challenge with 3.3 and 20 mM (n = 6 cells/group). Insulin secreted into the culture media was measured by ELISA (modified with permission from Elsevier [26])

Fig. 8.13  Transplantation of an engineered sheet of islet cells into the subcutaneous space. (a) Sheet of islet cells attached to a support membrane (Su) was transplanted into the subcutaneous space of the Lewis rat. Five minutes after transplantation, the islet cell sheet was found to be well attached to the surrounding tissue and the Su was removed. (b–d) Microscopic observation of the engineered monolayer sheet of islet cells (arrows) engrafted in the subcutaneous space after 7 days following the transplantation procedure. Transplanted tissues from Lewis rats were processed into 5-mm-thick sections, and either (b) stained for H&E, or immunostained for insulin (c) and glucagon (d). (e) PKH26 red fluorescent cell membrane labeling for viable islet cells within the engineered islet sheet. Scale bar 1 cm (a); 50 mm (b–e) (modified with permission from Elsevier [26])

contact, could contribute to improved survival and functionality. Pancreatic islets were isolated from 7- to 8-week-old Lewis rats and treated enzymatically to obtain single cell suspensions. One challenge was to find a suitable configuration of dish surface to allow both attachment and detachment of the cells. For successful detachment, the amount of PIPAAm had to be increased compared to commercially available culture dishes, and attachment was enhanced by adding laminin-5 [26]. Functional analysis was performed by exposing the cells to different concentrations of exogenous glucose. Insulin secretion increased at higher concentrations of glucose (Fig. 8.12). Transmission electron microscopy revealed intact cell-to-cell connections (desmosomes) as well as islet cell characteristics, such as secretion

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granules and exocytosis near the plasma membrane [26]. The cell sheet was transplanted using a support membrane to subcutaneous space on the animals back (Fig. 8.13). Within 5  min, the sheet was attached to host tissue and the support membrane could be removed without loss of cells. Cell sheets were observed to survive for at least 7 days [26].

8.4 Conclusions Cell sheet tissue engineering has shown many advantages to other regenerative medicine techniques. Compared to biodegradable scaffold techniques, cell sheet transplantation does not cause inflammation, enables improved integration and attachment to host tissue, and may be more suitable for cell-dense tissue like heart and liver. Single cell suspension injections are widely used for regenerative medicine. To create cell suspensions, however, enzymes such as trypsin are used to harvest the cells from the culture dishes. These enzymes damage many of the cell surface proteins important for cell function and tissue localization. When using temperature-responsive dishes described in this chapter, cells are harvested from the dish without the need of enzymes. This results in donor cells with sustained cell-to-cell contact, intact adhesion proteins, and appropriate ECM. In the tissues we describe, the cell sheet attaches easily and integrates into the host tissue without the need for suture or tissue glue. The use of cell sheets in human clinical trials of artificial cornea transplantation using oral mucosal cells, and using cell sheets to accelerate wound healing and alleviate stricture formation in esophageal surgery, are highly encouraging. These data suggest that cell sheets have great potential and may be useful in other clinical scenarios.

References 1. Yamato M, Akiyama Y, Kobayashi J, Yang J, Kikuchi A, Okano T (2007) Temperatureresponsive cell culture surfaces for regenerative medicine with cell sheet engineering. Prog Polym Sci 32(8–9):1123–1133 2. Matsuda N, Shimizu T, Yamato M, Okano T (2007) Tissue engineering based on cell sheet technology. Adv Mater 19(20):3089–3099 3. Yang J, Yamato M, Kohno C, Nishimoto A, Sekine H, Fukai F, Okano T (2005) Cell sheet engineering: recreating tissues without biodegradable scaffolds. Biomaterials 26(33):6415–6422 4. Yang J, Yamato M, Sekine H, Sekiya S, Tsuda Y, Ohashi K, Shimizu T, Okano T (2009) Tissue engineering using laminar cellular assemblies. Adv Mater 21(32–33):3404–3409 5. Okano T, Yamada N, Okuhara M, Sakai H, Sakurai Y (1995) Mechanism of cell detachment from temperature-modulated, hydrophilic-hydrophobic polymer surfaces. Biomaterials 16(4):297–303 6. Kushida A, Yamato M, Konno C, Kikuchi A, Sakurai Y, Okano T (1999) Decrease in culture temperature releases monolayer endothelial cell sheets together with deposited fibronectin matrix from temperature-responsive culture surfaces. J Biomed Mater Res 45(4):355–362

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7. Shimizu T, Yamato M, Kikuchi A, Okano T (2003) Cell sheet engineering for myocardial tissue reconstruction. Biomaterials 24(13):2309–2316 8. Kikuchi A, Okuhara M, Karikusa F, Sakurai Y, Okano T (1998) Two-dimensional manipulation of confluently cultured vascular endothelial cells using temperature-responsive poly(Nisopropylacrylamide)-grafted surfaces. J Biomater Sci Polym Ed 9(12):1331–1348 9. Nishida K, Yamato M, Hayashida Y, Watanabe K, Maeda N, Watanabe H, Yamamoto K, Nagai S, Kikuchi A, Tano Y, Okano T (2004) Functional bioengineered corneal epithelial sheet grafts from corneal stem cells expanded ex vivo on a temperature-responsive cell culture surface. Transplantation 77(3):379–385 10. Ohki T, Yamato M, Murakami D, Takagi R, Yang J, Namiki H, Okano T, Takasaki K (2006) Treatment of oesophageal ulcerations using endoscopic transplantation of tissue-engineered autologous oral mucosal epithelial cell sheets in a canine model. Gut 55(12):1704–1710 11. Hannachi I, Yamato M, Okano T (2009) Cell sheet technology and cell patterning for biofabrication. Biofabrication 1(2):022002 12. Shimizu T, Yamato M, Isoi Y, Akutsu T, Setomaru T, Abe K, Kikuchi A, Umezu M, Okano T (2002) Fabrication of pulsatile cardiac tissue grafts using a novel 3-dimensional cell sheet manipulation technique and temperature-responsive cell culture surfaces. Circ Res 90(3):e40 13. Miyagawa S, Sawa Y, Sakakida S, Taketani S, Kondoh H, Memon I, Imanishi Y, Shimizu T, Okano T, Matsuda H (2005) Tissue cardiomyoplasty using bioengineered contractile cardiomyocyte sheets to repair damaged myocardium: their integration with recipient myocardium. Transplantation 80(11):1586–1595 14. Kanzaki M, Yamato M, Hatakeyama H, Kohno C, Yang J, Umemoto T, Kikuchi A, Okano T, Onuki T (2006) Tissue engineered epithelial cell sheets for the creation of a bioartificial trachea. Tissue Eng 12(5):1275–1283 15. Cooperstein M, Canavan H (2010) Biological cell detachment from poly(N-isopropyl acrylamide) and its applications. Langmuir 26(11):7695–7707 16. Fukumori K, Akiyama Y, Kumashiro Y, Kobayashi J, Yamato M, Sakai K, Okano T (2010) Characterization of ultra-thin temperature-responsive polymer layer and its polymer thickness dependency on cell attachment/detachment properties. Macromol Biosci. doi:10.1002/ mabi.201000043 17. Dua H, Azuara-Blanco A (2000) Limbal stem cells of the corneal epithelium. Surv Ophthalmol 44(5):415–425 18. Gomes J, Geraldes Monteiro B, Melo G, Smith R, Pereira C, da Silva M, Lizier N, Kerkis A, Cerruti H, Kerkis I (2010) Corneal reconstruction with tissue-engineered cell sheets composed of human immature dental pulp stem cells. Invest Ophthalmol Vis Sci 51(3):1408–1414 19. Nishida K, Yamato M, Hayashida Y, Watanabe K, Yamamoto K, Adachi E, Nagai S, Kikuchi A, Maeda N, Watanabe H, Okano T, Tano Y (2004) Corneal reconstruction with tissue-engineered cell sheets composed of autologous oral mucosal epithelium. N Engl J Med 351(12): 1187–1196 20. Nakamura T, Endo K, Cooper L, Fullwood N, Tanifuji N, Tsuzuki M, Koizumi N, Inatomi T, Sano Y, Kinoshita S (2003) The successful culture and autologous transplantation of rabbit oral mucosal epithelial cells on amniotic membrane. Invest Ophthalmol Vis Sci 44(1): 106–116 21. Kinoshita S, Nakamura T (2004) Development of cultivated mucosal epithelial sheet transplantation for ocular surface reconstruction. Artif Organs 28(1):22–27 22. Wong V, Teoh A, Fujishiro M, Chiu P, Ng E (2010) Preemptive dilatation gives good outcome to early esophageal stricture after circumferential endoscopic submucosal dissection. Surg Laparosc Endosc Percutan Tech 20(1):e25–e27 23. Kanzaki M, Yamato M, Yang J, Sekine H, Takagi R, Isaka T, Okano T, Onuki T (2008) Functional closure of visceral pleural defects by autologous tissue engineered cell sheets. Eur J Cardiothorac Surg 34(4):864–869 24. Seely A, Ivanovic J, Threader J, Al-Hussaini A, Al-Shehab D, Ramsay T, Gilbert S, Maziak D, Shamji F, Sundaresan R (2010) Systematic classification of morbidity and mortality after thoracic surgery. Ann Thorac Surg 90(3):936–942, discussion 942

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25. Kanzaki M, Yamato M, Yang J, Sekine H, Kohno C, Takagi R, Hatakeyama H, Isaka T, Okano T, Onuki T (2007) Dynamic sealing of lung air leaks by the transplantation of tissue engineered cell sheets. Biomaterials 28(29):4294–4302 26. Shimizu H, Ohashi K, Utoh R, Ise K, Gotoh M, Yamato M, Okano T (2009) Bioengineering of a functional sheet of islet cells for the treatment of diabetes mellitus. Biomaterials 30(30):5943–5949 27. Vaidya B, Pearce S (2008) Management of hypothyroidism in adults. BMJ 337:a801 28. Arauchi A, Shimizu T, Yamato M, Obara T, Okano T (2009) Tissue-engineered thyroid cell sheet rescued hypothyroidism in rat models after receiving total thyroidectomy comparing with nontransplantation models. Tissue Eng Part A 15(12):3943–3949 29. Shapiro A, Ricordi C, Hering B, Auchincloss H, Lindblad R, Robertson R, Secchi A, Brendel M, Berney T, Brennan D, Cagliero E, Alejandro R, Ryan E, DiMercurio B, Morel P, Polonsky K, Reems J, Bretzel R, Bertuzzi F, Froud T, Kandaswamy R, Sutherland D, Eisenbarth G, Segal M, Preiksaitis J, Korbutt G, Barton F, Viviano L, Seyfert-Margolis V, Bluestone J, Lakey J (2006) International trial of the Edmonton protocol for islet transplantation. N Engl J Med 355(13):1318–1330

Chapter 9

Myocardial Repair and Restoration Sharon S.Y. Wong and Harold S. Bernstein

Abstract  Over 19 million people in the USA and Europe alone suffer with heart failure, causing 230,000 deaths each year incurring tremendous costs. Heart transplantation remains the definitive treatment for end-stage heart failure, but this therapy is invasive, costly, and excludes some patients who are not candidates for transplantation and others for whom an organ is not available. New therapies are needed to treat the millions of patients with debilitating heart failure worldwide. Myocardial engineering represents a realistic strategy for reversing the deleterious effects of what has until now been considered terminal damage to the heart. This chapter reviews potential sources of cardiac-specific stem cells, efforts to enhance their engraftment and survival in damaged tissues, their incorporation into tissue patches, and recent progress made in developing methods to assess functional improvement in engineered myocardium.

Abbreviations ABCG2 aMHC ANF BMP4

ATP-binding cassette subfamily G member 2 a-Myosin heavy chain Atrial natriuretic factor Bone morphogenetic protein 4

H.S. Bernstein (*) Cardiovascular Research Institute, San Francisco, CA, USA Eli and Edythe Broad Center of Regeneration Medicine and Stem Cell Research, University of California San Francisco, San Francisco, CA, USA Department of Pediatrics, University of California San Francisco, San Francisco, CA, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_9, © Springer Science+Business Media, LLC 2011

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bFGF CM cTnI cTnT DKK1 ECM FRET GFP hEB hESC HTK/GCV iPSC KDR MDR1 MEA MEF2 MHC MI miRNA MLC2a MLC2v MSC PGI2 RFP Sca-1 SP TMRM VEGF

Basic fibroblast growth factor Cardiomyocyte Cardiac Troponin I Cardiac Troponin T Dickkopf homolog 1 Extracellular matrix Fluorescence resonance energy transfer Green fluorescent protein Human embryoid body Human embryonic stem cell Herpes thymidine kinase/ganciclovir Induced pluripotent stem cell Kinase insert domain receptor (VEGF receptor 2) Multidrug resistance-like protein 1 Multielectrode array Myocyte enhancer factor-2 Major histocompatibility complex Myocardial infarction microRNA Atrial myosin light chain 2 Ventricular myosin light chain 2 Mesenchymal stem/stromal cell Prostaglandin I2 Red fluorescent protein Stem cell antigen-1 Side population Tetramethylrhodamine methyl ester perchlorate Vascular endothelial growth factor

9.1 Introduction Over 19 million people in the USA and Europe alone suffer with heart failure, resulting in approximately 230,000 deaths at a cost of over $140 billion/year [1, 2]. These patients also are more prone to sudden cardiac death, causing approximately 1.9 million deaths annually. In addition, 2.5 million children are born each year worldwide with congenital heart disease [3], the most common human birth defect, and many of these children eventually develop heart failure. Heart failure occurs when the myocardium is damaged and becomes unable to meet the metabolic demands placed on it. Unlike some organs, the heart has a severely limited, if any, capacity for repair after injury. Heart transplantation remains the ultimate approach to treating end-stage heart failure, but this therapy is invasive, costly, and excludes some patients who are not

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candidates for transplantation given their co-morbidities. Most importantly, there are not enough organs for transplanting the increasing number of patients with endstage disease, and the infrastructure necessary to transplant organs is not available in all countries. New, accessible therapies are needed to treat the millions of patients with debilitating heart failure worldwide. Myocardial engineering, including stem cell transplantation, may represent the first realistic strategy for reversing the deleterious effects of what has until now been considered terminal damage to the heart. This chapter reviews potential sources of cardiac-specific stem cells, efforts to enhance their engraftment and survival in damaged tissues, and their incorporation into tissue patches (summarized in Table  9.1), as well as the progress made in assessing functional improvement in transplanted animals and human patients (Table 9.2).

9.2 Cell Sources 9.2.1 Mesenchymal Stem Cells MSCs reside in the stroma of the bone marrow, and can differentiate into osteoblasts, chondrocytes, and adipocytes [4, 5]. In addition, MSCs can differentiate in vitro into spontaneously beating CMs after exposure to the demethylating agent, 5-azacytidine [6, 7]. Because of their cardiomyogenic differentiation potential, MSCs have been transplanted in animal models of MI. Collectively, these studies have demonstrated improved left ventricular function, reduced infarct scar size, attenuated left ventricular remodeling, greater vascular density, and increased survival following transplantation [7–19]. Besides being able to transdifferentiate into CMs, MSCs are relatively easy to grow due to their ability to adhere to tissue culture plastic, and can be expanded in culture to large enough numbers required for transplantation, making them an attractive candidate for cell-based therapies [20–22]. Another advantage of MSCs for the repair of damaged myocardium is their ability to suppress immune rejection and curb the inflammatory response. Despite expressing MHC class I and low levels of class II antigens as well as Fas ligand, MSCs fail to elicit an alloreactive lymphocyte response when added to mixed lymphocyte cultures [23–29]. MSC transplantation has been tested in the clinic. In a randomized, placebocontrolled study by Chen et al., patients with acute MI received either intracoronary injections of autologous MSCs or saline [30]. Evaluation of treated patients compared to a placebo control group and to that of pre-transplantation after injury, the MSC-treated group at 3  months postinjection showed increased wall movement velocity at the infarct site, left ventricular ejection fraction, and end-systolic pressure to end-systolic volume ratio. Cardiac mechanical and electrical properties were also significantly improved in MSC patients. Osiris Therapeutics, Inc. completed a

Expandable in culture; functional improvement Autologous; expandable in culture

Tissue integration; functional improvement Functional improvement Autologous; expandable in culture; functional improvement

Advantages Immunosuppressive/ anti-inflammatory; expandable in culture; functional improvement

Tumor formation; heterogeneous CM subtypes Tumor formation; heterogeneous CM subtypes; genetic/ epigenetic abnormalities Heterogeneous CM subtypes Toxicity

– –

–

Concerns Tumor formation; ossification; arrhythmias

– –

–

– [32, 86]

–

Human [30–32]

[88, 98, 123, 133–138, 154]

[72] [74, 76, 78–80, 83–85]

[59–64]

Functional trials Animal [7–19, 33–36, 194]

Autologous; expandable in culture – – Target multiple pathways; nutrient/ [98] – bioreagent delivery Hydrogels Structural support; nutrient/bioreagent Toxicity [78, 155–172] – delivery Traumatic delivery; electrome [174–176, 187, 190, 191] – Tissue patches/bioreactors Structural/mechanical support; chanical interference nutrient/bioreagent delivery; functional ECM; improved spatial architecture; fill large defects ECM extracellular matrix, hESCs human embryonic stem cells, iCMs induced cardiomyocytes, iPSCs induced pluripotent stem cells, MSCs mesenchymal stem/stromal cells, SP side population

iCMs Prosurvival factors

iPSCs

Pluripotent stem cells hESCs

SP cells Cardiospheres

Resident cardiac stem cells c-kit+Sca-1+ cells

Type MSCs

Table 9.1  Cell sources and tissue adjuvants

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Table 9.2  Cardiac functional assessment Method Parameters measured Indices derived Animal studies Hemodynamic Echocardiography: 2D HR, ESV, EDV, ESD, EDD, SWT, DWT EF, FS, mass [123, 193]   and M-mode MRI/CT HR, ESV, EDV, ESD, EDD, SWT, DWT, EF, FS, mass [84, 194, 195] tissue morphology Echocardiography: Blood flow velocity, strain rate Contractility [196]   Doppler tissue imaging Contractility/relaxation (SV, EF, cardiac [80, 197] Invasive P–V loop HR, ESV, EDV, Pmax, Tau, IRC, dP/dtmax, dP/dtmin work, stroke work, cardiac output) analysis Sonomicrometry SWT, DWT, wall motion, segment EF, segmental abnormalities [16] shortening, P–V relationships Electrophysiological – Patch clamp Cellular conduction (rate of upstroke, APD50, APD90 peak, APD, AP/min) FPdur (»AP duration), FPrise (»AP rise) – MEA mapping Cellular conduction (FPmin, FPmax, FP/min) Optical mapping Tissue conduction ex vivo VT induction, conduction velocity, APD50, APD90 [201, 202, 204] Surface electro‑ P duration/axis, PR interval, QRS duration/ Atrial depolarization, AV conduction, ventricular [203]   cardiography axis, QT interval, T duration/axis depolarization/repolarization Tissue conduction in vivo, arrhythmias [203] Invasive Chamber reconstruction, tagging of important electroanatomical mapping anatomic landmarks and lesions, activation mapping, voltage/scar mapping 2D two-dimensional, AP action potential, APD action potential duration, APD50 time to attain 50% repolarization, APD90 time to attain 90% repolarization, AV atrioventricular, CT computed tomography, DTI Doppler tissue imaging, dP/dtmax maximum rate of systolic LV pressure development, dP/dtmin decline during isovolumic relaxation, DWT diastolic wall thickness, EDD end-diastolic diameter, EDV end-diastolic volume, ESD end-systolic diameter, ESV end-systolic volume, FP extracellular field potential, FPdur the time interval between FPmin and FPmax, FPmax last positive field potential, FPmin size of the largest negative field potential, FPrise decay of extracellular field potential, HR heart rate, IRC isovolumic relaxation constant, MEA multielectrode array, MRI magnetic resonance imaging, Pmax maximum pressure generated, P–V pressure–volume, SV stroke volume, SWT systolic wall thickness, VT ventricular tachycardia

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phase 1 clinical trial to assess the safety and preliminary efficacy of intravenous allogeneic adult human MSC delivery into patients within 10 days after acute MI [31]. While MSC infusion proved safe in patients, with no ectopic tissue formation, the results of the trial also yielded improved outcomes for MSC-treated patients with respect to left ventricular function, cardiac arrhythmias, pulmonary function, and symptomatic global assessment. Osiris Therapeutics, Inc. is currently conducting a larger phase 2 clinical trial for evaluating MSC therapy in patients following acute MI. Similarly, Angioblast Systems, Inc. is advancing their “off the shelf” allogeneic MSC therapeutics in phase 2 clinical trials for acute MI and congestive heart failure. Despite ongoing clinical trials with MSCs [32], there are still some key considerations for future MSC application in cardiac repair and regeneration. The longterm safety and effects of cardiac MSC transplantation remain unresolved. Some studies have reported unwanted calcification and ossification inside murine and rodent ventricular tissue after transplantation [33, 34]. Increased nerve density and shortened epicardial refractory periods in MSC-treated hearts with MI as well as re-entrant arrhythmias have been observed, raising the possibility of an arrhythmic risk associated with MSC-based therapies [35–37]. Dai et al. have shown that the improvements in left ventricular stroke volume and ejection fraction after MSC injection were lost at 6 months, suggesting a lack of durability [10].

9.2.2 Resident Cardiac Stem Cells Since the 1960s, studies of CM proliferation in rodents had indicated that the adult mammalian heart was a terminally differentiated, postmitotic organ without the capacity for cellular regeneration [38–49]. Over the past 10 years, however, several findings have challenged this view. Observation of cell division, telomerase activity, telomere shortening, and CM apoptosis have provided evidence of CM turnover in adult human hearts [50–53]. In addition, pockets of mitotically active cells in hypertrophic myocardium and hearts of patients with end-stage heart failure have been described [54]. Using a genetic fate mapping approach, Hsieh et  al. showed that although adult CMs are not replaced in the uninjured heart during normal aging up to 1 year in the mouse, they are refreshed after myocardial infarction or pressure overload [55]. Recently, Porrello et al. showed that the hearts of 1-day-old neonatal mice can regenerate after partial surgical resection, but this capacity is lost by 7 days of age [56]. The most compelling evidence for CM renewal in maintaining cardiac homeostasis comes from a study in which the amount of 14C generated from aboveground nuclear testing between 1955 and 1963, before the implementation of the Limited Nuclear Test Ban Treaty, and integrated into the DNA of human myocardial cells, was measured to date the birth of the cells [57]. Carbon dating of myocardial cells indicated postnatal cell turnover that declines with age: about 1% per year at age 25 to 0.45% at 75. This provided strong evidence not only for low level turnover of CMs in the adult heart, but that adult myocardial tissue is also indeed capable of incorporating new muscle cells to preserve tissue mass and function.

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9.2.2.1 c-kit and Sca-1 as Markers of Cardiac Progenitors The presence of the Y chromosome in migrated undifferentiated cells and in differentiated myocytes found in female donor hearts transplanted into male patients not only suggests the presence of cardiac chimerism in humans, but also supports the existence of a pool of cardiac stem or progenitor cells that can give rise to new CMs [58]. This putative stem cell population was positive for surface antigens c-Kit/ CD117 and/or MDR1, or Sca-1. Cardiac stem cells have been identified and isolated by others based on expression of these markers [53, 59–61]. In the presence of 5-azacytidine or oxytocin, Sca-1+ cells exhibit spontaneous beating, express cardiacspecific genes, and when injected intravenously into mice following MI, they home to the infarct site, differentiate, and fuse with host myocardium [60, 61]. Likewise, self-renewing, clonogenic, multipotent c-Kit+ cells reconstitute the injured myocardium by forming new blood vessels and myocytes, comprising as much as 70% of the ventricle [59, 62, 63]. The structural regeneration of infarcted myocardium, however, is independent of cell fusion and accompanied by decreased left ventricular hypertrophy and improved cardiac performance. Remarkably, Bearzi et al. obtained similar results after isolating human c-Kit+ cells from myocardial tissue of patients who underwent cardiac surgery, expanding them in culture, and injecting them into the infarcted myocardium of immunodeficient mice and immunosuppressed rats [64]. 9.2.2.2 Side Population Cells SP cells, known for their ability to efflux vital dyes such as Hoechst 33342, were initially discovered as hematopoietic stem cells [65]. ATP-dependent transporters, including MDR1 and ABCG2, are believed to mediate dye exclusion and both ABCG2 and MDR1 have been cited as molecular determinants of cardiac SP cells in the adult myocardium [66–70]. Further characterization shows that cardiac SP cells widely express Sca-1 and CD31, but are largely negative for hematopoietic markers CD45 and CD34, cellular adhesion marker CD44, and c-Kit [71]. Moreover, functional differentiation and maturation of cardiac SP cells are also restricted to the Sca-1+/CD31− subpopulation. Direct injection of heart-derived Sca-1+/CD31− cells into the peri-infarct region immediately following myocardial infarction in mice limits left ventricular remodeling, attenuates contractile dysfunction, and improves myocardial energy metabolism [72]. The transplanted Sca-1+/CD31− cells promoted neo-angiogenesis and underwent in  vivo differentiation into CMs and endothelial cells. 9.2.2.3 Cardiospheres Cardiospheres present another potential source of endogenous cardiac stem cells. These small, round, highly refractive cells are derived from cultured explants of mouse hearts and human atrial or ventricular biopsy samples following gentle enzymatic digestion [73]. They are able to migrate over the adherent portion of the

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explanted tissues. When collected and cultured, they form clonal multicellular clusters termed cardiospheres because of their ability to self-renew as well as differentiate into CMs and vascular cells. Immunophenotypic characterization shows that newly formed cardiospheres express stem cell markers CD34, c-Kit, Sca-1, and endothelial markers CD31 and KDR/VEGFR2/Flk-1. Recently, Ye et al. have shown that the Sca-1+ subpopulation of cardiosphere cells preferentially include Isl-1expressing precursors that give rise to second heart field structures [74]. Multipotent cardiospheres that can give rise to neurons, Schwann cells, smooth muscle cells, and CMs have also been isolated from a subset of cardiac SP cells [75]. Since their isolation, alternative approaches and more efficient methods to generate cardiospheres have been developed [76, 77]. Of particular note, the yield of cardiosphere cells from injured hearts is greater than from uninjured hearts, and cardiospheres are easily isolated and expanded from “middle aged” hearts [74], supporting their feasibility in autologous cell transplantation. Clinical studies in animal models of MI have shown that cardiosphere injection into infarcted mice and pig hearts preserves ventricular function, improves hemodynamic indices, produces less adverse remodeling, and reduces infarct size [74, 76, 78–80]. Despite displaying some degree of overlap in marker expression, the origins and exact lineage relationships among these various adult cardiac stem cell pools remain unknown. In fact, the validity of these cell populations as resident cardiac stem or progenitor cells has been questioned. Pouly et al. reported that c-Kit+ cells harvested from either human right ventricular endomyocardial or right atrial appendage tissues lack other stem cell markers such as MDR1, and co-express CD45 suggesting they are of hematopoietic origin [81]. In addition, c-Kit+ cells stained positive for the mast cell lineage marker, tryptase, implying that these cells may not be CM precursors, but rather mast cells. Studies in which Sca-1 was used to identify adult stem cells from human hearts were based on immunoreactivity to nonhuman Sca-1 antibodies. To date, a human homologue of Sca-1 has not been found [82], although the preponderance of studies that have identified cells based on the expression of “human Sca-1” make an immunoreactive analogue likely. There also have been some discrepant findings with regard to cardiospheres. Studies by several groups have demonstrated that some explant migrating cells do not differentiate into functional CMs or yield any significant physiological benefit in the infarcted mouse heart [83–85]. Notwithstanding, the clinical utility of cardiospheres is currently being tested in phase 1 clinical trials in patients with MI [32, 86].

9.2.3 Pluripotent Stem Cells 9.2.3.1 Human Embryonic Stem Cells hESCs grow and divide indefinitely while maintaining the potential to develop into derivatives of all three embryonic germ layers. Under appropriate culture conditions, hESCs spontaneously differentiate into CMs with structural and functional properties characteristic of endogenous CMs [87]. To induce spontaneous differentiation,

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Fig. 9.1  Approaches to preparing hESC- and iPSC-derived CMs for tissue repair. Methods for directing differentiation of pluripotent stem cells to CMs have focused on chemical (e.g., 5-azacytidine, p38MAPK inhibitors) and biological (e.g., activin A, BMP, bFGF, VEGF, DKK1) factors, genetic (e.g., miRNAs) and epigenetic (e.g., miRNAs, chromatin remodeling) manipulation, and mechanical factors (e.g., hydrodynamics, surface tension). In transplantation experiments, these approaches have been complemented by purification methods that take advantage of the biochemical properties of CMs (e.g., Percoll density centrifugation, mitochondrial content), and selection strategies that rely on the expression of cardiac-specific genes (e.g., reporter lines, molecular beacons) and surface markers. Adapted from [91] with permission

hESCs are cultured in suspension with serum for a period of 7–10  days to form three-dimensional cell aggregates called hEBs. hEBs are then allowed to adhere to gelatin-coated plates, where further cultivation results in the appearance of spontaneously contracting areas. This has been adopted by most laboratories as the standard approach for spontaneous CM differentiation from hESCs. Nonetheless, there are limitations to this protocol, most notably the small number of CMs produced. With this method, beating areas are visible in only 5–15% of hEBs, as reported by many groups [87–90]. Consequently, much work has focused on directing the differentiation of hESCs into the cardiac lineage [91]. Innovative enrichment, purification, and selection strategies have been developed to guide cardiac differentiation to relatively pure homogeneity (Fig. 9.1). These efforts, discussed in more detail below, have presented hESCs as an attractive candidate for cell-based cardiac repair.

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Directed Differentiation of hESC-Derived CMs Defined culture media have been developed to direct hCM differentiation from hESCs. 5-Azacytidine treatment at days 6–8 of hESC differentiation significantly increases cardiac aMHC expression and enhances CM differentiation, suggesting that DNA demethylation is a key factor in directing tissue-specific differentiation [88]. Similarly, exposure to SB203580, a small molecule inhibitor of p38MAPK, has been shown to significantly improve CM differentiation of hESCs grown in medium conditioned by mouse END2 cells, supporting a role for p38MAPK signaling in regulating human CM differentiation [92]. SB203580-treated hEBs display an increase in expression of both early mesoderm markers (Brachyury T, Tbx6, Mesp1) and cardiac aMHC, as well as increased CM numbers. Gaur et al. subsequently showed that p38MAPK inhibition occurs in a dose and stage-dependent manner, that it also causes the accelerated differentiation of hESC-derived CMs using the standard hEB formation method, and that it appears to act at the ectoderm/mesoendoderm branchpoint during hESC differentiation [93]. In the original study with SB203580, cells were subjected to an adapted differentiation system in which hESCs were differentiated in a suspension culture using serum-free medium conditioned by the mouse END2 cell line [94, 95]. END2 conditioned medium alone exhibits CM-inducing activity during hESC differentiation [92], and biochemical as well as microarray analysis of END2 conditioned medium and END2 cells, respectively, identified PGI2, a product of prostaglandin synthase enzymes, as an inducing factor in hESC cardiac differentiation [96]. Two key enzymes involved in PGI2 synthesis are upregulated in END2 cells compared to control MES1 cells [97], which lack cardiogenic activity. PGI2 levels are between six- and tenfold higher in END2 conditioned medium compared to control conditioned medium from MES1 cells. Moreover, insulin, a common supplement in media formulations, was discovered to be an inhibitor of hESC cardiac differentiation. END2 conditioned medium supplemented with increasing concentrations of insulin results in a dramatic decrease in hESC CM differentiation. Thus, addition of PGI2 in combination with insulin-free, unconditioned medium yields effective cardiac induction that is similar to that produced by END2 conditioned medium. Cardiac differentiation is further augmented in the presence of SB203580. Taken together, these three components provide a basic, synthetic recipe for directing CM differentiation of hESCs. Another system utilizing sequential exposure of undifferentiated hESCs cultured on Matrigel to activin A followed by BMP4 within the first 5 days of differentiation has proved to be 50-fold more efficient in generating CMs than the conventional serum-induction of hEBs [98]. Both factors were selected based on previous work showing that mesoderm formation and cardiogenesis are mediated by activin A and BMP4 [99–104]. Likewise, Yao et al. have reported that hESCs seeded on Matrigel and treated with both activin A and BMP4 express specific CM markers (aMHC, cTnI, MEF2, GATA4, Nkx2-5, ANF) [105]. This list of media supplements has grown to include bFGF, VEGF, and the Wnt inhibitor, DKK1. By mimicking the signaling environment of the early mouse

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embryo, another group has established a three-staged protocol that supports cardiac development at high frequency in differentiating hESC cultures [106]. This protocol exposes hEBs to a combination of activin A, BMP4, and bFGF during the first 4 days of differentiation to induce primitive-streak formation, representing the onset of gastrulation. Between days 4 and 8, the differentiating hEBs are incubated in medium containing VEGF and DKK1 to induce cardiac mesoderm development and maturation. Previous studies demonstrated that Wnt inhibition is required for cardiogenesis from mesodermal cells [107, 108]. From day 8 to day 14, bFGF is added to VEGF and DKK1 to promote CM expansion. Gene expression analysis of hEBs at this stage displays expression of cTnT, MLC2a, and cardiac transcription factors Tbx5 and Tbx20. Mechanical Force and CM Differentiation Since cardiac muscle is one of the few tissues that develops under the effects of dynamic force, it is not surprising that conditions generated by the force of fluids in motion can enhance CM differentiation. Supplying a constant rotary orbital motion for 7 days to suspension cultures of differentiating mouse EBs results in a significantly increased number of beating mEBs compared to mEBs cultured in static suspension [109]. Analysis of gene expression shows higher levels of mesodermal and cardiac proteins (Brachyury, GATA4, Nkx2-5, MEF2c, aMHC, MLC2v) in rotary mEBs than in static mEBs. In addition, a greater proportion of rotary mEBs stain positive for a-sarcomeric actin compared to static EBs. The enhanced CM differentiation is independent of rotary speed ranging from 25 to 55 rpm as determined by the expression of cardiomyogenic genes [110]. Domian et al. have examined the effects of surface tension on cardiomyogenic differentiation of murine cardiac progenitors [111]. Embryonic- and mESC-derived progenitors are cultured on either fibronectin-coated slides or micropatterns of fibronectin alternating with a surfactant that blocks cell adhesion. When grown on these micropatterned surfaces, a population of cells form longitudinally aligned myocardial fibers. In addition, culturing this population on micropatterned surfaces results in a statistically significant increase in the proportion of CMs, supporting a role for microenvironmental forces in CM differentiation. Genetic and Epigenetic Regulation of CM Differentiation miRNAs are small, noncoding RNAs thought to regulate the expression of 30% of protein-coding genes [112]. Their biological importance in stem cell biology is underscored by recent studies demonstrating that mESCs lacking the miRNA processing enzyme Dicer display differentiation and proliferation defects [113–116]. MiR-1 and miR-133 specifically are expressed in the mouse heart [117, 118]. Targeted deletion or knockdown of these miRNAs results in dysregulation of cardiac morphogenesis, electrical conduction, cell-cycle, and cardiac hypertrophy

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[117–120]. Recently, Ivey et  al. showed that miR-1 and miR-133 regulate the differentiation of mESCs and hESCs into the cardiac lineage [121]. Both miRNAs are enriched in mESC-derived CMs. Lentiviral introduction of either miR-1 or miR-133 into mESCs enhances early mesoderm differentiation as evidenced by increased expression of Brachyury. miR-1 and miR-133 also reinforce mesoderm lineage decisions by repressing endoderm and neuroectoderm differentiation. When stimulated to differentiate into either endoderm or neuroectoderm lineages, mEBs expressing either miR-1 or miR-133 express lower levels of endodermal and neural markers compared to control mEBs. However, further differentiation revealed opposing roles of miR-1 and miR-133. miR-1 promotes differentiation of mesoderm into the cardiac and skeletal muscle lineages as determined by enhanced Nkx2-5 and myogenin expression, respectively, whereas miR-133 blocks induction of both markers. Importantly, the differentiation of hESCs in the presence of miR-1 behaves comparably to that of mESC differentiation. Overexpression of miR-1 in hESCs increases Nkx2-5 expression and yields more than a threefold higher number of beating hEBs compared to wild-type controls. While miRNAs direct cell lineage determination by controlling protein dosage, epigenetic regulation through chromatin remodeling has been shown to control cell fate as well. Takeuchi et  al. have identified a minimal set of factors necessary to execute the cardiac transcriptional program [122]. Baf60c, a cardiac-enriched subunit of the Swi/Snf-like BAF chromatin remodeling complex, in combination with cardiac transcription factors GATA4 and Tbx5, is able to induce cardiac differentiation in mouse embryos when ectopically expressed. With this combination, 90% of the transfected embryos display expression of the early cardiac marker, Actc1, and 50% of the transfected embryos exhibit beating tissue. GATA4 together with Baf60c is essential in initiating the cardiac gene program as assessed by expression of Actc1. None of the other transcription factors tested alone (Tbx5, Nkx2-5) or in concert with Baf60c are able to induce Actc1 expression. GATA4/Baf60c, however, is not sufficient for generating beating embryos: Tbx5 is required to achieve contracting CMs. Purification of hESC-Derived CMs While the efficiency of cardiac differentiation has markedly improved, it is vital that the final CM population is of high purity since differentiated hESC cultures still may consist of heterogeneous populations that include other cell types derived from hESCs. Manual dissection of the beating areas has been one way to achieve high purity [87]. A less labor intensive method utilizing Percoll gradient centrifugation has been described to purify hESC-derived CMs [88, 123]. Differentiating hESCs are applied to a discontinuous Percoll gradient consisting of 40.5% Percoll layered over 58.5% Percoll. Following centrifugation, the majority of CMs resides within the 58.5% Percoll layer and express cTnI, sarcomeric MHC, aMHC, bMHC, and N-cadherin. hCMs of £70% purity are obtained using this approach. It has been shown that Percoll-purified CMs can be further enriched by culturing the purified CM clusters in suspension for an additional week or longer [124]. These clusters,

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re-cultured for at least 8 days following purification, exhibit significant increases in cardiac aMHC and bMHC expression. Analysis by flow cytometry demonstrate that the cells comprising these clusters also express sarcomeric MHC, and that the percentage of sarcomeric MHC+ cells increase with time in culture. A third purification strategy is based on the observation that CMs have high mitochondrial content compared to that of nonmyocytes [125]. Using the fluorescent dye, TMRM, that freely diffuses into the mitochondrial matrix, Hattori et al. found that TMRM fluorescence in embryonic rat hearts increases with developmental stage suggesting that mitochondrial biogenesis is linked to myocardiogenesis. In whole rat embryos, TMRM fluorescence in the heart is more robust than in other tissues, and when analyzed by flow cytometry, flow-sorted populations with the highest TMRM fluorescence are observed to express cardiac a-actinin. TMRMlabeled CMs derived from mESCs are positive for both Nkx2-5 and a-actinin. The CM content in cultured cells sorted from day 12 to day 25 mEBs is greater than 99% as determined by Nkx2-5 and a-actinin expression. Most notably, greater than 99% CM purity is also obtained in cultured cells sorted from differentiating hEBs. Selection of hESC-Derived CMs and Cardiac Progenitors Besides achieving a high degree of CM purity, these separation methods have the added advantage of not requiring genetic manipulation of hESCs. A disadvantage, however, is that none of these techniques allow purification of cardiac progenitor cells. Mechanical dissection can only be performed toward later stages of differentiation when sufficient amounts of beating areas are visible. Percoll separation is less effective at earlier times of hESC differentiation [124]. hEBs used for TMRM purification experiments were between 50 and 90  days of differentiation [125]. Moreover, sorted TMRM-fluorescent cells from early mEBs fail to differentiate into CMs during subsequent culture. The use of genetic selection strategies has addressed both the issue of CM homogeneity and isolation of cardiac progenitors. Many laboratories have developed transgenic/reporter hESC lines to derive pure CM populations. This approach relies on a cardiac-restricted promoter to drive the expression of a reporter gene or selectable marker. Huber et  al. used lentiviral vectors to produce stable hESC lines in which enhanced GFP is expressed under control of the cardiac-specific human MLC2v promoter [126]. Xu et al. generated stable hESC lines using a reporter plasmid consisting of the cardiac-specific mouse a-myosin heavy chain promoter driving expression of the neomycin resistance gene [127]. Kita-Matsuo et al. designed a set of lentiviral vectors to generate multiple stable hESC lines with eGFP and mCherry reporters or with puromycin resistance downstream of the mouse aMHC promoter [128]. Ritner et al. generated a cardiac-specific hESC reporter line using a lentiviral construct consisting of a fragment of the mouse aMHC promoter upstream of eGFP. The specific promoter fragment used has allowed for the identification and analysis of early cardiac progenitors expressing Nkx2-5, but before the onset of cTnT or chamber-specific MLC expression [129].

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Collectively, fluorescence-activated cell sorting or antibiotic selection of these lines has yielded 85–99% pure CMs or cardiac progenitors that express cardiacspecific genes and exhibit action potentials characteristic of human embryonic CMs. eGFP-expressing cells derived from the MLC2v transgenic line form stable intracardiac cell grafts following transplantation [126]. Injection of neomycin resistance-selected hEBs into the hind limb muscles of SCID mice results in no teratoma formation after 23 weeks [127]. Contractile forces in puromycin resistance-selected CMs are similar to those generated by rat neonatal ventricular CMs [128]. While isolation of hESC-derived CMs from these transgenic/reporter lines is based on positive selection, Anderson et  al. implemented a negative selection strategy to deplete undifferentiated, proliferating hESCs from cultures of hESC-derived CMs [130]. Their transgenic hESC line utilized a HTK/GCV suicide gene system under the control of a constitutive phosphoglycerate kinase promoter. Following administration of the antiviral drug GCV, cells expressing HTK phosphorylated GCV, which then incorporated into nascent DNA chains of proliferating cells, causing chain termination and cell death. The increased number of a-actinin positive cells after GCV treatment lead to an almost sevenfold enrichment of CMs. An important caveat of this approach, however, is that other nonproliferating cell types would remain in the culture while proliferating hCMs would be depleted. The culture would still need to undergo a cardiac purification step and, as discussed below, the excluded proliferating hCMs and cardiac progenitors may be of greater benefit for transplantation than fully differentiated, nonproliferating CMs. As an alternative to genetically modified hESC lines for tracking and isolating human CMs and cardiac progenitors, King et  al. adapted dual-FRET “molecular beacon” technology for transient, real-time detection of gene expression during hESC differentiation [131]. Molecular beacons are single-stranded oligonucleotide probes that have been employed to assay gene expression in vitro, as in real-time PCR, and in vivo using microscopy [132]. These consist of short sequences capable of forming stem-loop structures bearing a fluorescent reporter group at one end and a fluorescent quencher at the opposite end [132]. In the absence of a target sequence, the oligonucleotide self-anneals, forming a stem that brings the reporter and quencher in close proximity, thereby quenching fluorescence. In the presence of a target sequence, the oligonucleotide anneals to the target, separating the reporter and quencher, thereby allowing fluorescence. These investigators showed that appropriately designed, dual-FRET molecular beacon pairs can identify the expression of specific mRNAs by microscopy and flow cytometry, and facilitate the collection of specific hESC populations by fluorescence-activated cell sorting, while leaving the hESC genome intact [131]. Application of hESC-Derived CMs in Animal Studies Transplantation with hESC-derived CMs in animal models of myocardial injury has yielded promising, albeit modest, results. Although the methods for generating hCMs and monitoring engraftment, as well as the number of transplanted hCMs, varied

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between studies, hESC-derived CM transplantation lead to measureable benefit [98, 123, 133–138]. Overall, hCM transplantation preserves left ventricular function, attenuates post-MI wall thinning, and even partially remuscularizes the infarct region. Despite these encouraging results, the cardiac-specific benefits determined from in vivo engraftment studies appear to be transient. Van Laake et al. have reported that while cardiac function was improved 4 weeks after myocardial infarction, the functional benefit was no longer significant at 12 weeks, even after tripling the number of transplanted hCMs in a repeat study [133–135]. These results call into question the utility of fully differentiated hCMs for cardiac repair. The integration of hESCderived CMs into existing muscle may be hampered by their limited developmental plasticity, whereas partially differentiated cardiac progenitors may retain the plasticity needed to enable extensive engraftment [129]. The transient improvement seen in rodent models also argues for the use of pre-clinical animal models with hemodynamics that more closely resemble human physiology. Porcine models provide an opportunity to study the effects of cell transplantation in an animal with more relevant cardiovascular physiology [139, 140]. A limitation to using hESC-derived CMs is that the transplanted cells will need to evade immune rejection. hESCs appear to have a lower immunostimulatory potential compared to adult cells. DNA microarray data of undifferentiated and differentiated hESCs indicate that almost half of the upregulated immunoregulatory genes in hematopoietic cells, lymphoid organs, and other tissues are not similarly expressed in hESCs, implying that hESCs are immunologically immature [141]. This observation suggests that immunosuppressive regimens for hESC-based therapeutics may not need to be as rigorous as conventional organ transplantation [141]. Nevertheless, transplanted hESC-derived CMs will be susceptible to immune rejection to some degree. 9.2.3.2 Induced Pluripotent Stem Cells The advent of iPSC technology [142, 143] offers a possible solution to immunorejection. This technology entails reprogramming terminally differentiated adult human fibroblasts to pluripotent stem cells through ectopic expression of four transcription factors: Oct4, Sox2, c-Myc, and Klf4. Reprogrammed cells exhibit many features characteristic of hESCs including morphology, feeder dependency, cell surface antigens, gene expression, promoter activities, telomerase activity, and proliferation rates [143]. They form teratomas and have the capacity to differentiate into cells of all three germ layers including CMs [143]. Recently, Zhang et al. have reported successful derivation of functional CMs from human iPSCs [144]. Detailed evaluation of these CMs showed robust expression of a full range of cardiac genes, sarcomeric organization, and action potentials characteristic of ventricular, atrial, and nodal cells. Since adult fibroblasts can be obtained directly from the patient, it is assumed that patient-specific iPSC-derived CMs can be transplanted without immunorejection. However, further work is needed to demonstrate that self-recognition is indeed retained in iPSCs. Another caveat is that the time needed to derive patient-specific

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iPSCs will not allow them to be used under urgent circumstances, such as in the ­setting of acute myocardial infarction or heart failure. While the prospect of iPSC banking may be useful in genetically homogeneous populations [145, 146], this approach may not be practical in more diverse communities [147]. It is also important to recognize that iPSCs have typically been generated by inducing the expression of reprogramming factors using retroviral, lentiviral, or adenoviral vectors, which carry the risk of permanent and harmful genomic integration. Indeed, some established iPSC lines are genetically unstable, exhibiting large-scale genomic rearrangements, copy number variations, and abnormal karyotype even in early passage stages [148, 149]. Alternative methods for reprogramming are significantly less efficient compared to viral integration [150, 151]. Furthermore, in some iPSC lines reprogrammed using nonintegrating viral method, high levels of mutational changes were still observed [148]. In addition to genomic changes, recent studies have also revealed that iPSCs contain epigenetic features that indicate either incomplete or aberrant reprogramming. In particular, iPSC DNA methylation patterns are frequently reminiscent of the somatic cell of origin [152], suggesting that iPSCs are not completely reprogrammed into the naïve pluripotent state seen in hESCs. It is unclear whether the observed genetic instability and epigenetic imprinting accrued during reprogramming or was present in the somatic cell of origin. Nevertheless, patient-specific iPSCs derived from somatic cells of older patients may be more likely to contain genomic mutations and disadvantageous epigenetic programs. Thus, the safety and efficacy of therapeutic iPSCs as currently derived remain to be tested. 9.2.3.3 Induced Cardiomyocytes A recently published study may provide another promising approach to bypass immune rejection. Ieda et  al. identified a minimal set of transcription factors to reprogram postnatal cardiac fibroblasts into functional CMs [153]. They showed that GATA4, MEF2c, and Tbx5 were sufficient for CM induction. “Induced CMs” expressed cardiac genes, displayed spontaneous contractile activity, exhibited calcium oscillations, and possessed action potentials resembling those of adult ventricular CMs. In vivo, cardiac fibroblasts transduced with GATA4, Mef2c, and Tbx5 transdifferentiated into CMs within 2 weeks after injection into immunosuppressed mouse hearts.

9.2.4 Survival and Engraftment A prevailing challenge to stem cell therapies is the poor rate of survival and engraftment in the damaged myocardium. In nude rats, injection of hESC-derived CMs into uninjured hearts resulted in a 90% engraftment success rate [154]. However, injection of human CMs into rat hearts infarcted by permanent coronary artery

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ligation yielded only an 18% engraftment success rate due to poor survival of the ­transplanted cells [98]. This discrepancy prompted the investigators to deliver the CMs in the presence of a prosurvival cocktail targeting multiple cell death pathways [98]. Not only did the use of a prosurvival formulation improve graft survival in infarcted hearts, but the engrafted myocardium demonstrated improved ventricular function 4 weeks after transplantation compared with controls. Tissue engineering is also being employed as a complementary approach to creating a prosurvival environment for cell engraftment. Hydrogels are three-dimensional networks of natural or synthetic hydrophilic polymers that can absorb water without dissolving. Because of their high water content, hydrogels create a moist environment, and are flexible and elastic, properties similar to natural tissue components in the body. Wall et al. engineered a bioactive hydrogel with tethered peptides that can interact with integrin receptors at the cell membrane to support cellular attachment and growth [155]. MSCs encapsulated in the hydrogel were directly injected into the infarct border zone of adult mice. The hydrogel augmented survival of transplanted cells and provided mechanical stability to the injured ventricular wall. Other engineered gels have been shown to promote cardiac differentiation, increase retention of transplanted cells, and subsequently enhance cardiac performance [78, 156–162]. These gel matrices consisted of polymers that include fibrin, collagen, Matrigel, alginate, chitosan, and self-assembling peptide nanofibers. Surprisingly, the biomaterials alone, devoid of cell encapsulation, can significantly attenuate cardiac dysfunction post-MI by modulating the remodeling response and by recruiting progenitor cells to the damaged region [156, 158, 159, 163–172]. Acellular biomimetic materials, therefore, offer potential treatments for MI and when delivered with cells, produce synergistically improved cardiac outcomes [157].

9.3 Tissue Patches There are several distinct advantages of employing polymer gels, including enhanced cell survival and engraftment, efficient nutrient transport, relatively atraumatic delivery, and biocompatibility [173]. The trade-off is that these gel-based carrier systems have low mechanical strength, are difficult to sterilize, and are limited in the size of tissue defect that can be repaired. In addition, they do not adequately mimic the native ECM of the myocardium, which provides numerous environmental cues essential in regulating cell behavior and function. Consequently, cardiac tissue engineering efforts have also focused on in vitro cultivation systems in which cells are grown on pre-formed scaffolds. These afford a matrix and culture parameters suitable for generating three-dimensional contractile and vascularized cardiac patches that resemble native cardiac muscle. Zimmermann et al. developed a technique to create engineered heart tissue whereby neonatal rat CMs were combined with liquid collagen I and Matrigel in circular molds subjected to mechanical strain [174]. Grafts implanted post-MI in rats demonstrated structural and electrical integration with the host myocardium as well as improved diastolic and systolic left ventricular function

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compared to controls. Caspi et al. constructed highly porous, multicellular scaffolds composed of poly-l-lactic acid and polylactic-glycolic acid in which hESC-derived CMs, human endothelial cells (either hESC-derived or from human umbilical vein), and embryonic fibroblasts were cultured [175]. The three cell-based culture strategy generated human cardiac tissue that was vascularized, and contracted spontaneously and synchronously with gap junctions mediating impulse conduction. Transplantation of the engineered cardiac tissue to healthy rat hearts resulted in the formation of stable tissue grafts and functional vascularization [176]. Bioreactors provide a way to tightly control environmental conditions for constructing functional cardiac tissue and are often designed to recapitulate the microenvironmental conditions within native cardiac muscle. Bioreactors that allow direct perfusion of oxygen-rich culture medium through the cultured cardiac biografts produce higher cell viability, enhance cardiac-specific marker expression, and improve the spatial architecture of the engineered tissue [177–180]. Human ventricular muscle is generally less than 1 cm thick [180]. Besides increasing the oxygen supply to cultured tissue greater than 100 mm thick [181, 182], bioreactors can supply other factors that play essential roles in engineering viable, thick, and well-differentiated cardiac constructs. Since cardiac tissue is subject to mechanical forces in vivo, pulsatile perfusion bioreactors that combine culture medium perfusion with mechanical conditioning have been developed. Neonatal CMs within cardiac tissue patches cultivated under pulsatile fluid flow exhibited defined Z-lines, highly ordered sarcomeres, intercalated discs, robust Connexin 43 expression, and better contractile properties than those cultivated in nonpulsatile medium flow [183, 184]. Application of electrical stimulation during cultivation of neonatal rat cardiac constructs enhanced ultrastructural differentiation and augmented contractile behavior compared to nonstimulated constructs [185, 186]. Using the rat omentum (the blood vessel-enriched fold of peritoneum that extends from the stomach to adjacent abdominal organs) as a natural bioreactor to prevascularize neonatal rat cardiac patches for a period of 7 days, Dvir et al. found that grafting omentum-generated patches onto infarcted rat hearts attenuated left ventricular remodeling and dysfunction, and structurally and electrically coupled with the host myocardium 4 weeks after implantation [187]. Decellularization of tissues to create native scaffolds is another approach taken to reconstitute whole tissues and organs (Fig. 9.2). Using detergent-based coronary perfusion to decellularize cadaveric rat hearts, Ott et  al. produced a whole heart scaffold with intact geometry, vasculature, and ECM [188]. They then repopulated the decellularized heart with cardiac and endothelial cells and mounted the recellularized heart in a bioreactor. By supplying oxygenated cell medium perfusion and electrical stimulation within the bioreactor, the organ displayed pumping capacity within 8  days of culture. This decellularization/recellularization technique holds tremendous promise for the generation of bioartificial myocardial tissue, as well as whole organs for transplantation. Beyond the use of conventional biodegradable scaffolds, Shimizu et  al. engineered cell sheets using thermo-responsive cell culture dishes to detach intact sheets, and then stacked them three-dimensionally to create thick cardiac constructs [189]. The temperature-sensitive culture surfaces were covalently grafted with

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Fig. 9.2  Generation of a bioartificial heart by perfusion decellularization/recelluarization of a whole heart. (a) Photographs of cadaveric rat hearts mounted on a Langendorff apparatus. Ao, aorta; LA, left atrium; LV, left ventricle; RA, right atrium; RV, right ventricle. Retrograde perfusion of cadaveric rat heart using sodium dodecyl sulfate over 12 h is shown. The heart becomes more translucent as cellular material is washed out from the right ventricle (left), then the atria (middle) and finally the left ventricle (right). (b) Hematoxylin/eosin-stained thin section of perfused heart showing no intact cells or nuclei. The protocol maintains large vasculature conduits (*). Bar, 200 mm. (c) Recellularized whole rat heart at day 4 of perfusion culture in a working heart bioreactor. Cross-sectional ring harvested for functional analysis (upper inset). Masson’s trichrome staining of a ring thin section showing cells throughout the thickness of the wall (lower inset). Bar, 100 mm. Adapted from [188] with permission

poly(N-isopropylacrylamide) upon which neonatal rat CMs are grown. On reducing the temperature below 32°C, confluent cells spontaneously lifted up as a single contiguous sheet with cell–cell junctions and ECM-deposited adhesive proteins preserved. In vitro, layered cardiac sheets pulsed simultaneously, and when implanted subcutaneously into nude rats displayed spontaneous macroscopic beating, synchronized electrical potentials, neovascularization, and cardiac structures including sarcomeres, desmosomes, and gap junctions. The subcutaneously grafted myocardial tissue has been shown to survive up to 1 year in vivo [190]. Transplantation of three-dimensional cardiac sheets onto the epicardial surface of injured rat hearts resulted in functional integration of electrical signals between the cardiac graft and host heart without any observed arrhythmias [191]. In addition, functionally synchronized cardiac tissue constructs of up to 1 mm thick can be achieved with multiple surgeries at 1- or 2-day intervals to layer additional CM sheets [192].

9.4 Functional Assessment in Animal Models 9.4.1 Hemodynamic Assessment While some tissues and organs require that transplanted cells perform specific metabolic tasks, it is essential that cells transplanted to the heart function mechanically and electrically in coordination with native tissue. Testing the efficacy of cell-based therapies in various experimental models of MI in mice, dogs, and pigs

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has created a need for accurate assessment of cardiovascular variables in the intact animal. Studies evaluating the feasibility of transplanting cells for myocardial repair have yielded beneficial effects on left ventricular remodeling and function. In addition to tissue measurements of infarct size, vascularity, cellular proliferation, and CM apoptosis postmortem [193], changes in left ventricular structure and function can be measured in vivo by a variety of techniques, of which serial echocardiography has been employed in the vast majority of studies (Fig.  9.3). Left ventricular end-systolic volume, left ventricular end-diastolic volume, left ventricular end-systolic diameter, left ventricular end-diastolic diameter, as well as anterior and posterior wall thickness during systole and diastole can be obtained with two-dimensional and M-mode echocardiography [123]. These measurements allow for calculation of left ventricular ejection volume, fractional shortening, and mass. Three-dimensional imaging using cardiac magnetic resonance or multidetector computerized tomography permits similar measurements but with a higher degree of precision and accuracy due to improved temporal and spatial resolution that also offers detailed examination of changes in tissue morphology [84, 194, 195]. Other indices for quantifying global and regional left ventricular function can be obtained with Doppler tissue imaging, which measures blood flow velocities and strain rates, both of which have been shown to correlate with the maximum rate of rise of left ventricular pressure, or dP/dtmax, an indicator of myocardial contractility [196]. Some studies have used more invasive procedures such as pressure–volume loop analysis to directly measure hemodynamic parameters [80, 197]. Left ventricular volume and pressure are recorded from a specially designed catheter with pressure sensor and conductance electrodes for simultaneous pressure–volume measurements. End-systolic volume, end-diastolic volume, heart rate, maximum pressure, contractile parameters (ejection fraction, stroke volume, dP/dtmax, stroke work), and relaxation parameters (Tau, isovolumic relaxation constant, dP/dtmin, maximum volume of power during cardiac cycle) can be obtained. Sonomicrometry has also been utilized by some researchers for determining wall  thickness, wall motion, segment shortening, end-systolic and end-diastolic

Fig. 9.3  Hemodynamic and histological assessment of cell therapy. (a) Infarct size in control (medium), bone marrow cell (BMC), and BMC extract groups is determined morphometrically in trichrome-stained sections. NS, not significant. (b) Effects of control versus treatment groups on left ventricular function are measured by echocardiography. Left ventricular ejection fraction (LVEF), end-systolic volume (ESV), and end-diastolic volume (EDV) are assessed over time. (c) Wall thickness of the peri-infarct region and the infarct scar are compared at the study end point. (d) Vascularity at the infarct border zone is assessed by quantitating CD31+ vessel density and number of CD31+/a-smooth muscle actin (SMA)+ arterioles. Bar, 300  mm (upper) or 100  mm (lower). (e) Effects of therapy on CM apoptosis are assessed by quantitation of caspase-3 (upper) and terminal deoxynucleotidyl transferase dUTP nick end labeling (TUNEL) (lower) staining cells. Both apoptotic CMs (white arrows) and non-CMs (yellow arrows) are seen. Bar, 50  mm. Adapted from [193] with permission

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p­ ressure volume relationships, stroke work, stroke volume, and ejection fraction [16]. This technique involves implanting small piezo-electric transducer crystals in the infarct region of the heart, which allows the distances between the transducers to be measured continuously through ultrasound waves transmitted and received by the crystals.

9.4.2 Electrophysiological Assessment Ventricular tachyarrhythmias and sudden cardiac death caused the suspension of clinical trials using autologous skeletal myoblasts transplanted into the myocardium of MI patients [198]. This underscored the importance of electrical integration of transplanted cells into the recipient myocardium. Thus, a fundamental operating property of transplanted cells is their electrical coupling with the host myocardium. In vitro and ex vivo electrophysiological assessment has included patch clamp, multielectrode array, and optical mapping analyses. For example, the patch clamp technique has been used to characterize action potentials from CMs derived from hESCs in vitro. Electrical recordings of dissociated hESC-derived CMs reveal fetal ventricular, atrial, and pacemaker action potentials, the majority of which are ventricular-like [95]. The cardiac action potentials are sensitive to the L-type calcium channel blocker, Verapamil, implicating the presence of L-type calcium channels. Moreover, the L-type calcium channels are responsive to adrenergic stimuli. More recently, as much as one-third of hESC-derived CMs have been shown to possess a mature electrical phenotype, as determined by patch clamp analysis [199]. MEA mapping can also be employed to acquire electrophysiological data. This technique consists of an array of embedded, substrate-integrated contact electrodes that allow simultaneous recording of extracellular field potentials from all electrodes over extended periods of time. The detected field potentials from the attached cells can be correlated to the shape and duration of the underlying action potential. Using MEAs to assess electrical conduction of hESC-derived CMs, Satin et  al. detected functional sodium channels that were sensitive to the sodium channel blocker, Tetrodotoxin [200]. Ritner et al. used MEAs to show that hESCs give rise to a heterogeneous combination of chamber-specific cell types [129]. Optical mapping uses voltage-sensitive dyes to record action potentials ex vivo. Voltage-sensitive dyes bind with high affinity to the cell membrane and when excited, emit light in direct proportion to the transmembrane voltage thereby producing an optical signal that mimics an action potential. Optical mapping not only maps membrane depolarization, but also the repolarization process. Verheule and co-workers utilized high-resolution optical mapping to study atrial conduction in canine models of chronic atrial dilatation and chronic rapid atrial pacing to determine whether the underlying mechanisms of atrial fibrillation inducibility are similar [201]. Based on the conduction profiles obtained, they found that the underlying substrate for atrial fibrillation in the chronic atrial dilatation model is distinctly ­different from that of the chronic rapid atrial pacing model. The same

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Fig. 9.4  Electrophysiology assessment of cell transplantation by high-resolution optical mapping. (a) Example of normal conduction at the site of hESC-derived CM engraftment. The resulting isochronal activation maps, with (middle) or without (left) the overlay of the image obtained by DiO labeling of transplanted cells, while pacing at 250 ms. At higher resolution, the corresponding conduction velocity vectorial map (right) shows slight conduction slowing at a transplantation site. (b) Example of conduction delay at the site of hESC-derived CM transplantation. Frame sequences from the optical mapping movie of the transplanted heart while pacing at 250 ms (left). Note the slight conduction delay at the site of hESC-derived CMs grafting. The resulting isochronal activation maps, with (right) or without (middle) the overlay of the DiO image. Note the slight conduction delay at the site of cell grafting. Adapted from [202] with permission

group used optical mapping to evaluate the electrophysiological effects of transplanted hESC-derived CMs on injured rat myocardium [202]. High-resolution analysis demonstrated a slight conduction delay at the site of hCM grafting (Fig. 9.4). The capacity of transplanted CMs to electrocouple with the host myocardium is often examined by postmortem immunohistochemical analysis, usually through Connexin 43 immunohistochemistry to demonstrate that the transplanted cells form gap junctions within the infarcted region. Besides identification of proteins involved in establishing electrocellular connections, additional evidence is clearly warranted to demonstrate electromechanical stabilization between donor cells and host myocardium. One study used pigs as a large animal model of atrioventricular heart block to show functional electrical integration. Injection of manually dissected, beating hEBs into the left ventricle of pig hearts with atrioventricular block resulted in successful pacing of the heart, manifested by the presence of a new ectopic ventricular rhythm measured by body surface electrocardiography and catheter-based electroanatomical mapping [203]. Another study injected microdissected, beating GFPlabeled hEBs subepicardially into the left ventricular anterior wall of guinea pigs. After atrioventricular nodal cryoablation to suppress the intrinsic heart rhythm, ex vivo optical mapping of control cryoablated hearts exhibited complete electrical silence, whereas hearts transplanted with hESC-derived CMs propagated spontaneous action potentials from the site of injection [204].

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9.5 Future Perspectives and Conclusions The last decade has improved our knowledge of stem cell biology and the development of the cardiovascular system. But, a more profound understanding of myocardiogenesis will be most certainly required for the development of stem cell therapeutics to repair or regenerate the damaged myocardium. hESCs offer an unprecedented opportunity to study human cardiac development. Two groups have independently identified multipotent cardiac progenitors derived from hESCs that will facilitate such further study. Flow cytometric analysis of hEBs detected three distinct KDR+/c-Kit+ populations at day 6 of the three-staged differentiation protocol described earlier: KDRhigh/c-Kit+, KDRlow/c-Kit−, KDR−/c-Kit+ [106]. Of the three, the KDRlow/c-Kit− population appears to contain cardiac progenitors that generate cells expressing markers of endothelial (CD31, CDH5, VE-cadherin, von Willebrand factor), vascular smooth muscle (calponin, smooth muscle actin, sarcomeric MHC, caldesmon), and cardiac (Nkx2-5, Isl-1, Tbx5, Tbx20, cTnT, MLC2a) differentiation. To establish clonality, hESC lines expressing GFP or RFP were employed in methylcellulose colony assays. Mixing of KDRlow/c-Kit− populations isolated from both lines resulted in colonies expressing either GFP or RFP, but not both. Expression analysis of colonies from the mixed GFP/RFP cultures confirmed the presence of cardiac, endothelial, and vascular smooth muscle lineages suggesting that the three cell types arise from a single cell. To track the fate of human Isl1+ cells and their progeny during hESC differentiation, Bu et al. used Isl1:cre hESCs transfected with a pCAG-flox-DsRed reporter plasmid to achieve irreversible DSRed expression in Isl1+ cells. In clonal assays of day 8 hEBs, about half of the DsRed+ (i.e., Isl1+) clones that were Nkx2-5+ expressed markers of all three major cardiac lineages: cTnT (CMs), PECAM1 (endothelial cells), and smooth muscle troponin (smooth muscle cells) [205]. Interestingly, KDR was not detected in DsRed+ cells from day 8 hEBs, but was present within 7 days after plating on mouse embryonic fibroblasts in the clonal assays, implying that Isl1+/Nkx2-5+/ KDR+ cells may represent a more restricted downstream cardiac progenitor. The path forward will likely include: (1) further investigations to delineate the human cardiac lineage tree, (2) methods to isolate specific cardiac progenitor pools or specialized CM subtypes, (3) strategies to ensure survival of transplanted cells, their functional integration with the host myocardium, and circumvention of immune rejection, (4) the development of accurate technologies to assess successful integration, (5) the determination of optimal parameters for efficacious engraftment such as delivery method, timing of transplantation post-MI, and cell preparations, and (6) large animal models of heart failure that closely resemble human cardiovascular physiology and disease for assessing cell engraftment, host immune response, and myocardial function. Acknowledgments The authors thank members of the Bernstein Laboratory for helpful discussion. H.S.B. is supported by grants from the National Institutes of Health, the California Institute for Regenerative Medicine, and the Muscular Dystrophy Association. S.S.Y.W. was supported by a fellowship from the National Institutes of Health.

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175. Caspi O, Lesman A, Basevitch Y, Gepstein A, Arbel G, Habib IH, Gepstein L, Levenberg S (2007) Tissue engineering of vascularized cardiac muscle from human embryonic stem cells. Circ Res 100(2):263–272 176. Lesman A, Habib M, Caspi O, Gepstein A, Arbel G, Levenberg S, Gepstein L (2010) Transplantation of a tissue-engineered human vascularized cardiac muscle. Tissue Eng Part A 16(1):115–125 177. Carrier RL, Rupnick M, Langer R, Schoen FJ, Freed LE, Vunjak-Novakovic G (2002) Perfusion improves tissue architecture of engineered cardiac muscle. Tissue Eng 8(2):175–188 178. Radisic M, Yang L, Boublik J, Cohen RJ, Langer R, Freed LE, Vunjak-Novakovic G (2004) Medium perfusion enables engineering of compact and contractile cardiac tissue. Am J Physiol Heart Circ Physiol 286(2):H507–H516 179. Dvir T, Benishti N, Shachar M, Cohen S (2006) A novel perfusion bioreactor providing a homogenous milieu for tissue regeneration. Tissue Eng 12(10):2843–2852 180. Radisic M, Marsano A, Maidhof R, Wang Y, Vunjak-Novakovic G (2008) Cardiac tissue engineering using perfusion bioreactor systems. Nat Protoc 3(4):719–738 181. Radisic M, Malda J, Epping E, Geng W, Langer R, Vunjak-Novakovic G (2006) Oxygen gradients correlate with cell density and cell viability in engineered cardiac tissue. Biotechnol Bioeng 93(2):332–343 182. Radisic M, Park H, Chen F, Salazar-Lazzaro JE, Wang Y, Dennis R, Langer R, Freed LE, Vunjak-Novakovic G (2006) Biomimetic approach to cardiac tissue engineering: oxygen carriers and channeled scaffolds. Tissue Eng 12(8):2077–2091 183. Dvir T, Levy O, Shachar M, Granot Y, Cohen S (2007) Activation of the ERK1/2 cascade via pulsatile interstitial fluid flow promotes cardiac tissue assembly. Tissue Eng 13(9):2185–2193 184. Brown MA, Iyer RK, Radisic M (2008) Pulsatile perfusion bioreactor for cardiac tissue engineering. Biotechnol Prog 24(4):907–920 185. Radisic M, Park H, Shing H, Consi T, Schoen FJ, Langer R, Freed LE, Vunjak-Novakovic G (2004) Functional assembly of engineered myocardium by electrical stimulation of cardiac myocytes cultured on scaffolds. Proc Natl Acad Sci USA 101(52):18129–18134 186. Tandon N, Cannizzaro C, Chao PH, Maidhof R, Marsano A, Au HT, Radisic M, VunjakNovakovic G (2009) Electrical stimulation systems for cardiac tissue engineering. Nat Protoc 4(2):155–173 187. Dvir T, Kedem A, Ruvinov E, Levy O, Freeman I, Landa N, Holbova R, Feinberg MS, Dror S, Etzion Y, Leor J, Cohen S (2009) Prevascularization of cardiac patch on the omentum improves its therapeutic outcome. Proc Natl Acad Sci USA 106(35):14990–14995 188. Ott HC, Matthiesen TS, Goh SK, Black LD, Kren SM, Netoff TI, Taylor DA (2008) Perfusion-decellularized matrix: using nature’s platform to engineer a bioartificial heart. Nat Med 14(2):213–221 189. Shimizu T, Yamato M, Isoi Y, Akutsu T, Setomaru T, Abe K, Kikuchi A, Umezu M, Okano T (2002) Fabrication of pulsatile cardiac tissue grafts using a novel 3-dimensional cell sheet manipulation technique and temperature-responsive cell culture surfaces. Circ Res 90(3):e40 190. Shimizu T, Sekine H, Isoi Y, Yamato M, Kikuchi A, Okano T (2006) Long-term survival and growth of pulsatile myocardial tissue grafts engineered by the layering of cardiomyocyte sheets. Tissue Eng 12(3):499–507 191. Furuta A, Miyoshi S, Itabashi Y, Shimizu T, Kira S, Hayakawa K, Nishiyama N, Tanimoto K, Hagiwara Y, Satoh T, Fukuda K, Okano T, Ogawa S (2006) Pulsatile cardiac tissue grafts using a novel three-dimensional cell sheet manipulation technique functionally integrates with the host heart, in vivo. Circ Res 98(5):705–712 192. Shimizu T, Sekine H, Yang J, Isoi Y, Yamato M, Kikuchi A, Kobayashi E, Okano T (2006) Polysurgery of cell sheet grafts overcomes diffusion limits to produce thick, vascularized myocardial tissues. FASEB J 20(6):708–710 193. Yeghiazarians Y, Zhang Y, Prasad M, Shih H, Saini SA, Takagawa J, Sievers RE, Wong ML, Kapasi NK, Mirsky R, Koskenvuo J, Minasi P, Ye J, Viswanathan MN, Angeli FS, Boyle AJ,

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Chapter 10

Skeletal Muscle Engineering: The Need for a Suitable Niche Frédéric Trensz, Anthony Scimè, and Guillaume Grenier

Abstract  There are currently no curative treatments available for people suffering from one of the many prevalent disease- and trauma-related muscle myopathies. One approach to ameliorate these conditions relies on the cell-based transplantation of myogenic stem cells or, more optimistically, the transfer of engineered skeletal muscle tissue. To date, clinical trials with myogenic stem cell transplantation have met with only modest success while engineered muscle tissue transplantation is at its earliest stages of development. The many studies on muscle tissue engineering underscore the importance of the myogenic stem cell niche that plays a pivotal role in transplantation success. More work is required to determine the components of the niche required for improving the integration and function of transplanted cells and engineered tissues in host muscle.

Abbreviations bFGF DMD ECM EGF HGF IGF1 PEG

Basic fibroblast growth factor Duchenne muscular dystrophy Extracellular matrix Epidermal growth factor Hepatocyte growth factor Insulin-like growth factor-1 Poly(ethylene glycol)

G. Grenier (*) Étienne-Lebel Clinical Research Center, Université de Sherbrooke, J1H 5N4, Sherbrooke, QC, Canada Department of Orthopedic Surgery, Université de Sherbrooke, J1H 5N4, Sherbrooke, QC, Canada e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_10, © Springer Science+Business Media, LLC 2011

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Polyglycolic acid Poly-l-lactic acid Vascular endothelial growth factor Wingless integration site

10.1 Introduction Adult skeletal muscle regeneration is a highly orchestrated process that involves the activation of myogenic stem cells, their specification and proliferation as myogenic precursor cells or myoblasts, and their subsequent differentiation and fusion into new or existing myofibers [1]. Duchenne muscular dystrophy results in a major functional impairment of myogenic stem cells, which leads ultimately to death of the individual [2]. Severe muscle trauma requires that a large number of muscle stem cells be actively engaged in muscle repair [1]. In addition, some cancer treatments involve surgery that requires the restoration of muscle mass. Lastly, the inability of myogenic stem cells to become activated and differentiate properly leads to extensive muscle wasting, as is the case in some diseases, as well as sarcopenia in the elderly [3]. In recent years, the field of regenerative medicine has expanded exponentially following the discovery of stem cells, which possess an exceptional capacity to differentiate into a variety of tissues. There are significant advantages to using stem cells with myogenic potential to treat myopathies and muscle trauma. The delivery of cells containing a normal genome can replenish the pool of dysfunctional myogenic stem cells and reduce muscle wasting. Due to their myogenic potential, stem cells are good candidates for engineering skeletal muscle tissue in vitro for subsequent use in vivo [4]. Of note is the cellular environment or microniche, which is key to ensuring the maintenance of functional myogenic stem cells that proliferate in perpetuity in the host tissue. In this chapter, we briefly review skeletal muscle structure and the molecular and cellular regulatory mechanisms that lead to the formation and repair of muscle fibers. We then discuss the availability of the main myogenic stem cell population, satellite cells, for use in transplantation therapies. We also describe the importance of a suitable microenvironment for successful muscle transplantation therapies and providing beneficial and sustainable results. Lastly, we evaluate the main advances and drawbacks of using matrices and scaffolds to produce engineered skeletal muscle tissue.

10.2 Skeletal Muscle Biology 10.2.1 Skeletal Muscle Organization Skeletal muscle is the largest tissue mass in the body, making up approximately 40% of the total body weight [5]. Its main function is to perform voluntary mechanical work by contracting. Skeletal muscle is composed of hundreds of cylindrical

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multinucleated cells known as muscle fibers or myofibers. Myofibers are surrounded by a basal lamina that separates them from the connective tissue through which blood vessels run to provide nutrition, nerves to induce fiber contraction, and muscle resident stromal cells to maintain tissue integrity. Muscle fiber is composed of an array of stacked myofibrils running the entire length of the cell constituting the contractile unit of the muscle. This high level of micro- and macroorganization permits optimal contraction, which allows the muscle to fulfill its function. Hence, it is imperative that this organization be maintained throughout the life of an individual. Lastly, it should be kept in mind that any intervention must be aimed at preserving or, at the very least, favoring this organization to ensure proper tissue replacement and repair.

10.2.2 Skeletal Muscle Regeneration Understanding the muscle regeneration process is fundamental for designing effective tissue engineering strategies using stem and progenitor cells. Following damage, as is the case after intensive exercise or trauma, skeletal muscle has the remarkable ability to initiate a rapid and extensive repair process. This strong regenerative potential is mainly attributed to a population of mononucleated myogenic stem cells called satellite cells [3]. These cells were first described by their anatomical localization between the plasma membrane of the myofiber and the basal lamina surrounding it [6]. This location provides a unique microenvironment that maintains the satellite cells in a quiescent state under normal conditions and ensures their rapid activation and differentiation following injury. Once activated, satellite cells actively proliferate before fusing with preexisting myofibers or they form entirely new myofibers as they undergo a terminal differentiation process [1].

10.3 Myogenic Cells for Transplantation Given that satellite cells are the main source of adult myogenic stem cells, myoblasts, their activated progeny has been tested for their regenerative potential following transplantation. Partridge et al. performed the first successful myoblast transplantation in mice [7]. The transplanted myoblasts fused to host myofibers and generated dystrophin-positive myofibers in mdx mice, a mouse model for muscular dystrophy [8]. Despite encouraging early studies, the effectiveness of in  vitro-expanded myoblasts injected into degenerating murine skeletal muscle has provided mixed results [9]. In fact, injected myoblasts are rapidly lost within 1 week with little or no incorporation into the satellite cell niche, which is essential for long-term therapy and further muscle regeneration. Recently, a new method to improve the efficacy of the transplantation of myogenic cells has been developed [10]. Rather than injecting myoblasts directly, satellite cells are purified and injected before they become activated. An elegant study by

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Sacco et al. using bioluminescence tracking showed that CD34−a7-integrin+ satellite cells (10–5,000) injected into the tibialis anterior muscle differentiated robustly. Following injury, these cells were able to generate massive waves of proliferation, self-renewal, and occupation of the satellite niche. They also retained the ability to differentiate, forming new myofibers or adding myonuclei to existing myofibers. Unlike purified satellite cells, transplanted control myoblasts displayed none of the functional capacity of the satellite cells [10]. Hence, these results underscore that the level of activation of transplanted cells are critical for the success of cell therapy. Notwithstanding satellite cells, other muscle resident cells have been identified and assessed for their myogenic potential. They are found between the fibers and often in close proximity to blood vessels. Among these different cells are mesoangioblasts that can be isolated from different species, including human [11, 12]. A number of studies have shown that mesoangioblasts are capable of improving the regenerative characteristics of different skeletal muscle pathologies by cell therapy [13, 14].

10.4 Current Challenge: Engineering Skeletal Muscle with a Viable Niche The satellite cell microenvironment or niche is a unique combination of cellular, biophysical, and biochemical components that can both preserve and promote satellite cell quiescence and activation. It is only through understanding the attributes and components of the microenvironment that it is possible to engineer viable and more functionally efficient muscle tissue. Satellite cells are located in an anatomically defined niche between the basal lamina and the myofiber plasma membrane [15, 16]. This location exposes them to signals from the basal lamina on their basal surfaces and adjacent myofibers on their apical surfaces (Fig. 10.1). Interestingly, the loss of myofibers, but not the basal lamina, by myotoxic drugs, leads to a great number of proliferating myogenic cells that are directed to regenerating damaged fibers or generating new ones [17]. In skeletal muscle, the basal lamina is made up of a thick network of ECM ­components, mainly laminin, fibronectin, type IV collagen, heparan sulfate, and other proteoglycans [18]. The ECM is known to modulate tissue homeostasis through its ability to locally bind, store, and release soluble bioactive effectors, such as growth factors [19, 20]. These factors are produced and secreted by muscle cells, other cell types such as stromal cells, and from the systemic environment through blood vessels. In muscle, the niche is enriched in proteoglycans that bind and sequester growth ­factors, such as bFGF [21], EGF [22], HGF [23], IGF1 [24], and diverse Wnt ligands [25, 26]. These factors are released primarily following tissue damage, thus influencing myogenic stem cell behavior. Apart from the basal lamina, the ECM in the interstitial space accommodates various cell types, such as endothelial and stromal cells, that also influence stem cell activity by releasing factors into the local environment [24, 27, 28]. Furthermore, satellite cells regulate their own behavior through autocrine and paracrine signaling by releasing the ligands for the Notch receptor family that are

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Fig. 10.1  Schematic representation of the satellite cell niche. Satellite cells are located between the plasma membrane of the myofiber and the basal lamina. This microenvironment or niche serves to maintain the cells in their quiescent state. However, the cells in the niche are exquisitely sensitive to changes in the microenvironment, which allows them to become quickly activated and proliferate when tissue repair is required. The stem-like behavior of satellite cells is made possible through the convergence of the numerous factors represented in this drawing: (1) the ECM, which includes the basal lamina and influences satellite cells through attachment (integrin), sequestration of cytokines and growth factors by specialized ECM proteins (e.g., proteoglycans), and mechanical sensing (e.g., stiffness). (2) The systemic environment is also a major contributor of soluble growth factors (e.g., Wnt ligands) that influence satellite cells. This is particularly true because of the close association between capillaries and satellite cells. (3) Cells in the perivascular region and interstitial space, including blood-borne cells (e.g., macrophages and immune cells) as well as muscle-resident stromal cells, play an important role in satellite cell function. These cells are essential for the degradation/production of the ECM and the secretion of cytokines and growth factors that can stimulate satellite cells within their niche. (4) Lastly, satellite cells interact with adjacent myofibers through M-cadherin junctions and communicate via the Delta-Notch pathway, which is involved in satellite cell activation

involved in their quiescence and self-renewal [29–31]. Lastly, other soluble factors, such as myostatin, IGF-1, and Wnt3a which originate from the systemic circulation, exert extrinsic control on satellite cell function [24, 27, 32]. The microenvironment is also important for providing optimal mechanical support for the satellite cell. The elastic stiffness of the ECM, which is often assessed by its

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elastic modulus, can strongly affect the proliferation and differentiation of satellite cells [33, 34]. Healthy skeletal muscles and cultured myotubes have a similar elastic modulus. However, changes in stiffness that occur during aging and in myopathies, such as muscular dystrophy, can alter the biophysical properties and functions of skeletal muscle. Subtle changes in matrix stiffness have been shown to alter the activation, proliferation, and differentiation of satellite cells and myogenic progenitors in vitro [35, 36]. Another important aspect of the niche in controlling satellite cells’ behavior is the elucidation of the mechanism involved in their symmetric and asymmetric self-renewal. In vivo, satellite cells are exposed to signals from the basal lamina on their basal surfaces and the adjacent myofibers on their apical surfaces. This spatial arrangement influences their self-renewal potential [16] and underscores the need to develop an in vitro 3D niche model that mimics the bipolar organization of the niche [28]. The importance of the niche for ensuring the full potential of satellite cells is highlighted by the fact that whole single myofiber transplantation is more effective in restoring skeletal muscle function than myogenic stem cell transplantation. Collins et al. have shown that a single transplanted myofiber containing between 7 and 22 satellite cells, depending on the muscle of origin, can contribute to generating hundreds of host myofibers by supplying thousands of myonuclei [15]. More importantly, the transplanted satellite cells on the myofiber were able to self-renew, expand, and repopulate the niche of the newly formed myofibers. Engineered satellite cell niches can be used to understand many aspects of the satellite cell microenvironment, including the niche components required for the maintenance of satellite cell quiescence and myogenic potential over prolonged culture. Following isolation, myoblasts can undergo only a finite number of divisions [37], and their inability to recapitulate the proliferative capacity in vitro is likely due to the loss of the highly specific niche that normally surrounds these cells. Thus, the development of an in  vitro system in which the molecular components are welldefined would make it possible to analyze their contribution and reproduce specific aspects of cell–cell and cell–ECM interactions. Recently, several in vitro biomaterial systems that mirror the regulatory characteristics of natural ECM have been developed to study the microenvironmental regulation of stem cell behavior. These include 2D and 3D hydrogels that have tunable elastic moduli and controlled ligand patterning and release [38–41]. PEG-based hydrogels are a promising in vitro-engineered cell culture model for the understanding of the in vivo maintenance of satellite cell quiescence because of their modifiable matrix stiffness and controlled ligand tethering that permit the control of cell adhesion [42]. PEG-based hydrogels can attain a wide range of elastic moduli, including the 10–15 kPa stiffness range of healthy resting muscle that is optimal for myogenic differentiation in vitro [35, 36]. Kloxin et al. recently developed a photodegradable PEG-based hydrogel by rapid polymerization of macromers that has the advantage of adapting or modulating gel characteristics once implanted [43]. This ability to temporally control biophysical properties without being toxic to cells is of great interest because it allows for the progression of biophysical modifications in diseases, such as those associated with skeletal muscle fibrosis [35, 44, 45].

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Another major advantage of PEG-based hydrogels is their ability to tether growth factors. Tethering enables sequestration of growth factors within the niche by myofibers and the basal lamina as normally occurs [46]. A number of studies highlighted that the surface tethering of growth factor ligands makes it possible to physically control receptor–ligand internalization and maintain receptor surface residency, activation, and proximal ligand concentrations [47–49]. Irvine et al. demonstrated that tethering can be spatially controlled on a nanoscale level, facilitating ligand clustering, which increases receptor activation [50]. Due to their tethered ligand presentation, nanoscale patterning, biophysical proprieties, and controlled microscale features, PEG-based hydrogels are a promising tool for engineered muscle tissue applications. This system is also of great value in identifying the precise roles of specific niche components in their tethered and soluble forms.

10.5 Engineered Skeletal Muscle Tissue The complex network of signals and mechanical properties in the niche underscores the need to develop an engineered muscle tissue that possesses biophysical and biochemical components similar to the in vivo environment. Over the years, a variety of strategies have been implemented to design tissues with functions and characteristics closer to those of native muscle. Basically, skeletal muscle tissue engineering consists of the proliferation and differentiation of myogenic cells into terminally differentiated myotubes within a specialized support material, matrix or scaffold, which makes up a 3D environment. The matrix or scaffold supports are a crucial component that play two main roles. First, they create an appropriate environment for the attachment, proliferation, alignment, and differentiation of myoblasts into myotubes. Second, they provide support for the engineered tissue that enhances proper development and function within a 3D space. The support material used in tissue engineering must be biocompatible and biodegradable for efficient integration into the host and must also have mechanical properties that enable the construct to support stretch and force production. Moreover, it should provide a large surface area for cell–environment interactions while maintaining a suitable rate of nutrient diffusion during in vitro culture to prevent necrosis [51]. It must also allow efficient vascularization, which is essential for resident cell viability. Indeed, oxygen diffusion, nutrient delivery, waste removal, protein transfer, and cell migration are critical factors that are governed by the porosity and permeability of the support material.

10.5.1 Gel Matrix Supports Matrigel in combination with collagen is a common biomaterial used for cultivating 3D muscle constructs. Matrigel is a matrix extracted from the Engelbreth-Holm-Swarm mouse sarcoma that contains ECM proteins and undefined concentrations of growth

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factors that enhance myogenic differentiation [52, 53]. Matrigel in combination with collagen has the ability to change the expression of genes in cells and allows myoblasts to differentiate in  vitro, resulting in the formation of an artificial skeletal muscle containing parallel arrays of myofibers expressing contractile proteins [53]. However, this artificial muscle, which is characterized by small-diameter myofibers, lacks many of the morphological attributes of native muscle. Fibrin is a natural compound that can be used to support muscle tissue growth within a gel matrix. It is a biocompatible, biodegradable, and nonimmunogenic compound that provides a large surface area for cell–matrix interactions. Additionally, it allows sufficient space for ECM generation and allows a minimum diffusion of nutrients in in vitro cultures [51, 54, 55]. Its porous structure can be further improved by incorporating thrombin at various concentrations to influence the diffusion of nutrition factors and, hence, cellular growth [56, 57]. Beier et al. used fibrin as a 3D platform for the growth and delivery of rat primary myoblasts into defective anterior gracilis muscles of female rats [58]. Interestingly, the myoblast-fibrin matrix enhanced the integration of the myoblasts into host muscle fibers without causing an inflammatory reaction [58]. While the myoblast-fibrin injection technique can be used to restore skeletal muscle tissue locally, it cannot be adapted to pathologies, such as muscular dystrophy, cachexia, or muscle disuse, which require a more systemic approach. The use of matrix supports has advanced our knowledge of myoblast behavior within a 3D environment. Although they can be used as vehicles for local myoblast delivery, their usefulness in tissue engineering is limited as they do not fulfill certain important requirements, such as ease of handling and maintaining an organized array of fibers.

10.5.2 Scaffold Supports One of the major shortcomings of using gel matrices as supports is the lack of structural organization of the newly formed constructs. Muscle fibers need to be oriented parallel to each other to ensure directed force production. Scaffolds able to support cell fusion and the formation of long continuous muscle fibers have been developed to overcome this major shortfall. These scaffolds are made from a variety of biomaterials, including biopolymers and synthetic polymers. Unlike simple gel matrix supports, engineered scaffold constructs are biocompatible, easily handled, and are suitable for surgical attachment to host tissue. Older muscle tissue scaffold technologies made use of microfibrous polymers composed of PLLA, a classic degradable polyester. Human skeletal muscle cells can easily differentiate into multinucleated myofibers on PLLA scaffolds coated with ECM proteins. In addition, ECM-coated PLLA scaffolds enable cells to form functional muscle by directing the organization of myofibers into a parallel orientation [59]. Another microfibrous polymer, PGA mesh, was originally used to fabricate muscle tissue. Differentiated neonatal rat myoblast PEG meshes implanted in rat

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peritoneal cavities displayed highly vascularized 3D structures with the ability to generate muscle-like tissue 6 weeks posttransplantation [60, 61]. More recently, Degrapol, an electrospun degradable polyesterurethane membrane, has been used to engineer soft tissues, including skeletal muscle. Degrapol promotes favorable cell/scaffold interactions with the myoblast cell lines C2C12 and L6 [62], which proliferate and differentiate into multinucleated myotubes expressing myogenic markers when grown in Degrapol [62, 63]. More importantly, Degrapol also promotes C2C12 cell adhesion and myotube alignment along the scaffold fibers and degrades within 180 days [62]. By applying an electrospinning process to Degrapol gel, cells become highly oriented. Hence, as well as serving as an architectural and mechanical support, Degrapol can also be arranged to provide necessary directional cues for cells during differentiation [63]. Despite these promising results, there are no reports in the literature to date describing the performance of Degrapol following transplantation. A shortcoming of scaffold-based tissue-engineered muscles is the lack of sufficient force generation. In most studies, the forces generated by fabricated constructs are 2–10% lower than those generated by adult mammalian skeletal muscle [64–67]. However, recent studies have shown that the force generated by engineered muscle can be significantly improved by adding IGF-1, a well-known skeletal muscle anabolic factor [68, 69]. The improvement in muscle force induced by IGF-1 is the result of an increase in myofiber size (40%), contractile protein synthesis, and sarcomeric organization [69]. Differentiated cells aligned within a construct that is supplemented with an anabolic factor may, thus, display increased force generation since the intensity of the contraction is directly related to the cross-sectional area of the myofibers [69–71]. It is also crucial that the fabricated muscle tissue be properly vascularized to maintain essential tissue nutrition and reconstitute the satellite cell niche. Unlike the peripheries of scaffold-engineered muscle grafts, the interiors are usually associated with poor nutrient diffusion. To circumvent this problem, Langer et al. induced the formation of blood vessel networks in engineered skeletal muscle tissue by seeding a porous biodegradable polymer scaffold made of PLLA and PGA with embryonic fibroblasts and endothelial cells together with myoblasts [72, 73]. This tissue construct produced a large number of blood vessels from endothelial cells that was stimulated by VEGF secreted by the embryonic fibroblasts [74]. When implanted in mice and rats, the long, thick donor myotubes aligned with the host myofibers. Donor vessels containing red blood cells also formed along the host muscle fibers, indicating that the vessels had anastomosed with the recipient vasculature. Despite this construct’s performance, its use for clinical applications requires further study with regard to biocompatibility, handling, viability, and contractility. Nonetheless, the addition of support cells within a construct may provide a number of advantages. Not only do support cells ensure proper nutrition for the engineered tissue, they also promote the formation of the niche. Indeed, stromal cells and other cell types secrete ECM, which provides a physiological architecture for the cells.

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10.6 Concluding Remarks In recent years, there have been many advances in the field of skeletal muscle tissue engineering, but its dissemination for use in muscle diseases and myopathies has not yet occurred. Satellite cells show great potential for contributing to skeletal muscle regeneration after transplantation. Despite this potential, other cell types might have advantages that should not be ignored. However, their successful use should not only be measured by the formation of viable myotubes, but also by the production of a viable niche closely resembling their in vivo environment. This is essential for the proper maintenance of satellite cell function, that is, the activation/ proliferation, differentiation, and self-renewal potential that ensure long-term engraftment. Hence, key factors, such as the ECM, neighboring cells, and the systemic environment, need to be evaluated and optimized in engineered muscle tissue. The use of stromal cells in combination with endothelial cells may favor a viable microenvironment in tissue-engineered constructs that allows myogenic stem cells to survive in perpetuity. Currently, numerous scaffolds and biomaterials offer promising avenues for fabricating functional skeletal muscle that are closer to imitating the biophysical and biochemical features of the myogenic stem cell microenvironment than in the past.

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Chapter 11

Restoring Blood Vessels Narutoshi Hibino, Christopher Breuer, and Toshiharu Shinoka

Abstract  In surgical repair for heart disease, it is sometimes necessary to fill or replace a pathological tissue or defect with autologous graft tissue or a foreign grafting material. To date, (1) autologous pericardium, (2) allograft, (3) xenograft, and (4) artificial graft (e.g., Dacron, Teflon, Gore-Tex) have been used as graft materials. These grafts, however, lack growth potential, are associated with increased risk of thrombosis and infection, and have limited durability, thus increasing the morbidity and mortality of their application. Vascular tissue engineering is a relatively new concept proposed in the latter half of the 1980s. It aims to produce neotissue from autologous cells with biodegradable polymer as a scaffold by the application of engineering and biological principles. The greatest advantage of tissue constructed by tissue engineering is that the scaffold polymer is completely biodegraded as cells fill the extracellular stroma, and foreign materials do not remain at later time points after transplant. In this review, we provide an overview of our work to demonstrate the advantages of tissue-engineered vascular grafts in animal models and in human clinical applications using autologous cells and biodegradable scaffolds.

Abbreviations Ac-LDL Acetylated low-density lipoprotein EPC Endothelial progenitor cell PTFE Polytetrafluoroethylene

T. Shinoka (*) Department of Cardiac Surgery, Interdepartmental Program in Vascular Biology and Therapeutics, Yale University School of Medicine, New Haven, CT, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_11, © Springer Science+Business Media, LLC 2011

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11.1 Introduction It is widely accepted that the ideal conduit has not yet been developed for surgery in cardiovascular disease. PTFE (or Gore-Tex) conduits are currently the most widely used vascular grafts [1]. Use of PTFE has essentially replaced the use of Dacron in most centers. Failure rates for Dacron grafts are limited but tend to be worse than PTFE failure rates described in the literature [2]. Homografts have also been used as conduits but to a much more limited extent compared to PTFE [3]. For synthetic conduits, thrombosis is the leading cause of graft failure and conduit replacement in the early postoperative period [2]. Synthetic conduits are a significant cause of thromboembolic complication due to the large surface area of synthetic material in contact with blood that can cause activation of the coagulation cascade [3]. Graft thrombosis frequently necessitates remedial surgery requiring replacement grafts, which in turn are associated with additional compromise of graft patency [4]. Other clinically available conduits, including biological grafts such as homografts and heterografts, are associated with significantly lower thromboembolic complication rates compared to synthetic grafts. However, they too lack growth potential and unfortunately have poor durability due to their propensity for accelerated calcific degradation and secondary graft failure [5–7]. These grafts tend to become stenotic and calcified, eventually needing to be replaced [8]. This process appears to be immune mediated and more aggressive in younger patients [9]. Early and midterm results for these grafts are variable with 5-year patency rates between 65 and 90%. Long-term data demonstrating graft failure rates between 70 and 100% at 10–15 years have been reported [4, 5, 7].

11.2 Animal Studies 11.2.1 Tissue-Engineered Grafts Seeded with Autologus Vascular Cells Using the classical tissue engineering paradigm, autologous cells can be seeded onto a biodegradable tubular scaffold. The scaffold provides sites for cell attachment and space for neotissue formation [10]. The resulting neotissue can be used for reconstructive surgical applications such as creation of a vascular graft for use in pediatric cardiothoracic operations. We were the first to describe the creation of a tissueengineered pulmonary artery conduit using tissue engineering methodology [11]. In this study, a conduit was created by seeding either autologous arterial or venous cells onto a biodegradable tubular scaffold fabricated from polyglactin woven mesh sealed with nonwoven polyglycolic acid mesh. The cells were obtained by explanting segments of autologous artery or vein and expanding the resulting mixed cell population in culture. An endothelial-enriched cell population was obtained by labeling the mixed cell population with an Ac-LDL marker and sorting the labeled cells using a

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fluorescence-activated cell sorter. The Ac-LDL-positive cell population (endothelial cell enriched) was then separately expanded in culture. Ten million Ac-LDL-negative cells were statically seeded onto the scaffold and maintained in culture for 1 week. The lumen of the Ac-LDL-negative seeded construct was then seeded with the Ac-LDL-positive cells and maintained in culture for one additional day. Seeded or unseeded constructs were used as interposition grafts replacing 2 cm sections of the main pulmonary artery in a juvenile lamb model. The tissue-engineered pulmonary conduits were serially monitored using echocardiography and angiography. Conduits were harvested 6  months after implantation. All seeded scaffolds were patent and demonstrated nonaneurysmal increase in size. The unseeded constructs formed thrombus and stenosed within 2 weeks of implantation. Histologically, none of the biodegradable polymer scaffold remained in any tissue-engineered graft by 11  weeks. Collagen content in the tissue-engineered grafts was 73.9% of native pulmonary artery 6  months after implantation. Histologically, elastic fibers were present in the media layer of the tissue-engineered vessel wall and endothelial specific factor VIII-related antigen was identified on the luminal surface. DNA assay showed a progressive decrease in numbers of cell nuclei over 11 and 24 weeks, suggesting remodeling. Calcium content of tissue-engineered grafts was elevated compared to native pulmonary artery but no macroscopic calcifications were found. We concluded that (a) living vascular grafts could be engineered from autologous cells and biodegradable polymers that functioned well in the pulmonary circulation as a pulmonary artery replacement grafts; (b) tissue-engineered grafts demonstrate an increase in diameter suggesting growth and development of the viable conduits over time; (c) grafts possess an endothelial cell lining and extracellular matrix with collagen and elastic fibers that resemble the native pulmonary artery [11]. In 2001, we reported our findings using a new biodegradable scaffold [12]. In this study, tissue-engineered vascular grafts were constructed by statically seeding 4–6 million autologous venous cells onto a biodegradable scaffold fabricated from a 50:50 copolymer of l-lactide and e-caprolactone reinforced with nonwoven polyglycolic acid fiber fabric. The mixed cell populations were obtained by explanting segments of the saphenous vein and expanding the cells in culture. The mixed cell population was seeded onto the lumen of the scaffold and maintained in culture for 1 week to allow for cell attachment. The tissue-engineered autografts were used as interposition grafts to replace the intrathoracic inferior vena cava in a dog model. Grafts were serially monitored using angiography. The tissue-engineered grafts were harvested over a 6-month time course. The grafts showed no evidence of stenosis or dilation. No thromboembolic complication occurred even without anticoagulation therapy. The scaffold was degraded by 3 months. Immunohistochemical staining revealed the presence of factor VIII-related antigen positive nucleated cells on the luminal surface of the tissue-engineered autograft demonstrating an endothelial monolayer lining the graft. Both smooth muscle cell a-actin and desmin, indicative of smooth muscle cells, were noted in the wall of the graft in addition to a robust extracellular matrix deposition including both collagen and elastin fibers. From these studies, we concluded that tissue engineering could be used to create a functional venous conduit [12].

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11.2.2 Autologous Bone Marrow-Derived Cells as an Alternative to Vascular Cells We realized that for future clinical application, use of autologous cells derived from the patient’s own tissue would require an additional surgical procedure, with associated risks inherent to any operation, including anesthetic complications, bleeding, and infection. In addition, explantation is not always successful and sometimes fails to produce cells for expansion in culture [13]. The ability to harvest cells for culture can be affected by the patient’s age or underlying medical condition [14]. The expansion of cells in culture is time-consuming (typically 8–12 weeks), which limits the availability and clinical utility of tissue-engineered grafts. Finally, growth of autologous cells in culture exposes the cells to various environmental factors including medium, serum, and environmental pathogens such as bacteria, viruses, and fungi. This exposure increases the risk of contamination and even cellular de-differentiation, thereby increasing the potential for infectious complications and the potential for malignant degeneration [13]. We postulated that another cell source for seeding, such as bone marrow cells, could substitute autologous cells from tissue. The identification of the bone marrow-derived EPC, which contributes to angiogenesis and vasculogenesis, provided a potential alternative cell source for cardiovascular tissue engineering [15–17]. Noishiki and colleagues reported that bone marrow cells implanted on the surface of an arterial graft led to earlier endothelialization in a large animal model [18]. The demonstration that these cells have the ability to participate in endothelialization lead to a series of studies evaluating the use of various cellular components of the bone marrow for cardiovascular tissue engineering. To avoid the limitations imposed by autologous vascular cell sources, we investigated the use of bone marrow cells for vascular tissue engineering. In 2003, we described a large animal study evaluating the use of autologous bone marrow cells to create a tissue-engineered vascular graft using a 50:50 copolymer of l-lactide and e-caprolactone reinforced with nonwoven polyglycolic acid fiber fabric [19]. Approximately, 750 autologous bone marrow cells/mm2 were statically seeded onto the luminal surface of the scaffold and maintained in culture for 2 h to allow for cell attachment. The seeded constructs were then used as interposition grafts replacing the intrathoracic inferior vena cava in an adult beagle model. The grafts were harvested over a 2-year time course. All seeded grafts were patent without evidence of thrombosis, stenosis, or aneurysm formation. Immunohistochemical analysis demonstrated that the seeded bone marrow cells, expressing endothelial cell lineage markers, such as CD34, Flk-1, and Tie-2, adhered to the scaffold. This was followed by proliferation and differentiation, resulting in expression of endothelial markers such as CD146, factor VIII, vWF, and CD31, and smooth muscle cell markers, such as smooth muscle cell a-actin, SMemb, SM1, and SM2. Vascular endothelial growth factor and angiopoietin-1 were also produced by the tissue-engineered autografts [19]. A subsequent study was designed to assess the use of bone marrow cells compared with the previous method using vascular wall cells to construct tissueengineered vascular grafts [20]. Biodegradable polymers seeded with different types

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of cells (cultured venous cells, bone marrow cells without culture, non-cell-seeded graft control) were implanted into the inferior vena cavae of dogs. The grafts were explanted at 4 weeks and assessed histologically and biochemically. Histologically, a regular layer of Masson-staining collagen fiber and a layer of factor VIII-stained endothelial and antismooth muscle cell a-actin-immunoreactive cells stained in the venous cell and bone marrow cell groups, similar to native vascular tissue, whereas no such staining was detected in the nonseeded control group [20]. Based on these results, we concluded that autologous bone marrow cells provided a practical source for creating tissue-engineered vascular grafts that enabled rapid cell harvest and seeding without the need for cell culture. In 2006, we reported our work characterizing tissue-engineered venous conduits constructed by seeding autologous bone marrow-derived mononuclear cells onto a biodegradable scaffold fabricated from poly (l-lactide-co-e-caprolactone) reinforced with polyglycolic acid fiber mesh [21]. The tissue-engineered autografts were used as interposition grafts replacing the intrathoracic inferior vena cava in the adult beagle model. Graft function was serially monitored in  vivo and the grafts were harvested over a 12-month time course. Animals were maintained without anticoagulation. All grafts were patent without evidence of significant stenosis or dilatation. Normal vessels dilate in response to acetylcholine, an agent that acts by inducing endothelial cells to release nitric oxide, a smooth muscle relaxant. Segments of the explanted grafts were stimulated with acetylcholine, and these produced nitrates and nitrites in a dose-dependent fashion. NG-nitro-l-arginine methylester, a competitive inhibitor of endothelial nitric oxide synthase, significantly inhibited the acetylcholine response. With stimulation by acetylcholine, factor VIIIpositive cells from the tissue-engineered grafts produced endothelial nitric oxide synthase proteins, and the ratio of endothelial nitric oxide synthase/s17 mRNA was similar to native inferior vena cava tissue. The tissue-engineered grafts had biomechanical properties and wall thicknesses similar to those of the native inferior vena cava within 6  months after implantation. The tissue-engineered grafts tolerated venous pressure without evidence of calcification. The number of inflammatory cells in the tissue-engineered grafts and expression levels of CD4/s17 mRNA were also noted to decrease significantly with time. We concluded that tissue-engineered vascular grafts could be successfully constructed by seeding autologous bone marrow-derived mononuclear cells onto a polyglycolic acid fiber fabric scaffold creating a conduit with functional endothelial cells and biomechanical properties similar to native inferior vena cava [21]. We have continued to investigate the use of tissue-engineered vascular grafts constructed by seeding autologous bone marrow-derived mononuclear cells onto a poly (l-lactide-co-e-caprolactone)/polyglycolic acid biodegradable tissue engineering scaffold as large caliber venous interposition grafts. We investigated the growth potential of tissue-engineered autografts with intrathoracic inferior vena cava replacement in a juvenile lamb model. This model was used because of its propensity to undergo accelerated calcific degradation, making it the preferred model for evaluating vascular grafts for use in congenital heart surgery. In this investigation, the tissue-engineered vascular grafts were serially monitored using magnetic resonance

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imaging to determine changes in graft size. Preliminary findings demonstrate excellent graft patency without evidence of thrombosis or aneurysmal dilation. Histologic evaluation of the conduits shows excellent neotissue formation with a robust extracellular matrix and endothelial lining. Immunohistolochemical staining and Western blot analysis showed that Ephrin B4, a determinant of normal venous development, is expressed in the seeded graft 6 months after implantation. Successful results of this ongoing work provide further support of the feasibility and safety of using this tissue engineering methodology to create a living graft for use as a large caliber venous graft for pediatric cardiovascular surgical applications [22].

11.3 Clinical Studies 11.3.1 First-in-Human Application of Tissue-Engineered Graft Based on these animal data, we initiated a clinical trial evaluating the use of tissueengineered grafts and patches for congenital heart operations at the Woman’s Medical University in Tokyo, Japan. We performed the first successful surgery using an autologous vascular graft in 1999 [23]. A 2-cm thrombosed segment of a pulmonary artery was replaced with a tissue-engineered graft in a 4-year-old girl who had undergone pulmonary artery angioplasty and extracardiac cavopulmonary conduit placement at the age of 3 years. The tissue-engineered graft was created from autologous cells obtained from a 2-cm segment of explanted peripheral vein. The cells from the explant were expanded in culture for 8 weeks and then 1.2 × 107 cells were statically seeded onto a tubular scaffold. The scaffold was fabricated from a 50:50 copolymer of e-polycaprolactone-polylactic acid reinforced with woven polyglycolic acid fibers. The biodegradable polymer conduit measured 10 mm in diameter, 20  mm in length, and 1  mm in thickness and was designed to degrade over an 8-week period. Ten days after seeding, the graft was implanted. No postoperative complications occurred. On follow-up angiography, the graft was noted to be completely patent. Seven months after implantation, the patient was doing well, without evidence of graft occlusion or aneurysmal changes on chest radiography [23].

11.3.2 Midterm Results of Clinical Trial Based on the results of previous animal experiments, we changed the cell source used to create the tissue-engineered vascular conduit for the clinical trial from autologous cells derived from explanted venous tissue to autologous bone marrow mononuclear cells. Our midterm clinical results of tissue-engineered autografts seeded with bone marrow cells used for reconstructive cardiovascular surgery are

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based on procedures performed between 2001 and 2004 [13]. During this time, 42 consecutive patients received tissue-engineered autografts. The tissue-engineered grafts were created from autologous mononuclear bone marrow cells seeded onto a biodegradable scaffold. The mononuclear bone morrow cells were obtained by aspirating 4–5 ml/kg body weight of bone marrow from the patient. The bone marrow mononuclear cells were obtained by density centrifugation with Histopaque-1077 (Sigma) and culture in the patient’s own plasma. The biodegradable scaffold was constructed from a woven fabric of either poly-l-lactic acid or polyglycolic acid fibers. The woven fabric were coated with a copolymer of lactic acid and e-caprolactone (50:50) and lyophilized. The scaffold was 1 mm thick and greater than 80% porous with pores measuring between 20 and 100  nm. Scaffolds of varying lengths and diameters measuring between 12 and 24  mm were utilized. The scaffolds were seeded with a concentrated suspension of 3 × 105 cells/cm2. The outer surface of the seeded biodegradable scaffolds was then sprayed with fibrin glue and placed in a incubator for 2–4 h at 37°C in 100% humidity and 5% CO2. Twenty-three patients had a tube graft as an extracardiac total cavopulmonary conduit implanted, while the other 19 patients had a sheet-type patch used for repair of congenital cardiac defects. Inclusion criteria for patients for this procedure were as follows: elective surgery, age younger than 30 years, and good quality of other organ function. All patients underwent a catheterization study, computed tomographic scan, or both after the operation. The patients received 3–6 months anticoagulation therapy. Mean follow-up after surgery was 490 ± 276 days (1.3–31.6 months, median 16.7 months). There were no complications such as thrombosis, stenosis, or obstruction of the tissue-engineered autografts. One late death at 3  months after extracardiac total cavopulmonary conduit was noted in a patient with hypoplastic left heart syndrome; this was unrelated to the tissue-engineered graft function. There was no evidence of aneurysm formation or calcification on cineangiography or computed tomography (Fig. 11.1). All tube grafts were patent, and the diameter of the tube graft increased with time (110 ± 7% of the implanted size) [13].

11.3.3 Late Term Results of Clinical Trial In 2010, we reported the late clinical and radiologic surveillance of a patient cohort who underwent implantation of tissue-engineered vascular grafts as extracardiac cavopulmonary conduits [24]. Autologous bone marrow mononuclear cells were seeded onto a biodegradable scaffold composed of polyglycolic acid and e-caprolactone/l-lactide. Twenty-five grafts were implanted as extracardiac cavopulmonary conduits in patients with single ventricle physiology. Patients were followed up by postoperative clinic visits and by telephone. In addition, ultrasonography, angiography, computed tomography, and magnetic resonance imaging were used for postoperative graft surveillance. There was no graft-related mortality during the follow-up period (range, 4.3–7.3 years; mean, 5.8 years). All patients underwent a

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Fig. 11.1  Three-dimensional computed tomography 1 year after tissue-engineered vascular graft implantation. The graft is patent and there is no aneurysmal dilation. Arrows indicate extracardiac total cavopulmonary conduit graft

catheterization-based angiographic study, computed tomography, or magnetic resonance imaging. There was no evidence of aneurysm formation, graft rupture, or ectopic calcification in any graft interrogated with any imaging modality (Fig. 11.2). Six (24%) patients had asymptomatic graft narrowing noted on routine surveillance imaging. Four of six patients underwent successful balloon angioplasty, including one patient who required repeat balloon angioplasty and stent placement in the stenosed segment of the graft. All contacted patients were attending school or work regularly. Seventeen (81%) patients were in New York Heart Association functional class I and three patients were in functional class II. Eight patients (40%) were not receiving any daily medications [24].

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Fig. 11.2  Tissue-engineered vascular graft angiography, 5  years (left) and 4  years (right) after implantation. There was no stenosis, aneurysm formation, or ectopic calcification in these tissueengineered vascular grafts

11.4 Conclusions The long-term failure and significant complication rates associated with currently used vascular grafts are a powerful impetus for exploring new and potentially improved alternatives. Tissue engineering provides a method for creating an improved vascular conduit. The feasibility of engineering large caliber, autologous vascular bioconduits for use as venous interposition grafts for congenital heart surgery has been demonstrated in both large animal investigations and in a human clinical trial. Early results suggest that the use of tissue-engineered vascular grafts is both safe and efficacious, however, these promising results have not been compared to conventional vascular grafts in a randomized, controlled study. In addition, this groundbreaking work has only been performed at one facility to date, Tokyo Women’s Medical University (with Institutional Review Board approval). The need for a carefully designed clinical trial under the supervision of the Food and Drug Administration is a necessary prerequisite for performance of this work in the USA. The development of a readily available vascular graft constructed from autologous tissue that also has growth potential has dramatic implications for the field of congenital heart surgery.

References 1. Wells W, Malas M, Baker CJ, Quardt SM, Barr ML (2003 Aug) Depopulated vena caval homograft: a new venous conduit. J Thorac Cardiovasc Surg 126(2):498–503 2. Giannico S, Hammad F, Amodeo A, Michielon G, Drago F, Turchetta A et al (2006) Clinical outcome of 193 extracardiac Fontan patients: the first 15  years. J Am Coll Cardiol 47(10):2065–2073

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3. Petrossian E, Reddy VM, McElhinney DB, Akkersdijk GP, Moore P, Parry AJ et al (1999) Early results of the extracardiac conduit Fontan operation. J Thorac Cardiovasc Surg 117(4):688–696 4. Homann M, Haehnel JC, Mendler N, Paek SU, Holper K, Meisner H et al (2000) Reconstruction of the RVOT with valved biological conduits: 25 years experience with allografts and xenografts. Eur J Cardiothorac Surg 17(6):624–630 5. Stark J (1998) The use of valved conduits in pediatric cardiac surgery. Pediatr Cardiol 19(4):282–288 6. Cleveland DC, Williams WG, Razzouk AJ, Trusler GA, Rebeyka IM, Duffy L et al (1992) Failure of cryopreserved homograft valved conduits in the pulmonary circulation. Circulation 86(5 Suppl):II150–II153 7. Jonas RA, Freed MD, Mayer JE Jr, Castaneda AR (1985) Long-term follow-up of patients with synthetic right heart conduits. Circulation 72(3 Pt 2):II77–II83 8. Bermudez CA, Dearani JA, Puga FJ, Schaff HV, Warnes CA, O’Leary PW et al (2004) Late results of the peel operation for replacement of failing extracardiac conduits. Ann Thorac Surg 77(3):881–887, discussion 8 9. Karamlou T, Ungerleider RM, Alsoufi B, Burch G, Silberbach M, Reller M et  al (2005) Oversizing pulmonary homograft conduits does not significantly decrease allograft failure in children. Eur J Cardiothorac Surg 27(4):548–553 10. Langer R, Vacanti JP (1993) Tissue engineering. Science 260(5110):920–926 11. Shinoka T, Shum-Tim D, Ma PX, Tanel RE, Isogai N, Langer R et al (1998) Creation of viable pulmonary artery autografts through tissue engineering. J Thorac Cardiovasc Surg 115(3):536–545, discussion 45–6 12. Watanabe M, Shin’oka T, Tohyama S, Hibino N, Konuma T, Matsumura G et al (2001) Tissueengineered vascular autograft: inferior vena cava replacement in a dog model. Tissue Eng 7(4):429–439 13. Shin’oka T, Matsumura G, Hibino N, Naito Y, Watanabe M, Konuma T et al (2005) Midterm clinical result of tissue-engineered vascular autografts seeded with autologous bone marrow cells. J Thorac Cardiovasc Surg 129(6):1330–1338 14. Poh M, Boyer M, Solan A, Dahl SL, Pedrotty D, Banik SS et al (2005) Blood vessels engineered from human cells. Lancet 365(9477):2122–2124 15. Asahara T, Murohara T, Sullivan A, Silver M, van der Zee R, Li T et al (1997) Isolation of putative progenitor endothelial cells for angiogenesis. Science 275(5302):964–967 16. Asahara T, Masuda H, Takahashi T, Kalka C, Pastore C, Silver M et al (1999) Bone marrow origin of endothelial progenitor cells responsible for postnatal vasculogenesis in physiological and pathological neovascularization. Circ Res 85(3):221–228 17. Shi Q, Rafii S, Wu MH, Wijelath ES, Yu C, Ishida A et al (1998) Evidence for circulating bone marrow-derived endothelial cells. Blood 92(2):362–367 18. Noishiki Y, Tomizawa Y, Yamane Y, Matsumoto A (1996) Autocrine angiogenic vascular prosthesis with bone marrow transplantation. Nat Med 2(1):90–93 19. Matsumura G, Miyagawa-Tomita S, Shin’oka T, Ikada Y, Kurosawa H (2003) First evidence that bone marrow cells contribute to the construction of tissue-engineered vascular autografts in vivo. Circulation 108(14):1729–1734 20. Hibino N, Shin’oka T, Matsumura G, Ikada Y, Kurosawa H (2005) The tissue-engineered vascular graft using bone marrow without culture. J Thorac Cardiovasc Surg 129(5):1064–1070 21. Matsumura G, Ishihara Y, Miyagawa-Tomita S, Ikada Y, Matsuda S, Kurosawa H et al (2006) Evaluation of tissue-engineered vascular autografts. Tissue Eng 12(11):3075–3083 22. Brennan MP, Dardik A, Hibino N, Roh JD, Nelson GN, Papademitris X et al (2008) Tissueengineered vascular grafts demonstrate evidence of growth and development when implanted in a juvenile animal model. Ann Surg 248(3):370–377 23. Shin’oka T, Imai Y, Ikada Y (2001) Transplantation of a tissue-engineered pulmonary artery. N Engl J Med 344(7):532–533 24. Hibino N, McGillicuddy E, Matsumura G, Ichihara Y, Naito Y, Breuer C et al (2010) Late-term results of tissue-engineered vascular grafts in humans. J Thorac Cardiovasc Surg 139(2):431–436, 6e1–6e2

Chapter 12

Engineering Functional Bone Grafts Sarindr Bhumiratana and Gordana Vunjak-Novakovic

Abstract  There is a strong medical need for biological tissue grafts that could reestablish the structure and function of bone lost to a major injury or disease. Routinely used prosthetic devices are most helpful in providing the necessary structure and mechanical support, but these devices often fail to fully integrate with the host tissues, and generally do not last longer than about 10 years. In addition, the important metabolic function of bone most certainly cannot be provided by prosthetic devices. Tissue engineering is now offering a potential to grow fully biological substitutes of native tissues, by an integrated use of living cells, biomaterial scaffolds, and culture systems (bioreactors). Today, tissue engineering modalities are designed based on the biological requirements and clinical constraints, and the progress is largely made at the interfaces between bioengineering, basic, and clinical sciences. This chapter is discussing the design criteria and parameters essential for engineering bone grafts, as well as the current status and future perspective of the field.

Abbreviations BMP FGF GM-CSF IGF Ihh IL M-CSF MSC

Bone morphogenetic protein Fibroblast growth factor Granulocyte-macrophage colony-stimulating factor Insulin-like growth factor Indian Hedgehog Interleukin Macrophage colony-stimulating factor Mesenchymal stem cell

G. Vunjak-Novakovic (*) Department of Biomedical Engineering, Columbia University, New York, NY, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_12, © Springer Science+Business Media, LLC 2011

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Platelet-derived growth factor Parathyroid hormone-related peptide Transforming growth factor-b Temporomandibular joint Vascular endothelial growth factor

12.1 Introduction Engineering of functional bone offers a promising alternative method for treating different types of diseases and complications as well as for esthetic purposes. In order to engineer a successful graft, engineers, scientists, and clinicians need to consider the specifics of the defect being repaired, the structural and functional requirements at the time of implantation, and the approach to providing integration with the host tissue, as well as supporting immediate survival and long-term function. These general requirements translate into the specific set of design considerations that are typically addressed in a highly interdisciplinary manner, by taking into account the biology of tissue development, repair and (if applicable) disease conditions, the implantation route, and the necessary properties of engineered tissue grafts. This chapter is focused on the biological, engineering and clinical aspects of bone tissue engineering, in the context of developing new modalities for tissue repair. We look into the complications that require the clinical use of bone grafts and the current treatment modalities. The biology of bone is briefly discussed to provide the background for selecting the optimal conditions for the cultivation of functional tissue grafts. Lastly, a few representative innovative systems for engineering and validation of bone grafts are discussed, in the context of the state of the art and remaining challenges that need to be addressed before tissue-engineered repair of bone becomes a clinical reality.

12.2 Bone Biology and Structure Bone is a hard connective tissue that provides mechanical and metabolic functions vital to survival and health, such as the support of the body’s framework, supply of blood cells for the entire body, and maintenance of mineral and fat reserves. Different types of bone are very different with respect to their geometry, cellularity, mechanical properties, and developmental pathways. Flat bones and the outer regions of long bones are comprised of compact (cortical) bone that contains ~80–90% mineralized tissue providing mechanical strength. The ends of long bones are made up primarily of trabecular (cancellous) bone, which contain ~15–25% mineralized tissue. The mechanical properties of bone generally depend on its structure and orientation. Compared to cancellous bone, cortical bone has much higher compressive stiffness

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(12–20 GPa versus 0.2–0.8 GPa) and strength (100–230 MPa versus 2–12 MPa) [1]. A successful bone graft should ideally match the physiological properties of the type of bone being replaced. Therefore, understanding bone biology allows tissue engineers to acquire basic scientific knowledge and apply it toward engineering functional bone grafts.

12.2.1 Bone Development Bone formation begins within the first month of development through two different developmental processes: intramembranous ossification and osteochondral ossification. Intramembranous ossification occurs when fibrous membranes are replaced by bone tissue to form flat bone such as cranium (skull), illium (pelvis), and rib cage. It involves direct differentiation of mesenchymal cells into preosteoblasts and osteoblasts [2]. In contrast, endochondral ossification occurs when cartilage is replaced by bone tissue to form long bones. The two processes result in distinctly different compositions and structures of the bone matrix [3], although recent studies have identified several shared molecular regulators [4]. In both processes, the major components for bone development are the formation and infiltration of vasculature and differentiation of stem cells into bone-forming cells. The key angiogenic regulators include members of the FGF, IGF, TGF-b, and VEGF families [5, 6]. In addition, the factors that play essential roles in bone development include growth hormone such as Ihh and PTHrP [7, 8], FGF-2 [9, 10], and TFG-b family members [11–13], especially BMPs [14, 15].

12.2.2 Bone Remodeling Bone remodeling is a process that occurs throughout a person’s life. It serves the needs of bone regeneration, maintenance, and homeostasis [16]. In the process, existing bone is resorbed by osteoclasts and new bone tissue is formed by osteoblasts. Remodeling is in fact the bone response to signals associated with bone growth, microdamage, and mechanical loading. Signaling pathways include the action of several hormones, such as PTHrP, vitamin D, and cytokines such as BMPs, FGF, IGF, TGF-b, PDGF (in bone formation) and GM-CSF, ILs, and M-CSF (in bone resorption) [16–18].

12.2.3 Bone Healing Healing of bone, unlike soft tissue, does not lead to scar formation and, if the defect is smaller than a critical size, results in the reestablishment of native bone anatomy and function. The healing is usually complete by 6–8 weeks after the initial injury.

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Fracture repair is characterized by the inflammatory phase, reparative phase (which includes intramembranous ossification), chondrogenesis with endochondral ossification, and remodeling [19, 20]. Examples of growth factors and extracellular matrix proteins involved in bone repair processes include TGF-b, FGF (I and II), PDGF, BMP (2, 4, and 7), osteonectin, osteocalcin, and collagens [21].

12.3 Bone Medical Complications Due to the multiple functions of bone in locomotion, mechanical support, and physical protection of various organs, bone injury can cause significant pain, discomfort, and physical disability. The injured bone can be replaced with a graft, which is a routine option for conditions following tumor resections and large fractures. Bone fracture is commonly caused by accident, high force impact or stress, or trivial injuries resulting from osteoporosis. Some fractures can lead to serious complications such as nonunion fracture where the fractured bone fails to heal, malunion where the fractured bone heals in a deformed manner, and compartment syndrome that may require amputation of the affected limb. Most fractures require an immediate treatment by securing the fracture in place to allow bone to undergo self-healing. Surgery is needed when conventional methods fail. Bone grafts can enhance the healing process by filling in the gap of a nonunion and facilitating bone integration.

12.4 Current Solutions to Bone Grafting Bone grafting requires a surgical procedure to replace defective or missing bone in situations in which the natural bone repair may be too slow or inadequate. After implantation, the graft needs to incorporate biologically and functionally and provide clinically functional load bearing. Three common types of bone grafts are autografts, allografts, and graft substitutes, each having advantages and disadvantages. Therefore, the treatment of choice depends on the patient’s general condition and the specific fracture or symptom.

12.4.1 Autografts Autografts, which are grafts taken from the patient’s own bone, represent a gold standard in bone grafting. The bone graft is harvested from nonessential bones such as illiac crest and shaped into the shape and size needed; thus, this method requires an additional surgical site and may pose additional postoperative pain and complications. Furthermore, the size and total amount that can be harvested are limited. Autografts are often preferred because they alleviate the risks of graft rejection and

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disease transmission. A special type of structural graft is the vascularized fibular graft, which requires a microsurgical approach to connect the nutrient fibular vessel to a vascular bundle adjacent to the site of the defect. The operation causes high morbidity at the donor site, and its usage is generally limited to reconstruction following tumor resection [22, 23].

12.4.2 Allografts Allografts are bone grafts harvested from human donors. Bone allografts include fresh or fresh-frozen bone, freeze-dried bone grafts, and demineralized freeze-dried bone grafts. Compared to autografts, allografts offer a better supply of bone in suitable shapes and sizes, and avoid donor site morbidity. However, allografts pose a risk of immune rejection and infection since the tissue is harvested from another individual and contains foreign cellular material. These risks are reduced when the cellular materials are destroyed during graft processing and storage, such as in demineralized bone allografts.

12.4.3 Graft Substitutes Graft substitutes include natural or synthetic materials, which can be organic, inorganic, or combination products. Graft substitutes are available in unlimited supply, and can be fabricated in any desired size and shape. The materials used as graft substitutes include collagen sponge, calcium phosphates (e.g., hydroxyapatite and tricalcium phosphate), bioactive glass, and metals such as titanium and its alloys. These materials provide osteoconductivity for bone healing formation and some of them are resorbed with time in vivo. Graft substitutes provide less risk of infection and graft immune rejection in comparison to allografts. However, stress shielding, which may cause bone atrophy, is a concern when applying this type of graft. Although graft substitutes offer great benefits, they are still inferior to autografts and allografts in terms of biological responses that enhance bone healing.

12.5 Key Components for Engineering of Functional Bone Grafts Ideal characteristics of bone grafts are high osteoinductive and angiogenic potential, biological safety, low patient morbidity, no size restrictions, ready access for surgeons, long shelf life, and reasonable cost [22, 24]. Successfully engineered functional bone grafts provide most or all of these characteristics. The process of engineering a functional bone graft tailored to the patient and specific defect is

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Fig. 12.1  Engineering of bone grafts. The process begins with 3D imaging of the defect to guide manufacturing an anatomical-shaped scaffold for the formation of bone. The scaffold is seeded with cells and cultured in a bioreactor (also manufactured with the aid of imaging) that provides environmental control and physical stimulation to create a functional bone graft

s­ ummarized in Fig. 12.1. In brief, the bone defect is scanned in order to obtain a 3D image of the defect. Using a computer program, the contour, size, and shape of the graft are designed, and the scaffold is fabricated. Cells are then seeded into the scaffold and cultured in a specially designed “anatomical” bioreactor to support the development of engineered bone grafts. The graft is then implanted into the patient and allowed to integrate with the native tissue. The process requires consideration of three major components: cells, scaffolds, and environmental factors provided by the bioreactor.

12.5.1 Cells Living cells are the drivers of bone development, remodeling, and healing. Thus, the cells used for bone tissue engineering application must be able to form bone, be compatible with the patient, and be available in sufficient amounts. Biology of bone has shown that bone cells originate from stem cells. Particularly, the stem cells that

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currently exhibit the most potential in engineering of bone graft are MSCs derived from bone marrow or adipose tissue. The ease of isolation [25, 26] and the ability to proliferate and differentiate in  vitro into osteoblasts [27–29] make these cells a strong candidate for bone tissue engineering applications.

12.5.2 Scaffolds Scaffolds need to resemble the extracellular matrix of bone tissue, and thereby provide infrastructure for the cells to reside, proliferate, differentiate, and assemble mechanically functional bone [30]. The scaffolds for bone tissue engineering should be biocompatible, degradable into nontoxic product, osteoconductive (to recruit bone cells from the recipient), osteoinductive (to differentiate stem cells into bone-forming cells), osteointegrative (to provide permanent and functional attachment to native bone), and exhibit mechanical properties similar to those of native bone. Chemists, material scientists, and engineers have extensively investigated scaffolds made of various materials including natural substances, protein- or organic-based polymers, ceramics, metals, as well as their combinations for bone tissue engineering. Nevertheless, there is no single perfect material. Different types of materials have advantages and disadvantages; for example, organic-based polymers such as silk and poly-lactic-glycolic acid, although much less stiff than native bone, are easily fabricated into desired size, shape, and porosity [31, 32]. On the other hand, ceramics such as hydroxyapatite have mechanical property similar to native bone but have limited flexibility in fabrication [33, 34]. Optimization of materials for bone tissue engineering is still in progress, with major advances over the past decade [35–37].

12.5.3 Environmental Factors Environmental factors include nutrients for maintaining cell viability, and biochemical and biophysical regulatory signals for bone tissue formation. In order to engineer a functional bone graft, the cells inside the scaffold must be viable and healthy pre- and postimplantation. To maintain cell viability and achieve homogenous tissue development inside large constructs, specially designed bioreactors were employed, including spinner flasks, rotating wall vessels, and perfusion bioreactors [38–40]. Mechanical signaling is also essential for bone formation. Two types of mechanical cues that affect bone formation physiologically are mechanical compression and shear stress [41]. In addition to mechanical cues, biological and chemical factors such as dexamethasone [42] and BMP-2 [43, 44] have been shown to play a significant role in stimulating MSC differentiation and osteogenic gene expression, as well as enhancing the cells to produce bone proteins and mineral. Other molecular cues regulating bone development and healing mentioned in the previous section can also enhance the development of engineered bone.

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12.6 Model Systems for Engineering Functional Bone Grafts The most promising cell source for engineering functional bone graft appears to be MSCs, which have been used as a standard cell type in many different studies, in conjunction with culture systems designed to induce and support bone formation in vitro.

12.6.1 In Vitro Culture Systems Several cultivation systems have been employed for bone tissue engineering. Perfusion bioreactors have exhibited the most promising results in terms of controllability and bone tissue formation. A typical bone bioreactor with medium perfusion consists of a medium reservoir, a culture chamber, and a perfusion loop with a pump and gas exchanger. The pump draws media from the reservoir through the scaffold, which resides in the culture chamber. Medium flow serves two purposes: provision of adequate nutrient supply to all the cells inside the porous construct, and mechanical stimulation through fluid shear load. Perfusion bioreactors have been shown to be superior to static culture [45] and spinner flasks [40]. In addition, the effect of fluid shear stress results in progressive deposition of mineralized matrix throughout the 3D engineered tissue constructs [46]. Perfusion bioreactors have showed an enhancement of bone-like tissue development in terms of production of bone matrix, i.e., collagen type I, osteocalcin, osteopontin, and bone sialoprotein with hMSC-seeded decellularized bovine trabecular bone [38]. Cellular content as high as that in native bone was achieved under optimal conditions in perfusion bioreactors [47].

12.6.2 In Vivo Models The utilization of animal models is an essential step in the testing of orthopedic implants. In order to determine the effectiveness and safety of bone implants, several animal models have been studied, from small species such as mouse, rat, and rabbit to large species such as dog, pig, and sheep. Animal models allow researchers to analyze graft biocompatibility, mechanical stability, and safety as well as osteoconductivity, osteoinductivity, osteointegrativity, and resorbability of the grafted cells and materials. However, there are differences between human and animal bone, and an appropriate animal model should be carefully selected [48]. One of the most common and most relevant bone defect models to test effectiveness of tissue-engineered bone grafts is the critical size defect model, which represents a bone lesion that would become scarred and lead to nonunion. The critical defect size can vary between species and locations; for example, critical size defects used in rat are 6 mm for long bones and 4 mm for calvaria while critical size defects used in dogs are 20 mm for calvaria and 45 mm for mandible [49].

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The effectiveness of tissue-engineered bone has been demonstrated in various critical size bone defect models [50, 51]. The result verified that tissue-engineered bone is superior to implanted hMSC-seeded scaffold and scaffold alone, in terms of bone formation and integration to the native tissue. Progress in engineering of functional bone graft has been evolving rapidly and animal models are an important component for translation into clinical application.

12.7 Requirements for Clinical Translation Translation of bone graft technology from the bench to the clinic requires combinatorial knowledge and skills from various professions including orthopedic surgeons, immunologists, biologists, and tissue engineers. Safety and efficacy are the major considerations when implanting grafts into humans. During implantation, the tissueengineered bone graft should allow for mechanically secure and stable fixation to the host tissue [52]. The graft must survive biologically and mechanically, and generate bone integration and formation after implantation, as well as provide the necessary function. Requirements for clinical translation of functional bone graft include (a) selection of appropriate cell source and cell function, (b) fabrication of anatomical shaped graft, and (c) graft preparation time.

12.7.1 Cell Sources and Cell Functions MSCs from bone marrow and adipose tissues have been the most characterized cell sources for the engineering of bone grafts. For clinical application, MSCs can be isolated from the patient to completely avoid rejection [53]. However, the quantity and quality of MSCs vary among individuals and decline with age [54]. Allogeneic MSCs have been shown to prolong allograft tissue survival and reverse severe graftversus-host disease [55–57]. Other potential cell sources for clinical use include MSCs from amniotic fluid, placenta, umbilical cord, skin, and thymus, as well as embryonic and induced pluripotent stem cells. However, more work needs to be done to characterize, proliferate, and differentiate cells to at least match the level of characterization of MSCs. Other important considerations include cell–host interactions (such as the inflammatory response to implanted cells), the ability of cells to home to the site of injury, and local regulation of differentiation of implanted cells [58].

12.7.2 Anatomical Shape Autografts, allografts, and graft substitutes all require shaping in order to fit into a specific defect, and meet functional and esthetic requirements. Tissue engineering offers a possibility of using scaffolds with precisely defined size and shape, to predefine

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the exact anatomical shape of the forming bone. For example, imaging techniques such as computed tomography can be used to reconstruct the 3D geometry of the defect and use these data to fabricate a scaffold exactly matching the defect. Maintaining cellularity in large anatomically shaped bone grafts generally requires perfusion (interstitial flow) of medium throughout the scaffold. For example, an engineered human-sized temporomandibular joint (TMJ) condyle was fabricated using a custom-designed scaffold and a custom-designed bioreactor [47]. By employing computer software to analyze fluid flow patterns, the medium perfusion was optimized to ensure nutrient transport within the forming tissue. As a result, engineered TMJ grafts contained high density of homogeneously distributed cells and a dense bone matrix. This study demonstrated the feasibility of engineered anatomical-shaped functional bone grafts for implantation.

12.7.3 Preparation Time “When would the bone graft be needed?” is the fundamental question that is typically answered on a case-by-case basis. A patient injured by an accident may need an immediate bone graft while a bone tumor patient could wait for weeks or months. Also, functional bone grafts can be used as a preventative measure in treating bone defects. Engineered bone using hMSCs typically requires a period of 5 weeks of cultivation to obtain compact and functional tissue [38, 40, 47, 59, 60]. In these grafts, cell growth is more or less completed within the first 2 weeks of culture. The optimal time point to implant engineered bone grafts still needs to be determined. When a bone graft is needed immediately, an off-the-shelf engineered functional graft would be most desirable. Since grafts are large and contain living cells, their preservation is a complex task. Cryopreservation may induce heat shock proteins or activate immune responses [61] and thereby reduce the utility of engineered bone grafts. The most effective method for preserving and maintaining bone grafts has yet to be devised.

12.8 Challenges and Current Needs 12.8.1 Integration into Host Tissues Several animal trials have shown promising results for implantation of tissueengineered bone [40, 59]. Human clinical trials addressing implantation of engineered bone graft have not been carried out in a systematic manner [62, 63]. Although there is evidence for the safety of these grafts and their integration into the native bone, optimal techniques to enhance the osteogenic, osteoinductive, osteoconductive, and osteointegrative properties of bone grafts are not yet available.

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12.8.2 Vascularization Following implantation, cell survival is the key issue. Incorporation of rapid revascularization strategies is essential, particularly for large-volume defects [64, 65]. Vascularization of an implanted graft can be accelerated by local delivery of angiogenic growth factors such as VEGF, PDGF, and FGF [66]. Growth factors can be incorporated by using two methods: (a) binding to the scaffolds or (b) encapsulation in controlled-release microspheres [67]. Alternatively, grafts can be prevascularized by incorporation of endothelial cells, shown to form capillary-like networks [68–70] with anastomosis to the host’s vasculature [71, 72]. MSCs were also shown to support the formation of vascular networks [73–75].

12.8.3 Long-Term Mechanical Function Providing physiological biomechanical function is the major goal of orthopedic tissue engineering [30]. Engineered bone grafts would ideally provide physical and biological signals to stimulate natural remodeling mechanisms, leading to complete integration of bone grafts and eventual replacement of engineered grafts with native bone tissue. Tissue-engineered bone grafts should ideally have mechanical properties similar to native bone immediately at the time of implantation, especially when constructs are implanted in load-bearing sites. Alternatively, scaffolds with inferior mechanical properties could potentially be used if they allow for fast and strong bone formation. A good criterion of engineering functional bone grafts is to match the degradation rate of the constructs with the cellular rate of bone formation so that healing at the site functions mechanically throughout the healing process [76].

12.9 Summary As autologous grafts are limited by their applicable size as well as concerns regarding injury to the harvest site, and suitable allogeneic tissue grafts are not in sufficient supply to meet clinical demand, the field of bone tissue engineering holds significant potential for providing clinically relevant tissue grafts for restoring joint function. These grafts can be personalized using techniques based on noninvasive patient-specific images, and provide grafts in a variety of shapes and sizes as methods to provide cell access to critical nutrients in culture are further optimized. In this interdisciplinary effort, stem cells will play an increasingly growing role as further insights into the role that physicochemical stimuli play in modulating stem cell differentiation and subsequent biosynthetic functions are realized. As our ability to grow mechanically functional bone grafts has progressed, more recent efforts are now focused on recapitulation of the interface region that spans the soft hydrated tissue and mineralized bone of osteochondral grafts [77].

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Acknowledgment  The authors gratefully acknowledge research support of the work described in this chapter (NIH grants DE016525, EB002520 and EB011869 and NYSCF grant CU09-3055).

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Chapter 13

Engineering Functional Cartilage Grafts Andrea R. Tan and Clark T. Hung

Abstract  Articular cartilage is the specialized connective tissue that covers diarthrodial joints (e.g., hip, knee, and shoulder) and serves a load-bearing and lubrication function. As the tissue is avascular, it exhibits a poor healing capacity when injured. Joint arthroplasty, comprised of metal and plastic prostheses, has a limited lifespan after implantation and are ideally reserved for cases of significant traumatic injury and pervasive arthritis. As such, there have been significant efforts to develop cell-based strategies for cartilage repair. Accordingly, there is great anticipation regarding the role that stem cells can serve as a cell source for generating functional articular cartilage grafts. There is a need for both the use of animal cells and models as well as parallel development using human cells to successfully translate these strategies to the clinic.

Abbreviations BMSC CAD CZ DZ ECM FGF GAG hESC IGF

Bone marrow-derived stem cell Computer-aided design Calcified cartilage zone Deep zone Extracellular matrix Fibroblast growth factor Glycosaminoglycan Human embryonic stem cell Insulin-like growth factor

C.T. Hung (*) Department of Biomedical Engineering, Columbia University, New York, NY, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_13, © Springer Science+Business Media, LLC 2011

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Magnetic resonance imaging Middle zone Osteoarthritis Proteoglycan Superficial zone Transforming growth factor

13.1 Introduction Cell-based strategies involving the design and use of a clinically relevant scaffold to encapsulate cells in a controlled three-dimensional environment offer a promising alternative for treating cartilage damage due to disease or degeneration. Engineering a functional graft capable of withstanding the complex loading schemes of the knee joint immediately following implantation and for long-term success requires an understanding of the biological development of native cartilage, the structural and functional requirements of the tissue at the time of implantation, and the challenges of host–graft integration. Here, we look first at the intricate structure of the tissue, how its natural design allows for function, and how changes to that intact form during injury or disease results in mechanical failure. As a model system, we explore the use of agarose hydrogels encapsulating juvenile bovine chondrocytes to replace damaged cartilage tissue for focal defects, and how both mechanical and chemical stimulation during in vitro growth may prepare the fledgling-engineered constructs to better handle the native environment after implantation. Lastly, we look at the potential for clinical translation and necessary alterations to the model system including cell source and anatomical fit, as well as some of the current challenges that must be overcome before widespread human clinical use.

13.2 Cartilage Biology and Structure Articular cartilage, a white, dense connective tissue, serves as the load-bearing material of joints and is characterized by excellent friction, lubrication, and wear properties [1]. Ranging from 1 to 7 mm thick, cartilage is composed primarily of two phases, a solid matrix (collagen fibrils and PG macromolecules) [2–4] and a mobile interstitial fluid phase (mostly water) [5, 6]. The polyanionic nature of the GAG chains of PG draw water into the tissue [6], resulting in a large osmotic pressure that expands against the constraining collagen network. The interplay between swelling pressure and tension in the collagen fibers results in a highly specialized tissue that is well suited to bear compressive loads within the joint. To maintain the necessary matrix composition, chondrocytes, which make up less than 10% of the

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tissue volume [7], balance extracellular degradation and matrix turnover by synthesizing and secreting ECM. The composition, structure, and material properties of articular cartilage are known to vary across the tissue’s depth [8] and can be divided into four discrete zones: the superficial, middle, deep, and calcified cartilage zones [9, 10]. The SZ is a thin region at the articulating surface marked by collagen fibers that are aligned parallel to the surface and it possesses the highest concentrations of water and collagen compared to the other zones, but has the lowest level of PG. Additionally, chondrocytes in this zone are flattened in morphology. In contrast, the MZ is rich in collagen whose orientation is random, but contains the highest concentration of PG with randomly dispersed cellular bodies. In the DZ, the chondrocytes are oriented perpendicular to the surface and arranged in a columnar structure with the lowest levels of collagen. Finally, the CZ separates the uncalcified layers of cartilage from the subchondral bone.

13.3 Cartilage Injury and Repair In most cases, damage to articular cartilage is a consequence of clinical OA and is marked by disability and pain [11]. Affecting 9% of the US population aged 30 and older, OA has total direct costs estimated at $28.6 billion dollars annually [12]. Due to its avascular nature, articular cartilage exhibits poor intrinsic healing response [13]. The hallmark characteristics of this debilitating disease include a loss of mechanical properties, increased collagen degradation, reduced PG synthesis, and decreased cellularity [14, 15]. Alternatively, damage to cartilage can be caused by physical injury to the articular surface. While physical injury to cartilage primarily occurs with traumatic loading of the joint (traumatic injury), it can also be a consequence of surgical procedures that include graft harvesting (iatrogenic injury). In native cartilage, the mechanotransduction resulting from injury can induce chondrocyte death as early as a few hours, and up to 7  days, postinjury [16]. The subsequent downstream effects are frequently marked by changes similar to those seen in OA.

13.4 Current Clinical Strategies for Cartilage Repair When damage to the tissue in the knee joint, for example, is widespread, a total knee replacement is often the solution to artificially replace the articulating surface and underlying bone. Such repairs, however, often require revision surgeries due to wear, subsidence, and loosening of the implant in the bony union [17–20]. For the repair of focal lesions and damage to the articular surface, more conservative approaches may be used. Clinical options include tissue adhesives [21, 22], enzymatic treatments [23], laser solder welding [24], autograft cell/tissue transfer via periosteal grafts [25], autologous osteochondral grafting such as mosaicplasty [26], and the Carticel method [27]. While these options offer temporary relief from

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Fig. 13.1  Schematic showing strategy for the development of cell-based therapies for cartilage repair using a large preclinical animal model (canine)

symptoms, they also introduce long-term problems. Primarily, the availability of healthy cartilage from which to harvest cells is limited for cell-based therapies or osteochondral graft harvesting. Furthermore, autologous osteochondral grafts are usually harvested from nonload-bearing regions which may provide tissue of suboptimal material properties for use in contact regions [28]. The harvest procedure itself can induce significant cell death in the surrounding region [29, 30], leading to further structural and biochemical breakdown of the donor site tissue. Alternatively, there have been recent attempts at developing cell-based therapies for cartilage repair, including tissue-engineered constructs of cultured cells on three-dimensional scaffolds [31–34], such as depicted in Fig. 13.1.

13.5 Model Systems for Cartilage Tissue Engineering 13.5.1 Agarose Hydrogel System Hydrogels have become the scaffold material of choice for cartilage tissue engineering due to their intrinsic hydrophilic nature and high water content, similar to native soft hydrated tissues [35]. Hydrogels being explored include

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polyethylene(glycol) [36], hyaluronic acid [37], silk [38], alginate, and collagen [39]. Agarose, a neutrally charged polysaccharide from seaweed, has been used extensively in cartilage biology for maintaining long-term chondrocyte suspension cultures due to its ability to promote and maintain the chondrocyte pheno-type [33, 40–45], as well as being nondegradable and noninteractive with chondrocytes. These advantages have prompted the use of agarose for cartilage tissue engineering applications, where agarose-embedded chondrocytes have been shown to successfully repair articular cartilage and tibial defects in in vivo models [46, 47]. The properties of agarose permit application of physiologic deformational loading immediately upon encapsulation such that constructs may be physically stimulated before extensive ECM development is present. Together, these characteristics of the agarose hydrogel system have allowed for the fabrication of the most reproducible and robust cartilage tissue growth in culture [48]. Furthermore, this system serves as an important tool to study tissue engineering strategies. Clinically, agarose is being used as a co-polymer with alginate as a hydrogel scaffold for autologous chondrocyte implantation (Cartipatch) for cartilage defect repair and has demonstrated good 2-year clinical follow-up [49, 50].

13.5.2 Mechanical Preconditioning Joint loading gives rise to a plethora of physical stimuli, including osmotic loading, hydrostatic pressure, electrokinetic phenomena, stress, and strain [51]. For functional tissue engineering, physiologically relevant stimuli are applied to encourage the development in vitro for tissues that can meet the in vivo functional demands, as well as mimic the composition of native tissue [52]. The choice of scaffold material and its inherent material characteristics dictate the nature of loading that can be applied on developing cartilage tissue constructs. Rotating wall bioreactors have been used to provide a hydrodynamic, low-shear environment supportive of enhanced nutrient transport [53, 54] and cartilage-like tissue growth [31, 55, 56]. However, these bioreactors do not reproduce the physiologic deformational loading and hydrostatic pressure environment of the chondrocyte [57]. In comparison, applying physiologic ­loading through a combination of applied physiologic hydrostatic pressure and perfusion [58] or through physiologic dynamic deformation loading [33, 45, 59] can achieve near-physiologic values for equilibrium modulus and GAG content. Applied deformational loading gives rise to enhanced convection of nutrients [60, 61] in a mechanism similar to how joint loading provides nutrients from the synovial fluid to avascular cartilage in situ. There is a growing body of literature suggesting that physical forces can be used to modulate chondrogenesis of mesenchymal stem cells [62–65], as reviewed by Huang et al. [66].

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13.5.3 Chemical Stimulation In addition to mechanical loading to promote tissue growth, many groups have focused on the application of a range of chemical cues such as growth factors TGF-b3, TGF-b1, IGF, FGF2 [67–70], corticosteroids [71, 72], and interleukins [73–75]. Through paracrine signaling and direct cell–cell contact, the exchange of these chemical factors has been found to promote ECM development. Physical loading may act to increase convective transport of growth factors, leading to ­possible synergistic interactions [68].

13.6 Requirements for Clinical Translation 13.6.1 Cell Sources Many studies have utilized isolated juvenile bovine chondrocytes encapsulated in hydrogel systems to recreate and even surpass the properties of native bovine tissue by the temporal application of chemical [67] or physical factors [76], or a combination of the two [48]. These studies capitalize on the significantly greater biosynthetic capacity of juvenile cells relative to their adult counterparts [77], also reported for human chondrocytes [78]. However, a juvenile cell source is limited for clinical applications due to the obvious challenges related to tissue procurement. As such, differentiated adult chondrocytes may be obtained from a patient’s own healthy, nonload-bearing cartilage, although this may lead to donor site morbidity and further tissue degeneration [79]. Chondrocytes from the diseased knee may be also harvested during preliminary debridement procedures. In both cases, though, obtaining sufficient cell number to produce constructs with sufficient mechanical properties is challenging due to the reduced biosynthetic activity of cells from patients with advanced stages of the disease [80, 81]. Alternatively, due to the immunoprivileged nature of diarthrodial joints, the implantation of allogeneic cells from other patients may be expanded for use [82], following from the clinical use of living osteochondral allografts [83]. Finally, undifferentiated cartilage precursor cells, including those isolated from patient bone marrow aspirate [84], from adipose tissue [85–88], or from the synovium [89–91] have been explored as alternative sources of cells, however, to date, constructs made with these human cells have not been capable of attaining mechanical properties at native tissue levels. Co-culture of hESCs with chondrocytes and fibrochondrocytes can modulate chondrogenic expression and matrix synthesis [92, 93]. Stem cells may have different cell density or cell–cell contact requirements than chondrocytes [94] in order to form cartilaginous tissue. Micromass, self-assembled constructs (scaffold-free), and suspension in hydrogel show potential for hESC chondrogenesis [92, 93].

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Chondrogenesis of hESCs that have been differentiated with growth factors in embyroid bodies and then dissociated and cultured in self-assembled (scaffoldless) constructs can be modulated by growth factor combinations as well as hypoxia [95, 96]. As with mesenchymal cells, these studies have not yet been ­successful in generating functional cartilage. More recently, induced pluripotent stem cells originating from somatic cells such as fibroblasts [97] or chondrocytes have been investigated as an alternative source for cartilage tissue engineering, as it is likely that these cells possess some degree of epigenetic memory of their previous phenotype [98] and can undergo directed differentiation into cartilage.

13.6.2 Cell Expansion The number of cells required to recreate engineered cartilage with functional properties surpasses the number that is readily available clinically. To address the limited supply of cells, one way to increase the cell number is by expanding cells through passaging under appropriate growth factors. Similar to the cocktail of growth factors used for the culture of chondrocytes encapsulated in an agarose hydrogel, these same growth factors encourage and maintain the chondrocyte phenotype. For adult chondrocytes that are less biosynthetically active, expansion in a cocktail of appropriate growth factors has been shown to rapidly expand cell number, prime the cells to reactivate rapid matrix synthesis when cultured in a three-dimensional scaffold environment, and prevent phenotype dedifferentiation which is typically seen in monolayer culture [41, 99, 100]. For undifferentiated stem cells, the application of a growth factor cocktail offers the ability to direct the differentiation of these plastic cells down the chondrocyte lineage [101–103].

13.6.3 Long-Term Storage of Grafts The use of fresh osteochondral allografts is limited by their short shelf life; cold ­storage (~4°C) is the current standard for osteochondral graft preservation and ­storage [104]. However, concerns over the decrease in cellular viability with storage time have limited their standard clinical use to within 28 days post harvest. Optimized chemically defined serum-free culture medium has been shown to preserve the mechanical and biochemical properties of allograft tissue for up to 56 days in culture [105], This culture medium, which was adapted from a well-established formulation known to foster chondrogenesis in BMSCs, also has been shown to promote de novo tissue matrix formation in tissue-engineered cartilage [106] without the associated decrease in cellular viability seen with allografts. With this medium formulation, tissue-engineered osteochondral tissues may prove to be more clinically favorable option as engineered constructs are capable of storage over a longer time period, thereby increasing the window of time for their implantation.

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13.7 Challenges and Current Needs 13.7.1 Anatomical Shape For an engineered construct to functionally bear the loads experienced in  vivo, recreating the natural topology of the articular surface is necessary to recapitulate the normal contact geometry and load distribution across the joint [107–113]. MRI has been used to construct stereolithographic models of the TMJ for anatomic studies [114], and to recreate iatrogenic defects on the articular surface of the femur [115] using CAD and rapid prototyping techniques. Furthermore, using MRI, chondral and osteochondral grafts with the anatomic form of the human patella have been fabricated [116, 117].

13.7.2 Size and Scale-up Tissue-engineered constructs have successfully been created to recapitulate the properties of native cartilage in Young’s modulus and GAG content [48] and implanted in  vivo in focal lesions in a canine model [118], however, these have limited utility due to their size. Increasing the size of these constructs, however, presents problems as the nutrient path length increases and the diffusion of soluble factors into the tissue is hindered. Efforts to increase uptake of soluble factors by dynamic loading [119] have only produced constructs ~20% stiffer than controls [48]. Alternatively, channels or holes placed within the construct to decrease the nutrient path length have shown to help facilitate the uptake of soluble factors, resulting in more homogenous tissue-engineered cartilage constructs with improved mechanical properties [120]. These strategies will be critical for cultivation of large anatomically shaped grafts.

13.7.3 Integration with Surrounding Tissue Integration of a functional cartilage construct with the surrounding tissue is a challenge arising from the natural surface topology of articular cartilage. Incongruities at the graft edge can cause stress concentrations upon articulation, and possibly lead to graft loosening [28]. Additionally, with inadequate structural bonding of the graft with the host tissue in the defect, there is poor cartilage–cartilage tissue integration [121–123]. To address these challenges, research groups have looked at the use of enzymatic treatments and chemical alterations coupled with in situ polymerization for increased chondral graft adherence [124, 125]. Alternatively, tissue-engineered osteochondral constructs [34, 126–130] consisting of a chondral region, chondral–bony interface, and bony substrate replacement may integrate with surrounding tissue better due to the neighboring vascularized bony

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region. The potential to replace entire articular surfaces with anatomic engineered cartilage, as described above, will focus graft integration to the underlying bone rather than to the surrounding cartilage.

13.8 Summary The field of cartilage tissue engineering holds significant promise for providing clinically relevant tissue grafts for repairing damaged cartilage due to injury or disease pathologies. We have shown here that the potential for functional replacement of native cartilage is subject to specific considerations including an appropriate cell source, necessary electro-physicochemical cues, as well as applicable size and fit into the surrounding tissue. This multi-disciplinary effort to optimize tissue development for future human clinical trials has increasingly begun to incorporate stem cells as a primary cell source as further insights into their directed differentiation pathways are realized. However, as evaluation of cartilage repair strategies in human patients is not generally practical, relevant large preclinical models such as the canine (Fig. 13.1), ovine, and caprine models, are needed to provide essential parallel systems to fully assess the outcome of these engineered grafts in vivo. Acknowledgment  The authors gratefully acknowledge research support of the work described in this chapter (NIH grants AR46568, AR52871, and AR060361).

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36. Rice MA et al (2008) Effects of directed gel degradation and collagenase digestion on the integration of neocartilage produced by chondrocytes encapsulated in hydrogel carriers. J Tissue Eng Regen Med 2(7):418–429 37. Burdick JA et al (2005) Controlled degradation and mechanical behavior of photopolymerized hyaluronic acid networks. Biomacromolecules 6(1):386–391 38. Chao PH et al (2010) Silk hydrogel for cartilage tissue engineering. J Biomed Mater Res B Appl Biomater 95(1):84–90 39. van Susante JL et al (1995) Culture of chondrocytes in alginate and collagen carrier gels. Acta Orthop Scand 66(6):549–556 40. Buschmann MD et  al (1995) Mechanical compression modulates matrix biosynthesis in chondrocyte/agarose culture. J Cell Sci 108(Pt 4):1497–1508 41. Benya PD, Shaffer JD (1982) Dedifferentiated chondrocytes reexpress the differentiated collagen phenotype when cultured in agarose gels. Cell 30(1):215–224 42. Buschmann MD et  al (1992) Chondrocytes in agarose culture synthesize a mechanically functional extracellular matrix. J Orthop Res 10(6):745–758 43. Lee DA, Bader DL (1995) The development and characterization of an in vitro system to study strain-induced cell deformation in isolated chondrocytes. In Vitro Cell Dev Biol Anim 31(11):828–835 44. Lee DA, Bader DL (1997) Compressive strains at physiological frequencies influence the metabolism of chondrocytes seeded in agarose. J Orthop Res 15(2):181–188 45. Mauck RL et al (2002) Influence of seeding density and dynamic deformational loading on the developing structure/function relationships of chondrocyte-seeded agarose hydrogels. Ann Biomed Eng 30(8):1046–1056 46. Rahfoth B et al (1998) Transplantation of allograft chondrocytes embedded in agarose gel into cartilage defects of rabbits. Osteoarthritis Cartilage 6(1):50–65 47. Cook JL et al (2003) Biocompatibility of three-dimensional chondrocyte grafts in large tibial defects of rabbits. Am J Vet Res 64(1):12–20 48. Lima EG et  al (2007) The beneficial effect of delayed compressive loading on tissueengineered cartilage constructs cultured with TGF-beta3. Osteoarthritis Cartilage 15(9): 1025–1033 49. Selmi TA et  al (2007) Autologous chondrocyte transplantation in combination with an ­alginate-agarose based hydrogel (Cartipatch). Tech Knee Surg 6(4):253–258 50. Selmi TA et  al (2008) Autologous chondrocyte implantation in a novel alginate-agarose hydrogel: outcome at two years. J Bone Joint Surg Br 90(5):597–604 51. Guilak F, Hung CT (2005) Physical regulation of cartilage metabolism. In: Mow VC, Hayes WC (eds) Basic orthopaedic biomechanics. Lippincott-Raven, Philadelphia, pp 179–207 52. Hung CT et al (2004) A paradigm for functional tissue engineering of articular cartilage via applied physiologic deformational loading. Ann Biomed Eng 32(1):35–49 53. Gooch KJ et al (2001) Effects of mixing intensity on tissue-engineered cartilage. Biotechnol Bioeng 72(4):402–407 54. Kaysen JH et al (1999) Select de novo gene and protein expression during renal epithelial cell culture in rotating wall vessels is shear stress dependent. J Membr Biol 168(1):77–89 55. Freed LE et  al (1997) Tissue engineering of cartilage in space. Proc Natl Acad Sci USA 94(25):13885–13890 56. Obradovic B et al (1999) Gas exchange is essential for bioreactor cultivation of tissue engineered cartilage. Biotechnol Bioeng 63(2):197–205 57. Ateshian GA, Hung CT (2003) Functional properties of native articular cartilage. In: Guilak F et al (eds) Functional tissue engineering. Springer-Verlag, New York, pp 46–66 58. Carver SE, Heath CA (1999) Increasing extracellular matrix production in regenerating cartilage with intermittent physiological pressure. Biotechnol Bioeng 62(2):166–174 59. Mauck RL et  al (2003) The role of cell seeding density and nutrient supply for articular cartilage tissue engineering with deformational loading. Osteoarthritis Cartilage 11(12): 879–890

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60. Mauck RL, Hung CT, Ateshian GA (2003) Modeling of neutral solute transport in a dynamically loaded porous permeable gel: implications for articular cartilage biosynthesis and tissue. J Biomech Eng 125(5):602–614 61. Albro MB et  al (2008) Dynamic loading of deformable porous media can induce active solute transport. J Biomech 41(15):3152–3157 62. Elder SH et al (2001) Chondrocyte differentiation is modulated by frequency and duration of cyclic compressive loading. Ann Biomed Eng 29(6):476–482 63. Huang CY et al (2004) Effects of cyclic compressive loading on chondrogenesis of rabbit bone-marrow derived mesenchymal stem cells. Stem Cells 22(3):313–323 64. Miyanishi K et al (2006) Effects of hydrostatic pressure and transforming growth factor-beta 3 on adult human mesenchymal stem cell chondrogenesis in vitro. Tissue Eng 12(6):1419–1428 65. Kisiday JD et  al (2009) Dynamic compression stimulates PG synthesis by mesenchymal stem cells in the absence of chondrogenic cytokines. Tissue Eng Part A 15(10): 2817–2824 66. Huang AH, Farrell MJ, Mauck RL (2010) Mechanics and mechanobiology of mesenchymal stem cell-based engineered cartilage. J Biomech 43(1):128–136 67. Byers BA et al (2006) Temporal exposure of TGF-B3 under serum-free conditions enhances biomechanical and biochemical maturation of tissue-engineered cartilage. Trans Orthop Res Soc 31:43 68. Mauck RL et al (2003) Synergistic action of growth factors and dynamic loading for articular cartilage tissue engineering. Tissue Eng 9(4):597–611 69. Mauck RL et al (2001) Transforming growth factor B1 increases the mechanical properties and matrix development of chondrocyte-seeded agarose hydrogels. Adv Bioeng 50:691–692 70. Thorp BH, Anderson I, Jakowlew SB (1992) Transforming growth factor-beta1, -beta2 and -beta3 in cartilage and bone cells during endochondral ossification in the chick. Development 114(4):907–911 71. Bian L et  al (2009) Effects of dexamethasone on the functional properties of cartilage explants during long term culture. Trans Orthop Res Soc 34:329 72. Awad H et al (2003) Effects of transforming growth factor beta1 and dexamethasone on the growth and chondrogenic differentiation of adipose-derived stromal cells. Tissue Eng 9(6):1301–1312 73. Ratcliffe A, Tyler JA, Hardingham TE (1986) Articular cartilage cultured with interleukin 1. Increased release of link protein, hyaluronate-binding region and other PG fragments. Biochem J 238(2):571–580 74. Aydelotte MB et al (1992) Influence of interleukin-1 on the morphology and PG metabolism of cultured bovine articular chondrocytes. Connect Tissue Res 28(1–2):143–159 75. Lima EG et al (2008) Differences in Interleukin-1 Response between Engineered and Native Cartilage. Tissue Eng Part A 14(10):1721–1730 76. Martin I et al (2000) Modulation of the mechanical properties of tissue engineered cartilage. Biorheology 37(1–2):141–147 77. Tran-Khanh N et al (2005) Aged bovine chondrocytes display a diminished capacity to produce a collagen-rich, mechanically functional cartilage extracellular matrix. J Orthop Res 23:1354–1362 78. Adkisson HDt et  al (2010) The potential of human allogeneic juvenile chondrocytes for restoration of articular cartilage. Am J Sports Med 38(7):1324–1333 79. Gilbert JE (1998) Current treatment options for the restoration of articular cartilage. Am J Knee Surg 11(1):42–46 80. Wang J et  al (2003) Homeostasis of the extracellular matrix of normal and osteoarthritic human articular cartilage chondrocytes in vitro. Osteoarthritis Cartilage 11(11):801–809 81. Bulstra SK et al (1989) Metabolic characteristics of in vitro cultured human chondrocytes in relation to the histopathologic grade of osteoarthritis. Clin Orthop Relat Res 242: 294–302 82. Glenn RE Jr et  al (2006) Comparison of fresh osteochondral autografts and allografts: a canine model. Am J Sports Med 34(7):1084–1093

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83. Garrett JC (1998) Osteochondral allografts for reconstruction of articular defects of the knee. AAOS Instr Course Lect 47:517–522 84. Tew SR et al (2008) Cellular methods in cartilage research: primary human chondrocytes in culture and chondrogenesis in human bone marrow stem cells. Methods 45(1):2–9 85. Zuk PA et  al (2001) Multilineage cells from human adipose tissue: implications for cellbased therapies. Tissue Eng 7(2):211–228 86. Zuk PA et al (2002) Human adipose tissue is a source of multipotent stem cells. Mol Biol Cell 13(12):4279–4295 87. Xu Y et  al (2007) In vitro expansion of adipose-derived adult stromal cells in hypoxia enhances early chondrogenesis. Tissue Eng 13(12):2981–2993 88. Xu Y et al (2007) Analysis of the material properties of early chondrogenic differentiated adipose-derived stromal cells (ASC) using an in vitro three-dimensional micromass culture system. Biochem Biophys Res Commun 359(2):311–316 89. Kim JH et  al (2011) Enhanced proliferation and chondrogenic differentiation of human ­synovium-derived stem cells expanded with basic fibroblast growth factor. Tissue Eng 17(7–8):991–1002 90. Arufe MC et al (2010) Chondrogenic potential of subpopulations of cells expressing mesenchymal stem cell markers derived from human synovial membranes. J Cell Biochem 111(4):834–845 91. Li J, Pei M (2010) Optimization of an in  vitro three-dimensional microenvironment to reprogram synovium-derived stem cells for cartilage tissue engineering. Tissue Eng Part A 17(5–6):703–712 92. Hwang NS, Varghese S, Elisseeff J (2008) Derivation of chondrogenically-committed cells from human embryonic cells for cartilage tissue regeneration. PLoS One 3(6):e2498 93. Hoben GM, Willard VP, Athanasiou KA (2009) Fibrochondrogenesis of hESCs: growth factor combinations and cocultures. Stem Cells Dev 18(2):283–292 94. Huang AH et al (2009) Transient exposure to transforming growth factor beta 3 improves the mechanical properties of mesenchymal stem cell-laden cartilage constructs in a densitydependent manner. Tissue Eng Part A 15(11):3461–3472 95. Koay EJ, Athanasiou KA (2008) Hypoxic chondrogenic differentiation of human embryonic stem cells enhances cartilage protein synthesis and biomechanical functionality. Osteoarthritis Cartilage 16(12):1450–1456 96. Koay EJ, Hoben GM, Athanasiou KA (2007) Tissue engineering with chondrogenically differentiated human embryonic stem cells. Stem Cells 25(9):2183–2190 97. Takahashi K, Yamanaka S (2006) Induction of pluripotent stem cells from mouse embryonic and adult fibroblast cultures by defined factors. Cell 126(4):663–676 98. Kim K et  al (2010) Epigenetic memory in induced pluripotent stem cells. Nature 467(7313):285–290 99. Ng KW et al (2010) Passaged adult chondrocytes can form engineered cartilage with functional mechanical properties: a canine model. Tissue Eng Part A 16(3):1041–1051 100. Francioli SE et  al (2007) Growth factors for clinical-scale expansion of human articular chondrocytes: relevance for automated bioreactor systems. Tissue Eng 13(6):1227–1234 101. Hwang NS et al (2006) Enhanced chondrogenic differentiation of murine embryonic stem cells in hydrogels with glucosamine. Biomaterials 27(36):6015–6023 102. Williams CG et  al (2003) In vitro chondrogenesis of bone marrow-derived mesenchymal stem cells in a photopolymerizing hydrogel. Tissue Eng 9(4):679–688 103. Guilak F et  al (2004) Adipose-derived adult stem cells for cartilage tissue engineering. Biorheology 41(3–4):389–399 104. Brighton CT et al (1979) Articular cartilage preservation and storage I. Application of tissue culture techiques to the storage of viable articular cartilage. Arthritis Rheum 1979:1093–1101 105. Bian L et  al (2008) Mechanical and biochemical characterization of cartilage explants in serum-free culture. J Biomech 41(6):1153–1159 106. Mauck RL, Yuan X, Tuan RS (2006) Chondrogenic differentiation and functional maturation of bovine mesenchymal stem cells in long term agarose culture. Osteoarthritis Cartilage 14(2):179–189

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107. Cooney WP, Chao EYS (1977) Biomechanical analysis of static forces in the thumb during hand functions. J Bone Joint Surg 59-A:27–36 108. Ateshian GA et al (1992) A biphasic model for contact in diarthrodial joints. Adv Bioeng ASME BED 22:191–194 109. Ateshian GA, Rosenwasser MP, Mow VC (1992) Curvature characteristics and congruence of the thumb carpometacarpal joint: differences between female and male joints. J Biomech 25(6):591–607 110. Ateshian GA et al (1995) Contact areas in the thumb carpometacarpal joint. J Orthop Res 13(3):450–458 111. Brown TD, Shaw DT (1983) In vitro contact stress distributions in the natural human hip. J Biomech 16(6):373–384 112. Eberhardt AW et  al (1990) An analytical model of joint contact. J Biomech Eng 112(4):407–413 113. Huberti HH, Hayes WC (1984) Patellofemoral contact pressures. The influence of q-angle and tendofemoral contact. J Bone Joint Surg Am 66(5):715–724 114. Undt G et al (2000) MRI-based stereolithographic models of the temporomandibular joint: technical innovation. J Craniomaxillofac Surg 28(5):258–263 115. Koo S et al (2010) Fabrication of custom-shaped grafts for cartilage regeneration. Int J Artif Organs 33(10):731–737 116. Hung CT et al (2003) Anatomically shaped osteochondral constructs for articular cartilage repair. J Biomech 36(12):1853–1864 117. Ateshian GA, Hung CT (2005) Patellofemoral joint biomechanics and tissue engineering. Clin Orthop Relat Res 436:81–90 118. Cook JL et  al (2005) In vitro and in  vivo evaluation of tissue-engineered constructs for articular cartilage regeneration. Trans Orthop Res Soc 30:1767 119. Albro MB et al (2010) Validation of theoretical framework explaining active solute uptake in dynamically loaded porous media. J Biomech 43(12):2267–2273 120. Bian L et al (2009) Influence of decreasing nutrient path length on the development of engineered cartilage. Osteoarthritis Cartilage 17(5):677–685 121. Ahsan T, Sah RL (1999) Biomechanics of integrative cartilage repair. Osteoarthritis Cartilage 7(1):29–40 122. Bobic V (1999) The utilization of osteochondral autografts in the treatment of articular cartilage lesions (part 1 of 3). Int Soc Arthrosc Knee Surg Orthop Sports Med 1–2 123. Hunziker EB (1999) Articular cartilage repair: are the intrinsic biological constraints undermining this process insuperable? Osteoarthritis Cartilage 7:15–28 124. Obradovic B et al (2001) Integration of engineered cartilage. J Orthop Res 19:1089–1097 125. Williams C et al (2002). Musculoskeletal tissue engineering and photopolymerizing hydrogels. In: Tissue engineering. Cold Spring Harbor Laboratory, New York 126. Kreklau B et al (1999) Tissue engineering of biphasic joint cartilage transplants. Biomaterials 20:1743–1749 127. Schaefer D et  al (2000) In vitro generation of osteochondral composites. Biomaterials 21(24):2599–2606 128. Sherwood JK et al (2002) A three-dimensional osteochondral composite scaffold for articular cartilage repair. Biomaterials 23(24):4739–4751 129. van Susante JL et al (1998) Chondrocyte-seeded hydroxyapatite for repair of large articular cartilage defects. A pilot study in the goat. Biomaterials 19(24):2367–2374 130. Lima EG et al (2008) The effect of devitalized trabecular bone on the formation of osteochondral tissue-engineered constructs. Biomaterials 29(32):4292–4299

Chapter 14

Adult Stem Cells and Regeneration of Adipose Tissue Daniel A. Hägg, Bhranti Shah, and Jeremy J. Mao

Abstract  Defects in soft tissue may arise from trauma, chronic diseases, tumor resection, and congenital anomalies. Current practices of autologous grafting and native or synthetic fillers including silicone gel or saline implants have limitations and, like most foreign substances, do not represent long-term solutions. In this chapter, we discuss some of the latest findings in the field of soft tissue regeneration and interventional adipogenesis. We suggest that the field can be advanced by the (1) identification of critical areas of clinical demand and corresponding strategies for different clinical entities; (2) translational approaches that promote angiogenesis and survival of adipose tissue grafts; (3) scale-up of bioengineered adipose tissue grafts; (4) development and adoption of large animal models for testing adipose tissue grafts; and (5) establishment of success criteria for adipose tissue regeneration including the maintenance of volume and shape over time.

Abbreviations ASC BAT bFGF ECM MSC PEG PEGDA

Adipose tissue-derived stem cell Brown adipose tissue Basic fibroblast growth factor Extracellular matrix Mesenchymal stem cell Poly(ethylene glycol) Poly(ethylene)glycol-diacrylate

J.J. Mao (*) Tissue Engineering and Regenerative Medicine Laboratory, Columbia University Medical Center, New York, NY, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_14, © Springer Science+Business Media, LLC 2011

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PGA PLA PLGA RGD WAT

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Poly(glycolic acid) Poly(lactic acid) Poly(l-lactic-co-glycolic acid) Arginine-glycine-aspartic acid White adipose tissue

14.1 Introduction Defects in soft tissue may arise from trauma, chronic diseases, tumor resection, and congenital anomalies. Soft tissue defects may severely hamper a patient’s social interactions and self-esteem, leading to a negative effect on emotional well-being [1]. Reconstructive surgery is often performed to correct soft tissue defects. In 2008, five million facial plastic and reconstructive surgery procedures were performed in the USA alone, with the vast majority (~75%) being corrections of the void following tumor removal [2]. Thus, the demand for reconstructive grafts is substantial. Ideally, reconstructive grafts should be immunocompatible with the patient, able to retain its shape and adapt to biological functions over time. Allografts, xenografts, and synthetic materials, despite their current use in clinical procedures, have limitations because of pathogen transmission, increased risk of infection and immune rejection. Autologous grafts face issues such as donor site morbidity, poor integration with surrounding tissues, and limited donor tissue availability [3–5]. Currently, such defects are predominantly corrected using synthetic filler materials or autologous soft tissue grafts [6, 7]. Autologous fat was reported as a grafting tissue as early as 1893 [8]. Autografts typically consist of vascularized flaps, where a piece of adipose tissue is transferred with its vascular tree intact, and pedicle flap transposition [9], which sometimes includes muscle and fascia tissue [10, 11]. This procedure is associated with both donor and recipient site morbidity. Poor integration between the recipient site blood vessels and the autograft vasculature may result in necrosis and thereby loss of autograft volume over time [12]. The early 1900s witnessed the practice of using nonviable materials as soft tissue fillers. Synthetic fillers included paraffin, mineral/vegetable oils, and beeswax, which introduced issues such as inflammation, infection, and migration. In the 1960s, silicone was developed by Thomas Cronin and Frank Gerow, and rapidly became wide spread as a filler material [13]. Saline and silicone fillers are both approved by the US Food and Drug Administration, and have been used as synthetic fillers for the reconstruction and augmentation of breast tissue. Despite their longstanding use, silicone gel and saline implants may have complications such as rupture, contraction of fibrous capsule, and potential volume loss, and some patients claim that synthetic fillers lack the natural feeling of breast tissue [14, 15]. In an effort to produce biological fillers, collagen, human dermal ECM, and adipose tissue extracts have become alternatives [16].

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In recent years, stem cells and biomaterials have emerged as tools to generate soft tissue grafts. Different strategies include the use of a biocompatible material that can be engineered to match the shape of the defect and can function as a vehicle to deliver cells and/or biologically active factors that promote tissue regeneration. Cell sources can be either autologous or allogeneic. Autologous sources (i.e., the patient’s own cells) may be preferable because these should avoid immune rejection or foreign body reaction [15], but allogeneic cells may serve as a universal donor cell source that eliminates the need of an invasive cell harvest procedure for the patient.

14.2 Adipose Tissue The human body has two major forms of adipose tissue, WAT and BAT. WAT stores energy in the form of triglycerides and plays important roles in metabolism. BAT, however, has very different physiological roles. BAT generates heat by consuming triglycerides, and is conventionally believed to be present predominantly in newborns [17]. Recent work shows active BAT in adult individuals [18], despite occurring in modest amounts compared to WAT. Obese individuals may have smaller amounts and/or less active BAT than lean individuals [19, 20]. WAT mass can range from as little as 3% (lean individuals) to as much as 70% (obese individuals) of body weight [21]. Contrary to the conventional view as a biologically inert tissue, WAT is now viewed as a highly metabolically active endocrine organ [22]. The metabolic properties of adipose tissue differ depending on anatomic location. WAT is predominantly located subcutaneously or intra-abdominally. In obese individuals, WAT can also accumulate intramuscularly and pericardially [23]. Several studies have linked intra-abdominal WAT to Type II diabetes and cardiovascular disease, whereas subcutaneous adipose tissue is associated with a decreased risk [24, 25]. Besides anatomic division, WAT may be divided into subcategories according to metabolic properties. Human adipose-derived stem cells from different subcutaneous depots show different adipogenic differentiation potential and lipolytic function [26]. Different biological and metabolic properties of the different WAT depots [27] can pose as an obstacle for autologous grafts, as biological noncompatibility factors may hamper the graft’s integration, which can lead to functional loss or volume reduction over time [16].

14.3 Adipose Tissue Engineering Work in recent years has shown that adipose tissue can regenerate or be bioengineered in vivo. Typically, a bioengineered adipose graft has several key components: cells, blood vessels, and a minimal amount of ECM. Angiogenesis is critical to the success of regenerated adipose tissue, as adipocytes have high demand for vascular

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supply. Adipocytes, while being the majority of cells in adipose tissue, are certainly not the only cells. Adipose stem cells, vascular-endothelial cells, and likely other cells are present among adipocytes.

14.3.1 Cell Sources There are many reasons not to use mature adipocytes for the regeneration of adipose tissue. Mature adipocytes are difficult to handle in cell culture, and undergo very little proliferation, which limits the ability to obtain large cell numbers. Instead, stem cells are frequently used for regeneration of adipose tissue by creating cell niches that provide a continuous supply of gradually maturing adipocytes. Stem cells used for regeneration can be autologous (from the same individual), allologous (from the same species), or xenologous (from different species). While nonhuman primary cells and cell lines may be useful research tools, postnatal human stem cells or pre-adipocytes are of more direct clinical relevance. Adult stem cells, either MSCs from bone marrow or ASCs, can both differentiate into multiple cell types including bone, cartilage, and adipose. Adipogenic potential of bone marrow MSCs has been well documented in vitro and in vivo [28–30]. ASCs share some but not all properties with MSC, and are more abundant and easily obtainable through liposuction or biopsy [31]. ASCs and MSCs have both been used in vitro and in vivo [32, 33] with promising clinical results [34].

14.3.2 Biomaterials Several biological and synthetic materials have been used to enhance soft tissue engineering. The materials function both as a vehicle for cell delivery and can be modified toward desired shape and mechanical properties. Adipose tissue scaffolds are designed to deliver both stem cells and their niche, and accommodate additional cell behaviors including proliferation and differentiation in vitro or in vivo. 14.3.2.1 Synthetic Scaffolds Synthetic polymers can be engineered to meet specific needs such as degradation rate as well as mechanical and chemical properties [14]. In vitro and in vivo studies have been performed using polymers based on lactic and glycolic acids, specifically PGA [35], PLA [33, 36–38], and PLGA [39–42]. These materials support adipogenesis and may become vascularized. They can also be altered to affect degeneration rate, ranging from weeks to several months in  vivo [43]. However, results regarding long-term success of engineered adipose tissue are lacking, as the adipose tissue typically disappears within a few months. Other synthetic materials include

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silicone, which is biocompatible but not degradable [44], and PEG, which retains shape and dimension but is not yet studied over longer time periods [45]. 14.3.2.2 Native Scaffolds Native scaffolds are natural components of the ECM. Native scaffolds, given their ability to accommodate cells [46, 47], are preferred in many ways but may suffer from low mechanical stiffness. Natural polymers include collagen, which is present in native adipose tissue ECM but degrades quickly after transplantation [28, 35, 48–56]. Hyaluronic acid has reasonably good mechanical properties and supports in vitro adipogenic differentiation [57–62]. Fibrin has been previously utilized for adipose tissue regeneration and supports adipogenesis in vivo, although it is difficult to create 3D scaffolds in desired shapes and dimensions [63–68]. 14.3.2.3 Vehicles for Drug and Growth Factor Delivery Several methods for drug and growth factor delivery have been used in adipose tissue regeneration. Frequently, drug or growth factors are adsorbed in the scaffold or encapsulated in spherical biomaterial particles. bFGF has been shown to promote both adipogenesis and angiogenesis [50, 69, 70]. Several studies have shown increased adipogenesis and/or angiogenesis by delivering bFGF in microspheres and scaffolds made of gelatin [50, 51, 71, 72], matrigel [69, 70], and fibrin glue [64, 67]. Similar to bFGF, insulin-like growth factor 1 and insulin are frequently used to promote adipogenesis [73–75].

14.3.3 Animal Models A number of studies have explored adipose tissue regeneration in vivo in various animal models (Tables 14.1–14.3). The vast majority of adipose regeneration studies have been performed in rodents. Only a few studies have been performed in larger animal models. Rodents are easy to handle and have low genetic variance, with the added benefit of availability of antibodies and other molecular probes. However, perceived differences in metabolic function and adipose mass between different strains, as well as the sex of the animals, hamper comparisons between studies [76]. Humans live 30–50 times longer than mice, and experience ~100,000 times more cell divisions [77]. Mice have a basal metabolic rate about seven times higher than humans [78]. In addition, the use of a rodent model only permits studies of small defects, and not large physiological relevant defects such as those created by breast cancer surgery. One of the critical aspects for graft survival is vascularization of the newly engineered construct in vivo. Repair of smaller defects may not replicate the formation of new vessels needed in larger animals or humans [79]. Long-term

ddY

C57BL/6

C57BL/6

C57BL/6

BALB/c BALB/c BALB/c C57BL/6

BALB/c BALB/c BALB/c BALB/c BALB/c BALB/c BALB/c

Chambers in contact with existing fat pad Sca-1+ CD34+ Lin-Mouse ASC De-differentiated mature adipocytes from mice

Human ASC Human ASC Mouse ASC None None Human ASC 3T3-F442A or 3T3-L1 cell line 3T3-FA42A cell line Mouse ASC Rabbit MSC Chambers in contact with existing fat pad None

SC

None

Matrigel

SC

Femoral area

Polycarbonate chambers filled with either matrigel or PLGA Silicone chambers filled with matrigel

Matrigel Fibrin glue PLGA spheres Silicone chambers filled with collagen

6 months

PGA with collagen and polypropylene PLGA spheres Collagen sponge Collagen sponge Matrigel Matrigel Fibrin None

3 weeks

3 weeks

[95]

[94]

[69]

[93]

6 weeks 6 weeks

[90] [91] [92] [51]

[41] [50] [53] [72] [71] [68] [89]

[33]

Reference [88]

6 weeks 8 weeks 2 weeks 6 weeks

8 weeks 6 weeks 8 weeks 15 weeks 6 weeks 36 weeks 15 weeks

Duration 12 weeks

Scaffold None

Femoral area

SC SC SC Femoral area

SC SC SC SC SC SC SC

Table 14.1  Adipose tissue regeneration in mouse models Strain Cell source Site A-Zip Sca-1+ CD34+ CD29+ Parametrial CD24+ Lin-Mouse ASC fat pad BALBc Human ASC SC

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Cell source

Site

Scaffold

Duration

Reference

KSN None SC Matrigel 2 weeks [70] Ncr Mouse ASC SC None 3 months [96] NMRI 3T3-L1 SC PGA 24 weeks [35] NMRI 3T3-L1 cell line SC PCA 3 weeks [97] NMRI Human ASC SC Collagen sponge 12 weeks [54] NMRI Human ASC SC Collagen sponge 8 weeks [98] NMRI Human ASC SC Hyaluronan sponge 12 weeks [62] SCID Human MSC SC PEGDA 4 weeks [45] SCID Human ASC SC Gelatin sponge 4 weeks [99] SCID Human ASC or fat graft Femoral groin area Silicone chambers filled with matrigel 9 weeks [100] SCID 3T3-FA42A cell line SC None 4 weeks [101] Human ASC SC PGA/PLLA scaffolds filled with fibrin 6 weeks [64] Unspecified athymic Human ASC SC Fibrin 6 weeks [67] Unspecified athymic Human MSC SC PEG with collagen 4 weeks [84] Unspecified athymic Ob17 cell line SC None 14 weeks [102] Unspecified athymic Human ASC SC Fluoropolymers 4 weeks [103] Unspecified athymic Mouse ASC SC Matrigel 4 weeks [104] Unspecified nude Abbreviations: SC subcutaneous; ADC adipose derived cells; MSC mesenchymal cells; PEG poly(ethylene glycol); PEGDA PEG diacrylate; PGA poly(glycolic acid); PLLA poly(l-lactic acid); PLGA poly(l-lactic-co-glycolic acid); PP polypropylene

Strain

14  Adult Stem Cells and Regeneration of Adipose Tissue 257

Scaffold Duration Reference PGA 8 [38] PLGA 42 weeks [43] PLGA 5 weeks [80] None 6 months [105] None 6 months [106] Collagen-chitosan 2 [107] Polycarbonate chambers filled with PLGA 6 [42] PLGA/PEG microspheres with insulin, 12 [74] IGF-1 and bFGF Sprague Dawley None SC PLGA/PEG microspheres with insulin, 4 [73] IGF-1 and bFGF Sprague Dawley None Femoral groin area Polycarbonate chambers filled with PLGA 8 [108] Wistar Rat pre-adipocytes Femoral groin area Collagen-polypropylene 6 [49] Wistar Rat ASC Spleen None 3 weeks [109] Wistar Rat ASC SC None 6 months [110] Wistar Rat pre-adipocytes Abdominal muscle Fibrin 48 [111] Wistar None SC Styrenated gelatin microspheres with 6 [75] insulin, IGF-1 and bFGF Wistar Rat pre-adipocytes Abdominal muscle Fibrin 48 [112] Unspecified athymic None Femoral groin area Silicone chamber with matrigel and PGA 20 [44] Not specified Rat pre-adipocytes SC Hyaluronic sponge 12 [60] Abbreviations: SC subcutaneous; ADC adipose derived cells; MSC mesenchymal cells; PEG poly(ethylene glycol); PEGDA PEG diacrylate; PGA poly(glycolic acid); PLLA poly(l-lactic acid); PLGA poly(l-lactic-co-glycolic acid); PP polypropylene

Table 14.2  Adipose tissue regeneration in rat models Strain Cell source Site Lewis Rat ASC SC Lewis Rat ASC SC Lewis Rat ASC SC Lewis Rat ASC Epididymal fat pad Lewis Rat adipocytes SC and epididymal fat pad Lewis Rat ASC SC Sprague Dawley Existing fat pad Femoral groin area Sprague Dawley Rat fat grafts SC

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14  Adult Stem Cells and Regeneration of Adipose Tissue Table 14.3  Adipose tissue regeneration in large animal models Animal model Cell source Site Scaffold New Zeeland Autologous SC None white rabbits adipose tissue New Zeeland Autologous SC None white rabbits adipose tissue New Zeeland Autologous SC None white rabbits adipose tissue Pig Pig ASC SC Hyaluronic gel Sheep Sheep ASC SC Alginate Abbreviations: SC subcutaneous; ADC adipose derived cells

259

Duration 12 months

Reference [113]

12 months

[114]

4 months

[115]

6 weeks 12

[61] [87]

studies in animal models are necessary for relevant comparison of the data to human applications. Importantly, few previous studies have measured the starting and final volume of the regenerated adipose tissue grafts (Tables 14.1–14.3), making it impossible to compare volume retention. Volume loss is one of the key deficiencies of existing soft tissue reconstruction [4, 5, 43, 80, 81], and therefore a factor that must be satisfactorily addressed by adipose tissue regeneration prior to clinical translation.

14.3.4 Critical Issues in Adipose Tissue Engineering Metabolically active adipose tissue has a high demand for blood supply. Slow revascularization of autologous grafts and suboptimal angiogenesis of adipose tissue grafts are an obstacle to the success of current procedures for soft tissue regeneration [30, 82]. Since adequate nutrition and oxygen supply is crucial for the survival of the engineered grafts [83] and vascular diffusion is limited to 200 mm [30, 83, 84], proper vascularization of engineered tissue is a top priority, yet remains the area within the field of tissue engineering where the least progress has been made during the last decade [85]. Clearly, vascularization and scale-up are two critical factors that limit the translation of adipose tissue regeneration toward clinical application [86].

14.4 In Vivo Soft Tissue Engineering Approaches While a great deal of meritorious work has collectively moved the field of adipose tissue regeneration forward, we highlight a few previous studies as examples that especially address scale-up and vascularization, two of the limiting factors for the translation of adipose tissue regeneration to the clinic.

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Fig. 14.1  Preparation of tissue grafts: ASCs were isolated from human lipectomy samples and expanded in cell culture. (a) A dome-shaped core of PEGDA was generated. (b) The PEGDA core was covered in an alginate layer containing ASCs (c) and differentiated to adipocytes in vitro for 4 weeks (d) before transplantation. The grafts were transplanted subcutaneous in the dorsum of athymic rats (e) and harvest after 4 weeks (f). At that time, the graft had integrated with the surrounding tissue (g). Hematoxylin and eosin staining of the implants showed that in constructs seeded with ASCs, large numbers of adipocytes (green arrows) formed within the alginate layer (h). (i) is a higher magnification of (h), and (j) is a higher magnification of (i). Blue arrow heads indicate erythrocytes and endothelial cells lining blood vessels. Asterisk indicates alginate residue. Constructs without ASCs showed little adipose tissue, but fibrous tissue in the alginate layer (k). (l) is a higher magnification of (k), and (m) is a higher magnification of (l). Adapted from [32] with permission from the publisher

14.4.1 Hybrid Implants One of the major barriers in soft tissue engineering is the small size of engineered grafts. A recent study attempted to scale-up regenerated adipose tissue implants by a hybrid construct with an acellular core of PEGDA coated with adipocytes in alginate [32] (Fig.  14.1a–d). The PEGDA core was used as a biocompatible

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filler material with the purpose of not only increasing the volume of the implant, but also maintaining the shape and form of the construct, whereas the cellular surface was designed to facilitate the implant’s integration with host tissue. The construct was implanted subcutaneously in the dorsum of athymic mice. After 4 weeks, the construct was well integrated into the surrounding host tissue (Fig. 14.1e–g). Adipose tissue was formed in the alginate layer of the construct (Fig. 14.1h–j), whereas the inner PEGDA core maintained the overall shape and dimensions of the implant (Fig. 14.1k–m). This study represents an alternative to cellular scale-up by providing an acellular inner core biomaterial of limitless size with an exterior cellular layer of adipose tissue, in this case, formed by adipose stem cells.

14.4.2 Microchannels and/or Angiogenic Cues Another major challenge in soft tissue engineering is vascularization. Scale-up of regenerated adipose tissue, just as with native adipose tissue, requires appropriate vascularization. A biophysical approach using microchannels was combined with a biological approach using angiogenic growth factors in an attempt to induce vascularization in regenerated adipose tissue [30]. PEG was used as scaffold material with multiple 1 mm diameter microchannels with or without adsorbed bFGF along with human MSC-derived adipogenic cells (Fig.  14.2a–c). The constructs were transplanted subcutaneously in the dorsum of immunodeficient mice, and harvested after 4 weeks. The combination of delivered cells, microchannels, and bFGF promoted both adipogenesis and angiogenesis (Fig. 14.2f), whereas microchannels or bFGF alone promoted angiogenesis but no adipogenesis. This study provides a successful example of vascularizing adipose tissue grafts using microchannels and/ or angiogenic growth factor(s).

14.4.3 Arteriovenous Pedicles Another strategy to address the challenge of poor vascularization in regenerated adipose tissue is to utilize preexisting blood vessels. In one study, the superficial epigastric vessels in the groin area of mice were stripped of existing fat and wrapped using a silicone tube (Fig. 14.3a–c) [69]. The tubes were filled with matrigel and supplemented with bFGF. Group one consisted of completely sealed chambers. Group two had partially open chambers in contact with surrounding adipose tissue. Group three consisted of sealed chambers with some of the removed adipose tissue placed inside the chamber. After 6 weeks, the chambers were removed. New blood vessels formed in all three groups, and the group with sealed chambers had virtually no adipogenesis. In contrast, adipose tissue formed in open chambers, especially in the open end. The group with autologous fat graft was filled with newly formed adipose tissue throughout the chamber (Fig. 14.3f–i).

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Fig. 14.2  Engineering of scaffolds using PEG hydrogel (a) with channels to facilitate vascularization (b, c). The scaffolds were supplemented with the angiogenic and adipogenic growth factor, bFGF (b, c), and with human MSCs that had been predifferentiated toward adipocytes in vitro for 1 week (c). The empty scaffolds (a, d) showed no vascularization or infiltration of host cells. Scaffolds with microchannels and supplemented with bFGF (b, e) showed infiltration of host cells and neovascularization. Scaffolds with micro channels supplemented with bFGF and with seeded cells (c, f) demonstrated both vascularization and formation of adipose tissue. Adapted from [30] with permission from the publisher

14.4.4 Adipose Tissue Regeneration in a Large Animal Model In a sheep model, alginate was used as an injectable material for adipose tissue regeneration [87]. Small alginate beads were fortified with a short RGD-peptide that facilitates cell adhesion. The beads were seeded with autologous ASCs before injected subcutaneously in the neck of sheep. After 3 months, the tissue was harvested. Alginate and alginate-RDG beads promoted infiltration of host tissue and neo-vascularization, but little adipose tissue formation (Fig. 14.4a, b). Alginate and alginate-RDG beads seeded with ASCs led to the formation of adipose tissue (Fig. 14.4c, d), although there was no determination whether regenerated adipose tissue derived from transplanted cells and/or host endogenous cells. This study addressed one of the important issues in soft tissue engineering in a large animal model, although the tissue grafts created were still somewhat small. It also highlights the need for distinguishing between the contributions of transplanted versus endogenous cells in adipose tissue regeneration.

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Fig. 14.3  The superficial epigastric vessels of mice were stripped of native adipose tissue and silicone tube was wrapped around the vessels. The tube was filled with matrigel supplied with bFGF, and was either completely sealed (a), partially opened and in contact with surrounding adipose tissue (b), or sealed and supplemented with 1–5 mg adipose tissue (c). Pictures show the implants at the day of surgery (d) and after 6 weeks (e). Staining with hematoxylin and eosin showed that the group with the sealed chamber had good vascularization but poor adipogenesis (f). The group with the open chamber showed good vascularization and adipogenesis close the open side, depicted left (g). When an autologous fat graft was added to the chamber, the adipogenesis were spread throughout the chambers, 2–5 mg fat (h), 1 mg fat (i). Adapted from [69] with permission from the publisher

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Fig. 14.4  Omental adipose tissue was harvested from domestic sheep. Stroma-vascular cells were isolated and expanded in vitro, and seeded onto alginate beads. Four groups were used; alginate beads (a), alginate-RGD beads (b), alginate beads with autologous adipose-derived cells (c), alginateRGD beads with autologous adipose-derived cells (d). The grafts were harvested after 3 months. Grafts consisting of alginate (a) and alginate-RGD (b) beads showed vascular and tissue ingrowth, but very little evidence of adipose tissue. In grafts made of alginate (c) and alginate-RGD (d) beads seeded with autologous adipose-derived cells, there was adipose tissue formed within the grafts, however, it was unclear if the tissue was from transplanted cells or cells migrating from the surrounding host tissue. Adapted from [87] with permission from the publisher

14.5 Conclusions Adipose tissue is made of abundant cells with little ECM that nonetheless houses vitally important blood vessels for the survival of adipocytes and progenitor cells. Clinical demands for adipose tissue regeneration are substantial and significantly under-addressed. Current practices using autologous grafting and native or synthetic fillers such as silicone gel or saline implants have limitations and, like most foreign substances, do not provide long-term solutions. The following suggestions are made as food for thought for soft tissue regeneration: (1) identify critical areas of clinical demand and formulate individual strategies for different clinical entities; (2) promote angiogenesis and survival of adipose tissue grafts; (3) include the use of microchannels, angiogenic growth factors, and hybrid implants to improve graft survival and

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function; (4) continue to scale-up bioengineered adipose tissue grafts and pursue experiments in large animal models; and (5) identify success criteria for adipose tissue regeneration to including the maintenance of volume and shape over time. Acknowledgments  We thank Michael Diggs, Qiongfen Guo, and Kening Hua for administrative assistance. The work described in this chapter is supported by the Swedish Society for Medical Research to D.A.H. and NIH grants RC2DE020767 and R01EB006261.

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Part IV

Synthetic Organs

Chapter 15

Hollow Organ Engineering Anthony Atala

Abstract  Applications of regenerative medicine technology may offer novel therapies for patients with injuries, end-stage organ failure, or other clinical problems. Currently, patients suffering from diseased and injured organs can be treated with transplanted organs. However, there is a severe shortage of donor organs that is worsening yearly as the population ages and new cases of organ failure increase. Scientists in the field of regenerative medicine and tissue engineering are now applying the principles of cell transplantation, material science, and bioengineering to construct biological substitutes that will restore and maintain normal function in diseased and injured tissues. In particular, efforts to engineer hollow organs, such as the urinary bladder, urethra, and vagina, have been particularly successful to date. Some therapies arising from these tissue engineering endeavors have already entered the clinical setting successfully, indicating the promise regenerative medicine holds for the future.

Abbreviations AFPS ECM hESC iPSC PCL PGA PLA

Amniotic-fluid and placental-derived stem Extracellular matrix Human embryonic stem cell Induced pluripotent stem cell Polycaprolactone Polyglycolic acid Polylactic acid

A. Atala (*) Wake Forest Institute for Regenerative Medicine, Wake Forest University School of Medicine, Winston-Salem, NC, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_15, © Springer Science+Business Media, LLC 2011

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PLGA Poly (lactic-co-glycolic acid) SCNT Somatic cell nuclear transfer SIS Small intestinal submucosa

15.1 Introduction Regenerative medicine follows the principles of cell transplantation, materials science, and engineering toward the development of biologic constructs that can restore and maintain normal organ function. Regenerative medicine strategies usually fall into one of three categories: cell-based therapies such as direct cell injection, the use of biomaterial scaffolds alone (in which the body’s natural ability to regenerate is used to orient or direct new tissue growth), or the use of scaffolds seeded with cells to create tissue substitutes. Such strategies have been particularly successful for the construction of hollow organs, such as the urinary bladder, intestine, urethra, blood vessels, esophagus, and trachea. Most hollow organs are organized in a similar fashion, consisting of epithelium or endothelium (which lines the lumen of the organ) surrounded by a collagen-rich connective tissue and muscle layer. Compared to the complex structure of solid organs such as the kidney or the brain, the layered structure of the hollow organs is relatively simple to replicate in the laboratory if the appropriate components are available. This chapter discusses these components and introduces the tissue engineering techniques used to construct hollow organs using specific examples.

15.2 The Basic Components of Hollow Organ Construction Most hollow organs are organized in a similar fashion, consisting of epithelium or endothelium (which lines the lumen of the organ) that is surrounded by a collagenrich connective tissue and muscle layer. The epithelial or endothelial layer serves as a barrier that prevents the contents of the lumen from escaping into the body cavity. The collagen-rich layers of muscle tissue surrounding the epithelium/ endothelium maintain the structural integrity of the organ and provide physiological functions such as contractility. The cells composing these layers interact with each other and with proteins and other factors in the ECM to regulate cellular differentiation and function. Thus, multiple cell types are required to create a hollow organ with the appropriate “layered” structure, and a biomaterial “scaffold” must be used as an artificial ECM in order to support and direct the growth of these cells. Since each of these cell types favors different conditions for optimal growth and differentiation, an ideal biomaterial must provide an environment in which corresponding cell types could interact with each other to guide appropriate regulation that governs adhesion, proliferation, migration, and differentiation. This section briefly discusses the components needed to construct a hollow organ using these techniques.

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15.2.1 Cells 15.2.1.1 Stem Cells Stem cells used for regenerative medicine can be autologous or heterologous, and in general, there are three broad categories of stem cells obtained from living tissue that are used for cell therapies. Embryonic stem cells are obtained through the aspiration of the inner cell mass of a blastocyst or, more recently, a single cell from this mass. Fetal and neonatal amniotic fluid and placenta may contain multipotent cells that may be useful in cell therapy applications. Finally, adult stem cells are usually isolated from organ or bone marrow biopsies. Stem cells are defined as having three important properties: the ability to self-renew, the ability to differentiate into a number of different cell types, and the ability to easily form clonal populations (populations of cells derived from a single stem cell).

Embryonic Stem Cells In 1981, pluripotent cells were found in the inner cell mass of the human embryo, and the term “human embryonic stem cell” was coined [1]. These cells are able to differentiate into all cells of the human body, excluding placental cells (only cells from the morula are totipotent; that is, able to develop into all cells of the human body). These cells have great therapeutic potential, but their use is limited by both biological and ethical factors. The political controversy surrounding stem cells began in 1998 with the creation of hESCs derived from discarded embryos. hESCs were isolated from the inner cell mass of a blastocyst (an embryo 5 days postfertilization) using an immunosurgical technique. Given that some cells cannot be expanded ex vivo, hESCs could be an ideal resource for regenerative medicine because of their fundamental properties: the ability to self-renew indefinitely and the ability to differentiate into cells from all three embryonic germ layers. Skin and neurons have been formed, indicating ectodermal differentiation [2, 3]. Blood, cardiac cells, cartilage, endothelial cells, and muscle have been formed, indicating mesodermal differentiation [4–6]. Pancreatic cells have been formed, indicating endodermal differentiation [7]. 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 [8]. These cells have demonstrated longevity in culture and can maintain their undifferentiated state for at least 80 passages when grown using current published protocols [9, 10]. However, an objection to hESC research is that it results in the destruction of embryos, and because of the ethical issues associated with this, research on hESCs is limited in many countries. In addition, the clinical application of hESCs is limited because they represent an allogenic resource and thus have the potential to evoke an immune response.

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Therapeutic Cloning SCNT, or therapeutic cloning, entails the removal of an oocyte nucleus in culture, followed by its replacement with a nucleus derived from a somatic cell obtained from a patient. Activation with chemicals or electricity stimulates cell division up to the blastocyst stage, and then ES cells which are genetically identical to the patient can be obtained from the inner cell mass. At this point, it is extremely important to differentiate between the two types of cloning that exist – reproductive cloning and therapeutic cloning. Both involve the insertion of donor DNA into an enucleated oocyte to generate an embryo that has identical genetic material to its DNA source. However, the similarities end there. In reproductive cloning, the embryo is then implanted into the uterus of a pseudopregnant female to produce an infant that is a clone of the donor. A world-famous example of this type of cloning resulted in the birth of a sheep named Dolly in 1997 [11]. However, there are many ethical concerns surrounding such practices, and as a result, reproductive cloning has been banned in most countries. While therapeutic cloning also produces an embryo that is genetically identical to the donor, this process is used to generate blastocysts that are explanted and grown in culture, rather than in utero. Embryonic stem cell lines can then be derived from these blastocysts, which are only allowed to grow up to a 100-cell stage. At this time the inner cell mass is isolated and cultured, resulting in ES cells that are genetically identical to the patient. It has been shown that nuclear transferred ES cells derived from fibroblasts, lymphocytes, and olfactory neurons are pluripotent and can generate live pups after injection into blastocysts. This shows that cells generated by SCNT have the same developmental potential as blastocysts that are fertilized and produced naturally [12]. In addition, the ES cells generated by SCNT are perfectly matched to the patient’s immune system and no immunosuppressants would be required to prevent rejection should these cells be used in regenerative medicine applications. Although promising, SCNT has certain limitations that require further improvement before its clinical application, in addition to the ethical considerations regarding the potential of the resulting embryos to develop into cloned embryos if implanted into a uterus. In addition, this technique has not been shown to work in humans to date. The initial failures and fraudulent reports of nuclear transfer in humans reduced enthusiasm for human applications [13], although it was recently reported that nonhuman primate ES cell lines were generated by SCNT of nuclei from adult skin fibroblasts [14, 15]. In addition, before SCNT-derived ES cells could be used as clinical therapy, careful assessment of quality of the lines must be determined. For example, some cell lines generated by SCNT have contained chromosomal translocations and it is not known whether these abnormalities originated from aneuploid embryos or if they occurred during ES cell isolation and culture. In addition, the low efficiency of SNCT (0.7%) and the inadequate supply of human oocytes further hinder the therapeutic potential of this technique. Still, these studies renew the hope that ES cell lines could one day be generated from human cells to produce patient-specific stem cells with the potential to cure many human diseases that are currently untreatable.

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Induced Pluripotent Stem Cells Recently, reports of the successful transformation of adult cells into pluripotent stem cells through a type of genetic “reprogramming” have been published. Reprogramming is a technique that involves de-differentiation of adult somatic cells to produce patient-specific pluripotent stem cells, eliminating the need to create embryos. Cells generated by reprogramming would be genetically identical to the somatic cells (and thus, the patient who donated these cells) and would not be rejected. Yamanaka was the first to discover that mouse embryonic fibroblasts and adult mouse fibroblasts could be reprogrammed into an “induced pluripotent state” [16]. These iPSCs possessed the immortal growth characteristics of self-renewing ESCs, expressed genes specific for ESCs and generated embryoid bodies in vitro and teratomas in vivo. When iPSCs were injected into mouse blastocysts, they contributed to a variety of cell types. However, although iPSCs selected in this way were pluripotent, they were not identical to ESCs. Unlike ESCs, chimeras made from iPSCs did not result in full-term pregnancies. Gene expression profiles of the iPSCs showed that they possessed a distinct gene expression signature that was different from that of ESCs. In addition, the epigenetic state of the iPSCs was somewhere between that found in somatic cells and that found in ESCs, suggesting that the reprogramming was incomplete. These results were improved significantly by Wernig and Jaenisch in July 2007 [17]. In this study, DNA methylation, gene expression profiles, and the chromatin state of the reprogrammed cells were similar to those of ESCs. Importantly, the reprogrammed cells from this experiment were able to form viable chimeras and contribute to the germ line such as ESCs, suggesting that these iPSCs were completely reprogrammed. It has recently been shown that reprogramming of human cells is possible [18, 19]. Yamanaka generated human iPSCs that are similar to hESCs in terms of morphology, proliferation, gene expression, surface markers, and teratoma formation. Thompson’s group showed that retroviral transduction of the stem cell markers OCT4, SOX2, NANOG, and LIN28 could generate pluripotent stem cells. However, in both studies, the human iPSCs were similar but not identical to hESCs. Although reprogramming is an exciting phenomenon, our limited understanding of the mechanism underlying it currently limits the clinical applicability of the technique, but the future potential of reprogramming is quite exciting. Amniotic-Fluid and Placenta-Derived Stem Cells The amniotic fluid and placental membrane contain a heterogeneous population of cell types derived from the developing fetus. Cells found in this heterogeneous population include mesenchymal stem cells [20]. In addition, the isolation of multipotent human and mouse AFPS cells that are capable of extensive self-renewal and give rise to cells from all three germ layers was reported in 2007 [21]. AFPS cells represent approximately 1% of the cells found in the amniotic fluid and placenta. The undifferentiated stem cells expand extensively without a feeder cell layer and double every

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36 h. Unlike hESCs, the AFPS cells do not form tumors in vivo. Lines maintained for over 250 population doublings retained long telomeres and a normal complement of chromosomes. AFPS cell lines can be induced to differentiate into cells representing each embryonic germ layer, including cells of adipogenic, osteogenic, myogenic, endothelial, neural-like, and hepatic lineages. In addition to the differentiated AFPS cells expressing lineage-specific markers, such cells can have specialized functions. Cells of the hepatic lineage secreted urea and a-fetoprotein, while osteogenic cells produced mineralized calcium. In this respect, they meet a commonly accepted criterion for multipotent stem cells, without implying that they can generate every adult tissue. AFS cells represent a new class of stem cells with properties somewhere between those of embryonic and adult stem cell types, probably more agile than adult stem cells, but less so than embryonic stem cells. Unlike ESCs and iPSCs, however, AFPS cells do not form teratomas, and if preserved for self-use, avoid the problems of rejection. The cells could be obtained either from amniocentesis or chorionic villous sampling in the developing fetus, or from the placenta at the time of birth. They could be preserved for self-use, and used without rejection, or they could be banked. A bank of 100,000 specimens could potentially supply 99% of the US population with a perfect genetic match for transplantation. Such a bank may be easier to create than with other cell sources, since there are approximately 4.5 million births per year in the USA. Since the discovery of the AFPS cells, other groups have published on the potential of the cells to differentiate to other lineages, such as cartilage [22], kidney [23], and lung [24]. Muscle differentiated AFPS cells were also noted to prevent compensatory bladder hypertrophy in a cryo-injured rodent bladder model [21]. Adult Stem Cells Adult stem cells, especially hematopoietic stem cells, are the best understood cell type in stem cell biology [25]. However, adult stem cell research remains an area of intense study, as their potential for therapy may be applicable to a myriad of degenerative disorders. Within the past decade, adult stem cell populations have been found in many adult tissues other than the bone marrow and the gastrointestinal tract, including the brain [26], skin [27], and muscle [28]. Many other types of adult stem cells have been identified in organs all over the body and are thought to serve as the primary repair entities for their corresponding organs [29]. The discovery of such tissue-specific progenitors has opened up new avenues for research. A notable exception to the tissue-specificity of adult stem cells is the mesenchymal stem cell, also known as the multipotent adult progenitor cell. This cell type is derived from bone marrow stroma [30]. Such cells can differentiate in  vitro into numerous tissue types [31] and can also differentiate developmentally if injected into a blastocyst. Multipotent adult progenitor cells can develop into a variety of tissues including neuronal [32], adipose [28], muscle [28], liver [33], lungs [34], spleen [35], and gut tissue [30], but notably not bone marrow or gonads.

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Research into adult stem cells has, however, progressed slowly, mainly because investigators have had great difficulty in maintaining adult nonmesenchymal stem cells in culture. Some cells, such as those of the liver, pancreas, and nerve, have very low proliferative capacity in vitro, and the functionality of some cell types is reduced after the cells are cultivated. Isolation of cells has also been problematic, because stem cells are present in extremely low numbers in adult tissue [36]. While the clinical utility of adult stem cells is currently limited, great potential exists for future use of such cells in tissue-specific regenerative therapies. The advantage of adult stem cells is that they can be used in autologous therapies, thus avoiding any complications associated with immune rejection. 15.2.1.2 Native Targeted Progenitor Cells In the past, one of the limitations of applying cell-based regenerative medicine techniques to organ replacement was the inherent difficulty of growing certain human cell types in large quantities. Native targeted progenitor cells, or native cells, are tissue-specific unipotent cells derived from most organs. The advantage of these cells is that they are already programmed to become the cell type needed, without any extra-lineage differentiation. By noting the location of the progenitor cells, as well as by exploring the conditions that promote differentiation and/or self-renewal, it has been possible to overcome some of the obstacles that limit cell expansion in vitro. One example is the urothelial cell. Urothelial cells could be grown in the laboratory setting in the past, but only with limited success. It was believed that urothelial cells had a natural senescence that was hard to overcome. Several protocols have been developed over the last two decades that have improved urothelial growth and expansion [37, 38]. A system of urothelial cell harvesting was developed that 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 that initially covers a surface area of 1 cm2 to one covering a surface area of 4,202 m2 (the equivalent area of one football field) within 8 weeks [37]. An additional advantage in using native cells is that they can be obtained from the specific organ to be regenerated, expanded, and used in the same patient without rejection, in an autologous manner [37, 39–43]. Bladder, ureter, and renal pelvis cells can equally be harvested, cultured, and expanded in a similar fashion. 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 biologic properties can be studied [37, 44, 45]. Major advances in cell culture techniques have been made within the past decade, and these techniques make use of autologous cells possible for clinical application. However, even now, not all human cells can be grown or expanded in vitro. Liver, nerve, and pancreas are examples of human tissues where the technology is not yet advanced to the point where these cells can be grown and expanded. When cells are used for tissue reconstitution, donor tissue is dissociated into individual cells, which is either implanted directly into the host or expanded in

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culture, attached to a support matrix, and reimplanted after expansion. The implanted tissue can be heterologous, allogeneic, or autologous. Ideally, this approach allows lost tissue function to be restored or replaced in total and with limited complications [46].

15.2.2 Biomaterials for Constructing Hollow Organs Biomaterials that function as an artificial ECM are designed to imitate the biologic and mechanical functions of native ECM found in tissues. They facilitate the localization and delivery of cells and/or bioactive factors (e.g., cell adhesion peptides, growth factors) to desired sites in the body, define a three-dimensional space for the formation of new tissues with the appropriate shape and structure, and guide the development of new tissues with appropriate function. As discussed, hollow organs consist of a “layered” structure consisting of an epithelial or endothelial layer surrounded by a collagen-rich, smooth muscle layer. Thus, an ideal biomaterial must provide an environment in which appropriate cell types, including epithelial/endothelial cells and muscle cells, can interact with each other to guide appropriate regulation that governs adhesion, proliferation, migration, and differentiation. In addition, there are clear advantages to using degradable, biocompatible materials that can function as cell delivery vehicles and/or provide the structural parameters needed for tissue replacement. Direct injection of cell suspensions without biomaterial matrices has been used in some cases [47], but it is difficult to control the localization of transplanted cells. In addition, most mammalian cell types are anchorage-dependent and will die if not provided with a cell-adhesion substrate. 15.2.2.1 Design and Selection of Biomaterials The design and selection of a biomaterial are critical in the development of engineered tissues for use in hollow organ reconstruction. Multiple cell types are required to create a hollow organ with the appropriate “layered” structure, and each of these cell types favors different conditions for optimal growth and differentiation. Thus, these factors must be taken into account when considering ideal strategies for hollow organ tissue engineering. The biomaterial must also be capable of controlling the structure and function of the engineered tissue in a predesigned manner by interacting with host cells. Generally, the ideal biomaterial should be biocompatible, promote cellular interaction, and tissue development, and possess proper mechanical and physical properties. The selected biomaterial should be biodegradable and bioresorbable to support the reconstruction of a completely normal tissue without inflammation. Such behavior of the biomaterials avoids the risk of inflammatory or foreign-body responses that may be associated with the permanent presence of a foreign material in the body.

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The degradation rate and the concentration of degradation products in the tissues surrounding the implant must be at a tolerable level [48]. The biomaterials should provide an appropriate regulation of cell behavior (e.g., adhesion, proliferation, migration, differentiation) in order to promote the development of functional new tissue. Cell behavior in engineered tissues is regulated by multiple interactions with the microenvironment, including interactions with celladhesion ligands [49] and with soluble growth factors [50]. Cell adhesion-promoting factors (e.g., Arg-Gly-Asp) can be presented by the biomaterial itself or incorporated into the biomaterial in order to control cell behavior through ligand-induced cell receptor signaling processes [51]. The biomaterials provide temporary mechanical support sufficient to withstand in vivo forces exerted by the surrounding tissue and maintain a potential space for tissue development. The mechanical support of the biomaterials should be maintained until the engineered tissue has sufficient mechanical integrity to support itself [46]. This potentially can be achieved by an appropriate choice of mechanical and degradative properties of the biomaterials. The biomaterials need to be processed into specific configurations. A large ratio of surface area to volume is often desirable to allow the delivery of a high density of cells. A high-porosity, interconnected pore structure with specific pore sizes promotes tissue ingrowth from the surrounding host tissue. Several techniques, such as electrospinning, have been developed that readily control porosity, pore size, and pore structure [52, 53]. 15.2.2.2 Types of Biomaterials Generally, three classes of biomaterials have been used for tissue engineering applications: naturally derived materials, such as collagen and alginate; acellular tissue matrices, such as bladder submucosa and small-intestinal submucosa; and synthetic polymers, such as PGA, PLA, and PLGA. These classes of biomaterials have been tested in regard to their biocompatibility with primary human urothelial and bladder muscle cells [54]. 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. Collagen is the most abundant and ubiquitous structural protein in the body, and it may be readily purified from both animal and human tissues with an enzyme treatment and salt/acid extraction [55]. Collagen has long been known to exhibit minimal inflammatory and antigenic responses [56], and it has been approved by the US Food and Drug Administration for many types of medical applications, including wound dressings and artificial skin [57]. Intermolecular cross-linking reduces the degradation rate by making the collagen molecules less susceptible to an enzymatic attack. Intermolecular cross-linking can be accomplished by various physical (e.g., ultraviolet radiation, dehydrothermal treatment) or chemical (e.g., glutaraldehyde, formaldehyde, carbodiimides) techniques [55]. Collagen contains cell-adhesion domain sequences (e.g., Arg-Gly-Asp) that exhibit specific cellular interactions.

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This may help to retain the phenotype and activity of many types of cells, including fibroblasts [58] and chondrocytes [59]. This material can be processed into a wide variety of structures such as sponges, fibers, and films [60]. Alginate, a polysaccharide isolated from seaweed, has been used as an injectable cell delivery vehicle [61] and a cell immobilization matrix owing to its gentle gelling properties in the presence of divalent ions such as calcium. Alginate is a family of copolymers of d-mannuronate and l-guluronate. The physical and mechanical properties of alginate gel are strongly correlated with the proportion and length of the polyguluronate block in the alginate chains [61]. Efforts have been made to synthesize biodegradable alginate hydrogels with mechanical properties that are controllable in a wide range by intermolecular covalent cross-linking and with celladhesion peptides coupled to their backbones [62]. Recently, natural materials such as alginate and collagen have been used as “bioinks” in a newly developed bioprinting technique based on inkjet technology [63]. Using this technology, these scaffold materials can be “printed” into a desired scaffold shape using a modified inkjet printer. In addition, several groups have shown that living cells can also be printed using this technology [64]. This exciting technique can be modified so that a three-dimensional construct containing a precise arrangement of cells, growth factors, and ECM material can be printed [65]. Such constructs may eventually be implanted into a host to serve as the backbone for a new tissue or organ. Acellular tissue matrices are collagen-rich matrices prepared by removing cellular components from tissues. The matrices are often prepared by mechanical and chemical manipulation of a segment of bladder tissue [66, 67]. The matrices slowly degrade after implantation and are replaced and remodeled by ECM proteins synthesized and secreted by transplanted or ingrowing cells. Acellular tissue matrices have been proved to support cell ingrowth and regeneration of genitourinary tissues, including urethra and bladder, with no evidence of immunogenic rejection [67]. Because the structures of the proteins (e.g., collagen, elastin) in acellular matrices are well conserved and normally arranged, the mechanical properties of the acellular matrices are not significantly different from those of native bladder submucosa [66]. Polyesters of naturally occurring a-hydroxy acids, including PGA, PLA, and PLGA, are also widely used in regenerative medicine. These synthetic polymers have gained Food and Drug Administration approval for human use in a variety of applications, including sutures. The degradation products of PGA, PLA, and PLGA are nontoxic, natural metabolites that are eventually eliminated from the body in the form of carbon dioxide and water [68]. Because these polymers are thermoplastics, they can easily be formed into a three-dimensional scaffold with a desired microstructure, gross shape, and dimension by various techniques, including molding, extrusion [69], solvent casting [70], phase separation techniques, and gas foaming techniques [71]. More recently, techniques such as electrospinning have been used to quickly create highly porous scaffolds in various conformations [52, 53]. As an approach toward incorporating cell recognition domains into these materials, copolymers with amino acids have been synthesized [51]. Other biodegradable synthetic

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polymers, including poly (anhydrides) and poly (ortho-esters), can also be used to fabricate scaffolds for regenerative medicine with controlled properties [72]. Composite scaffolds have also been created using both natural materials such as collagen and synthetic materials such as PLGA or PCL. Composite scaffolds are particularly useful for engineering hollow organs, as they can be designed to facilitate formation of endothelium on the luminal side of a construct and formation of smooth muscle or another type of barrier on the outer surface of the same construct. This concept has been introduced for bladder tissue engineering by Eberli and ­colleagues [73]. In this study, composite scaffolds were fabricated by bonding a collagen matrix to PGA polymers using threaded collagen fiber stitches. Urothelial and bladder smooth muscle cells were seeded on the composite scaffolds, and implanted in mice for up to 4 weeks and analyzed. Both cell types readily attached and proliferated on the scaffolds and formed bladder tissue-like structures in vivo. Composite scaffolds composed of collagen and PCL have also been created to improve engineering of blood vessels [74]. Using electrospinning, this group was able to create a layered vascular scaffold in which the inner layer was composed of small pores to support formation of an endothelial layer, while the outer layer had large pores that allowed infiltration of smooth muscle cells and formation of an aligned muscle layer. This study suggested that these bilayered scaffolds led to improved in vivo vessel formation.

15.3 Regenerative Medicine Strategies for Reconstructing Several Specific Hollow Structures 15.3.1 Urinary Bladder The urinary bladder is a quintessential hollow organ designed to store urine. Currently, gastrointestinal segments are commonly used as tissues for the replacement or repair of bladders damaged by disease or injury. However, gastrointestinal tissues are designed to absorb specific solutes, whereas bladder tissue is designed for the excretion of solutes. When gastrointestinal tissue is in contact with the urinary tract, multiple complications may ensue, such as infection, metabolic disturbances, urolithiasis, perforation, increased mucus production, and malignancy [75]. Because of the problems encountered with the use of gastrointestinal segments, numerous investigators have attempted alternative reconstructive procedures for bladder replacement or repair. These include autoaugmentation [76] and ureterocystoplasty [77]. In addition, alternate methods for bladder reconstruction have been explored, such as the use of regenerative medicine with cell transplantation. Over the last few decades, several bladder wall substitutes have been attempted with both synthetic and organic materials. Synthetic materials that have been tried in experimental and clinical settings include polyvinyl sponge, Teflon, collagen matrices, Vicryl (PGA) matrices, and silicone. Most of these attempts have failed

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because of mechanical, structural, functional, or biocompatibility problems. Usually, permanent synthetic materials used for bladder reconstruction succumb to mechanical failure and urinary stone formation, and use of degradable materials leads to fibroblast deposition, scarring, graft contracture, and a reduced reservoir volume over time [39]. However, it was hypothesized that biomaterials based on natural substances may perform better. 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 [41, 78, 79]. The matrices can serve as vehicles for partial bladder regeneration, and relevant antigenicity is not evident. Another naturally derived material, SIS, is a biodegradable, acellular, xenogeneic collagen-based tissue-matrix graft, and it was first described by Badylak and colleagues in the 1980s as an acellular matrix for tissue replacement in the vascular field [80]. It has been shown to promote regeneration of a variety of host tissues, including blood vessels and ligaments [81]. The matrix is derived from the small intestine of pig in which the mucosa is mechanically removed from the inner surface, and the serosa and muscular layer are removed from the outer surface. Animal studies have shown that the SIS matrix used for bladder augmentation is able to regenerate some bladder tissue in vivo [82, 83]. However, while the transitional layer was the same as that of the native bladder tissue, the muscle layer was not fully developed. A large amount of collagen was interspersed among a smaller number of muscle bundles. In vitro contractility ­studies performed on SIS-regenerated dog bladders showed a decrease in maximal contractile response by 50% from those of normal bladder tissues. Similarly, when bladder augmentation was performed on minipigs with porcine bowel acellular ­tissue matrix, human placental membranes, or porcine SIS, the grafts contracted to 70, 65, and 60% of their original sizes, respectively, and histologically the grafts showed predominantly only mucosal regeneration [84]. These studies indicate that even when naturally derived materials are used as grafts for cystoplasty, the urothelial layer is able to regenerate normally, but the muscle layer, although present, does not fully develop [78, 85]. Thus, it was hypothesized that building a three-dimensional tissue construct in vitro, using both cells and a biomaterial, before implantation, would facilitate the eventual terminal differentiation of the cells after implantation in vivo and would minimize the inflammatory response toward the matrix, thus avoiding graft contracture and shrinkage. This transplantation technique has been studied as a means to create functional new bladder segments [86]. Various cell sources have been explored for bladder regeneration, but native cells are currently preferable due to their autologous nature, as they can be used without rejection [37]. Human urothelial and muscle cells can be expanded in vitro, seeded onto polymer scaffolds, and allowed to attach and form sheets of cells. Histologic analysis indicates that these cells are able to self-assemble back into their respective tissue types, and they retain their native phenotype [40]. In order to determine the effects of implanting these engineered bladder tissues in continuity with the urinary tract, animal models of bladder augmentation have been used [41]. In one important experiment, partial cystectomies

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were performed in dogs. The animals were divided into two experimental groups. One group had their bladders augmented with a nonseeded bladder-derived collagen matrix, and the second group had their bladder augmented with a cell-seeded construct. The bladders augmented with matrices seeded with cells showed a 100% increase in capacity compared with bladders augmented with cell-free matrices, which showed only a 30% increase in capacity. This study demonstrated a major difference between matrices used with autologous cells (tissue-engineered matrices) and those used without cells [41]. Matrices implanted with cells for bladder augmentation retained most of their implanted diameter, as opposed to matrices implanted without cells for bladder augmentation, in which graft contraction and shrinkage occurred. The results of these initial studies showed that the creation of artificial bladders may be achieved in vivo; however, it could not be determined whether the functional parameters noted were caused by the augmented segment or by the intact native bladder tissue. To better address the functional parameters of tissue-engineered bladders, an animal model was designed that required a subtotal cystectomy with subsequent replacement with a tissue-engineered organ [42]. Cystectomy-only and nonseeded controls maintained average capacities of 22 and 46% of preoperative values, respectively. However, an average bladder capacity of 95% of the original precystectomy volume was achieved when cell-seeded tissueengineered bladder replacements were used. In addition, the subtotal cystectomy reservoirs that were not reconstructed and those reconstructed with the polymer only showed a marked decrease in bladder compliance (10 and 42% total compliance, respectively), but the compliance of the cell-seeded tissue-engineered bladders was almost identical to the preoperative values that were measured when the native bladder was present (106%). Histologically, use of nonseeded scaffolds resulted in a pattern of normal urothelial cells with a thickened fibrotic submucosa and a thin layer of muscle fibers, while the tissue-engineered bladders showed a normal cellular organization, consisting of a trilayer of urothelium, submucosa, and muscle. Immunocytochemical analyses confirmed the muscle and urothelial phenotype. S-100 staining indicated the presence of neural structures [42]. These studies, performed with PGA-based scaffolds, have been repeated by other investigators, showing similar results in large numbers of animals for long term [85, 87]. However, it is important to note that not all types of scaffolds perform well if a large portion of the bladder needs replacement. In a study using SIS for subtotal bladder replacement in dogs, both the unseeded and cell seeded experimental groups showed graft shrinkage and poor results [88]. Thus, the type of scaffold used is critical for the success of tissue engineering technologies. A clinical experience involving engineered bladder tissue for cystoplasty reconstruction was conducted starting in 1998. A small pilot study of seven patients was reported. In some patients, a collagen scaffold seeded with cells obtained from each patient was used to reconstruct the bladder, either with or without omental coverage. In other patients, a combined PGA-collagen scaffold seeded with autologous cells and omental coverage was used (Fig.  15.1). The patients reconstructed with the engineered bladder tissue created with the PGA-collagen cell-seeded scaffolds with

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Fig. 15.1  Construction of a tissue-engineered human bladder. (a) Scaffold material seeded with cells for use in bladder repair. (b) The seeded scaffold is anastamosed to native bladder with running 4–0 polyglycolic sutures. (c) Implant covered with fibrin glue and omentum

omental coverage showed increased compliance, decreased end-filling pressures, increased capacities, and longer dry periods over time [89]. It is clear from this experience that the engineered bladders continued their improvement with time, mirroring their continued development. Although the experience is promising in terms of showing that engineered tissues can be implanted safely, it is just a start in terms of accomplishing the goal of engineering fully functional bladders. Phase 2 studies are now being completed. From the above studies, it is evident that the use of cell-seeded matrices is superior to the use of nonseeded matrices for the creation of engineered bladder tissues. Although advances have been made with the engineering of bladder tissues, many challenges remain. Current research in many centers is aimed at the development of biologically active and “smart” biomaterials that may improve tissue regeneration.

15.3.2 Urethra The urethra is a hollow tube that carries urine from the bladder to the outside of the body. Various strategies have been proposed over the years for the regeneration of urethral tissue. Woven meshes of PGA without cells [90] or with cells [91] have been used to regenerate urethras in various animal models. Naturally derived collagenbased materials such as bladder-derived acellular submucosa [67], and an acellular urethral submucosa [90], have also been tried experimentally in various animal models for urethral reconstruction. The bladder submucosa matrix [67] proved to be a suitable graft for repair of urethral defects in rabbits. The neourethras developed a normal urothelial luminal lining and organized muscle bundles. These results were confirmed clinically in a series of patients with a history of failed hypospadias reconstruction. In this study, urethral defects were repaired with human bladder acellular collagen matrices [92]. The neourethras were created by anastomosing the matrix in an onlay fashion to the urethral plate. The size of the created neourethra ranged from 5 to 15 cm. After a

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Fig. 15.2  Tissue engineering of the urethra using a collagen matrix. (a) Representative case of a patient with a bulbar stricture. (b) During the urethral repair surgery, strictured tissue is excised preserving the urethral plate on the left side, and matrix is anastamosed to the urethral plate in an onlay fashion on the right. The boxes in both photos indicate the area of interest, including the urethra, which appears white in the left photograph. In the left photograph, the arrow indicates the area of stricture in the urethra. On the right, the arrow indicates the repaired stricture. Note that the engineered tissue now obscures the native white urethral tissue in an onlay fashion in the right photograph. (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

3-year follow-up, three of the four patients had a successful outcome in regard to cosmetic appearance and function. One patient who had a 15-cm neourethra created developed a subglandular fistula. The acellular collagen-based matrix eliminated the necessity of performing additional surgical procedures for graft harvesting, and both operative time and the potential morbidity from the harvest procedure were decreased. Similar results were obtained in pediatric and adult patients with primary urethral stricture disease using the same collagen matrices [93] (Fig. 15.2). Another study in 30 patients with recurrent stricture disease showed that a healthy urethral bed (2 or fewer prior urethral surgeries) was needed for successful urethral reconstruction using the acellular collage-based grafts [94]. More than 200 pediatric and adult patients with urethral disease have been successfully treated in an onlay manner with a bladder-derived collagen-based matrix. However, the above techniques, using nonseeded acellular matrices, were applied experimentally and clinically in a successful manner for onlay urethral repairs. In some cases, an entire section of the urethra must be removed, necessitating the insertion of a tubular graft. When such tubularized urethral repairs were attempted experimentally using matrices without cells, adequate urethral tissue regeneration

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was not achieved, and complications ensued, such as graft contracture and stricture formation [95]. To determine whether addition of autologous cells would improve results, autologous rabbit bladder epithelial and smooth muscle cells were grown and seeded onto preconfigured tubular matrices. Entire urethral segments were resected and urethroplasties were performed with tubularized collagen matrices either seeded with cells, or without cells. The tubularized collagen matrices seeded with autologous cells formed new tissue which was histologically similar to native urethra. The tubularized collagen matrices without cells lead to poor tissue development, fibrosis, and stricture formation. These findings were confirmed clinically as well. A clinical trial using tubularized nonseeded SIS for urethral stricture repair was performed in eight evaluable patients. Two patients with short inflammatory strictures maintained urethral patency, but stricture recurrence developed in the other six patients within 3 months of surgery [96]. The reasons for this can be explained by the results of studies on the normal wound healing response. At the time of tissue injury, cell ingrowth is initiated from the wound edges in order to cover the tissue defect. The cells from the edges of the native tissue are able to traverse short distances without any detrimental effects. However, if the wound is larger than a few millimeters in distance or depth, increased collagen deposition, fibrosis, and scar formation ensue. Matrices implanted in wound beds are able to lengthen the distances that cells can traverse but the maximum distance that adjacent cells from the wound edge can travel over a biologic matrix is approximately 1 cm [97]. Tissue defects greater than 1 cm that are treated with a matrix alone, without cells, usually have increased collagen deposition, increased fibrosis, and scar formation. On the other hand, cell-seeded matrices implanted in wound beds are able to further lengthen the distance for normal tissue formation, without initiating an adverse fibrotic response.

15.3.3 Uterus Congenital malformations of the uterus may have profound implications clinically. Patients with cloacal exstrophy and intersex disorders may not have sufficient uterine tissue present for future reproduction. We investigated the possibility of engineering functional uterine tissue using autologous cells [98]. 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 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.

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Fig. 15.3  Appearance of tissue-engineered neo-vaginas. (a) Tubular polymer scaffold after cell seeding and 1 week in  vitro culture, prior to implantation in  vivo. (b, d, and f) indicate gross appearance and (c, e, and g) show vaginography of cell-seeded constructs 1, 3, and 6 months, postimplantation, respectively. (h) Unseeded control scaffold prior to implantation. (i, k and l) gross appearance of unseeded construct at 1, 3, and 6 months postimplantation. (j) Vaginography of unseeded graft at 1 month

15.3.4 Vagina Similarly, 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. Acellular materials have been used experimentally for vaginal reconstruction in rats [99]. The feasibility of engineering vaginal tissue with cells in vivo was also investigated [100]. Vaginal epithelial and smooth muscle cells of female rabbits were harvested, grown, and expanded in culture. These cells were seeded onto biodegradable polymer scaffolds, and the cell-seeded constructs were then implanted into mice. Functional studies in the tissue-engineered constructs showed similar properties to those of normal vaginal tissue. When these constructs were used for autologous total vaginal replacement in a rabbit model, patent functional vaginal structures were noted in the tissue-engineered specimens, while the noncell-seeded structures were noted to be stenotic [101] (Fig.  15.3). These studies indicated that a regenerative medicine approach to clinical vaginal reconstruction would be a realistic possibility. Clinical trials are currently being conducted.

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15.4 Summary and Conclusions Regenerative medicine efforts are currently underway experimentally for virtually every type of tissue and organ within the human body. As regenerative medicine incorporates the fields of tissue engineering, cell biology, nuclear transfer, and materials science, personnel who have mastered the techniques of cell harvest, culture, expansion, transplantation, as well as polymer design are essential for the successful application of these technologies to extend human life. Various tissues are at different stages of development, with some already being used clinically, a few in preclinical trials, and some in the discovery stage. Recent progress suggests that engineered tissues may have an expanded clinical applicability in the future and may represent a viable therapeutic option for those who would benefit from the lifeextending benefits of tissue replacement or repair. Acknowledgment  The author would like to thank Dr. Jennifer Olson for editorial assistance with this manuscript.

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Chapter 16

Engineering Complex Synthetic Organs Joan E. Nichols, Jean A. Niles, and Joaquin Cortiella

Abstract  At this time there is a substantial, and as yet unmet, demand for organs to replace nonfunctional tissues resulting from congenital defects, or to repair damaged or degenerated tissues. The field of regenerative medicine hopes to provide engineered replacement tissues in situations where our body’s regenerative capability or nonbiological mechanical devices cannot adequately replace lost physiological functions. This technology holds the promise to supply customized organs to overcome the severe shortages we currently face. Engineering synthetic organs is a complex process which necessitates careful (1) selection of cells or controlled proliferation of stem or progenitor cells to achieve appropriate numbers of cells for seeding onto biodegradable scaffolds to create cell-scaffold constructs, (2) design and selection of appropriate biodegradable or biomodifiable scaffold materials, and (3) design and construction of bioreactors to support generation of functional tissue replacements. To be successful, ongoing efforts to understand and engineer multicellular systems must continue, and new efforts to induce vascularization and integration of engineered tissues into the body will need to be developed. Current studies lead to improved understanding of how tissue systems can be integrated, as well as development of biomedical technologies not traditionally considered in tissue engineering, such as development of biohybrid organs or “bionic” devices.

J.E. Nichols (*) Laboratory of Regenerative and Nano-Medicine, University of Texas Medical Branch, Galveston, TX, USA Departments of Internal Medicine and Infectious Diseases, University of Texas Medical Branch, Galveston, TX, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_16, © Springer Science+Business Media, LLC 2011

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Abbreviations COPD ECM ESRD FGF-2 FPC mESC PF-127 PGA RAD SLPC SP-A SP-C TTF1

Chronic obstructive pulmonary disease Extracellular matrix End-stage renal disease Fibroblast growth factor-2 Fetal pulmonary cell Murine embryonic stem cell Pluronic F-127 Polyglycolic acid Renal tubule assist device Somatic lung progenitor cell Surfactant protein A Surfactant protein C Thyroid transcription factor-1

16.1 Introduction Variable degrees of success have been achieved in the generation of a variety of synthetic organ systems as a result of determined efforts by multidisciplinary groups in the field of biotechnology, engineering, and regenerative medicine. The design and development of bioengineered tissues/organs has been a slow and stepwise process to produce tissues that possess functional microvasculature, tissue-specific morphology, and specified physiological functions. It is important to realize that our knowledge of the field of regenerative medicine and synthetic organ production is evolving. Just as early scientists made discoveries in the development of biomaterials and techniques to support cell and tissue formation, so we too, will continue to discover new biomaterials, cell sources, and methods of growing functional tissues and organs. This chapter is intended to be a concise summary of the current status of complex synthetic organ production, providing a foundation from which we can expand our understanding of the design and construction of tissues for use in regenerative medicine therapies in the future. There are currently two major approaches to the production of engineered tissues for transplantation. One approach to create complex tissues focuses on building modular microtissue equivalents formed from repeated functional cell sheets or tissue units. This approach of modular tissue engineering focuses on fabricating tissue building blocks with specific microarchitectural and physiologic features and on the use of repeated groupings of such modular units to engineer functional biological tissues. Excellent reviews of this “bottom-up” approach have been published [1] and although there have been some successes in the development of functional tissue constructs through the generation of microtissue equivalents which has allowed us to learn a great deal about organ or tissue-specific physiological functions, development of microequivalents has not yet led to the production of fully biologically

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based synthetic organs. This approach has led to the design and development of biohybrid organs or “bionic” devices or constructs which combine engineering of microtissue equivalents performing specific physiologic functions with mechanical systems which can be implanted as a mechanical medical device, or as a device designed to remain external to the body. Medical devices that utilize nanotechnology may actually be the first step leading to development of artificial organ systems to replace kidney or pancreatic function. The second approach for organ engineering relies on the use of acellular natural scaffolds consisting of the ECM secreted by the resident cells of the tissue or organ from which they are prepared. These biological scaffolds already possess the correct anatomical, chemical, and morphological structure of the natural tissue. Biological scaffold materials composed of ECM have been shown to facilitate the constructive remodeling of many different tissues in both preclinical animal studies and in human clinical applications.

16.2 Kidney The human kidney is a complex organ containing well over two dozen cell types organized into three-dimensional structures containing a number of cell types that interact with each other to form the glomeruli, tubules, and nephrons [2]. The kidney maintains the body’s metabolic and endocrine functions. The most basic structures of the kidneys are the nephrons, which filter the blood, reabsorb water and nutrients, and secrete wastes, producing urine. After injury, these specialized cells are able to regenerate but with advancing age or conditions of severe damage, regeneration becomes limited. For those who have chronic diseases in which the kidney is unable to maintain normal function, the condition is referred to as ESRD. ESRD continues to have an exceedingly high mortality rate, despite advances in dialysis technology. Data from the U.S. Renal Data System show that as of 2008 there are about two million patients on dialysis or who had received a working kidney [3]. Current dialysis therapies also replace only the filtration function and not the critical transport, metabolic, and endocrine functions of the kidney. One of the most important considerations in kidney engineering is the development of artificial membranes suitable for use in bioartificial kidneys [2, 4–7]. Initial approaches for developing kidney replacements were based on replacement of the filtration component of kidney function. These include use of nanotechnology to produce membranes with increased numbers of nanoscale pores to be used in nonbiological renal replacement devices. In systems such as this, stacks of membranes mimic functions of the glomerulus in terms of filtration and of the renal tubules by selectively reclaiming designated solutes [4, 5]. Work continues to develop implantable or wearable dialysis devices with and without the addition of renal cells that provide functions other than filtration. Human nephron function is provided using a filter design that contains two membranes with different pore sizes between them that are capable of filtering solute in a manner similar to the nephrons and renal tubules. Development of artificial membranes that simulate nephron function has been a major

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hurdle in the development of synthetic kidneys or biohybrid devices. A few recent novel approaches include the creation of nanofibers by electro-spinning to create bioactive synthetic membranes [8], and the use of silk porous scaffolds [9]. In one study, fibrous supramolecular membranes were incorporated into electro-spun nanofibers to create a synthetic basement membrane capable of supporting and maintaining renal cellular monolayers. In another approach, mouse embryonic kidney cells grown on silk porous scaffolds containing collagen-Matrigel and maintained in bioreactor culture for over 2 weeks led to the development of kidney-like structures [9]. To engineer any type of tissue there are two main considerations: selection and culture of an appropriate cell population and production of support scaffolds to provide a platform for cell and tissue development. These are critical considerations for the design and construction of synthetic kidneys or biohybrid kidney devices. For cell selection, researchers have already demonstrated that it is possible to engineer kidney tissue using cultured embryogenic kidney precursor cells [2, 4, 5]. One important in vitro study looked at the interaction between the ureteric bud, which is an embryonic tissue, and the metanephric mesenchyme in order to understand how the nephron develops [6]. Using a hanging drop method of culturing cell aggregates from embryonic and adult renal epithelial cells originating from the uteric bud and medullary collecting duct cells, one group was able to show the development of collecting ducts and tubulogenesis. In vivo studies showed that after 5  weeks these implants contained host-derived glomerular architecture [6]. Technology to miniaturize and automate dialysis using micromechanical systems is also currently being developed [7] with researchers working toward the design and development of biologically based biohybrid devices combining culture of renal cells and membrane-based filtration systems. In order to develop an implantable biohybrid kidney device several key factors must be addressed. There must be adequate filtration to support removal of appropriately sized molecules from the blood, combined with the development of cellular components capable of maintaining both the metabolic and endocrine functions of the kidney [2]. A RAD containing living renal proximal tubule cells has been successfully engineered and has demonstrated the differentiated absorptive, metabolic, and endocrine functions of normal kidney in an animal model [10]. The addition of RAD containing human cells to conventional therapies has been shown in preclinical and clinical studies to advance acute kidney therapy, from enhancing renal clearance to true renal replacement therapy. Recent developments in the field of kidney engineering also include the use of decellularized natural kidney matrices which have the appropriate morphology, structure, and ECM signals to support and encourage site-specific differentiation [11]. Age-matched decellularized kidney fragments layered with fetal kidney explants demonstrated the capacity of the decellularized scaffold to support cell attachment and migration by fetal cells from the explanted tissue. This study supports the premise that the decellularized ECM of the kidney retains structural and functional properties necessary to develop a functional kidney. A combination of both natural scaffold and nanoengineered membranes that have appropriate pore size for filtration may be a reasonable strategy to engineer a kidney in the future.

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16.3 Esophagus The majority of esophageal procedures performed in infants and children are to repair congenital defects where portions of the esophagus are missing (esophageal atresia) or to repair acquired caustic scarring that results in strictures [12, 13]. In adults, esophageal atresia, esophageal burns, and cancer are the major conditions requiring esophageal reconstruction [12, 13]. Esophageal replacement using current surgical techniques presents major challenges in both adult and pediatric patients, and is often associated with severe complications [14–16]. The esophagus is composed of three distinct layers: the mucosa, submucosa, and muscularis externa [12]. In the mucosa, the epithelial lining serves as a barrier or protective layer and is constantly damaged and replenished following the mechanical stresses resulting from swallowing food. The submucosa of the esophagus consists of dense connective tissue. Esophageal glands are found in the upper portion of the esophagus and these glands serve to lubricate the lumen. The muscularis externa is composed of skeletal muscle proximally and smooth muscle distally. These extreme variations in tissue structure pose a major challenge for those attempting to engineer replacement of esophageal tissues. There have been many attempts to produce esophageal replacements using scaffolds made of synthetic materials [17], natural biomaterials [18–20], or even decellularized natural scaffolds [21–23], but most have been limited by clinical complications such as stenosis or leakage. Many research groups allow for development of the two major esophageal components, epithelium and muscle, using a hybrid tissueengineering approach where epithelium and muscle cells are seeded together on natural or synthetic polymers to form a composite construct prior to transplantation into animal models. In vivo esophageal repair has been performed with patch or circumferential implantation of synthetic and natural scaffolds. Cell constructs formed using collagen-I or -III sponge matrices layered on a silicone stent were used to replace a 5-cm cervical esophagus in a canine model [18]. Removal of the stents at 4 weeks and development of host-derived esophageal tissue was documented in the implanted collagen scaffolds that supported animal feeding. Unfortunately, when the same scaffolds were used to replace a 5-cm thoracic esophageal segment, there was limited production of the muscularis mucosae layer. Esophageal patch technology [23, 24] has led to the development of circumferential grafts. Circumferential grafts have been associated with complications that range from stricture formation to graft dilation, with little or no muscle regeneration in these implants [18]. It has been suggested that graft coverage is often possible in cervical replacements of the esophagus where, unlike the thoracic cavity, adjacent tissues can provide nutrition through diffusion [25]. Studies have shown that epithelial coverage after circumferential graft placement is a slow process requiring weeks for complete epithelialization of 2–5 cm grafts, and that formation of epithelium impacts muscle regeneration [18]. There have been some recent successes in the design and development of engineered esophagus, including intrathoracic replacement, using a tissue-engineered

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esophagus composed of oral keratinocytes and fibroblasts on a human amniotic membrane scaffold [25]. An explant technique to increase the numbers of isolated rodent esophageal epithelial cells prior to Matrigel application has been used to improve scaffold attachment in  vitro [17]. Enhanced vascularization was seen in scaffolds seeded with rodent esophageal epithelial cells sutured over stents and implanted into the omentum, effectively using it as an in situ bioreactor for epithelium generation [17]. Similar procedures were used in the development of an ovine model that utilized an omental wrap to facilitate vascularization [26]. Attempts have also been made to replace esophagus using conduits of acellular natural biomaterials, such as human dermal matrix, porcine elastin-based aorta patch, and porcine-derived small intestine or urinary bladder submucosa. Decellularized scaffolds such as porcine small intestinal submucosa and urinary bladder submucosa have also been employed as esophageal replacements in large animal models, but formation of strictures was the major complication resulting in severe morbidity in early studies [21, 22]. Later studies demonstrated the production of near normal esophageal tissue and the prevention of stenosis [23, 24].

16.4 Trachea Severe airway stenosis caused by long-term intubation, tracheomalacia, infectious disease, trauma, or neoplasm resistant to chemotherapy and radiotherapy has been managed in the past by tracheal resection [27–29]. The potential extent of tracheal resection or the maximum defect size was first considered to be up to 2 cm [27] but it is currently about half the tracheal length in adults and one third in small children [27]. Despite advances in surgical techniques, resection of airway stenosis or stricture can result in anastomotic tension, tissue ischemia, and failure to heal. Early tracheal reconstruction efforts using cadaveric trachea in children who suffered from congenital tracheal stenosis were promising [30], and the demand for a prosthetic tracheal replacement or engineered trachea has steadily increased over the years. Tracheal replacements must be strong and flexible, and the conduit must be airtight and, if possible, allow for development of a surface of ciliated epithelium. The two main challenges to engineering of tracheal tissues include (1) the need for complete epithelialization and (2) the development of hyaline cartilage that possesses suitable mechanical properties that maintain structural integrity of the airway. Tissue engineered airway prostheses have been under development since the early 1990s using both canine and sheep models. Use of foreign materials to produce a solid tracheal prosthesis have been previously reviewed [29, 31, 32]. Problems regarding the durability and functionality of regenerated cartilage led to the use of porous frameworks or biosynthesized prosthetics. Early work by Kojima and ­colleagues using nasal chondrocytes to produce a bioengineered structure that would be fully absorbed by the body proved disappointing due to the inability of the final product to maintain the mechanical properties required of a functional trachea. This work

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showed that it was possible to bioengineer a replacement trachea, but the type of collagen and extracellular matrix necessary to provide the mechanical support and strength could not be provided by nasal chondrocytes. Artificial tracheas have also been developed using a mixture of sheep nasal chondrocytes and fibroblasts. Sheep receiving trachea grafts survived for 2–7 days after implantation, but collapse of constructs led to tracheal collapse [33–35]. A Marlex mesh tube covered by collagen sponge was used successfully in a canine model as a patch graft, and on the basis of successful experimental results, was later applied to repair the trachea of a 78-year-old woman [35]. The right half of three rings of the trachea was resected and the scaffold was covered with epithelial growth after 2  months, and was followed for 2  years without complications. Several groups have used scaffold-free approaches to trachea design. One of the most successful attempts using this technique involved the use of rabbit auricular chondrocytes to produce biocompatible, autologous scaffold-free collagen sheets that formed a vascularized, tracheal replacement that remained both rigid and flexible after 1 year of implantation [36, 37]. Initially, cadaveric trachea or other tissues that were chemically fixed, frozen, or lyophilized were used in experimental animal and human clinical trials. These nonviable, biologically based prosthetic treatments involved the implantation of nonviable autograft, allograft, or xenograft scaffold materials. The repair of tracheal defects with autogenous tissue, such as periosteal [38], jejunal [39], muscular [40], bronchial [41], aortic [42], and esophageal [43] grafts has met with limited success due to mechanical issues resulting in difficulty maintaining a functional airway. Biologic scaffolds derived from tracheal tissue have been evaluated for use in tracheal replacement due to the mechanical and biochemical similarity of the material to native trachea. Initial use of cryopreserved, lyophilized or glutaraldehydetreated trachea or aortic allograft was promising. Hydrated decellularized porcine tracheal matrix as a scaffold for tracheal reconstruction provided sufficient strength to support a prosthetic construct for 8 weeks [44]. Some tissue development was seen but existing cartilage structures were severely degraded and there was little indication of new generation of cartilage, perhaps due to lack of appropriate tissue vascularization [44, 45]. Small intestinal submucosae have been used as patches to repair small tracheal defects but did not fully support the development of functional tissue [46, 47]. Recent use of urinary bladder ECM scaffold in a canine model provided a much better result, and promoted healing of tracheal defects without development of stricture, and with coverage by ciliated epithelium and formation of dense collagenous tissue [48]. Recently, a cadaveric acellular human trachea seeded with autologous mesenchymal stem cells was successfully transplanted in a woman with left main bronchus stenosis [31, 49]. Current improvements in the decellularization procedure now allow for production of a bioengineered trachea in approximately 3 weeks [32]. Careful examination of these acellular tracheas indicates that both mechanical and architectural components of natural trachea are maintained, as are critical angiogenic factors necessary to promote production of a vascularized tracheal replacement [32].

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16.5 Lung COPD is ranked as the fourth leading cause of death in the USA, affecting over 16 million individuals. The World Health Organization estimated in 2000 that 2.74 million people died of COPD worldwide [50]. Although current medical treatments provide some benefit, surgical interventions in the form of bullectomy or lung volume reduction merely remove diseased or damaged tissue in an attempt at symptomatic relief (i.e., increase ease of breathing). For many, the only option is lung transplantation. Production of functional lung tissue equivalents, lobes, or even whole lungs could be used as therapeutic treatments for COPD, as well as many other pulmonary diseases including pulmonary hypoplasia and severe hyaline membrane disease. A problem that has hampered the development of engineered lung has been the lack of design or production of suitable scaffolding material that supports but does not restrict lung function. The main requirement for any scaffold used in regenerative medicine is biocompatibility of the material, and this is especially true for the lung. Use of materials that do not possess degradation profiles similar to that of normal lung ECM, or that produce immunogenic intermediates, can induce inflammation and result in fibrosis [51, 52]. Of critical importance in scaffold selection is the elasticity of the material. For development of lung tissue, the scaffolding must remain long enough to provide the framework necessary to support cell functions without impeding the elasticity or altering the elastic recoil of the engineered tissue, or other areas in close proximity to the implanted tissue [51, 52]. Restrictive disease that impedes pulmonary function could be exacerbated by scaffolding that lacks the appropriate elasticity or does not sufficiently degrade. The nature of complex organs such as the lung may eventually require the development of hybrid scaffolds formed from more than one material to meet all of the above requirements. Both natural and synthetic polymers have been used in lung tissue engineering, a topic that has been extensively reviewed by us elsewhere [52, 53]. Natural materials include collagen [54–58], Matrigel (BD Biosciences) [59–61], Gelfoam (Pfizer) [61], and Englebreth-Holms tumor basement membrane [58, 62]. Collagen (generally type I) is commercially available and has been used as a scaffold for engineering a variety of  tissues, including lung, and a number of scaffolds based on it are available for clinical use. Matrigel, a scaffold composed of basement membrane proteins is commercially available and has also been used to culture a wide variety of cell types. Gelfoam is a compressed sponge of porcine skin gelatin and was originally developed as a hemostatic device to arrest bleeding and promote clotting. A collagen-matrigel microcapsule has been used to create alveolus-like structures in vitro with promising results, with the development of SP-C-producing alveolar type II cells [63]. There are limitations to the use of natural scaffolds such as these due to their mechanical properties and variation in degradation rates [64]. There is also the possibility that natural scaffolds may be immunogenic and invoke an immune reaction leading to inflammation, or that natural materials may harbor bacteria or viruses if adequate steps have not been taken to ensure the cleanliness of the materials produced.

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Synthetic polymers can be produced with a much wider range of mechanical and chemical properties than that of natural materials. To maintain the elastic nature of the environment, it is generally accepted that a degradable or an extensively biomodifiable material would be the best overall choice for engineering lung tissue. Degradable synthetic matrices that have been used to engineer lung tissue include PGA [53] in the form of a felt sheet, often combined with PF-127 [53], poly(lactic-co-glycolic acid) [53], or poly-l-lactic-acid [53]. PGA degrades by acid hydrolysis to lactic acid and glycolic acid [53]. The degradation rate of PGA is controlled by the molecular weight of the polymers as well as the ratio of glycolic acid to lactic acid subunits. PGA is also a widely used polymer and is better known as the suture material, Dexon (Tyco). Polymers often referred to by the trade name Pluronic (BASF) are liquids at low temperatures (below 15 °C) and gel at higher temperatures such as at body temperature. The firmness or density of the gel increases as the concentration of the hydrogel is increased. Poloxamer hydrogels such as Pluronic, when used in cell culture systems, maintain cells in a 3D structure that enables them to secrete ECM and engage in cell signaling. Currently, there have been few instances of in vivo use of scaffolding material in engineering distal lung tissue. The first was performed in both a large (sheep) and small (nude mouse) animal models by Cortiella and coworkers [51], with ovine SLPC scaffold constructs produced using a combination of PGA and/or PF-127. These constructs were implanted onto either the backs of nude mice or in sheep. Autologous ovine endogenous SLPC constructs were implanted directly into the right upper lobe of the lung into a pocket created by a wedge resection. Implanted constructs were well tolerated and tissue assembly was facilitated in vivo by the use of the synthetic polymer scaffolds. In this same study, autologous SLPC/PGA constructs were implanted into the thoracic cavities of three adult sheep with attachment of the construct to the right main stem bronchus site after a full pneumonectomy. When harvested after 3 months, the implants did not appear to support lung epithelia development but did form soft, fleshy, well-vascularized tissue fragments (Fig. 16.1). Another in vivo use of scaffolds in lung engineering involved the use of Gelfoam delivered into the lung by injection of the constructs (sponge with fetal rat lung-derived cells) directly into lung parenchyma [65]. The sponge degraded over several months and was also well tolerated. Most of the newly formed alveolar-like structures, however, were found close to the border between the sponge and the surrounding normal tissue with few found within the sponge itself. It is unclear why this occurred because the porosity of the Gelfoam scaffold should have provided cells with an adequate environment to support cell movement and tissue development. Mondrinos and colleagues also developed constructs using FPCs and a Matrigel scaffold [59]. The FPC/Matrigel construct was injected subcutaneously into the anterior abdominal wall of a mouse. Matrigel was shown to support both the development of lung epithelia and vascularization of the construct. Although each of these materials was adequate for the development of tissue, the degradation of the scaffold material is also an important consideration. Andrade and

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Fig. 16.1  (a, b) In vivo tissue-engineered ovine lung using SLPC/PGA construct. Surgery showing implantation of construct after a pneumonectomy at the right main stem bronchus site (arrow shows SLPC/PGA construct implanted at pneumonectomy site). (c) Fleshy tissue growth produced after 3 months. (d) Sections of tissue stained with hematoxylin and eosin; magnification 400×

colleagues [65] demonstrated that Gelfoam was an excellent supporting material for lung cell attachment, and the timely degradation of the material left the newly formed alveoli in place once degradation was complete. This has also been true of PGA/PF-127 in both in vitro and in vivo use of this combination matrix, which also degraded as the development of epithelial tissues progressed [66]. FGF-2-loaded Matrigel produced highly vascularized tissue with few structures reminiscent of alveolar forming units [59], possibly due to either the structure of the scaffold or slow degradation of the scaffold material. Recently, decellularized natural lung has been used by several groups for in vitro and in vivo engineering of lung tissues [66–68]. Each of these groups took a slightly different approach to the process of producing and recellularizing natural acellular rat scaffolds. Use of perfusion decellularization techniques significantly reduced the time required for decelluarization of the rat lungs [67, 68] over bioreactor-based techniques [66], although all methods slightly altered the ECM composition compared to natural lung. Culture times for recellularization of lung scaffolds were significantly

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Fig. 16.2  (a) Freshly isolated rat lung prior to decellularization. (b) Bioreactor showing chamber containing lung in 1% SDS. (c) Fully decellularized lung being removed from the bioreactor chamber. (d) Acellular lung prior to reseeding with C57B6 murine embryonic stem cells. Photographs by Kenneth D. Frohne

different between these studies, and although there was some indication that lung tissue was developed by 8 [68] or 9 [67] days of bioreactor culture, significantly better scaffold coverage and indications of mature, complex lung and endothelial tissue formation was seen after 14 or 21 days of culture [66]. One approach to construction of engineered lung used whole acellular lung from Sprague–Dawley rats derived using a bioreactor-based decellularization (Fig. 16.2) process were used as a scaffold to support lung lineage differentiation, and tissue development by mESCs [66]. Comparisons were made related to the influence of acellular lung, Gelfoam, Matrigel, and a collagen I hydrogel matrix on mESC attachment, differentiation, and subsequent formation of complex tissue. In these studies, acellular lung allowed for better retention of cells with more differentiation of mESCs into epithelial and endothelial lineages. In 14- or 21-day constructs produced on whole acellular lung, there were indications of organization of differentiating mESCs into three-dimensional structures reminiscent of complex tissues, with expression of TTF1, an immature lung epithelial cell marker; pro-SP-C and

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Fig. 16.3  (a) Control for b. (b) Tracheal region of recellularized lung showing cytokeratin positive ciliated cells found in patches throughout the trachea. (c) Higher magnification of boxed area in (b). (d) Control for e. (e) Expression of Clara Cell Protein 10 (CC10; red granules) and cytokeratin (green). (f) Control for (g). (g) Alpha-actin expresion by cells underlying regions of CC10 positive cells in the trachea. Magnification 630× in a, b, d–g

SP-A, type II pneumocyte products; PECAM-1/CD31, an endothelial cell marker; cytokeratin 18; a-smooth muscle actin; CD140a or platelet-derived growth factor receptor-a; and Clara cell protein 10 [66]. There was also evidence of site-specific differentiation in the trachea with the formation of sheets of cytokeratin-positive cells (Figs. 16.3 and 16.4), a-smooth muscle actin-positive cells, and Clara cell protein 10-expressing cells in the trachea (Fig.  16.3), and production of type II pneumocytes as indicated by production of SP-C (Fig. 16.3c). Differences in ECM identified in trachea and distal lung using two-photon microscopy, included a rigid, dense, parallel deposition of collagen in trachea (Fig. 16.4a), and more wavy, less dense collagen formation in the distal lung (Fig.  16.4e). These data support the utility of acellular lung as a matrix for engineering lung tissue, and highlight the critical role played by matrix and scaffold-associated cues in guiding ESC differentiation toward lung-specific lineages [66]. In a separate study, Sprague–Dawley rat matrix was produced by perfusion decellularization. Acellular lung was then seeded with human umbilical vein endothelial cells and rat fetal lung, and cultured for up to 9  days [67]. Examination of the engineered lungs showed interspersed type II pneumocytes, confirmed by the expression of SP-A, SP-C, and TTF1 (also known as Nkx2-1) [67]. Most TTF1positive cells were found localized to luminal spaces bordered by rings of cuboidal cells expressing nuclear TTF1 protein. The defining functional property of lung tissue is the capacity to allow gas exchange between circulating blood and inhaled air. By day 5, constructs could be perfused with blood and ventilated using physiologic pressures, and they generated gas exchange comparable to that of isolated

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Fig. 16.4  (a, e) Two-photon microscopy using autofluorescence and second harmonic generation to visualize fibrilar collagen and relative extracellular matrix makeup of decellularized rat lung. Images of trachea (a–d) and distal lung (e–h). (b–d, f–h) Recellularized lung after 21 days of bioreactor culture. (b) Phase contrast image of differentiated mESCs in the trachea showing sheets of cells lining the trachea; magnification 100×. (c) Isotype control for (d). (d) Confocal analysis of expression of cytokeratin-18 in cells lining the trachea. (f) Isotype control and confocal analysis of expression of pro-SP-C (g) and SP-A (h) (red) in distal lung; magnification 630×

native lungs. Implantation of engineered lungs was performed following left lung pneumonectomy. Blood gas analysis performed 6 hours after transplantation, with rats breathing room air, revealed significantly higher arterial blood oxygen tension in recipients of regenerated lungs compared to pneumonectomized controls. Histological evaluation revealed blood perfused vasculature without evidence of interstitial hematoma, airway bleeding, or thrombus formation, but with the presence of proteinaceous fluid in alveolar spaces and distal airways, suggestive of pulmonary edema [67]. Finally, neonatal rat epithelial cells were used to repopulate acellular lung scaffolds produced by perfusion decellularization from adult Fisher 344 rats [68]. A novel bioreactor was used to culture pulmonary epithelium and vascular endothelium on the acellular lung scaffold for up to 8 days. The engineered epithelium displayed hierarchical organization within the scaffold, and endothelial cells were shown to repopulate the vascular compartments. In vitro, the mechanical characteristics of the engineered lungs were similar to those of native lung tissue, and when implanted into rats for short time intervals (45–120  min), the engineered lungs participated in limited gas exchange [68]. Although these results represent an initial step toward the ultimate goal of generating fully functional lungs, they confirm that repopulation of acellular lung scaffold is a viable strategy for lung regeneration.

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16.6 Conclusions Clearly, engineering a complex organ such as lung or kidney presents so many challenges that development of clinically applicable replacement tissues have not yet been realized. Problems to be faced in the development of any complex tissue, including those mentioned above, depend on the development of better systems to promote angiogenesis, the selection of tissue- or organ-appropriate cell sources, the reproducible differentiation of selected cell types along organ-specific lineages, and the development of appropriate scaffolds or matrices to enhance and support 3D production of tissues and subsequent physiologic functions. Because complex organs are composed of more than one cell type, each with specific tasks, an understanding of the factors involved in the physiological functions to be reproduced by a synthetic organ is essential. Another obstacle in the engineering of any tissue for clinical application is selecting human cell sources with the potential to provide sufficient numbers of cells for the development of biosynthetic tissues or organs. While a great deal of progress has been made in developing many of these complex tissues, it is important to understand aspects of the anatomy and physiology of the normal organs as well as of the diseased or damaged tissues to be replaced before considering the design process to engineer tissue or organ replacements. Development of new materials designed to meet the anatomic and physiologic needs of specific organs must occur before we can begin to realize the goal of engineering functional synthetic organs. Better understanding of factors promoting cell adhesion, migration, differentiation, and vascularization of grafts and tissue regeneration as a whole are also needed. Advances in the development of mathematical models to examine the conditions that promote tissue morphogenesis and tissue growth for computational investigations of tissue development will also be necessary if we are to realistically evaluate the production of synthetic organs strictly from the engineering perspective. It is obvious that engineering synthetic organs or biohybrid devices will require a multidisciplinary approach if we are to eventually succeed in our attempts to construct synthetic organs or biohybrid assist devices worthy of clinical application in the future. Acknowledgment  The authors would like to thank Kenneth D. Frohne for his photographic and editorial assistance during the preparation of this manuscript.

References 1. Nichol JW, Khademhosseini A (2009) Modular tissue engineering: engineering biological tissues from the bottom up. Soft Matter 5(7):1312–1319 2. Perin L, Giuliani S, Sedrakyan S, Da Sacco S, De Filippo R (2008) Stem cell and regenerative science application in the development of bioengineering of renal tissue. Pediatr Res 63:467–471 3. US Renal Data System. Annual data report. http://www.usrds.org/atlas.htm. Accessed 6 Aug 2011 4. Nissenson AR (2009) Bottom-up nanotechnology: the human nephron filter. Semin Dialysis 22(6):661–664

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5. Reisner Y (2007) Growing organs for transplantation from embryonic precursor tissues. Immunol Res 38:261–273 6. Rosines E, Johkura K, Zhang X, Schmidt HJ, DeCambre M, Bush KT, Nigam SK (2010) Constructing kidney-like tissues from cells based on programs for organ development: towards a method of in vitro tissue engineering of the kidney. Tissue Eng Part A 16(8):2441–2455 7. Fissell WH, Roy S (2009) The implantable artificial kidney. Semin Dialysis 22(6):665–670 8. Dankers PYW, Roomker JM, Huizinga-van der Vlag A, Wisse E, Appel WPJ, Smedts FMM, Hamsen MC, Bosman AW, Meijer W, van Luyn MJA (2010) Bioengineering of living renal membranes consisting of hierarchical, bioactive supramolecular meshes and human tubular cells. Biomaterials 32(3):723–733. doi:10.1016/j.biomaterials.2010.09.020 9. Subramanian B, Rudym D, Cannizzaro C, Perrone R, Zhou J, Kaplan DL (2010) Tissueengineered three-dimensional in  vitro models for normal and diseased kidney. Tissue Eng Part A 16(9):2821–2831 10. Humes HD, Buffington DA, MacKay SM, Funke AJ, Weitzel WF (1999) Replacement of a renal function in uremic animals with a tissue-engineered kidney. Nat Biotechnol 17:451–455 11. Nakayama KH, Batchelder CA, Lee CI, Tarantal AF (2010) Decellularized rhesus monkey kidney as three-dimensional scaffold for renal tissue engineering. Tissue Eng Part A 16(7):2207–2216 12. Squier CA, Kremer MJ (2001) Biology of oral mucosa and esophagus. J Natl Cancer Inst Monogr 29:7–15 13. Pace F, Antinori S, Repici A (2009) What is new in esophageal injury (infection, drug-induced, caustic, stricture, perforation)? Curr Opin Gastroenterol 25(4):372–379 14. Davenport M, Hosie GP, Tasker RC, Gordon I, Kiely EM, Spitz L (1996) Long-term effects of gastric transposition in children: a physiological study. J Pediatr Surg 31(4):588–593 15. Erdogan E, Emir H, Eroglu E, Danişmend N, Yeker D (2000) Esophageal replacement using the colon: a 15-year review. Pediatr Surg Int 16(6):546–549 16. Borgnon J, Tounian P, Auber F, Larroquet M, Boeris Clemen F, Girardet JP, Audry G (2004) Esophageal replacement in children by an isoperistaltic gastric tube: a 12-year experience. Pediatr Surg Int 20(11–12):829–833 17. Beckstead BL, Pan S, Bhrany AD, Bratt-Leal AM, Ratner BD, Giachelli CM (2005) Esophageal epithelial cell interaction with synthetic and natural scaffolds for tissue engineering. Biomaterials 26(31):6217–6228 18. Takimoto Y, Nakamura T, Yamamoto Y, Kiyotani T, Teramachi M, Shimizu Y (1998) The experimental replacement of a cervical esophageal segment with an artificial prosthesis with the use of collagen matrix and a silicone stent. J Thorac Cardiovasc Surg 1169(1):98–106 19. Sato M, Ando N, Ozawa S et al (1994) An artificial esophagus consisting of cultured human esophageal epithelial cells, polyglycolic acid mesh, and collagen. ASAIO J 40(3):M389–M392 20. Natsume T, Ike O, Okada T, Shimizu Y, Ikada Y, Tamura K (1990) Experimental studies of a hybrid artificial esophagus combined with autologous mucosal cells. ASAIO Trans 36(3):M435–M437 21. Badylak S, Meurling S, Chen M, Spievack A, Simmons-Byrd A (2000) Resorbable bioscaffold for esophageal repair in a dog model. J Pediatr Surg 35(7):1097–1103 22. Badylak SF, Vorp DA, Spievack AR et al (2005) Esophageal reconstruction with ECM and muscle tissue in a dog model. J Surg Res 128(1):87–97 23. Nieponice A, Gilbert TW, Badylak SF (2006) Reinforcement of esophageal anastomoses with an extracellular matrix scaffold in a canine model. Ann Thorac Surg 82(6):2050–2058 24. Nieponice A, McGrath K, Qureshi I, Beckman EJ, Luketich JD, Gilbert TW, Badylak SF (2009) An extracellular matrix scaffold for esophageal stricture prevention after circumferential EMR. Gastrointest Endosc 69(2):289–296 25. Nakase Y, Nakamura T, Kin S, Nakashima S, Yoshikawa T, Kuriu Y, Sakakura C, Yamagishi H, Hamuro J, Ikada Y, Otsuji E, Hagiwara A (2008) Intrathoracic esophageal replacement by in situ tissue-engineered esophagus. J Thorac Cardiovasc Surg 136(4):850–859 26. Saxena AK, Baumgart H, Komann C, Ainoedhofer H, Soltysiak P, Kofler K, Höllwarth ME (2010) Esophagus tissue engineering: in situ generation of rudimentary tubular vascularized esophageal conduit using the ovine model. J Pediatr Surg 45(5):859–864

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2 7. Grillo HC (2002) Tracheal replacement: a critical review. Ann Thorac Surg 73:1995–2004 28. Grillo HC (1994) Slide tracheoplasty for long-segment congenital tracheal stenosis. Ann Thorac Surg 58:613–621 29. Weber TR, Connors RH, Tracy TF Jr (1991) Acquired tracheal stenosis in infants and children. J Thorac Cardiovasc Surg 102:29–35 30. Elliott MJ (1967) Prosthetic replacement of the trachea. Ann Thorac Surg 4:1–11 31. Bader A, Macchiarini P (2010) Moving towards in situ tracheal regeneration: the bionic tissue engineered transplantation approach. J Cell Mol Med 14(7):1877–1889 32. Kalathur M, Baiguera S, Macchiarini P (2010) Translating tissue-engineered tracheal replacement from bench to bedside. Cell Mol Life Sci 67(24):4185–4196 33. Kojima K, Bonassar LJ, Roy AK, Vacanti CA, Cortiella J (2002) Autologous tissue-engineered trachea with sheep nasal chondrocytes. J Thorac Cardiovasc Surg 123(6):1177–1184 34. Kojima K, Bonassar LJ, Ignotz RA, Syed K, Cortiella J, Vacanti CA (2003) Comparison of tracheal and nasal chondrocytes for tissue engineering of the trachea. Ann Thorac Surg 76(6):1884–1888 35. Omori K, Nakamura T, Kanemaru S, Asato R, Yamashita M, Tanaka S, Magrufov A, Ito J, Shimizu Y (2005) Regenerative medicine of the trachea: the first human case. Ann Otol Rhinol Laryngol 114(6):429–433 36. Weidenbecher M, Tucker HM, Gilpin DA, Dennis JE (2009) Tissue-engineered trachea for airway reconstruction. Laryngoscope 119(11):2118–2123 37. Gilpin DA, Weidenbecher MS, Dennis JE (2010) Scaffold-free tissue-engineered cartilage implants for laryngotracheal reconstruction. Laryngoscope 120(3):612–617 38. Cohen RC, Filler RM, Konuma K, Bahoric A, Kent G, Smith C (1986) A new model of tracheal stenosis and its repair with free periosteal grafts. J Thorac Cardiovasc Surg 92:296–304 39. Jones RE, Morgan RF, Marcella KL, Mills SE, Kron IL (1986) Tracheal reconstruction with autogenous jejunal microsurgical transfer. Ann Thorac Surg 41:636–638 40. Papp C, McCraw JB, Arnold PG (1985) Experimental reconstruction of the trachea with autogenous materials. J Thorac Cardiovasc Surg 90:13–20 41. Murakami S, Sato H, Yokoi K, Tunezuka N, Hayashi Y, Shimizu J, Watanabe Y (1994) An experimental study of tracheal reconstruction using a freed piece of the right bronchus. Thorac Cardiovasc Surg 42:76–80 42. Carbognani P, Spaggiari L, Solli P, Corradi A, Cantoni AM, Barocelli E et al (1999) Experimental tracheal transplantation using a cryopreserved aortic allograft. Eur Surg Res 31:210–215 43. Kato R, Onuki AS, Watanabe M, Hashizume T, Kawamura M, Kikuchi K, Kobayashi K, Ishihara T (1990) Tracheal reconstruction by esophageal interposition: an experimental study. Ann Thorac Surg 49:951–954 44. Remlinger NT, Czajka CA, Juhas ME, Vorp DA, Stolz DB, Badylak SF, Gilbert S, Gilbert TW (2010) Hydrated xenogeneic decellularized tracheal matrix as a scaffold for tracheal reconstruction. Biomaterials 31(13):3520–3526 45. Nakanishi R (2009) Cryopreservation of the tracheal grafts: review and perspective. Organogenesis 5(3):113–118 46. Gubbels SP, Richardson M, Trune D, Bascom DA, Wax MK (2006) Tracheal reconstruction with porcine small intestine submucosa in a rabbit model. Otolaryngol Head Neck Surg 134:1028–1035 47. Park JW, Pavcnik D, Uchida BT, Timmermans H, Corless CL, Yamakado K, Yamada K, Keller FS, Rosch J (2000) Small intestinal submucosa covered expandable Z stents for treatment of tracheal injury: an experimental pilot study in swine. J Vasc Interv Radiol 11:1325–1330 48. Gilbert TW, Gilbert S, Madden M, Reynolds SD, Badylak SF (2008) Morphologic assessment of extracellular matrix scaffolds for patch tracheoplasty in a canine model. Ann Thorac Surg 86(3):967–974 49. Macchiarini P, Walles T, Biancosino C, Mertsching H (2004) First human transplantation of a bioengineered airway tissue. J Thorac Cardiovasc Surg 128(4):638–641 50. Hurd S (2000) The impact of COPD on lung health worldwide. Chest 117(2):1S–4S

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51. Cortiella J, Nichols JE, Kojima K, Bonassar LJ, Dargon P, Roy AK, Vacanti MP, Niles JA, Vacanti CA (2006) Tissue-engineered lung: an in vivo and in vitro comparison of polyglycolic acid and pluronic F-127 hydrogel/somatic lung progenitor cell constructs to support tissue growth. Tissue Eng 12:1213–1225 52. Nichols JE, Niles JA, Cortiella J (2009) Design and development of tissue engineered lung: progress and challenges. Organogenesis 5(2):57–61 53. Nichols JE, Cortiella J (2008) Engineering of a complex organ: progress toward development of a tissue-engineered lung. Proc Am Thorac Soc 5(6):723–730 54. Sugihara H, Toda S, Miyabara S, Fujiyama C, Yonemitsu N (1993) Reconstruction of alveoluslike structure from alveolar type II epithelial cells in three-dimensional collagen gel matrix culture. Am J Pathol 142(3):783–792 55. Chakir J, Pagé N, Hamid Q, Laviolette M, Boulet LP, Rouabhia M (2001) Bronchial mucosa produced by tissue engineering: a new tool to study cellular interactions in asthma. J Allergy Clin Immunol 107:36–40 56. Mondrinos MJ, Koutzaki S, Lelkes PI, Finck CM (2007) A tissue-engineered model of fetal distal lung tissue. Am J Physiol Lung Cell Mol Physiol 293(3):L639–L650 57. Chen P, Marsilio E, Goldstein RH, Yannas IV, Spector M (2005) Formation of lung alveolar-like structures in collagen-glycosaminoglycan scaffolds in vitro. Tissue Eng 11(9–10):1436–1448 58. Blau H, Guzowski DE, Siddiqi ZA, Scarpelli EM, Bienkowski RS (1988) Fetal type 2 pneumocytes form alveolar-like structures and maintain long-term differentiation on extracellular matrix. J Cell Physiol 136:203–214 59. Mondrinos MJ, Koutzaki SH, Poblete HM, Crisanti MC, Lelkes PI, Finck CM (2008) In vivo pulmonary tissue engineering: contribution of donor-derived endothelial cells to construct vascularization. Tissue Eng 14:361–368 60. Mondrinos MJ, Koutzaki S, Jiwanmall E, Li M, Dechadarevian JP, Lelkes PI, Finck CM (2006) Engineering three-dimensional pulmonary tissue constructs. Tissue Eng 12:717–728 61. Lin YM, Zhang A, Rippon HJ, Bismark A, Bishop AE (2010) Tissue engineering of lung: the effect of extracellular matrix on the differentiation of embryonic stem cells to pneumocytes. Tissue Eng Part A 16(5):1515–1526 62. Shannon JM, Mason RJ, Jennings SD (1987) Functional differentiation of alveolar type II epithelial cells in vitro: effects of cell shape, cell-matrix interactions and cell-cell interactions. Biochim Biophys Acta 931:143–156 63. Zhang WJ, Lin QX, Zhang Y, Liu CT, Qui LY, Wang HB, Duan CM, Liu ZQ, Zhou J, Wang CY (2010) The reconstruction of lung alveolus-like structure in collagen-matrigel/microcapsules scaffolds in vitro. J Cell Mol Med Oct 3. doi:10.1111/j.1582-4934.2010.01189.x 64. Lavik E, Langer R (2004) Tissue engineering: current state and perspectives. Appl Microbiol Biotechnol 65:1–8 65. Andrade CF, Wong AP, Waddell TK, Keshavjee S, Liu M (2007) Cell-based tissue engineering for lung regeneration. Am J Physiol Lung Cell Mol Physiol 292(2):L510–L518 66. Cortiella J, Niles J, Cantu A, Brettler A, Pham A, Vargas G, Winston S, Wang J, Walls S, Nichols JE (2010) Influence of acellular natural lung matrix on murine embryonic stem cell differentiation and tissue formation. Tissue Eng Part A 16(8):2565–2580 67. Ott HC, Clippinger B, Conrad C, Schuetz C, Pomerantseva I, Ikonomou L, Kotton D, Vacanti JP (2010) Regeneration and orthotopic transplantation of a bioartificial lung. Nat Med 16(8):927–933 68. Petersen TH, Calle EA, Zhao L, Lee EJ, Gui L, Raredon MB, Gavrilov K, Yi T, Zhuang ZW, Breuer C, Herzog E, Niklason LE (2010) Tissue-engineered lungs for in vivo implantation. Science 329(5991):538–541

Chapter 17

Liver Regeneration and Tissue Engineering Ji Bao, James Fisher, and Scott L. Nyberg

Abstract  The liver is one of the largest and most complex organs in the human body. It is a vital organ weighing about 1,500 g and it continuously performs over 500 different functions. As a result, the liver is a vital organ and liver failure in the absence of liver transplantation often results in death to the patient. The shortage of donor livers for transplantation, the demand from industry to develop new drugs and new systems to test their safety, along with the need to better understand the many metabolic pathways of the liver, have been the major driving forces behind liver tissue engineering and advances to create livers synthetically, either in culture or ex vivo or by using animals as in vivo incubators. The liver is also a privileged organ in its ability to regenerate spontaneously in response to acute injury. For these reasons, it has been a major focus of research in tissue engineering and regenerative medicine. However, the ideal source of liver cells (hepatocytes) for synthetic livers has not yet been identified, and numerous research efforts are underway. The ultimate goal of these efforts is to produce an abundant, high quality and readily available source of primary human hepatocytes or synthetic liver tissue constructs for discovery, therapeutic, and diagnostic applications. This chapter explores many recent advances in the field of liver regeneration and liver tissue engineering, and new areas of research and future development.

Abbreviations BAL ECM ESC

Bioartificial liver Extracellular matrix Embryonic stem cell

S.L. Nyberg (*) Mayo Clinic, Rochester, MN, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_17, © Springer Science+Business Media, LLC 2011

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Induced pluripotent stem cell Poly (dl-lactide-co-glycolide) Poly (l-lactic acid) Spheroid reservoir bioartificial liver

17.1 Introduction The liver is the largest internal organ within the body, accounting for about onefiftieth of body weight in the adult and around one-twentieth of the weight in a neonate. The liver serves a central role in metabolic homeostasis consisting of more than 500 functions, including the metabolism, synthesis, storage and redistribution of nutrients, carbohydrates, fats, and vitamins. The liver is the main site of protein synthesis, which produces large numbers of serum proteins including albumin, acute-phase proteins, enzymes, and cofactors. In addition, the liver is the most important detoxifying organ of the body inactivating toxins and xenobiotics absorbed by the intestine and removing wastes from the blood by metabolic conversion and biliary excretion. The human liver has a remarkable capacity to regenerate after physical or toxic injury, returning to its original mass even if less than 20% of the original cell number remains [1]. Liver failure causes 40,000 deaths annually and it is the eighth most frequent cause of death in the USA [2]. Currently, liver transplantation is the only clinically proven effective cure for patients with liver failure; unfortunately, less than 7,000 organs are available per year in the USA. The scarcity of donor organs is a major limitation of liver transplantation. This organ shortage has spurred research utilizing novel hepatocyte-based therapeutic options for end-stage liver diseases, which offer the potential to augment or replace whole-organ transplantation. The field of liver-tissue engineering includes several approaches to fulfill these aims. These approaches include transplantation of individual hepatocytes or implantable liver tissue constructs, partial liver transplantation, and/or extracorporeal BAL systems. These therapies also offer the potential to extend the life of those awaiting transplantation. Hepatocyte culture technology is also used for pharmaceutical industry screening of the effects of new drugs prior to animal and human studies [3, 4], as well as elucidating fundamental characteristics of liver biology. The development of hepatic tissue engineering poses unique challenges stemming largely from the complexity of liver structure and function. The current state-of-theart in each of these aspects is reviewed in this chapter. We discuss liver architecture and functions as they relate to liver cell culture techniques in vitro, as well as relevant critical issues pertaining to cell culture, cell sources, bioreactor design, implantable hepatocyte-based devices, and animal models.

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17.2 Liver Structure and Function 17.2.1 Structure The liver is the largest gland in the human body, accounting for 2% of the weight of an adult (approximately 1,500 g). It lies in the right upper quadrant of the abdominal cavity just inferior and attached to the diaphragm. Anatomically, the liver is divided into two lobes (right and left) which are further divided into eight segments based on independent vascular and biliary supplies. Blood is supplied to the liver by the portal vein (80%) and hepatic artery (20%). Bile drains from the liver via right and left hepatic ducts and is stored and concentrated in the gallbladder. To engineer an optimal microenvironment for hepatocytes in  vitro, one must consider the normal architecture of the liver. On a microstructural level, liver lobes contain repeating functional units called lobules. The lobule is centered on a central vein surrounded peripherally by portal triads containing portal venules, arterioles, and bile ductules as illustrated in Fig. 17.1. The intrahepatic circulation consists of sinusoids, specialized capillaries of fenestrated liver sinusoidal endothelial cells that are separated from the hepatocyte compartment by a thin reticular basement membrane region termed the space of Disse. The hepatocytes surrounding the sinusoid are arranged in unicellular plates. Blood flows through the sinusoid at an average flow rate of 144  mm/s delivering approximately 2,000 nmol/mL of oxygen to the surrounding shell of hepatocytes [5]. Bile canaliculi are located on the lateral surface of adjoining hepatocytes (opposite to the surface in contact with the space of Disse). These canaliculi allow for bile excretion from the hepatocytes to bile ductules formed by biliary ductal cells.

17.2.2 Function The main functional cell type of the liver is the parenchymal cell, or hepatocyte, constituting approximately 80% of the liver mass. They extract and process nutrients and other materials from the blood including the glucuronidation of bilirubin allowing for secretion of bile into the intestinal tract, the synthesis of albumin, complement proteins, glycogen (allowing for storage of glucose) and coagulation proteins, and detoxification of ammonia via conversion to urea. The liver also serves as a core role for carbohydrate, lipid, and amino acid regulation. Hepatocytes synthesize a wide variety of liver-specific enzymes that carry out the many synthetic functions and metabolize drugs and toxins from the gut via the portal circulation. The other 20% of cellular mass is comprised of the nonparenchymal cells that modulate hepatocyte function. These nonparenchymal cells include stellate cells,

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Fig. 17.1  Schematic of the hepatic sinusoid demonstrating microscopic orientation of major vascular, biliary, and cellular structures of normal liver

cholangiocytes (bile duct cells), sinusoidal endothelial cells, and Kupffer cells. Kupffer cells are the liver resident macrophage, which reside in the sinusoids and are essential for the phagocytosis of bloodborne toxicants and particles, as well as the production of cytokines. Stellate cells serve as vitamin A and other fat-soluble vitamin-storing pericytes. These stellate cells are found well positioned within the sinusoids for communication with hepatocytes. Stellate cells are also the main ECMproducing cell in the liver, synthesizing collagen when activated.

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17.3 In Vitro Hepatic Culture Techniques 17.3.1 Hepatocyte Isolation The first method for isolation of hepatocytes from the liver involved collagenase perfusion in laboratory animals. This method was first described by Berry and Friend in 1969 [6] and further refined by Seglen in 1976 [7]. The Seglen technique relies on a two-step in situ collagenase perfusion, followed by mechanical segregation of the tissue, and a purification step based on cell density. Isolation in this manner allows for high viability (>90%) and, due to the high density of hepatocytes, a relatively pure population (>95%). Following this methodology, 100–500 million hepatocytes can be routinely isolated from a single rat liver.

17.3.2 Hepatocyte Cryopreservation Cryopreservation of hepatocytes can potentially avoid the costs of long-term hepatocyte culture, and reduces the risk of contamination during prolonged cell culture. However, hepatocytes are highly susceptible to a loss of viability and functionality by the freeze–thaw process. In an effort to counteract the detrimental effects of the freeze–thaw process, research has identified methods of improving hepatocyte viability and functionality following cryopreservation. Inhibition of caspase proteins [8], storage in University of Wisconsin solution [9], and encapsulation have been shown to reduce injury of hepatocytes during the cryopreservation process [10, 11].

17.4 Ex Vivo Hepatocyte Culture Techniques Isolated primary hepatocytes show very little capacity for proliferation ex vivo, no matter which architectural configuration is employed. The most significant problem with cultured hepatocytes is that they rapidly lose their differentiated structures and liver-specific functions following isolation.

17.4.1 Small Scale The most common primary hepatocyte culture technique is to seed the cells as a single layer on collagen gel-coated dishes in conditioned medium. When primary hepatocytes are cultured on a single collagen layer, they produce albumin and urea, and show cytochrome P450 activity, but their liver-specific functions steadily decline within the first week of culture. To mimic the matrix surrounding the hepatocytes in

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the sinusoid, a second layer of collagen is added on top of the cultured hepatocytes, termed a collagen “sandwich” configuration [12]. This scheme maintains hepatocyte function, polarity and induces distinct apical and lateral membrane formation. In an effort to recreate the interactions between parenchymal and nonparenchymal cells, several liver nonparenchymal cells such as fibroblasts, stellate cells, Kupffer cells, and endothelial cells have been co-cultured with hepatocytes and showed remarkable liver-like structure and function. Currently, high-throughput experimental systems, hepatocyte culture microarrays, are available to meet the needs for genome-wide screening or large pharmaceutical and chemical library characterizations, which allow parallel control of environmental or chemical stimuli while measuring hepatocyte responses. The most common four kinds of microarrays include (1) microtiter plates, the current industry standard for high-throughput drug screening; (2) high-density spotted arrays used to study small-molecule libraries; (3) microfabricated arrays that are ideal for creating single-cell and multicell colonies of controlled shape; and (4) microfluidics, ideal for creating high-throughput closed-volume perfusion bioreactors [13].

17.4.2 Large Scale: Bioreactor Generating a system for culturing large numbers of hepatocytes may be required for clinical applications such as the development of a BAL or hepatic tissue engineering. These systems include the flat membrane configuration, the hollow fiber system, encapsulation technology, and the use of cell aggregates. 17.4.2.1 Flat Membrane System The use of a flat membrane bioreactor allows for control of the internal flow distribution and perfusion of all hepatocytes under a stable oxygen and hormone gradient in  vitro. Hepatocytes cultured using this technique demonstrated specific in  vivo zonal differentiation characteristics such as the expressions of phosphoenolpyruvate carboxykinase in the upstream oxygen-rich region, and cytochrome P450 2B in the downstream oxygen-poor region. Recently, this system was applied to study the effects of acetaminophen toxicity on metabolically zoned hepatocytes [14]. Two main drawbacks of this configuration are the potentially large unused volume and the low surface area-to-volume ratio. With regard to an extracorporeal liver support system, it would be especially difficult to build a system containing a sufficient cell concentration utilizing this culturing technique. 17.4.2.2 Hollow Fiber System Most devices tested clinically employ hollow fiber cartridges containing either porcine hepatocytes or human hepatoblastoma cells. The cells in these devices are separated

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from blood or plasma by a semipermeable membrane with a defined molecular weight cutoff. In most cases, hepatocyte aggregates or hepatocyte-seeded microcarriers are placed on the outside of hollow fibers while oxygenated blood, plasma, or culture medium flow through the hollow fiber lumens. In contrast, we have utilized a device where hepatocytes in a supporting matrix were seeded in the intrafiber space of hollow fibers allowing for oxygenated plasma to flow over the outer surface of the fibers [15]. No matter which configuration is used, capillary hollow fiberbased bioreactors have inherent physical limitations including substrate transport limits associated with the fiber wall, limited total diffusion surface area, increased diffusion distance, and decreased capacity for hepatocyte mass [16]. 17.4.2.3 Encapsulation Technology Hepatocyte microencapsulate techniques within synthetic semipermeable membranes have been developed to provide physical separation that protect xenogeneic cells from the recipient’s immune system within a support system. In theory, the biomaterial excludes high molecular weight components of the immune system while allowing low molecular weight nutrients, oxygen, and stimulus freely across the semipermeable membranes. The most common protocol for microencapsulating hepatocytes is to envelope hepatocytes within a collagen matrix in an ultrathin sodium alginate copolymer membrane, which allows molecules such as oxygen, albumin, and clotting factors to exchange freely while preventing hepatocytes, antibodies, and complement from exchanging. Other encapsulation systems utilize biomaterials such as poly-l-ornithine, chitosan, or agarose, resulting in better biocompatibility and mechanical stability. Encapsulated xenogeneic hepatocytes have been used in liver support systems perfused directly by human plasma or blood. Nevertheless, capsules may potentially break down when meeting high shear stress or due to deterioration of the biomaterial, which can release the cells from the capsules [17]. Encapsulation technology also has substrate transport limitations associated with the capsule biomaterial, and complement breakdown products may be small enough to cross the membranes leading to immune responses against cells. 17.4.2.4 Aggregate Culture Spherical aggregates (spheroids) of hepatocytes, which are nonadherent multicell aggregates of greater than 50 mm diameter, provide a three-dimensional tissue construct that forms spontaneously. Several methods are used for spheroid formation from animal hepatocytes, such as culturing rat hepatocytes on nonadherent plastic surfaces for self-assembly [18], or rotational culture via spinner flasks [19]. More recently, our group reported preliminary observations that hepatocytes form spheroids when rocked (oscillatory motion) in suspension as illustrated in Fig. 17.2 [20]. Rocking promotes mixing, oxygenation, and increased frequency of collisions between freshly isolated hepatocytes, which in turn accelerates their aggregation

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Fig. 17.2  Spheroid formation by rocked suspension technique. (a) Multishelf rocker with blended gas inflow for controlled oxygenation at high cell density. Figures (b–h) show microscopy of spheroids: b – phase microscopy demonstrating size and shape of newly formed spheroids, c – immunohistochemical staining showing brown staining of hepatocyte nuclei for hepatocyte nuclear factor 4a (HNF4a), d – confocal microscopy showing immunofluorescence staining of actin (red) in cell membranes and DAPI staining nuclei (blue), e – confocal microscopy immunofluorescence staining of calcium-dependent surface adhesion molecules E-cadherin (green), f – DAPI staining demonstrating morphology of hepatic nuclei (blue) and a rare example of condensed chromosomes during prometaphase of mitosis (**), g – scanning electron micrograph demonstrating microvilli on surface of hepatocyte spheroids and pits resembling canalicular structures, h – transmission electron micrograph of a hepatocyte spheroid containing several well-defined sharp border nuclei, organized chromatin, and abundant densely stained mitochondria

into clusters and the formation of spheroids. Spheroid formation allows recapitulation of the cuboidal geometry of primary hepatocytes with relatively stable long-term differentiated function [21]. Reports of structural polarity and bile canaliculi formation by primary rat hepatocytes in spheroid aggregates provide further evidence that hepatic spheroids mimic the hepatocellular microanatomy of the liver [22]. Due to internal mass transfer limitations, spheroids with large diameters can lead to the formation of necrotic cores. That is, 100 mm hepatic spheroids have been shown to represent the size at which maximal oxygen consumption and albumin secretion

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occurs while still containing viable hepatocytes. Systems utilizing hepatic spheroids and encapsulated hepatocytes allow suspension culture of primary hepatocytes at high cell density under oxygenated bioreactor conditions, scalable to a size able to sustain the patient’s life.

17.5 In Vivo Incubators: Humanized Animal Liver A large demand exists for an abundant, routinely available, high quality source of human hepatocytes for therapeutic and diagnostic applications. Unfortunately, primary adult hepatocytes have limited proliferation potential and lose function and viability ex vivo. In contrast, hepatocytes exhibit a remarkable regenerative capacity in vivo. Based on this observation animal models have been developed to serve as hepatocyte incubators. For example, immunodeficient urokinase-type plasminogen activator transgenic mice have been bioengineered in which an albumin promoter directs high-level toxic expression of urokinase-type plasminogen activator. The hepatotoxicity creates a permissive environment for the expansion of transplanted hepatocytes, allowing 12 cell divisions on average [23]. Recently, Azuma et  al. introduced a method whereby primary human hepatocytes were efficiently expanded to near complete (>90%) hepatocyte replacement in livers of mice mutant for Fah, Rag2, and the common g-chain of the interleukin receptor in the absence of the protective drug, 2-(2-nitro-4-fluoromethylbenzoyl)-1,3-cyclohexanedione [24]. The human liver chimeric mice have provided a model for hepatitis B and C virus infection and treatment research [25]. The unique capacity of normal human hepatocytes to expand in the liver of a bioengineered mouse is based on a strong selective advantage for the transplanted cells to survive compared to the host cells. However, a limitation in the repopulated FAH deficient mouse is related to the absolute number of primary human hepatocytes that can be obtained. Thus, the generation of bioengineered Fah-null homozygous pigs is underway for much larger scale expansion of human hepatocytes.

17.6 Sources of Hepatocytes 17.6.1 Human Donor Livers Primary human hepatocytes would be a desirable option for hepatocyte transplantation as well as internal and external liver assist devices; however, human hepatocytes from a human source are not a good option because human livers are preferentially allocated for transplant. Alternatively, at least one group has reported isolating human hepatocytes from livers unsuitable for organ donation [26].

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17.6.2 Immortalized Human Hepatocyte Lines In an attempt to bypass the limitations associated with terminally differentiated hepatocytes, several groups have attempted to immortalize hepatocytes via spontaneous transformation, telomerase introduction, and retroviral transfection of the simian virus 40 large T antigen. To date, the hepatoblastoma C3A line, a subclone of the HepG2 cell line, is the only human-based cell line to be tested clinically in BAL device trials [27]. Limitations of C3A cells are their reduced levels of cytochrome P450 activity, ammonia removal, and amino acid metabolism compared to primary hepatocytes. Reduced ammonia removal by C3A cells is due to reduced expression of urea cycle genes [28].

17.6.3 Primary Pig Hepatocytes In place of primary human hepatocytes, xenogeneic cells could be a potential cell source for bioartificial systems. Because of differentiated metabolic functions, unlimited supply, and a high yield of cells, primary porcine hepatocytes have been most commonly used in several liver support devices undergoing preclinical and clinical evaluation [29]. Concerns regarding the use of porcine hepatocytes in human treatment include the risk of humoral and cellular immunologic response, transmission of porcine endogenous retrovirus, and potential function mismatch between porcine proteins and their human counterparts.

17.6.4 Stem Cells In recent years, great advances have been made in the production of stem cellderived hepatocytes. A large number of studies have utilized liver-derived stem cells including fetal liver stem cells (hepatoblasts) and adult liver stem cells (oval cells) to generate primary hepatocytes. But these hepatic progenitor cells are rare within liver tissue, with hepatoblasts comprising only 0.1% of fetal liver mass, and oval cells comprising 0.3–0.7% of adult liver mass. Other studies have induced ESCs to differentiate into hepatocyte-like cells. Induction involves exposure of ESCs to Wnt3a signaling to mimic events within the developing embryo, however, this results in limited functionality and incomplete maturity. In addition, ethical concerns may limit clinical application of ESCs. Hepatocytes can also be generated from hematopoietic stem cells and either bone marrow or adipose tissue mesenchymal stem cells through both transdifferentiation and fusion [30]. However, stem cell-derived hepatocytes have not yet been shown to fully function as primary hepatocytes.

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Reprogramming of adult somatic cells to iPSCs with the introduction of a defined set of transcription factors addresses the concerns of embryo destruction to produce ESC cells [31, 32]. Recent progress in iPSC technology may provide an individualized approach to the treatment of genetic disorders by providing cells that are theoretically identical to the patient for treatment, thereby avoiding immune rejection without immunosuppressive drug therapy [33, 34].

17.7 Application of Synthetic Liver 17.7.1 Diagnostic: Drug Discovery and Toxicity Testing The pharmaceutical industry uses high-throughput assessment of drug metabolism, toxicology, distribution, and pharmacokinetics from in  vitro hepatocyte culture assays. The microtiter plate (96-, 384- and 1,536-well), which is a miniaturized and parallel version of a conventional tissue culture dish, is most commonly used as the industry standard for high-throughput liver cell assays for drug discovery and toxicology. To provide better, faster, and more efficient prediction of in vivo toxicity and clinical drug performance, microfabricated cell arrays and microfluidics were recently developed to optimize liver-specific function of primary hepatocytes [13]. Microfabricated hepatocyte cultures exhibit characteristic patterns of gene expression phase I/II metabolism, canalicular transport, secretion of liver-specific products, and susceptibility to hepatotoxins [35]. Advances in microfluidics include the development of in vitro models of physiologically based pharmacokinetics. These models are designed to mimic physiological architecture and dynamics to allow for extrapolation of key in vivo drug parameters from in vitro cell culture assays and animal studies.

17.7.2 Therapeutic Liver transplantation is the only available treatment for severe end-stage liver disease. But the organ shortage and the need for life-long immunosuppression still limit its application. Thus, there are demands for alternative treatments for liver failure, including hepatocyte transplantation, hepatic tissue engineering, and extracorporeal BALs, which utilize isolated hepatocytes based on the belief that these cells should express full functionality in substituting for the normal liver. 17.7.2.1 Extracorporeal Devices BAL devices are extracorporeal temporary liver support systems that are analogous in concept to kidney dialysis machines, but use living hepatocytes within a bioreactor. These living hepatocytes can provide synthetic functions, regulation, and selective

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detoxification of water-soluble and protein-bound waste substances. BALs are expected to stimulate regeneration of injured liver, increase spontaneous recovery, and/or bridge to liver transplantation with the goal of improving patient survival [27]. To fulfill those goals, the BAL should contain high cell densities, maintain longterm well-differentiated hepatocyte function, and reduce mass transfer limitations. 17.7.2.2 Oxygenation Oxygen is an important component in the hepatic microenvironment since primary hepatocyte energy production is highly dependent on oxidative phosphorylation. Thus, BALs must overcome the issue of the high oxygen uptake rate of hepatocytes and the relatively low solubility of oxygen in aqueous media [4]. To ensure a sufficient supply of oxygen, some BAL designs use an in-line oxygenator in the extracorporeal perfusion circuit, while other designs incorporate an oxygenator into the bioreactor. To improve oxygen delivery, some BAL designs have employed oxygen carriers such as emulsified fluorocarbon and hemoglobin or additional fibers to carry gaseous oxygen directly into the bioreactor [36]. 17.7.2.3 Immunologic and Membrane Considerations To protect hepatocytes from the host immune system, most BAL designs employ selective membranes to prevent direct contact between patient blood and hepatocytes. In these designs, mass transfer is determined by the molecular weight cutoff of the membrane. Mass transfer refers to the transport of immunoglobulins and toxins out of the patient’s circulation and transport of hepatic proteins from the BAL into the patient’s circulation. Hollow fiber membranes with a molecular weight cutoff between 100 and 200 kDa appear optimal in providing immunoprotection of allogeneic/xenogeneic hepatocytes and serving as a barrier against zoonoses, without impairing removal of toxins [37]. 17.7.2.4 Cell Mass On the basis of data from human liver surgical resection, the maintenance of normal human liver function occurs with 10–30% of the liver mass. With an average liver mass of 1,500 g and hepatocytes comprising 90% of liver cell mass, this corresponds to 100–400 g of hepatocytes. This mass serves as the adequate mass for BALs to treat patients with liver failure. Hepatocyte mass of BALs that have undergone clinical testing has ranged from 75 g of cryopreserved porcine hepatocytes [29] up to 400 g of immortalized human C3A cells [27]. On the horizon is the SRBAL, a novel extracorporeal device equivalent to 40% of the hepatocyte mass of a normal human liver (Fig. 17.3). Anchorage-independent spherical aggregates of hepatocytes (i.e., spheroids) engineered by a novel rocked mixing technique serve as the source of detoxification activity in the SRBAL [21].

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Fig. 17.3  Features of the Mayo SRBAL. (a) Schematic of generic modular extracorporeal cellbased liver support device. (b) Components of SRBAL set-up. (c, d) Photograph and schematic of spheroid reservoir demonstrating fenestrated funnel configuration used to keep hepatocyte spheroids in suspension during continuous perfusion of the reservoir

17.7.2.5 Transplantation of Hepatocytes Hepatocyte transplantation is also a potential therapy for the treatment of numerous liver disorders. Hepatocytes are infused directly into the portal vein or indirectly into the spleen and then undergo blood flow-mediated translocation into the hepatic sinusoids. Intraportal injection of hepatocytes can cause transient portal hypertension; thus, intrasplenic delivery is often preferred for hepatocyte transplantation. On the basis of animal and human data, transplantation of hepatocytes corresponding to <5% of total liver mass can provide therapeutic benefit [38]. To date, human hepatocyte transplantation has been attempted in patients with acute liver failure, liver cirrhosis, and several metabolic disorders [38].

17.7.3 Implantable Constructs: Tissue Engineering Hepatocytes are attachment-dependent cells that maintain their liver-specific functions and viability through their attachments to the ECM and cell-to-cell contacts. Thus, hepatocyte transplantation would be much enhanced with a transplantable scaffold. Tissue engineering uses artificial three-dimensional, porous, and biodegradable scaffolds to provide a platform for hepatocyte attachment and a template for new tissue formation.

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17.7.3.1 Scaffolding Scaffolds can serve as synthetic analogs of the ECM to provide the necessary microenvironment for hepatocyte adhesion, differentiation, and survival. Many synthetic polymers or copolymers such as PLLA, PLGA, PLLA–PLGA, and other biomaterials including alginate, agarose, and collagen gels, have been employed as scaffolds in hepatic tissue engineering. To improve the attachment and differentiation of hepatocytes, the scaffolds are coated with ECM molecules and growth factors. Decellularized natural liver ECM represents an ideal material for hepatic tissue engineering as it retains relevant aspects of microstructure, chemical composition, and extracellular cues. Another benefit of decellularized matrices is that they have the potential to provide a cell number for transplantation equivalent to a whole organ. Recently, a new technology to engineer artificial liver tissue by cultured uniform hepatocyte sheets in a scaffold-independent manner has been introduced [39]. 17.7.3.2 Vascularization A major roadblock to successful application of BAL tissue is the need for a functional vascular network within the engineered liver tissue. Researchers have attempted various approaches in polymer biochemistry and scaffold design to build this vascular network. Strategies have included the co-seeding of endothelial cells with supportive fibroblasts that spontaneously form capillary-like networks [40, 41]. In addition, some groups have attempted to induce angiogenesis by incorporating angiogenic molecules into the scaffold and engineering cells to express these molecules [42]. Other strategies have included prevascularizing the implant site [39] or a threedimensional synthetic self-assembling peptide hydrogel to promote angiogenesis [43]. However, these efforts have fallen short of producing a scaffold that contains a natural vascular tree with centralized inlet and outlet vessels and a pervasive nutrient and gas exchange that are suitable for transplantation. More recently, two groups have developed transplantable recellularized liver grafts using the scaffold obtained via perfusion-decellularization, that contain a perfusable vascular tree that facilitates in  vitro perfusion and reconnection to the blood torrent that can greatly enhance nutrient delivery and waste removal in the engineered liver construct [44, 45]. Furthermore, Bao et al. utilized a layer-by-layer self-assembly heparin deposition technique that prevented thrombosis of the decellularized liver matrix leading to successful implantation of the recellularized liver graft into the rat portal system for 72 h (Fig. 17.4). 17.7.3.3 Bioengineered Xenografts Pigs may serve as potential liver donors for human patients because of unlimited availability; liver size similar to human livers; breeding characteristics; and physiologic and immunologic similarities to humans [46, 47]. A significant step toward avoiding

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Fig. 17.4  Example of a transplantable tissue-engineered liver graft from decellularized liver matrix in rat. (a–d) Sequential photographs of the right median lobe of a rat liver during tissue engineering process: a – cadaveric, b – decellularized, c – recellularized, d – transplanted graft. (e–h) Representative microscopy of hematoxylin and eosin staining during tissue engineering process: e – cadaveric, f – decellularized, g – recellularized, h – transplanted graft. (i–l) Representative scanning electron micrographs during tissue engineering process: i – cadaveric, j – decellularized, k – recellularized, l – transplanted graft (H = hepatocyte)

hyperacute rejection in xenotransplantation was the production of an a1,3galactosyltransferase gene-knockout animal that does not express Gala1,3 antigens [48]. These xenoantigens are a major target of natural antibodies of humans. Recently, a1,3-galactosyltransferase gene-knockout pigs with genetically modified human complement-regulatory protein, such as human CD46, improved protection of porcine xenografts from antibody-mediated injury [49].

17.8 Animal Models of Liver Disease An essential element in the development of clinically relevant cell-based liver therapies is the use of animal models to examine therapeutic effects as well as understand safety considerations. Animal models of hepatic failure include partial or total hepatectomy, ischemia- and chemical-induced injury, or a combination of insults [50]. Both small and large animal models have been used in preclinical studies. Small animal models (rats, mice) are cheaper and best suited for elucidation of the molecular mechanisms, while large animal models (pigs, dogs) are more suitable for evaluating novel therapies in preclinical trials [51].

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17.9 Current Challenges and Opportunities The field of tissue engineered liver therapies is in transition from development to clinical application. A significant practical barrier for all hepatocyte-based liver therapies is the identification of a reliable cell source. Such a cell source may come from genetically modified animals (e.g., the Fah knockout pig) that serve as in vivo incubators for robust expansion of healthy human hepatocytes from stem cells or iPSCs. To become clinically feasible, tissue-engineered liver therapies, such as the BAL or recellularized liver matrix, must provide a large number of hepatocytes with high functionality and human characteristics. Solutions to these barriers will open the doors to a new era of liver tissue engineering.

References 1. Taub R (2004) Liver regeneration: from myth to mechanism. Nat Rev Mol Cell Biol 5(10):836–847 2. Everhart J, Ruhl C (2009) Burden of digestive diseases in the United States part III: liver, biliary tract, and pancreas. Gastroenterology 136:1134–1144 3. Gomez-Lechon MJ et al (2003) Human hepatocytes as a tool for studying toxicity and drug metabolism. Curr Drug Metab 4(4):292–312 4. Kidambi S et  al (2009) Oxygen-mediated enhancement of primary hepatocyte metabolism, functional polarization, gene expression, and drug clearance. Proc Natl Acad Sci USA 106(37):15714–15719 5. McCuskey RS et al (2003) Hepatic microvascular development in relation to the morphogenesis of hepatocellular plates in neonatal rats. Anat Rec A Discov Mol Cell Evol Biol 275(1):1019–1030 6. Berry MN, Friend DS (1969) High-yield preparation of isolated rat liver parenchymal cells: a biochemical and fine structural study. J Cell Biol 43(3):506–520 7. Seglen P (1976) Preparation of isolated rat liver cells. Methods Cell Biol 13:29–83 8. Yagi T et al (2001) Caspase inhibition reduces apoptotic death of cryopreserved porcine hepatocytes. Hepatology 33(6):1432–1440 9. Kunieda T et al (2003) Cryopreservation of primarily isolated porcine hepatocytes with UW solution. Cell Transplant 12(6):607–616 10. Aoki T, Koizumi T, Kobayashi Y, Yasuda D, Izumida Y, Jin Z, Nishino N, Shimizu Y, Kato H, Murai N, Niiya T, Enami Y, Mitamura K, Yamamoto T, Kusano M (2005) A novel method of cryopreservation of rat and human hepatocytes by using encapsulation technique and possible use for cell transplantation. Cell Transplant 14(9):609–620 11. Mai G et al (2005) Treatment of fulminant liver failure by transplantation of microencapsulated primary or immortalized xenogeneic hepatocytes. Xenotransplantation 12(6):457–464 12. Dunn J, Thompkins R, Yarmush M (1992) Hepatocytes in collagen sandwich: evidence for transcriptional and translational regulation. J Cell Biol 116:1043–1053 13. Yarmush ML, King KR (2009) Living-cell microarrays. Annu Rev Biomed Eng 11:235–257 14. Allen JW, Khetani SR, Bhatia SN (2005) In vitro zonation and toxicity in a hepatocyte bioreactor. Toxicol Sci 84(1):110–119 15. Nyberg S et al (1993) Evaluation of a hepatocyte-entrapment hollow fiber bioreactor: a potential bioartificial liver. Biotechnol Bioeng 41:194–203 16. Catapano G (1996) Mass transfer limitations to the performance of membrane bioartificial liver support devices. Int J Artif Organs 19(1):18–35

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17. David B et al (2004) In vitro assessment of encapsulated C3A hepatocytes functions in a fluidized bed bioreactor. Biotechnol Prog 20(4):1204–1212 18. Landry J, Bernier D, Ouellet C, Goyette R, Marceau N (1985) Spheroidal aggregate culture of rat liver cells: histotypic reorganization, biomatrix deposition, and maintenance of functional activities. J Cell Biol 101:914–923 19. Sakai Y et al (1996) Large-scale preparation and function of porcine hepatocyte spheroids. Int J Artif Organs 19:294–301 20. Nyberg S et al (2005) Rapid, large-scale formation of porcine hepatocyte spheroids in a novel spheroid reservoir bioartificial liver. Liver Transpl 11:901–910 21. Brophy CM et al (2009) Rat hepatocyte spheroids formed by rocked technique maintain differentiated hepatocyte gene expression and function. Hepatology 49(2):578–586 22. Abu-Absi S et al (2002) Structural polarity and functional bile canaliculi in rat hepatocytes spheroids. Exp Cell Res 274:56–67 23. Mercer DF et al (2001) Hepatitis C virus replication in mice with chimeric human livers. Nat Med 7(8):927–933 24. Azuma H et al (2007) Robust expansion of human hepatocytes in Fah−/−/Rag2−/−/Il2rg−/− mice. Nat Biotechnol 25(8):903–910 25. Bissig KD et al (2010) Human liver chimeric mice provide a model for hepatitis B and C virus infection and treatment. J Clin Invest 120(3):924–930 26. Gerlach JC et al (2003) Use of primary human liver cells originating from discarded grafts in a bioreactor for liver support therapy and the prospects of culturing adult liver stem cells in bioreactors: a morphologic study. Transplantation 76(5):781–786 27. Millis J, Losanoff J (2005) Technology insight: liver support systems. Nat Clin Pract Gastroenterol Hepatol 2:398–405 28. Mavri-Damelin D et al (2008) Cells for bioartificial liver devices: the human hepatoma-derived cell line C3A produces urea but does not detoxify ammonia. Biotechnol Bioeng 99(3):644–651 29. Demetriou AA et al (2004) Prospective, randomized, multicenter, controlled trial of a bioartificial liver in treating acute liver failure. Ann Surg 239(5):660–667, discussion 667–670 30. Vassilopoulos G, Wang P-R, Russell D (2003) Transplanted bone marrow regenerates liver by cell fusion. Nature 422:901–904 31. Yu J et al (2007) Induced pluripotent stem cell lines derived from human somatic cells. Science 318(5858):1917–1920 32. Yu J et al (2009) Human induced pluripotent stem cells free of vector and transgene sequences. Science 324(5928):797–801 33. Si-Tayeb K et  al (2010) Highly efficient generation of human hepatocyte-like cells from induced pluripotent stem cells. Hepatology 51(1):297–305 34. Raya A et al (2010) A protocol describing the genetic correction of somatic human cells and subsequent generation of iPS cells. Nat Protoc 5:647–660 35. Khetani SR, Bhatia SN (2008) Microscale culture of human liver cells for drug development. Nat Biotechnol 26(1):120–126 36. Gerlach JC et  al (1994) Bioreactor for a larger scale hepatocyte in  vitro perfusion. Transplantation 58(9):984–988 37. Nedredal GI et al (2009) Optimization of mass transfer for toxin removal and immunoprotection of hepatocytes in a bioartificial liver. Biotechnol Bioeng 104(5):995–1003 38. Dhawan A et  al (2010) Human hepatocyte transplantation: current experience and future challenges. Nat Rev Gastroenterol Hepatol 7:288–298 39. Ohashi K et al (2007) Engineering functional two- and three-dimensional liver systems in vivo using hepatic tissue sheets. Nat Med 13:880–885 40. Kaihara S et al (2000) Silicon micromachining to tissue engineer branched vascular channels for liver fabrication. Tissue Eng 6(2):105–117 41. Nahmias Y et al (2006) Endothelium-mediated hepatocyte recruitment in the establishment of liver-like tissue in vitro. Tissue Eng 12(6):1627–1638 42. Soto-Gutierrez A et al (2006) Reversal of mouse hepatic failure using an implanted liver-assist device containing ES cell-derived hepatocytes. Nat Biotechnol 24(11):1412–1419

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Part V

Immune Response

Chapter 18

Immune Modulation for Stem Cell Therapy Gaetano Faleo and Qizhi Tang

Abstract  The remarkable strides made by stem cell biologists and tissue engineers have brought us ever so close to the promised land of having unlimited supplies of cells, tissues, and even organs to cure end-stage organ failure and reverse the course of degenerative diseases. Will our immune system perceive these engineered cells as foreign and target them for destruction as it does for conventional transplants? If so, how do we manage the immune response to prevent rejection, or better yet, to teach the immune system to accept the transplanted stem cells as self? In this ­chapter, we review how the immune system recognizes transplant antigens and analyze current data on immunogenicity of the various types of stem cells. We summarize current strategies for controlling transplant rejection and speculate on future directions in inducing transplant tolerance with the exciting possibilities of using stem cells to reeducate the immune system.

Abbreviations CTLA-4 ESC HLA IFN Ig iPSC IL LFA-1

Cytotoxic lymphocyte antigen-4 Embryonic stem cell Human leukocyte antigen Interferon Immunoglobulin Induced pluripotent stem cell Interleukin Leukocyte function antigen-1

Q. Tang (*) Department of Surgery, University of California, San Francisco, CA, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_18, © Springer Science+Business Media, LLC 2011

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Major histocompatibility complex Mesenchymal stem cell Natural killer T cell receptor T helper Tumor necrosis factor Regulatory T cells

18.1 Introduction Recent progress in stem cell biology and tissue engineering has generated great expectations for an unlimited supply of cells, tissues, and even organs to cure various debilitating illnesses such as spinal cord injury, Parkinson’s disease, and type 1 diabetes mellitus. In October 2010, stem cell therapy made headline news around the world when the first spinal cord injury patient received treatment with cells derived from human ESCs. The remarkable strides made by stem cell biologists and tissue engineers also bring forward the challenge of controlling rejection of the transplanted cells. The immune system is the body’s main defense against infectious organisms and malignancy. An effective immune response against these agents depends on orchestrated actions of two interconnected parts: the innate and adaptive immune systems. Two cardinal features of the immune system are its ability to distinguish self from non-self antigens and to form immunological memory so that a more effective immune response can be mounted upon later encounter with the same antigens. This highly sophisticated and efficient system, shaped by 450 million years of evolutionary pressure, is also the barrier to transplanted cells, tissues, and organs. In this chapter, we first briefly review the components of the immune system with emphasis on immune responses to transplanted foreign tissues. We then summarize the current literature on immune responses to various stem cells. Lastly, we point out currently available approaches to manage the immune response to stem cell-derived grafts and speculate on future development in this field.

18.2 A Brief Review of the Immune System The innate immune system is an evolutionarily conserved system that provides the first line of protection against invading microbial pathogens. It consists of a humoral module, mainly the complement system, and a cellular module that work together to provide protection against pathogens. The innate cells include phagocytes (macrophages, neutrophils, and dendritic cells), mast cells, eosinophils, basophils, and NK cells. These innate immune cells are equipped with various cell surface receptors that allow them to sense molecular structures that are foreign. The complement system, dendritic cells, and NK cells are particularly relevant to immune responses to transplanted tissues; therefore, they are reviewed in more detail below.

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The complement system consists of over 25 small proteins abundantly present in serum. Once activated, the complement system can lyse target cells, attract phagocytes, and mark cells for phagocytosis. The complement system can be activated by three pathways. The classical pathway is triggered by antibody–antigen complexes; therefore, the specificity of this response is dictated by the antibodies. The alternative pathway is constantly active, producing protein fragments that bind to the surface of pathogens so that they can be recognized by phagocytes. Host cells express cell surface complement regulatory proteins that deactivate complement to avoid being targeted by the alternative pathway. The third way to activate the complement system is by the immobilization of mannose-binding lectin on the surface of mannose-expressing bacteria and fungi. Thus, the complement system relies on antibodies, inhibitory selfproteins, and carbohydrate structures to distinguish between self and foreign tissues. Dendritic cells are normal residents in all peripheral tissues, where they constantly ingest antigens from their surroundings, and enzymatically process the antigens into smaller peptides that are presented at the cell surface by MHC proteins. Dendritic cells express abundant amounts of MHC class II molecules that present antigens to CD4+ T cells, and MHC class I molecules that present antigens to CD8+ T cells. Under physiological conditions, dendritic cells are immature, presenting self-antigens to T cells. The consequence of this steady-state interaction between dendritic cells and T cells is the acquisition of self-tolerance. Dendritic cells express pattern recognition receptors, such as toll-like receptors, that allow them to recognize microbial products [1, 2]. Once engaged, these receptors activate the dendritic cells to undergo functional maturation by increasing the expression of ligands for T cell co-stimulatory molecules and pro-inflammatory cytokines. Mature dendritic cells migrate via lymphatic channels to the draining lymph nodes, where they activate T cells to initiate the adaptive immune response [3]. Thus, these tissue dwellers are sentinels of the immune system, silencing or activating the adaptive immune system depending on the inflammatory milieu of the tissue they reside in. NK cells were discovered initially for their “natural” ability to kill tumor cells. Now it is clear that NK cells do not naturally (spontaneously) kill and they possess many other functions in addition to killing [4, 5]. NK cells respond to proinflammatory cytokines such as IL-2, IL-12, IL-15, and IL-18 by producing large quantities of IFNg. IFNg, in turn, induces MHC expression on antigen-presenting cells, thus, indirectly enhancing T cell activation. Cytotoxic killing activity of NK cells is tightly controlled by the balance of signals transduced by myriad activating and inhibitory receptors expressed by NK cells [6]. The predominant ligands for inhibitory NK receptors are self-MHC class I molecules; thus, transplanted allogeneic cells that do not express host MHC are promptly rejected by NK cells. NK killing of nonself cells is particularly evident when hematopoietic cells are transplanted, likely because of their constitutive expression of ligands for activating NK receptors. Ligands for activating NK receptors are generally expressed by transformed cells and stressed cells that are exposed to radiation, heat shock, viral infection, or genotoxic agents [7]. High level of expression of ligands for activating NK receptors can override the suppressive function of the inhibitory NK receptors, leading to the killing of tumor and stressed cells. Overall, the balance of activating and inhibitory receptors

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on NK cells endows the cells with an ability to distinguish healthy self-tissue and nonself or altered self-tissues, making NK cells a significant barrier to the engraftment of transplanted cells. Although innate immune cells such as dendritic cells and NK cells contribute to immune response against infections and tumors, these cells by themselves are insufficient to confer full immune protection against these insults without the adaptive immune system. This fact is vividly illustrated by severe infections and a high incidence of tumors in people who have primary immune deficiency or acquired immune deficiency syndrome, and in transplant patients who are chronically immunosuppressed [8–10]. T and B cells form the adaptive branch of the immune system. The vast arrays of TCRs expressed by T cells and antibodies secreted by B cells enable these cells to recognize enormous varieties of antigens that an individual is likely to encounter in a lifetime. During T cell development in the thymus and B cell development in the bone marrow, individual immature T and B cells randomly rearrange the genomic sequence of the TCR and Ig gene loci, respectively. Thus, a relatively small number of exons at these loci can generate greater than 1014 different TCRs and antibodies that are uniquely expressed by each individual T and B cell. When they encounter antigens, T and B cells that express receptors specific for the antigen are activated to proliferate and express effector functions to eliminate the source of the antigen. The random generation of TCR and antibodies before antigenic exposure is a powerful defense against rapidly evolving pathogens. Conversely, the random nature of the rearrangement creates, by chance, a population of T and B cells that are selfreactive, and for this reason, they have to go through a strict selection during their development. Inactivation of self-reactive T and B cells during their development is referred to as central tolerance and this process depends on the presence of selfantigens in the thymus and bone marrow. Thus, antigens ubiquitously expressed by all cell types and those transported to thymus or bone marrow via circulation, either as free antigens or presented by circulating dendritic cells, can induce central tolerance. The thymus has an additional source of antigen through the action of the transcription factor AIRE, which allows the expression of tissue-specific antigens such as insulin in the thymus [11]. The main consequence of self-antigen engagement during T and B cell development is apoptosis. T cells with moderate self-reactivity may be instructed to become Tregs. Tregs are not merely escapees of central tolerance, but a vital component of self-tolerance. When activated through their TCR, Tregs can suppress functions of other immune cells, including dendritic cells, macrophages, NK cells, T cells, and B cells. Deficiency in Tregs due to mutation of the Treg-specific transcription factor, Foxp3, leads to severe autoimmune diseases and early lethality [12, 13]. T cells recognize short peptide fragments presented by MHC expressed on the surface of antigen-presenting cells, and dendritic cells are most efficient at presenting antigens to T cells because of their high constitutive expression of MHC class I and II, and their extensive cell surface area [3]. Optimal activation of naïve T cells requires the provision of co-stimulatory signals by the antigen-presenting cells. The most potent co-stimulatory molecule, among numerous molecules in this role, is CD28 expressed by T cells, and their ligands CD80 and CD86, which are expressed

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by activated antigen-presenting cells [14]. T cells receiving a TCR signal without CD28 co-stimulation not only fail to proliferate, but also enter an unresponsive state called anergy, refractive to subsequent optimal activation. Co-stimulation through CD28 induces the expression of another potent co-stimulatory receptor, CD154. CD154 interacts with its ligand CD40 on B cells to induce B cells to switch from producing IgM to IgG and other classes of antibodies [15], leading to the amplification of humoral immune response. Thus, blocking co-stimulatory molecules can be an effective means to prevent T and B cell activation, and is the basis for many new immunosuppressive drugs. When properly activated, CD4+ Th cells can be induced to express various effector functions depending on the microenvironment the cells are in [16]. Various CD4+ Th cells can activate functions of other immune cells. For example, Th1 can promote CD8+ T cell proliferation by secreting IL-2 and activate macrophages via IFNg ­production. Th2 cells help B cell proliferate and produce antibodies, and Th17 cells activate innate immune cells such as neutrophils to expel pathogens. Activation of CD8+ T cells induces them to express cytotoxic proteins such as perforin and granzymes. These proteins enable CD8+ T cells to kill cells that express antigen they can recognize. Because CD8+ T cells recognize antigens presented by MHC class I molecules that are expressed on all nucleated cells, virtually all cells can become targets of CD8+ T cells. During T cell activation, some of the antigen-specific T cells develop into short-lived effector cells that carry out the functions described above, and others become long-lived memory cells. Memory T cells can act quickly when confronted with the same antigens at later times, and their activation is relatively less dependent on co-stimulation than their naïve counterparts. Therefore, they have a lower threshold for activation and are more resistant to immunosuppression. B cells are activated when their cognate antigens bind to membrane-bound IgM that serves as a B cell antigen receptor. Activated B cells differentiate into plasma cells that produce copious amounts of antibodies. During the first antigenic encounter, plasma cells are short lived and primarily produce IgM antibodies. With the help of T cells, particularly through CD154 and CD40 interactions, activated B cells further modify the genomic sequence at the immunoglobulin locus to “class switch” and make IgG, IgA, IgE, and IgD antibodies, and to generate antibodies with higher affinity for the antigen. These class-switched B cells can differentiate into long-lived plasma cells, providing long-term protection to the host that can last for decades. Antibodies protect the host by neutralizing viruses to prevent them from entering cells, coating bacteria so that phagocytes can eliminate them, and marking infected cells for complement-mediated killing.

18.3 Immune Response to Transplantation Antigens All vertebrates can discriminate between self-tissues and allogeneic tissues from other members of the same species [17]. Antigens that distinguish self-tissue from others are the ABO blood group antigens, MHC class I and II, and minor histocompatibility antigens [18].

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ABO blood group antigens constitute two types of antigenic carbohydrate moieties, types A and B, on glycolipids expressed by most cell types throughout the body, in addition to red blood cells [19]. Individuals can express A, B, both, or neither, thus forming four different blood types, A, B, AB and O, respectively. People who do not express A or B antigens develop antibodies specific for the nonexpressed antigens. Transplanting cells that express A or B antigen in these individuals will lead to antibody binding to the transplanted cells and immediate killing by complement-mediated lysis. To circumvent ABO antigen-mediated graft rejection, donor and recipient should be matched for the ABO antigens or type O donors that do not express either A or B antigens should be used, which is easily achievable as 40% of the world population is type O. The major antigens for allorecognition are the polymorphic MHC class I and class II molecules. Human MHC molecules are also referred to as human leukocyte antigens, and there are three MHC class I genes, HLA-A, B, and C, and six MHC class II genes, the a and b chains of HLA-DR, DP, and DQ. Among these genes, the HLA-A, B, and DRb are the most polymorphic, together contributing to more than 1,300 allelic variations; better matching of these MHC alleles between donors and recipients correlates with better graft survival, less need for immunosuppression, and improved transplant outcome [20, 21]. Allogeneic MHC proteins expressed by donor cells can be taken up, processed, and presented by host antigen-presenting cells as described above for self and microbial antigens. One unique feature of the alloimmune response is that host T cells can also directly recognize intact donor MHC molecules, and this mode of allo-recognition is referred to as the direct pathway, in contrast to the conventional indirect recognition of processed antigenic peptides presented by host antigen presenting cells [22] (Fig. 18.1). Healthy individuals can have 1% of their T cells capable of directly recognizing a mis-matched MHC antigen. This frequency is orders of magnitude higher than that for a nominal antigen. The high frequency of direct allo-reactive T cells underlies the exceptional strong immune response elicited by a transplant and the challenge in controlling transplant rejection. Although the indirect alloreactive T cells are present at lower frequencies, they are not insignificant in the rejection process. T cells of the indirect pathway expand more vigorously after transplant and can reject an allograft independent of the cells of direct pathway [23, 24]. As described earlier in this chapter, activation of naïve T cells requires costimulatory signals in addition to TCR engagement by MHC–antigen complexes; therefore, it is a specialized task for the professional antigen-presenting cells such as dendritic and B cells. Dendritic cells are present in all tissues and are transplanted along with the donor graft as “passenger leukocytes.” Upon transplantation, some of the graft-resident donor dendritic cells migrate to draining lymph nodes via lymphatic channels and activate T cells with direct donor specificity. Donor MHC antigens can also reach the draining lymph node in cell-free form and be presented by host dendritic cells to T cells with indirect specificity. In the case of vascularized grafts, MHC antigens can disseminate systematically via blood circulation and activate indirect T cells throughout the body. Since all nucleated cells express MHC class I, all cell types, including stem cells, can be a source of allogeneic MHC antigens for

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Fig. 18.1  Allogeneic transplants can be rejected directly or indirectly. In the direct pathway, donor antigen-presenting cells (APC) transplanted with the graft travel to the lymph nodes and come in contact with recipient T cells that recognize the mismatched major histocompatibility complex molecules on donor APC. In the indirect pathway, antigens shed from transplanted cells or tissues are taken up and processed by recipient APC, which then present the foreign antigen to recipient T cells in the lymph nodes. Both pathways lead to the activation of donor-reactive recipient T cells, which are capable of destroying the graft

indirect presentation. Once activated, donor-specific direct and indirect T cells traffic to the graft, where they are further activated to express their effector functions and attack the graft tissue. CD8+ T cells with direct specificity can directly kill donor cells. CD4+ T cells with direct and indirect specificity can be activated by donor and host dendritic cells (or other MHC class II-expressing cells), respectively, to express effector molecules that are either directly toxic to the graft or indirectly activate other infiltrating immune cells to destroy the graft [24–27]. In lymphoid organs, direct and indirect CD4+ T cells can help B cells to produce alloantibody [28], which can in turn inflict graft injury by activating the complement system after binding to allogeneic MHC on the graft tissues. Minor histocompatibility antigens are polymorphic proteins encoded on the Y chromosome, mitochondrial DNA, and autosomes. Protein encoded by these genes of donor origin are processed and presented as antigenic peptides by host antigenpresenting cells and can lead to activation of host CD4+ and CD8+ T cells, and B cells. They are less immunogenic than HLA antigens, but minor histocompatibility antigens alone are sufficient at inducing graft rejection in animal models and have been implicated in the loss of HLA-matched allografts in humans [18, 29]. Another significant barrier to transplantation is autoimmune-mediated destruction of the graft tissue. Autoimmune disease is a consequence of one’s immune system

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failing to refrain from attacking healthy self-tissue. Some autoimmune diseases affect a specific organ, as in the case of diabetes mellitus type 1 that targets pancreatic islet b cells, and multiple sclerosis that destroys the myelin sheath in the central nervous system. Other autoimmune diseases cause widespread systemic tissue damage, such as systemic lupus erythematosus, scleroderma, and rheumatoid arthritis. Autoimmune diseases often lead to irreversible tissue damage and organ failure, and treatment of these conditions using transplantation should include adjunct immunotherapy to control destruction of the new graft by autoreactive immune cells. Autoimmune responses in these patients may be more difficult to control than alloimmune responses [30], likely because autoreactive cells often have a memory phenotype due to chronic self-antigen stimulation, and therefore, have a lower triggering threshold. In summary, the immune system has evolved to sense and eliminate cells with minute molecular aberrations using a multitude of redundant pathways and positive feedback mechanisms. Its sensitivity in detecting foreign antigens exceeds that of the polymerase chain reaction. In addition, some of the patient populations targeted for stem cell therapy have preexisting autoimmune conditions that pose additional barriers to engraftment. Therefore, it is naïve to think that a foreign tissue, such as a stem cell-derived graft, can completely evade immune response without active intervention. Because immune responses to stem cell grafts are likely to be distinct from those of a conventional transplant, responses of various immune components to different forms of stem cell-derived grafts should be carefully analyzed with a goal of designing optimal immunotherapy to achieve stable engraftment.

18.4 Immune Response to Stem Cells An alloimmune response to stem cell-derived grafts is likely to be distinct from that described above for solid organ and tissue transplants [31, 32]. First, the in vitro generated stem cell-derived grafts are devoid of professional antigen-presenting cells. In addition, stem cell-derived grafts of nonhematopoietic origin likely do not  express high levels of MHC and co-stimulatory molecules. Therefore, stem ­cell-derived grafts carry less alloantigen load and most likely cannot directly activate naïve T cells. However, CD4+ T cells of indirect specificity can still be activated by host antigen-presenting cells when mismatched HLA are presented. Once activated, the indirect CD4+ T cells can potentially provide help to activated CD8+ T cells, NK cells, and B cells to destroy the graft. Most of the assays currently used to assess T cell alloimmune response, such as in  vitro mixed lymphocyte reaction, measure activity of the direct pathway because the cells of the indirect pathway are present at lower frequencies, making them difficult to detect. These factors combined have created the impression that stem cells are not immunogenic and can evade immune rejection in  vivo. In this section, we analyze the current data on immune responses to ESCs, MSCs, and iPSCs, taking into account all pathways of the alloimmune response.

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Immunogenicity of cellular transplants directly correlates with the amount of allogeneic MHC, the cells express. ESCs express low amounts of MHC class I and no MHC class II in their undifferentiated state and therefore have low immunogenicity for T cells [33, 34]. However, the low level of MHC I expression may make the undifferentiated ESC susceptible to rejection by NK cells, especially since some fetal tissues express high levels of ligands that activate NK receptors [35, 36]. Furthermore, the expression of embryo-specific antigens by ESCs can trigger an immune response [37]. In vitro and in vivo differentiation of ESCs can lead to an increase in MHC class I expression, which can be further augmented upon ESC exposure to IFNg [38–40]. Even after differentiation, most ESC derivatives do not express co-stimulation molecules, and therefore are not competent at directly activating naïve T cells. It is not surprising that ESCs and their differentiated progeny cannot induce allogeneic T cells to proliferate in vitro in a mixed lymphocyte reaction, because such an assay is designed to detect directly alloreactive T cells stimulated by professional antigen-presenting cells [41]. However, the lack of ability to induce T cell proliferation in vitro does not mean that the cells are not immunogenic when injected in vivo. Mouse ESCs transplanted into allogeneic hosts can induce host T and B cell activation and trigger rejection [42–45], even when only a single minor histocompatibility antigen disparity exists between the host and the donor [46]. Human ESCs transplanted into immunocompetent mice are promptly rejected [12, 47, 48], but it is not clear whether such a xenogeneic response reflects what may happen when allogeneic ESCs are transplanted into humans. ESCs and their differentiated derivatives are found to persist in an immunodeficient mouse reconstituted with allogeneic human peripheral blood leukocytes, a system that readily rejects adult allogeneic skin tissue [47]. This result should be interpreted with caution. Although the reconstituted mice have human T cells, B cells, and NK cells, the antigen-presenting dendritic cells, MHC molecules, co-stimulatory ligands, adhesion proteins, and immune cytokines necessary for a productive immune response are of mouse origin and may not be fully compatible with the receptors and ligands expressed by human T and B cells. A human ESC graft would have to rely on this incompatible system to present the allogeneic HLA. In contrast, human adult skin graft contains donor human dendritic cells and can directly present alloantigens to the T cells. Therefore, the lack of rejection of human ESCs in this system may be an artifact of the experimental setup, not the lack of allo-immunogenicity of the ESC graft. Humanized mice are a promising tool for in  vivo analysis of human alloimmune responses; however, considerable reconstruction of the mouse immune system is still needed for it to closely represent that of the human immune system [48]. MSCs were originally isolated from bone marrow and have a multipotent potential for developing into cells of mesoderm lineage, such as bone, cartilage, and fat cells. MSCs have also been widely studied for their ability to escape immune recognition and suppress tissue inflammation [49]. MSCs have been found to be minimally immunogenic in vitro and in vivo [50, 51]. In vitro, MSCs can inhibit the activation and proliferation of T cells [52], suppress antibody production by B cells [53], and inhibit IFNg production by NK cells. When injected in vivo, MSCs preferentially

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traffic to the site of tissue injury and inflammation [54, 55], and can reduce the severity of acute inflammation in a variety of conditions, such as ischemic brain and heart injury, graft-versus-host disease after allogeneic bone marrow transplantation, and autoimmune diseases [56–58]. It seems that the local inflammatory milieu, especially IFNg, activates MSCs to express their immunosuppressive function through transforming growth factor b, prostaglandine E2, indolamine 2,3-dioxygenase, heme oxygenase 1, nitric oxide, IL-10, and hepatocyte growth factor [59]. Some of the anti-inflammatory mediators produced by MSCs can disseminate and alleviate tissue damage at a distal site [60]. Thus, MSCs may have limited tissue regenerative potential when compared to ESCs, but may offer an approach to controlling acute inflammation. One way to circumvent the alloimmune response may be to use iPSCs derived from the patient to be treated. However, due to the nascent nature of this field, less experimental data are available to support or contradict this claim. Feasibility and cost of developing individualized cell-based therapy aside, iPSC applications can encounter several potential immunological barriers. First, if iPSC-derived cells are to be used in patients to replace organs destroyed by autoimmune diseases, such as type 1 diabetes mellitus, the graft may still be targeted by recurrent autoimmunity for rejection. Second, if iPSCs are generated using nuclear transfer technology by replacing donor oocyte nuclei with those from a patient’s own cells, minor histocompatibility antigens encoded by oocyte mitochondrial genes can potentially elicit immune-mediated destruction [61]. Lastly, using iPSCs to replace cells lost due to faulty genes would require repairing the genetic defects in the patient’s cells [62]. It is important to remember lessons learned in the mid-1990s in gene therapy – that new genes might become the targets of the immune system [63]. Although stem cells in general are hypoimmunogenic even after differentiation, it is clear that they are not impervious to the immune system. In assessing the immunogenicity of various stem cells, we should consider all facets of alloimmune responses. Clearly, MHC expression is an important parameter. ABO antigens are not to be forgotten. Minor histocompatibility antigens are not to be ignored and using female stem cells is an easy solution for avoiding Y-chromosome-associated minor histocompatibility antigens. Expression of ligands for activating and inhibitory NK receptors at various stage of differentiation on stem cells should be monitored. Lastly, expression of complement regulatory proteins on stem cells may affect their immunogenicity and should be documented. A comprehensive analysis of these immune parameters of various stem cells and their derivatives can help the design of immunotherapy tailored for stem cell transplantation.

18.5 Modulating Immune Response for Stem Cell Therapy As stem cell clinical trials begin enrolling patients, the question of how to control immune rejection of stem cell-derived grafts becomes one of the forefront considerations. Luckily, the transplant field has four decades of experience managing immune

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responses to various solid organ and cellular transplants, and many of the current practices can be applied to stem cell transplants. Moreover, research in immunology in the past two decades has offered great insight into the intricate regulatory network for ensuring immune self-tolerance. Immune tolerogenic strategies that tap into this newfound knowledge to coerce the immune system to accept transplant as self are being developed and tested in clinical trials [64]. Some of these strategies can be applicable to stem cell transplantation. In addition, stem cell transplants may not simply be passive beneficiaries of the myriad new immune tolerogenic treatments, but may offer exciting possibilities to be part of the tolerogenic regimen in reeducating the immune system to accept transplanted tissues. In this section, we review immunosuppressive strategies that are currently approved or in development, and speculate on ways to use stem cells to induce immunological tolerance. Currently, a variety of immunosuppression drugs are used in combination for controlling rejection after allogeneic transplantation. Considerable center-to-center variations exist in selecting immunosuppressive regimens. Regimens also differ for different transplanted organs and patient populations and often change over time to adjust to the health status of the patient and the organ. Thus, transplantation and its associated immunosuppression truly is personalized medicine. Immunosuppression for stem cell transplantation will likely to be similar in that it would need to be specifically tailored for each cell type and each patient. The key concern when designing immunosuppressive regimens are the risks of over- and under-immunosuppression. Insufficient immunosuppression risks graft injury. Some grafts, such as liver, can tolerate short-term injury and recover most graft function when immune rejection is controlled. The resilience of liver grafts is likely because of their ability to regenerate. Other grafts, such as kidneys that have limited regenerative potential, irreversibly lose part of the graft function with each episode of rejection; thus, graft injury can accumulate over time with repeated insults, leading to complete graft demise. Islet grafts are even less tolerant of rejection episodes due to their small size and inability to regenerate. Alternatively, overimmunosuppression can expose patients to opportunistic infections and malignancies [65]. For example, some immunosuppressive regimens have been associated with fatal cases of progressive multifocal leukoencephalopathy due to activation of latent JC virus from the central nervous system. Over-immunosuppressed transplant patients can also develop low-grade lymphomas called post-transplant lymphoproliferative disorder [10]. Thus, the challenge to managing immunosuppression after transplant is to do it just right, in the absence of clear a priori criteria or blood tests that define the right level of immunosuppression. Another consideration in designing of immunosuppressive regimens for transplant is the tolerance of the graft type for the specific drug. For example, islet transplantation was not successful initially because the widely used multidrug immunosuppressive regimen containing steroids and calcineurin inhibitors was toxic to islets. Avoidance of steroids and calcineurin inhibitors made islet transplants a possible treatment option for type 1 diabetes [66]. The current paradigm for immunosuppression was established based on experience with solid organ grafts. How commonly used immunosuppression drugs such as the calcineurin inhibitors cyclosporine A and

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tacrolimas, as well as sirolimus and mycophenolate mofetil, affect stem cell and stem cell-derived grafts should be carefully evaluated. All of these drugs target T cell activation and proliferation required for mediating graft rejection. Calcineurin inhibitors antagonize the signaling pathway triggered by intracellular calcium mobilization. Sirolimus targets mTOR signaling pathways and thus has an effect on protein translation, cellular metabolism, and growth. Mycophenolate mofetil reduces T cell proliferation by inhibiting purine synthesis through blocking inosine monophosphate dehydrogenase. While T and B cells are sensitive to these immunosuppressive drugs, none of them are selective for the immune system and may have off-target effects. Therefore, calcineurin inhibitors may not be suitable for stem cell grafts that rely on calcium signaling for function, and sirolimus and mycopheonolate mofetil are likely to be detrimental to grafts that actively proliferate. Antibody-based immunosuppressive agents more specifically target the immune system [67]. Therapeutic antibodies function by killing cells they target, blocking functional interactions between proteins, and neutralizing soluble mediators. Antithymocyte globulins are rabbit polyclonal antibodies obtained after immunizing rabbits with human thymocytes. The antibodies induce transient depletion of circulating T cells and NK cells. Monoclonal antibodies target specific proteins; they therefore have fewer off-target effects and minimal toxicity. They can be humanized to reduce their immunogenicity, prolong half-life and biologic effect, and thus only require intermittent administration. A monoclonal antibody to one of the CD3 chains of TCR was first approved for prevention and treatment of renal allograft rejection in 1986. Like antithymocyte globulins, anti-CD3 induces a sharp drop in circulating T cell counts followed by their gradual return. Rituximab is a monoclonal antibody directed against CD20 expressed on B cells. Treatment with Rituximab induces profound depletion of B cells for up to a year, and has marked efficacy in treating autoimmune disease and graft-versus-host disease after allogeneic hematopoietic stem cell transplantation [68]. Currently, these depleting antibodies are used as “induction” therapy for patients at higher risk of rejection at the time of transplantation to acutely reduce immune responses to new grafts. Induction therapy is then followed by maintenance immunosuppression typically containing calcineurin inhibitors, mycophenolate mofetil, and/or steroids. Blocking antibodies used in transplantation target co-stimulatory molecules. A variation of blocking antibody is the use of soluble receptors that bind to ligands of co-stimulatory molecules. For example, a fusion protein of the CTLA-4 extracellular domain with IgG heavy chain produces a recombinant protein CTLA-4Ig that can compete with CD28 for its ligands CD80 and CD86 on antigen-presenting cells. Blocking CD28 co-stimulation using CTLA-4Ig induces long-term graft ­survival in animal models of transplantation and is now in clinical trials for solid organ transplantation in humans. Intercepting another co-stimulation pathway between CD40 on antigen-presenting cells and CD154 on activated T cells using anti-CD154 antibodies induces graft tolerance in multiple models of transplantation in preclinical models. Its translation into the clinic has met an unexpected obstacle because CD154 is also expressed on platelets and anti-CD154 triggers thrombosis in patients. An alternative approach that targets CD40 may help

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to  circumvent this problem. Adhesion molecules such as LFA-1 and very late ­antigen-4 are also effective targets of blocking antibodies. These molecules strengthen the interaction between T cells and antigen-presenting cells during T cell activation and also contribute to leukocyte entry into sites of inflammation such as a newly transplanted graft. Therefore, targeting adhesion molecules can potentially control antigraft response during T cell priming and at the graft site. Human ESCs transplanted into mouse testis are accepted by blocking co-stimulation through CD28, CD154, and LFA-1 [69]. These blocking antibodies are usually used as a maintenance treatment, repeatedly administered to maintain constant therapeutic levels. Neutralizing cytokines with antibodies, soluble receptors, and receptor antagonists is an effective approach to control tissue inflammation in autoimmune ­diseases. Antibodies to TNFa showed remarkable efficacy in alleviating symptoms of rheumatoid arthritis and inflammatory bowel disease [70, 71]. IL-1 receptor antagonists can successfully control disease flares in patients with familial Mediterranean fever and systemic juvenile idiopathic arthritis [72, 73]. Although anti-TNFa and IL-1 antagonism can benefit patients with graft-versus-host disease after bone marrow transplantation, their use in solid organ transplantation has been limited, most likely because of the high efficacy of conventional immunosuppressive regimens. Possibly, a less immunogenic stem cell graft may not require the heavyhanded treatment of conventional immunosuppression and would benefit from more targeted therapies that block specific inflammatory mediators. None of the treatments mentioned above are aimed at drug-free graft acceptance, i.e., transplantation tolerance. Hematopoietic chimerism that replaces the host immune system with that from the donor is an effective strategy to induce donorspecific tolerance [74]. T cells that emerge in the recipient after this procedure are donor-derived, but have gone through thymic selection in the recipients and are, therefore, devoid of donor and host reactivity; therefore, organs and cells from the same donor are accepted with no need for immunosuppression. However, the procedures for establishing hematopoietic chimerism are relatively toxic, therefore, it will likely not be the first choice for promoting stem-cell graft tolerance. Treg cellular therapy has shown promise in preclinical models in inducing transplant tolerance. Clinical trials are ongoing to test its safety and efficacy in graft-versus-host disease and autoimmune diabetes [75]. In animal models of solid organ and cellular transplants, Treg cell therapy alone was not sufficient at preventing rejection due to the exceptionally vigorous alloimmune response. By combining Treg therapy with strategies that reduce the frequency of donor-reactive T cells, drug-free graft survival may be achieved. In the setting of stem cell transplant, Tregs can be isolated from the recipient before antithymocyte globulin induction and then infused back together with the stem cell transplant. This way, patients will have greatly reduced levels of effector T cells and a greatly increased ratio of Tregs to effector T cells, favoring tolerance induction. Transplanting graft into immune privileged sites may be another approach to achieve graft tolerance. There are areas of the body that are physiologically sequestered from the immune system and are called immune privileged sites.

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Previous studies have shown that transplantation of allogeneic tissue grafts into these sites, such as the brain, eye, hair follicles, testis, and uterus, does not trigger an immune response [76–81]. If a stem cell graft does not need to be transplanted in a specific tissue or organ, then placing it in one of the immune-privileged sites may help to avoid immunosuppression. Along the same line, engineered cell encapsulation devices that create a physical and chemical barrier between the host and the stem cell graft may shelter the graft from the immune system. Additional modification of the capsule by adding immunosuppressive mediators may help to reduce and even eliminate the need for immunosuppression. Stem cells themselves may be applied to induce graft-specific tolerance. Stem cells can stably integrate genetic modifications to their genomes, which provide opportunities to modify the immunological makeup of the cells before differentiation and transplantation. Reducing expression of MHC and ligands for activating NK receptors and forcing expression of host MHC class I may help the cells to escape from T and NK cell-mediated killing. Forced expression of immunosuppressive cytokines such as IL-10 and transforming growth factor b may help to create a local immunosuppressive milieu that may protect the grafts [82]. Transgenic expression of ligands on the surface of the stem cell for negative signaling receptors such as Fas ligand and programmed death ligand-1 on T cells may also help to limit the antigraft response [83, 84]. Membrane-bound single-chain monoclonal antibodies can also be used to selectively engage a negative signaling receptor [85]. However, these tolerogenic strategies need to be successfully tested in vivo since they may suppress or promote immune response depending on the cellular context and levels of expression [86–88]. Lastly, stem cells may be differentiated into immune regulatory cells and used to induce donor-specific tolerance. For example, ESCs may be differentiated into hematopoietic stem cells for generating hematopoietic chimeras for tolerance induction. ESCs may also be differentiated into thymic epithelial cells to delete T cells that are reactive to stem cells from the same source, in a process analogous to central tolerance induction in the thymus. These immune modulatory stem cellderived cells can be used as a preconditioning regimen to create a tolerant environment for the subsequent transplant of the desired cell types.

18.6 Conclusions To deliver the promise of stem cell medicine, we must first overcome the inevitable immunological barriers that will limit the use of stem cell-derived grafts. Immunosuppression currently used to prevent rejection of allogeneic transplants most likely will be effective in controlling stem cell rejection, but requires patients to risk the morbidity associated with long-term immunosuppression. A new paradigm is emerging in immunology that it is possible to reeducate the adult immune system to tolerate transplanted foreign tissues as one’s own so that life-long immunosuppression can be eliminated. In this regard, the plasticity of stem cells may hold

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the key to transplant tolerance. The crossroads at which we are arriving between stem cell biology and immunology may signal the beginning of a productive collaborative journey for both fields.

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Chapter 19

Regenerative Medicine and the Foreign Body Response Kerry A. Daly, Bryan N. Brown, and Stephen F. Badylak

Abstract  The host response, and in particular the innate immune response, is critical to the successful application of tissue engineering to the reconstruction of injured or missing tissues. Cell-based, scaffold-based, and signal molecule-based strategies are utilized in regenerative medicine and each of these approaches elicits a distinct host immune response that has a significant impact upon the downstream outcome. Modulation, but not suppression of the immune component of wound healing appears to be essential for constructive remodeling of tissues and organs. Promotion of a pro-wound healing and anti-inflammatory response, and avoidance of the foreign body reaction is associated with a constructive functional remodeling outcome. While macrophages play a pivotal role in this response, other immune cells and the interactions between all cell types involved in tissue remodeling are also clearly important. The objective of this chapter is to provide an overview of the host response to biomaterials including both the pro-inflammatory and resultant foreign body reaction, and the pro-wound healing, anti-inflammatory response that is associated with constructive remodeling.

Abbreviations bFGF C5 CCL CD CXCL DAMPs

Basic fibroblast growth factor Complement cascade component 5 Chemokine C–C ligand Cluster of differentiation Chemokine C–X–C ligand Damage-associated molecular patterns

S.F. Badylak (*) McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_19, © Springer Science+Business Media, LLC 2011

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Extracellular matrix 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide; carbodiimide Epidermal growth factor Immunoglobulin isotype G Interleukin Inducible nitric oxide synthetase Classically activated macrophage Alternatively activated macrophage Matrix metalloproteinase Platelet-derived growth factor Poly-lactic co-glycolic acid Reactive nitrogen intermediates Reactive oxygen intermediates Small intestinal submucosal ECM Transforming growth factor alpha Transforming growth factor beta T helper cell (either type 1 or 2) Tissue inhibitor of metalloproteinase Tumor necrosis factor alpha Vascular endothelial growth factor

19.1 Introduction Regenerative medicine approaches for the reconstruction and replacement of injured or missing tissues typically include cell-based, scaffold-based, or signaling moleculebased strategies. Regardless of the initial strategy, all approaches are intended to promote the eventual formation of a functional and structurally near normal tissue or organ. Each of these approaches is ideally associated with a period of tissue remodeling during which the cells and extracellular components will assume a vascularized, innervated, and functional configuration that is seamlessly integrated with the surrounding native tissue. This period of remodeling is critical to the success or failure of the regenerative medicine approach and importantly, this period invariably involves components of the host inflammatory response, including the innate and adaptive immune response. There are several criteria by which success or failure of any regenerative medicine approach can be measured but without question, the manner in which the host responds to the selected intervention will be a critical determinant of outcome. Cellbased approaches will elicit a host response that depends in large part upon the autologous, allogeneic, or xenogeneic source of cells being introduced to the recipient. Signaling molecules such as chemokines and growth factors may have varying degrees of immunogenicity and are typically delivered to the site of interest upon a carrier or scaffold. Scaffold-based approaches involve a wide spectrum of foreign materials, each of which will elicit a distinctive host tissue response.

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This chapter is predominantly devoted to an important subset of the host innate immune response, and its potential impact upon the success or failure of regenerative medicine attempts to reconstruct injured or missing tissues. Specifically, the subject matter is focused upon the role of the macrophage component of the innate immune response and the associated foreign body response. It should not be inferred that other components of the inflammatory response and adaptive immunity are less important, but the subject matter herein is limited for the sake of brevity and a reductionist approach to understanding factors that influence outcomes in regenerative medicine.

19.2 Mechanisms of Tissue Repair Following Injury or Biomaterial Implantation The success of tissue engineering and regenerative medicine approaches to tissue reconstruction may be attributed in large part to their ability to promote the formation of site-specific functional tissue instead of encapsulation or scar tissue formation. The mechanisms by which tissue-engineered constructs promote this type of “constructive” remodeling are only partially understood. It is increasingly evident, however, that modulation of the default host response to tissue injury is essential for success. In the following sections, we discuss the default host response that occurs following tissue injury and the foreign body response that occurs following the implantation of non degradable materials.

19.2.1 Host Response to Tissue Injury The default mammalian host response following tissue injury is a well-documented series of events that typically result in the deposition of dense fibrous connective tissue within the site of injury [1–3]. Very few tissues in adult mammals have the ability for regeneration; among them are the bone marrow, liver, intestinal epithelium, and epidermis of the skin. The default response to tissue injury has been described as occurring in four stages: hemostasis, inflammation, proliferation, and remodeling [2]. 19.2.1.1 Hemostasis Following tissue injury and resultant damage to the vasculature, platelets contact the damaged tissues resulting in the release of clotting factors that initiate hemostasis. A provisional matrix forms consisting largely of fibrin and entrapped erythrocytes. The provisional matrix provides a substrate for further cell migration into the site of injury and a medium for cell signaling [4]. In addition to their role in hemostasis and provisional matrix formation, platelets also release cytokines

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including PDGF, TGF-b, chemokine C-X-C ligand 4 (CXCL4), and interleukin 1 (IL-1) [5–7]. These factors contribute to initial repair process via recruitment of multiple cell types including neutrophils, macrophages, fibroblasts, and other tissuespecific cells to the injury site [7]. 19.2.1.2 Inflammation Neutrophils are the first inflammatory cell type to arrive at the wound site. Neutrophils phagocytose and destroy foreign material, bacteria, or dead cells that may have entered the wound site as a result of the injury and also provide further signaling molecules that recruit macrophages to the injury site [6]. Mast cells also participate in the early stages of wound healing by releasing granules containing enzymes, histamine, and other factors that modulate the inflammatory response [2, 8]. By 48–72 h post injury, macrophages begin to dominate the cell population at the site of injury [9]. Activated macrophages secrete cytokines and chemokines that promote the further recruitment of leukocytes to the site of injury [6, 7]. Macrophages also clear apoptotic neutrophils, the phagocytosis of which leads to a change toward a more reparative macrophage phenotype and the resolution of the inflammatory phase of wound healing [10]. The T lymphocyte population plays an important late regulatory role in the resolution of the inflammatory process through local secretion of cytokines and chemokines [11]. 19.2.1.3 Proliferative Phase The proliferative phase of wound healing involves cellular proliferation, angiogenesis, new ECM deposition, and the formation of granulation tissue – processes that are largely mediated via the effects of the local microenvironment including pH and oxygen tension, and cytokines secreted by macrophages, T lymphocytes, and other cells within the wound site [7, 12, 13]. These cytokines include EGF, bFGF, TGF-a, TGF-b, VEGF, and others depending on the nature of the injured tissue [7]. 19.2.1.4 Remodeling Phase Following the deposition of significant amounts of ECM (predominantly collagens type I and III) during the proliferative phase, the remodeling phase of wound healing begins. This phase is characterized by MMP- and TIMP- mediated degradation and remodeling of the newly deposited collagen, generally resulting in scar tissue formation/maturation [2, 14]. In some cases, prolonged remodeling leading to fibrosis or hypertrophic scar formation may occur due to dysregulation of the healing process [2]. Although these wound-healing events are described as part of the default response to tissue injury, many of these events also occur as part of a regenerative process. The selective activation of components of the inflammatory, proliferation,

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Fig. 19.1  Default host remodeling response at 28 days postinjury of abdominal wall musculature. Spindle-shaped cells, likely fibroblasts, within increasingly dense connective tissue can be observed forming at the periphery of the site of remodeling (top). A mononuclear cell response is observed within the area of the injured musculature (middle) and angiogenesis is observed. Uninjured muscle tissues can also be observed (bottom). Magnification = 20×, scale bar = 100 mm

and remodeling phase can result in a constructive and functional outcome as opposed to scar tissue formation. In particular, the role of macrophages and T cells in promoting a constructive remodeling outcome following implantation of biomaterial scaffolds is discussed in further detail in the following sections. In Fig.  19.1, an example of the default host remodeling response following injury of the abdominal wall musculature in a rat is shown. By 28 days post injury an increasingly dense layer of collagenous connective tissue can be seen forming over the injured musculature. This connective tissue contains spindle-shaped cells, likely fibroblasts. Numerous mononuclear cells, consisting of macrophages and T cells can also be seen within the area of the disrupted musculature and an angiogenic process can be observed, indicating that the inflammatory process is not entirely complete at this stage of tissue remodeling. With time, the collagenous tissue being deposited at the site of injury will mature into dense scar tissue and the presence of inflammatory cells will subside, signaling complete repair of the wound by the host. However, this repair generally does not result in the restoration of the function of the injured tissue. This default response to tissue injury is important and necessary to understand if one hopes to appreciate a variant of

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the inflammatory response; specifically, the foreign body response that occurs following the implantation of a non degradable or slowly degradable biomaterial, such as polypropylene.

19.2.2 Host Response Following Biomaterial Implantation The host response following the implantation of a non degradable synthetic or metallic biomaterial involves a series of overlapping processes that include (1) blood-material interaction with deposition of a protein film on the biomaterial, (2) provisional matrix formation, (3) acute inflammation, (4) chronic inflammation, (5) granulation tissue formation, (6) foreign body reaction, and (7) fibrosis and capsule development [15, 16]. While many of these processes are similar to those described above for the default host response to tissue injury, there are a number of key differences. The surgical implantation of a biomaterial is invariably associated with tissue damage and disruption of the vasculature at the surgical site. Release of blood into the wound site results in degranulation of platelets, formation of a provisional matrix, and signaling that recruits inflammatory cells (i.e., neutrophils and macrophages) to the surgical site. Blood contact also results in adsorption of proteins to the surface of the biomaterial within seconds of implantation [17]. The proteins that adsorb to a biomaterial may include components of the coagulation system (fibrinogen and tissue factors), complement cascade (C5), and other plasma-derived proteins (albumin and IgG) [16, 18]. These proteins provide a substrate through which the inflammatory cells arriving at the site of injury interact with the surface of the biomaterial. The specific proteins that attach, and the behavior of the attached cells, are dependent on the nature of the biomaterial surface and on an adsorption/desorption process that is governed by the affinity of the proteins for the biomaterial surface (known as the Vroman Effect) [17, 18]. As described briefly below, interactions of cells with the proteins adsorbed to the surface of the biomaterial may lead to a variety of cellular responses including adherence, activation, or triggering of phagocytic pathways, among others, depending on the cell type and the proteins involved [19, 20]. Acute inflammation, consisting of the emigration of neutrophils from the vasculature into the implant site, follows formation of the provisional matrix and the release of chemoattractant factors by platelets and other cells within the inflammatory site, much like the process described above for default wound healing. However, upon arrival within the wound site, neutrophils interact with the proteins adsorbed onto the biomaterial surface through integrin receptors specific for the adsorbed proteins [16]. For example, the adsorption of fibronectin and IgG plays significant roles in the Mac-1-mediated attachment of neutrophils and macrophages to biomaterial surfaces during the acute phase of inflammation [21]. Complement and serum immunoglobulin adsorption to a pathogen (termed opsonization) leads to phagocytosis by neutrophils and/or macrophages, or destruction of the pathogen via the complement pathway. In comparison, an opsonized biomaterial elicits

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either phagocytosis from neutrophils (and later macrophages), or will be subjected to frustrated phagocytosis, depending on the nature of the biomaterial and its size [16]. The process of frustrated phagocytosis involves the extracellular release of microbicidal contents at the surface of a foreign body. This release may cause the erosion of implanted materials, and may eventually lead to failure of the material to perform as intended. The chronic inflammation phase associated with the implantation of a biomaterial is typically characterized by the presence of activated macrophages. This process of macrophage accumulation may occur for a period of days to months depending on the nature of the implanted material and the adsorbed proteins. A meshwork of new ECM usually is deposited around the biomaterial and the accompanying angiogenic process is prominent. The continued presence of macrophages at the site of biomaterial implantation is often the precursor to the formation of granulation tissue, the foreign body giant cell response, and the eventual encapsulation of the biomaterial within a dense layer of collagenous connective tissue. Chronic inflammation can progress to a granulation tissue phase, in which the deposition of new ECM and the growth of vasculature into the implantation site through the process of angiogenesis are conspicuous. The persistence of granulation tissue combined with the presence of a non degradable biomaterial eventually leads to the formation of foreign body giant cells [16]. The classic histologic description of a foreign body reaction consists of macrophages and foreign body giant cells, formed through fusion of macrophages, which are typically located at the surface of the biomaterial. There are a number of factors including the chemical composition and surface topography that play a role in determining the degree to which a material elicits a foreign body giant cell response, predominantly through the modulation of protein adsorption [22–24]. As previously stated, macrophages generally interact with protein adsorbed surfaces through cell surface integrin receptors, the ligation of which induces intracellular signaling cascades that regulate macrophage behavior. Depending on the type of signaling elicited and the immunologic microenvironment, macrophages may undergo fusion, thus forming foreign body giant cells. The exact mechanisms of foreign body giant cell formation are highly complex and have yet to be fully described. An in-depth discussion of the process of foreign body giant cell formation is beyond the scope of this chapter, however, the topic of foreign body giant cell formation as it relates to biomaterials has been reviewed elsewhere [16]. In the final stage of the host response following the implantation of a biomaterial, an increasingly dense layer of collagenous connective tissue is deposited around the surface of the material, thus isolating or “encapsulating” it from the surrounding healthy tissue. Many approaches have been investigated to coat or modify the surface of a biomaterial to minimize the potentially detrimental processes of platelet activation, coagulation, and protein adsorption [25, 26]. Similarly, a number of approaches have attempted to develop material coatings or other strategies that reduce the foreign body giant cell and fibrotic responses to non degradable biomaterials intended for long-term implantation [27, 28]. All attempts to date have resulted in modest success at best.

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19.2.3 Role of the Host Response in Tissue Engineering and Regenerative Medicine The biomaterials most suitable for tissue engineering and regenerative medicine applications are typically degradable, manufactured from either synthetic or naturally occurring raw materials and are commonly combined with bioactive molecules or living cells. Therefore, the host tissue response tends to be distinctly different from that just described for metallic or polymeric biomaterials intended as permanent implants. While an in-depth consideration of how each of these components (degradable materials, bioactive factors, and cells) of tissue-engineered constructs affects the host response is beyond the scope of this chapter, a short discussion of each and examples of the host response elicited by commonly used degradable materials are provided below. In general, tissue engineering strategies utilize naturally or synthetically derived materials that are degradable, either in the short term or the long term, following implantation. These materials include degradable polymers such as poly-lactic coglycolic acid (PLGA), poly-caprolactone, and poly-ester urethane urea, among many others. While these materials are subject to the same processes of protein adsorption as the non degradable materials described above, their transient nature can affect both the severity and the duration of the host response. An example of this process is provided in Fig. 19.2. In this instance, a Vicryl (Polyglactin 910) mesh has been placed in the abdominal wall musculature of a rat following creation of a surgical defect. By 14 days post implantation, the Vicryl mesh is still largely intact and is subject to an intense foreign body giant cell response at the surface of the material. This multinucleated cell response is accompanied by the deposition of large amounts of granulation tissue consisting of newly deposited ECM, mononuclear cells, and blood vessels. By 35 days post implantation, the Vicryl mesh, which is largely degraded, is still present within the implantation site. The portion of the mesh that remains is surrounded by foreign body giant cells, with increasingly dense collagenous tissue formation at the periphery. The foreign body giant cell response, while reduced as compared to the response observed at 14 days, will only resolve after the material has been completely degraded. However, resolution of the inflammatory response will not result in constructive remodeling in this case. Rather, dense collagenous connective tissue resembling the scar tissue will remain at the site of implantation. One of the advantages to the use of polymeric materials is the ability to form the material into specific shapes and sizes using a variety of techniques such as weaving, extrusion, or electrospinning, among others. The use of such processes allows for highly accurate tuning of many factors associated with the biomaterial construct. These factors include three-dimensional configuration, mechanical and material properties, vascular networks, porosity, and degradability. These factors are known to have important effects upon the migration, proliferation, and differentiation of tissue-specific cells when cultured during the creation of tissue-engineered constructs in vitro. However, the impact of these factors upon the eventual success of a biomaterial

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Fig. 19.2  Host remodeling response following implantation of a degradable synthetic biomaterial (Vicryl mesh). The material can be observed within the site of remodeling at 14 days postimplantation (panels a and b) and is surrounded by a multinucleate giant cell population indicating a foreign body response. Angiogenesis, a robust mononuclear cell response, and the deposition of new ECM (granulation tissue) can be observed further from the surface of the implanted material. At 35 days postimplantation (panels c and d), a significant portion of the material has been degraded. However, remnants of the material can still be observed within the site of implantation and a chronic inflammatory response including foreign body giant cells persists. Increasingly dense connective tissue can be observed separating the implant from the native musculature (bottom). Magnification (a and c) = 10×, magnification (b and d) = 40×, scale bars = 100 mm

device following in vivo implantation appears to be limited to the immediate time period surrounding the surgical implantation of the construct. Surface chemistry, surface topography, and porosity have all been shown to have an effect upon the in vitro response of macrophages to biomaterials but few studies have investigated the effects of these factors upon the in vivo host response [22, 24, 29–31]. Other approaches to tissue engineering and regenerative medicine utilize naturally derived biologic materials such as those isolated from the ECM of mammalian tissues or those of plant origin. These materials may elicit a distinctly different type of host response than those of a synthetic origin due to differences in the surface topology and ligand landscape. Naturally derived materials likely experience adsorption of a different repertoire of molecules than do synthetically derived materials and also often possess inherent surface functionality related to the function of the biologic structure from which they were isolated. There are a number of factors

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Fig. 19.3  Host remodeling response following implantation of a degradable biologic scaffold material (urinary bladder matrix). At 14 days postimplantation (panels a and b) a robust mononuclear cell response accompanied by angiogenesis and deposition of new ECM are observed within and surrounding the degrading scaffold material. Of note, no foreign body giant cells are observed at the surface of the material. By 35 days postimplantation (panels c and d), the material is no longer identifiable in a histologic section and has been replaced by organized collagenous tissue, blood vessels, and bundles of skeletal muscle – a response that can be characterized as constructive remodeling. Magnification (a and b) = 10×, magnification (c and d) = 40×, scale bars = 100 mm

that influence the host response to naturally derived materials including the rate at which the material degrades and the molecular weight or composition in the case of biopolymers such as chitosan. For example, degree of deacetylation of chitosan has been shown to affect the rate of degradation and the host response it elicits upon implantation [32]. Figure  19.3 shows an example of a four-layer biologic material composed of ECM derived from porcine urinary bladder, which is capable of promoting a constructive remodeling response. The material is surrounded by mononuclear cells, which also infiltrate the degrading material, at 14 days post implantation into an abdominal wall musculature defect. This cell response is accompanied by the deposition of new ECM within the implantation site, as well as angiogenesis, similar to the example provided in Fig. 19.2, although notably no foreign body giant cells are present. However, the host response to the material at 35 days post implantation is quite different. By 35 days post implantation, the material has degraded and is replaced with well-organized collagen, blood vessels, and new islands of skeletal muscle.

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While the response shown in Fig. 19.3 represents only the beginning of the constructive remodeling process, studies have shown that similar ECM scaffold materials are capable of promoting the formation of functional, innervated muscle tissue at later time points [33, 34]. The two examples provided here are intended to show the diversity of responses that may be elicited by the materials used in tissue engineering and not to imply that synthetic-based approaches do not ever result in constructive remodeling. Indeed, there are numerous examples of tissue-engineered constructs containing polymeric components that have facilitated what can be described as constructive remodeling, some of which have been described in preceding chapters. Similarly, there are numerous examples of biologically derived scaffold materials that do not promote constructive remodeling. For example, chemical crosslinking of ECM-based scaffold materials similar to those used in Fig. 19.3 has been shown to result in fibrous tissue encapsulation as opposed to the constructive remodeling outcome that was achieved using non crosslinked forms of the same material [35]. It is also interesting to note that, although the outcome of tissue remodeling was different for the materials used in the examples provided in Figs. 19.2 and 19.3, there are many similarities in the early host response to both materials. That is, both materials elicited a robust mononuclear cell response and both elicited the formation of new ECM within the wound site. Phenotypic differences in the cells that make up the observed mononuclear cell population and their roles in determining the ability of a tissue-engineered construct to promote constructive remodeling are discussed in more detail in the sections that follow. Classic approaches to tissue engineering often dictate that cells be seeded into a carrier material and then cultured until reaching desired mechanical and biochemical properties, which are similar to the tissue of interest. Other strategies involve the culture of scaffold free constructs containing only cells, which can then be stacked or otherwise shaped prior to implantation. These strategies have achieved varying levels of success in a number of applications, many of which are described in earlier chapters. However, cell-based approaches that result in new tissues that include the originally implanted cells are few. There may be important and beneficial paracrine effects that initially promote the integration of tissue-engineered constructs or the formation of new tissue, but it is highly unlikely that the entire implanted cell population will survive in the host long term. The process of cell death and subsequent phagocytosis of cellular debris by neutrophils and macrophages can modulate the host response to the tissue-engineered construct. In particular, the mechanism of cell death (i.e., necrosis versus apoptosis) may play a role in determining how immune cells respond to cells implanted as part of a tissue-engineered construct. The specific effects of cellular debris and, in particular, DAMPs, upon the immune cell response to tissue-engineered constructs are discussed in more detail below. The cells used in tissue engineering strategies are generally of an autologous or allogeneic nature, with few approaches utilizing xenogeneic cell sources. These cells may be recognized by the adaptive immune system through mechanisms similar to those that have been described in detail for organ transplantation and rejection.

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Recent evidence suggests that certain cell sources may be capable of evading the adaptive immune system or elicit less immune activation [36, 37]. A number of strategies for the provision of cells that produce therapeutic molecules, such as insulin production by transplanted islet cells, have utilized hydrogels as encapsulating materials to prevent the recognition of cells by the host immune system [38, 39]. However, these approaches do not promote the integration of such cells into the tissue or organ of interest and are, therefore, often only effective in the short term. Such recognition of non self by the host will obviously have deleterious effects upon the ability of the tissue-engineered construct to integrate within the host tissues and otherwise perform as intended. The inclusion of bioactive factors such as cytokines, chemokines, and peptides within scaffold materials have been shown to have wide ranging effects upon the ability of the material to promote migration, proliferation, and differentiation of the cells that are seeded onto the scaffold and cultured prior to implantation. These same factors, when included on scaffolds prior to implantation, have been shown to have similarly wide ranging effects, which are often dependent on the tissue or organ into which the constructs are implanted. Each of these factors, especially those that have immunomodulatory properties, will logically have an effect upon the host response to the construct and, thus, affect the subsequent tissue remodeling outcome associated with its implantation. Regardless of the implant’s components (cells, biomaterials, signaling molecules), the host response will play a much more important role in the ultimate functionality than the properties of the construct at the time of surgery. Therefore, strong emphasis should be placed upon those factors that influence the host recognition of the engineered construct and subsequent cellular response. Following implantation, the engineered construct will acquire a surface coating of adsorbed proteins, the nature of which will affect the subsequent host response and the interaction with the first responding cell types (neutrophils and macrophages). Further, the type of cells that are included may elicit recognition by the host immune system, and the bioactive factors may alter the nature of this host response. The complex interplay between these events and others will determine whether the implanted construct is infiltrated with cells, degraded, integrated with the surrounding native tissue, or encapsulated.

19.3 Host Immune System Response to Biomaterials Both the adaptive and innate immune systems are important in natural wound healing [40] and in the constructive remodeling response induced by an engineered bioscaffold [41–45]. The neutrophils, macrophages, and T cells that respond to both the implantation procedure and biomaterial itself (as described in the previous sections), are pivotal in the clinical success or failure of the many commercially available biomaterials. As such many biomaterials are treated (e.g., surface functionalization) with agents to reduce their immunogenicity, however, such approaches typically fail to consider the complexity of the host immune response. As mentioned

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Table 19.1  Th1 and Th2 polarized immune responses to biomaterials Markers Pro-inflammatory Anti-inflammatory (wound healing) Humoral – activates B cells, Immunity Cell mediated – activates macrophages, some effects on macrophages activates B cells to produce opsonizing antibody isotypes Cytokines IL-1, IL-6, TNF-a, IFN-g IL-4, IL-10, TGF-b Chemokines IL-8, CCL5, CCL2, CXCL5 CCL24, CCL22 References [48, 49, 51, 55, 56] [41, 42, 46]

in the previous sections, the immune response is an integral part of wound repair, and modulation of the immunogenicity of biomaterials often prolongs the chronic inflammatory phase and leads to scar tissue formation or encapsulation. Similar groups of effector cells respond to biomaterials that induce very different remodeling outcomes. Distinctly different polarized immune responses of these effector cells have been associated with the success or failure of biomaterial scaffolds [16, 43, 45]. A pro-inflammatory or Th1 polarized and M1 reaction to biomaterials has been associated with deleterious outcomes, such as a prolonged chronic inflammatory phase, increased fibrous tissue deposition, encapsulation, and the presence of foreign body giant cells [43, 45]. An anti-inflammatory or predominantly Th2 and alternatively activated (M2) response is associated with constructive remodeling of the scaffold [43, 45, 46]. The exact mechanisms by which biomaterials influence the immune response are unclear. However, the composition of the scaffold [16, 27, 28, 35, 47–49], rate of degradation [44], and the presence of intact cells or cell debris [45] may all play a role. The details on this immune response, with particular focus on the role of polarized macrophages and the foreign body response, are described below.

19.3.1 T Cells and the Host Reaction to Biomaterials All biomaterials induce an initial robust immune response both in  vitro [16, 47, 49–51] and in  vivo [35, 43, 48, 52, 53]. Cells of both the adaptive (T cells) and innate (macrophages and neutrophils) immune system are among the first responders to implantation of biomaterials in  vivo [43, 48, 52]. Roles for other immune cells (e.g., dendritic cells) have been suggested by recent studies [47, 54]. The critical role of macrophages has been highlighted in several recent studies [43, 44] and will be discussed later. This section serves as an overview of the role of T cells and, in particular polarized Th cells in host immune responses to biomaterials. T cell interactions with both the biomaterial and other responding cells are also an important component of the host immune reaction [41, 42]. T cells are classified as either cytotoxic (CD8+) or helper (CD4+) T cells. Furthermore, Th are phenotypically divided into Th1 or Th2 effector cells, which have distinct immune responses in terms of cytokine and chemokine expression, and effector action (Table 19.1). In the realm of wound healing and regenerative

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responses, Th1 polarization is typically associated with pro-inflammatory responses, whereas Th2 are correlated with anti-inflammatory effects, wound healing, and constructive remodeling [41, 42] (Table 19.1). Polarization of the Th cell response toward biomaterials and an associated correlation with constructive remodeling [41] has been observed, however, the exact mechanisms behind this are unclear. Biomaterials derived from porcine small instestinal submucosal ECM (SIS) and syngeneic grafts elicited Th2 cytokine expression but not Th1 cytokines, and were constructively remodeled in a murine abdominal wall defect model [41]. Similarly, a clinical study of SIS usage in human patients determined that a Th2 polarized cytokine and antibody isotype profile was associated with tissue acceptance, rather than rejection of the scaffold [46]. This Th2 cytokine panel was found to be T celldependent in the murine model, although remodeling of the graft site was not [41], implying a role for other immune cells in constructive remodeling process. In contrast, expression of the Th1 cytokine, IFN-g, in response to xenogeneic implants was associated with chronic pro-inflammation and rejection of the graft [41]. Strong differential pro-inflammatory Th1 cytokine panels have also been elicited by several synthetic biomaterials [48, 49, 51] (Table 19.1) and has been typically associated with several synthetic scaffolds in particular [48, 49, 51, 55, 56]. In a recent study, polyethylene was found to be more pro-inflammatory compared to other synthetic scaffolds as it induced high levels of Th1 cytokines and chemokines but low levels of pro-wound healing cytokines (Table 19.1) [48]. The synthetic scaffolds used in this study were not able to be constructively remodeled due to use of a cage implant system, and consequently, long-term cytokine expression panels (56 weeks postimplantation) to all test devices were characterized by a chronic pro-inflammatory response [48]. IL-4 (a Th2 cytokine) expression in all synthetic scaffolds tested was found to be low over the entire study period, which is consistent with the inability of this model to allow healing, the presence of a chronic inflammatory response, and the apparent correlation of IL-4, and constructive remodeling of biologic scaffolds [41].

19.3.2 The Involvement of Innate Immunity and Macrophages in the Host Response to Biomaterials Regardless of the marked inflammatory response that biomaterials induce in vivo, there are no systemic effects; i.e., this host immune response appears to be a local tissue phenomenon [42]. While a localized response may have important clinical implications, it also highlights the role that the innate immune system plays in the host response to biomaterials. An innate immune response is consistent with the rapid cell and tissue reaction that is found to both synthetic and biologic scaffolds in vivo [35, 43, 48]. An intense mononuclear infiltrate surrounds both synthetic and biologic scaffolds within days of implantation [35, 48, 57, 58]. Several recent studies have suggested that macrophages facilitate constructive remodeling in response to biomaterial implantation [27, 44]. In fact, wound healing [59, 60], myogenesis [40], and biologic scaffold degradation [44] are all delayed or inhibited when there are deficiencies in macrophage number and/or their function.

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Monocyte/macrophage adhesion, foreign body giant cell formation, and subsequent pro-inflammatory cytokine production (as described in the previous sections) are seen frequently in response to many synthetic scaffolds and some modified biologic scaffolds (most notably in those that are chemically crosslinked). Even in degradable synthetic scaffolds, a fibrous connective tissue host response with encapsulation of the implant can persist longer than the synthetic device and the macrophage and foreign body giant cells [57, 58]. Consequently, synthetic biomaterials are frequently modified to delay this adhesion, or decrease and/or alter the cytokine production so as to improve the host response and remodeling outcome [27, 28]. Hydrophilic and anionic synthetic biomaterials have been associated with a decrease in monocyte/macrophage adhesion, macrophage activation, fewer foreign body giant cells, and a decrease in pro-inflammatory and increase in wound healing cytokine production by adherent macrophages [27, 28, 51, 52]. Macrophages are also part of the host response to both chemically crosslinked and non crosslinked biologic scaffolds [35, 43, 44]. However, in crosslinked biomaterials these macrophages are unable to infiltrate and degrade the device, remaining adherent to the periphery of the scaffold, with subsequent foreign body giant cell formation [35, 54]. Such crosslinked biomaterials are frequently associated with deleterious outcomes in vivo, such as fibrous tissue encapsulation and a foreign body reaction [35, 54]. 19.3.2.1 The Influence of Macrophage Polarization on Constructive Remodeling and the Host Response Clearly, macrophages are a vital component of the host response to a wide variety of different biomaterials – including both crosslinked and non crosslinked biologic scaffolds [35, 44] and synthetic biomaterials [27]. These biomaterials have distinctly different wound healing and constructive remodeling outcomes. The phenotype of these responding macrophages has been found to be an important determining factor in the ultimate success of an implanted biomaterial [43, 45]. Macrophages have recently been classified as either classically activated and pro-inflammatory (M1) or alternatively activated and anti-inflammatory (M2) [61], in a similar fashion to the Th1/Th2 dichotomy (Table 19.2). M2 or alternatively activated macrophages are further divided into M2a to c dependent upon their method of activation and function (Table  19.2). Polarized macrophages secrete a distinct panel of cytokines and chemokines (Table 19.2). In addition, there is a plasticity in macrophage phenotype and a given population of cells may have a range of both M1 and M2 polarized cells. Much of the recent work on macrophage polarization in response to biomaterials has focused on biologic scaffolds and consequently there is a focus on these scaffolds in the following section. However, characterization of macrophage phenotype in response to synthetic scaffolds will be an interesting area of future work. Macrophage phenotype is not static in response to biologic scaffolds, as early host immune reactions are commonly a mix of M1 and M2 phenotypes [45]. This mix of phenotype is important during the initial stages of response to injury as both sides of the macrophage spectrum have been reported to play roles in repair and regeneration of tissue [40] (M1 in phagocytosis, M2 in myogenesis in muscle

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Table 19.2  M1/M2 markers M2b (immune complexes and TLR agonists) IL-10 IL-1 TNFa IL-6

M2c (IL-10 and glucocorticoid hormones) IL-10 TGF-b IL-1RA

M1 IL-12 Il-1b IL-23 TNFa IL-6

M2a (IL-4/IL-13) IL-10 IL-1RA

Chemokines

CXCL8 CXCL9 CXCL10 CXCL11 CXCL16 CCL2 CCL3 CCL5

CCL17 CCL18 CCL22 CCL24

CCL1

CCL16 CCL18 CXCL13

Surface markers

CD80 CD86 CCR7 IL-1R1

CD23 CD163 CXCR1 CXCR2

CD80 CD86

CD14 CD150 CCR2

Other markers

iNOS ROI RNI

Arginase

Arginase

Arginase

Cytokines

Adapted from [61]

injury), and a deficiency of all macrophages prevents degradation and constructive remodeling of biologic scaffolds [44]. After the initial mixed immune response to a scaffold, macrophage polarization can be correlated with the success of the biomaterial (Fig. 19.4). A predominant M2 population has been associated with the degradation of the scaffold and constructive remodeling in response to biologic scaffolds [43, 44] (Fig. 19.4c). In contrast, a M1 skewed phenotype is typically correlated with persistence of the scaffold, fibrous encapsulation, the presence of foreign body giant cells, and a deleterious host tissue remodeling response in biologic and synthetic biomaterials [43, 44] (Fig. 19.4b). Furthermore, there appears to be a spatial distribution to macrophage phenotype associated with the chronic pro-inflammatory reaction found in response to certain biomaterials, which can be seen in Fig. 19.4b. In the host response to synthetic biomaterials, M1 macrophages and foreign body giant cells are found in higher densities immediately surrounding the device and are associated with a deleterious host response (Fig. 19.4b). In such a response, those M2 cells present will be found further away from the scaffold material, generally within the forming fibrous capsule. As stated previously, there has been less focus on macrophage phenotype in response to synthetic scaffolds. As there are no studies to date that have examined macrophage phenotype markers with use of synthetics, polarization can be inferred from cytokine and chemokine expression panels. In light of this, studies utilizing

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Fig. 19.4  Host macrophage response at 14 days following tissue injury (panel a), at 14 days following implantation of a degradable synthetic biomaterial (Vicryl mesh, panel b), and at 14 days following implantation of a degradable biologic scaffold (urinary bladder matrix, panel c). A mix of M1 (CCR7+, orange) and M2 (CD206+, green) macrophages can be seen within the area of necrosis associated with injured skeletal muscle (panel a). CD68 is used as a pan-macrophage marker (red), and nuclei are counterstained with DRAQ5 (blue). A predominance of M1 macrophages and multinucleate giant cells is observed at the surface of the degrading synthetic mesh with a lesser number of M2 cells present further from the surface of the material. A predominance of M2 cells can be seen within and surrounding the degrading biologic scaffold material. A small number of M1 cells are also observed. Magnification = 40×, scale bars = 100 mm

synthetic scaffolds both in vivo and in vitro (with monocyte and/or macrophage cell lines), describe cytokine and chemokine profiles (Table  19.1) that are consistent with a heterogeneous macrophage population [48, 50, 52]. Macrophages adherent to synthetic biomaterials express a range of M1 (IL-8, IL-6, and IL-1a) and M2 (IL10, IL-13 and TGF-a) cytokines, suggesting a mixed M1/M2 response [52]. This cytokine response changes over time, with less cytokines being expressed, and of those fewer pro-inflammatory cytokines (M1) and more pro-wound healing (M2) [48, 52] (Fig. 19.4c). 19.3.2.2 The Factors Present Within Biomaterials that Influence the Host Immune and Remodeling Responses This correlation between macrophage phenotype and success or failure of the biomaterial does not address the issue of what within the biomaterial is directing this host immune reaction. This question will likely be an area of future research due to the profound implications that it may have on design of successful biomaterial implants. The immunomodulatory factors within the biomaterial themselves are responsible for directing the host response to the implant. For synthetic scaffolds, hydrophilic and anionic surface molecules have decreased monocyte/macrophage adhesion, decreased activation of macrophages, decreased pro-inflammatory cytokines (mainly M1), increased pro-wound healing cytokines (mostly M2) by adherent macrophages, and have been associated with small numbers of foreign body giant cells [27, 28, 52]. Macrophage adhesion and its subsequent downstream effects are dependent upon the absorption of proteins to the biomaterials surface that occurs immediately upon implantation. Consequently, synthetic biomaterials are frequently modified to prevent macrophage adhesion or induce apoptosis, and to elicit a less pro-inflammatory and more pro-wound healing host response [27, 28, 52].

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Factors present within biologic scaffolds and the surface characteristics of the scaffolds influence the host immune response. Cellular remnants, such as DNA or DAMPs, have been suggested to contribute to a more pro-inflammatory and M1-mediated reaction [45]. Inadequate decellularization protocols of the source tissues for ECM scaffolds may be partially responsible for increased cellular remnant content. There is a wide variety of source materials and variability in decellularization methods for biologic scaffolds. Effectiveness of decellularization, often measured by DNA content, has been shown to vary between different source tissues and different decellularization methods [62]. In addition, most commercial biologic scaffolds are derived from xenogeneic sources and contain xenogeneic antigens such as the aGal epitope [63]. This particular antigen elicits increased production of anti-porcine and anti-Gal antibodies in response to porcine SIS products [46, 64]. However, there have been no deleterious effects noted in the host remodeling response in conjunction with these humoral effects, and antibody production diminished with the degradation of the implants [46, 64].

19.4 The Host Immune and Constructive Remodeling Response to a Biomaterial May Be Affected by the Anatomical Site of Implantation The host response described in this chapter, while representing the general host reaction, may not be applicable in all anatomical sites. The host immune and constructive remodeling response to an implanted biomaterial is influenced by not only the extent of the tissue injury but also the anatomical site. Vascularization, mechanical properties, the immune response, and inherent properties of immune cells (such as macrophages) may differ between different tissues and organs, and consequently may impact the ultimate success or failure of a biomaterial implant. Subcutaneous and intramuscular defects are commonly used models for preclinical studies of biomaterials; however, care must be taken when extrapolating such results to other tissues and organs. In fact, even between these two commonly used models, there is some suggestion that there are different levels of “reactivity” of host response. Intramuscular implantation has been found to be more host “reactive,” with a quicker pro-inflammatory response to synthetic scaffolds than subcutaneous implantation [65]. Implantation of biomaterials at sites with a rich vascular component or blood interface, such as at heart valves, provides unique challenges. Biomaterials will be covered in absorbed proteins immediately following implantation, and these proteins may influence the binding of immune cells (such as macrophages) and their polarization of immune response. In the case of the cardiovascular system, an adverse immune or thrombogenic response can have serious consequences for the success of the graft. A recent study of Synergraft (a porcine decellularized heart valve scaffold) in pediatric patients showed failure of the graft due to pro-inflammatory and foreign body reactions, with disastrous consequences for the patients [66]. Pro-inflammatory responses with intense neutrophil and macrophage infiltration

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were seen as early as 2 days postimplantation, beginning on the outside of the implant [66]. By 1 year after surgery, the host response had progressed to severe degeneration of the implant with minimal cellular infiltration into the device [66]. In comparison to rodent abdominal wall studies [35, 43, 45], this pro-inflammatory response occurred in a much shorter time frame. In vitro, porcine acellular heart valves have been found to be highly thrombogenic [67], to have immunoglobulin deposit (this has also been found for SIS [63]), to activate the classical complement pathway, and to allow adhesion of activated neutrophils [68]. Indeed, the immune responses of human plasma to porcine acellular heart valves have been shown to be similar to that of fresh porcine endothelial tissue [67], suggesting either an inadequacy of decellularization or that the increased vascularity and mechanical forces in the heart heighten host immune responsiveness. In other sites, such as the cornea, there is low vascularity and diminished immune responsiveness. Porcine decellularized corneal implants were used in rabbits, where ingrowth of host cells into the scaffold was found despite a diminished inflammatory response [69]. This result would appear to conflict with previous work that found porcine SIS device remodeling was prevented in the absence of macrophages [44]. A similar study using an EDAC-crosslinked porcine type I collagen implant in mice elicited corneal opacity due to a foreign body response to the graft [54]. A typical chronic pro-inflammatory reaction, involving macrophages and neutrophils, was noted [54]. The difference in response between these two different scaffolds in a similar anatomic site highlights the importance of considering the anatomic site when deciding upon the most appropriate biomaterial for a given clinical application.

19.5 Conclusions Ultimate success of tissue engineering approaches to the reconstruction of injured or missing tissues is dependent upon the host response and, in particular, the innate immune response. Cellular, scaffold, and signaling molecule-based strategies are utilized in regenerative medicine; however, each of these approaches elicits a distinct host immune response that may have a significant impact upon the ability of the scaffold or tissue-engineered construct to promote constructive remodeling. Scaffold biomaterials encompass a wide range of materials (both biologic and synthetic), and consequentially may elicit a variety of innate and adaptive immune responses. Modulation, but not suppression of the immune component of wound healing appears to be essential for constructive remodeling. In particular, promotion of a pro-wound healing and anti-inflammatory response (i.e., Th2 and M2 polarized macrophages) and avoidance of the foreign body reaction is associated with constructive remodeling. While macrophages play a pivotal role in this response, other immune cells and the interactions between all cell types involved in tissue remodeling are also clearly important, and are discussed elsewhere. A thorough understanding of the mechanisms that underlie the immune response to the materials, cells, and

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bioactive factors used in tissue engineering and regenerative medicine will lead to the design of strategies that promote the restoration of functional, site-appropriate tissue as opposed to inflammation and encapsulation or scar tissue formation.

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43. Badylak SF, Valentin JE, Ravindra AK, McCabe GP, Stewart-Akers AM (2008) Macrophage phenotype as a determinant of biologic scaffold remodeling. Tissue Eng Part A 14(11):1835–1842. doi:10.1089/ten.tea.2007.0264 44. Valentin JE, Stewart-Akers AM, Gilbert TW, Badylak SF (2009) Macrophage participation in the degradation and remodeling of extracellular matrix scaffolds. Tissue Eng Part A 15(7):1687–1694 45. Brown BN, Valentin JE, Stewart-Akers AM, McCabe GP, Badylak SF (2009) Macrophage phenotype and remodeling outcomes in response to biologic scaffolds with and without a cellular component. Biomaterials 30(8):1482–1491 46. Ansaloni L, Cambrini P, Catena F, Di Saverio S, Gagliardi S, Gazzotti F et al (2007) Immune response to small intestinal submucosa (surgisis) implant in humans: preliminary observations. J Invest Surg 20(4):237–241 47. Shokouhi B, Coban C, Hasirci V, Aydin E, Dhanasingh A, Shi N et al (2010) The role of multiple toll-like receptor signalling cascades on interactions between biomedical polymers and dendritic cells. Biomaterials 31(22):5759–5771 48. Schutte RJ, Xie L, Klitzman B, Reichert WM (2009 Jan) In vivo cytokine-associated responses to biomaterials. Biomaterials 30(2):160–168 49. Bonfield TL, Colton E, Marchant RE, Anderson JM (1992 Jul) Cytokine and growth factor production by monocytes/macrophages on protein preadsorbed polymers. J Biomed Mater Res 26(7):837–850 50. Schutte RJ, Parisi-Amon A, Reichert WM (2009 Jan) Cytokine profiling using monocytes/ macrophages cultured on common biomaterials with a range of surface chemistries. J Biomed Mater Res A 88(1):128–139 51. Chen H, Li P, Yin Y, Cai X, Huang Z, Chen J et al (2010) The promotion of type 1 T helper cell responses to cationic polymers in vivo via toll-like receptor-4 mediated IL-12 secretion. Biomaterials 31(32):8172–8180 52. Brodbeck WG, Voskerician G, Ziats NP, Nakayama Y, Matsuda T, Anderson JM (2003 Feb 1) In vivo leukocyte cytokine mRNA responses to biomaterials are dependent on surface chemistry. J Biomed Mater Res A 64(2):320–329 53. Rogers TH, Babensee JE (2010 Feb) Altered adherent leukocyte profile on biomaterials in Toll-like receptor 4 deficient mice. Biomaterials 31(4):594–601 54. Liu L, Kuffova L, Griffith M, Dang Z, Muckersie E, Liu Y et al (2007 Sep) Immunological responses in mice to full-thickness corneal grafts engineered from porcine collagen. Biomaterials 28(26):3807–3814 55. Anderson JM, Ziats NP, Azeez A, Brunstedt MR, Stack S, Bonfield TL (1995) Protein adsorption and macrophage activation on polydimethylsiloxane and silicone rubber. J Biomater Sci Polym Ed 7(2):159–169 56. Jones JA, Chang DT, Meyerson H, Colton E, Kwon IK, Matsuda T et al (2007 Dec 1) Proteomic analysis and quantification of cytokines and chemokines from biomaterial surface-adherent macrophages and foreign body giant cells. J Biomed Mater Res A 83(3):585–596 57. Tiainen J, Soini Y, Tormala P, Waris T, Ashammakhi N (2004) Self-reinforced polylactide/ polyglycolide 80/20 screws take more than 1(1/2) years to resorb in rabbit cranial bone. J Biomed Mater Res B Appl Biomater 70(1):49–55 58. Tiainen J, Soini Y, Suokas E, Veiranto M, Tormala P, Waris T et al (2006) Tissue reactions to bioabsorbable ciprofloxacin-releasing polylactide-polyglycolide 80/20 screws in rabbits’ cranial bone. J Mater Sci Mater Med 17(12):1315–1322 59. Lucas T, Waisman A, Ranjan R, Roes J, Krieg T, Muller W et al (2010 Apr 1) Differential roles of macrophages in diverse phases of skin repair. J Immunol 184(7):3964–3977 60. Mirza R, DiPietro LA, Koh TJ (2009) Selective and specific macrophage ablation is detrimental to wound healing in mice. Am J Pathol 175(6):2454–2462 6 1. Benoit M, Desnues B, Mege JL (2008) Macrophage polarization in bacterial infections. J Immunol 181(6):3733–3739 62. Gilbert TW, Freund JM, Badylak SF (2008) Quantification of DNA in biologic scaffold materials. J Surg Res 152(1):135–139. doi:10.1016/j.jss.2008.02.013

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Part VI

Animal Models

Chapter 20

Small Animal Models of Tissue Regeneration Fernando A. Fierro, J. Tomas Egana, Chrisoula A. Toupadakis, Claire Yellowley, Hans-Günther Machens, and Jan A. Nolta

Abstract  In order to develop new therapies for tissue regeneration, experimentation in nonhuman animals is fundamental. To a great extent, animal models serve as the link between in vitro stem cell manipulation and synthesis of products, such as synthetic matrices and organs, and clinical application. Animal models are indeed the basic tool to evaluate biosafety and efficacy of new technologies prior to their application in human patients and these models contribute greatly to our understanding of the underlying mechanisms involved in the regenerative process. In this chapter, we outline important rodent models used in the field of tissue repair, highlighting key immune-deficient mouse strains for transplantation of human stem cells. We also present examples of methods to be used in mice to mimic tissue damage and repair in a reproducible and quantifiable manner. These examples relate to the repair of bone fracture, skin wounds, and limb ischemia.

Abbreviations ALDH BM DFU FDA GUSB IV MPB MPSVII

Aldehyde dehydrogenase Bone marrow Diabetic foot ulcer Food and Drug Administration b-Glucuronidase Intravenous Mobilized peripheral blood Mucopolysaccharidosis Type VII

J.A. Nolta (*) Department of Internal Medicine, Stem Cell Program and Institute for Regenerative Cures, University of California, Davis, CA, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_20, © Springer Science+Business Media, LLC 2011

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Mesenchymal stem cell Natural killer NOD/SCID IL2Rg−/− Severe combined immune deficient Scaffolds for dermal regeneration Umbilical cord blood

20.1 Introduction With significant advances in physiology made during the first half of the nineteenth century, primarily in Germany and France, the establishment of large-scale use of animals in research began. Since then, experiments on living animals have become an essential part of most biomedical research disciplines [1]. In the field of cellular therapies and tissue engineering, different species have been employed, including rats (Rattus norvegicus), New Zealand white rabbits (order Lagomorpha), and minipigs (Sus scrofa). But by far, the most common animal used in the fields of cellular therapy and tissue engineering is the mouse (Mus musculus). Over the past century, the mouse became the leading mammalian model due to its fast and inexpensive breeding and because of its close genetic and physiological similarities to humans. About 40% of the human and mouse genomes can be directly aligned with one another, while 80% of human genes have at least one corresponding gene in the mouse genome [2]. Mice can develop diseases common to humans such as cancer, atherosclerosis, hypertension, diabetes, osteoporosis, and glaucoma. Other diseases, such as cystic fibrosis and Alzheimer’s disease, that normally afflict only humans, can also be induced in mice by manipulating the mouse genome and environment [3].

20.2 Common Mouse Strains Used in Tissue Engineering The ideal mouse strain to study tissue regeneration depends on the type of damage and therapeutic method applied. The most common inbred mouse strains used in the laboratory are C57BL/6, BALB/c, or substrains derived from them. The major advantage of these strains is their low propensity to develop spontaneous tumors and their well-studied genome. For tissue engineering, C57BL/6 and BALB/c are only recommended for studies where either chemicals are tested or autologous cells and tissues are transplanted. For studies, such as when xenografts are tested, other strains are recommended. The development of improved immune-deficient mouse strains has allowed the field to test human stem cells and tissue grafts. This technique is currently often required by the United States FDA prior to investigators seeking application for clinical trials of cellular therapy or stem cell-engineered products. In this context, nude mice, a BALB/c substrain characterized by a mutation on the foxn1 gene, which makes them athymic and consequently lack

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T cells, has been used for decades. However, their immune response to human cells is too strong, rejecting xenografts due to very high levels of NK cells. In 1983, C17/SCID mice were generated, which lack both T and B cells [4]. The immune defficiency of C17/SCID, however, is lost during aging [5]. Consequently, new strains were developed, such as the nonobese diabetic (NOD)/SCID mouse [6], which stably lacks both B and T lymphocytes. Since then, NOD/SCID b2M null and NSG mice, both commercialy available, were created exhibiting enhanced immunodeficiency compared with the parental NOD/SCID strain due to reduced NK cell function. Consequently, these mice are very supportive for xenografts, providing the highest engraftment with human cells to date [7–9]. Therefore, we recommend NSG mice for tissue repair studies, because of their more robust engraftment capacity, compared with the standard NOD/SCID strain, or nude mice which have high levels of human stem cell rejection [10–13]. In addition, mice with MPSVII, which is a lysosomal storage disease caused by a deficiency in GUSB activity, have been backcrossed onto the NOD/SCID strain, resulting in NOD/SCID/MPSVII mice. NOD/SCID/MPSVII mice provide a unique murine xenotransplantation model because human cells that carry normal levels of GUSB can easily be visualized against the background mouse tissues that are null for the enzyme [14–16]. The NOD/SCID/MPSVII mouse [15] thus allows the sensitive detection of individual unmarked donor cells (mouse, human, canine, etc.) in engrafted tissues by normal levels of the enzyme, without reliance on the continued expression of human cell surface markers or in situ hybridization. A simple substrate reaction is used on tissue slides. This mouse is highly useful for tissue repair studies, because the transplanted cells can be found in the tissues, by flow cytometry or on slides, without the use of antibody technology. This reduces background and errors in interpretation. Another benefit is that the transplanted cells do not need to be first manipulated to tag them with a vector, membrane dye, or other particle that would identify them later; they are detected based only on their innate expression of GUSB, absent in the mice. Immune-deficient mouse models of tissue damage provide systems in which human stem cell migration to sites of damage and contribution to repair can be carefully evaluated. Popular strains have been discussed in review papers [17]. Immunedeficient mice can be used to test tissue repair strategies, since most of the cytokines, chemokines, and inflammatory modulators involved with stem cell recruitment and activation are conserved from mouse to human. Hematopoietic stem cells from adult sources such as BM, MPB and UCB have been shown, in our laboratory and others, to promote tissue repair in immune-deficient mouse models through secretion of factors that enhance revascularization and tissue remodeling. Different populations of stem cells have been described that contribute to the regeneration of muscle [18, 19], liver [20–23], heart [24–27], and to regenerate vasculature [28–30]. Following intravenous injection, human stem cells can be shown to home to injured areas, in particular to hypoxic and/or inflammed areas, and to release trophic factors that hasten endogenous repair. These secreted bioactive factors suppress the local immune system, enhance angiogenesis, inhibit fibrosis and apoptosis, and stimulate recruitment, retention, proliferation, and differentiation of tissue-residing stem cells.

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Data from our laboratory and others show that IV-injected adipose and bone marrow-derived MSC lodge in multiple tissues following various routes of administration into sublethally irradiated immune-deficient mice [16]. Yet in models of acute local injury, MSC appear to preferentially home to, or accumulate in, the damaged tissue [25]. Our group has studied the trafficking of human BM and UCBderived ALDH+, CD34+, and more highly purified subpopulations, and of bone marrow and adipose tissue-derived MSC to damaged heart, liver, muscle, pancreas, and to the ischemic limbs of immune-deficient mice. Cell trafficking and functional outcomes from intravenous injection of the stem cell populations versus nonstem cell controls (i.e., ALDHhi versus ALDHlo sorted populations) were assessed. Trafficking was determined using fluorescent nanoparticle labeling and in  vivo dynamic imaging. Endpoints were determined by FACS analysis and by tissue staining using the NOD/SCID/MPSVII mouse, where unmarked donor-derived cells can be easily identified using a simple enzymatic substrate reaction [15]. Within each study, mice were assessed in groups large enough to achieve statistical significance. The groups that had received the stem cell populations demonstrated improved regional blood flow and tissue function, compared with the nonstem cellinjected control groups, at various timepoints after transplantation [31–34]. There was a paucity of human cells remaining in the tissue months after repair in models of acute injury, as opposed to long-term stable engraftment in chronic models (up to 18 months, as reviewed in ref. [10]). In the case of MSCs, the factors that control retention of the cells in the area of tissue damage are not well understood. However, the relatively short retention of the human cells in the acute models suggest that the tissue improvements observed are caused by the human stem cells initiating or catalyzing cascades of angiogenic activity and tissue repair from endogenous murine cells. While the combinations of factors involved are still being established, our data and that of others indicate that in acute wound or ischemic injury models, adult human stem cells do not become a significant part of the damaged tissue, but rapidly home to and persist only temporarily at a site of hypoxia or inflammation to exert significant trophic effects on tissue repair. In the following sections, we describe methods used in mice to mimic tissue damage and repair in a reproducible and quantifiable manner. The first method, a bone defect model, is currently done in our laboratory using immune competent C57BL/6 mice, because the immune response may play a critical role in the mechanisms underlying the normal bone regeneration process. This model has also allowed our laboratory and others to study the effect of small molecules and drugs on bone repair. The second method, a skin defect model, is currently applied in nude mice because the hairless animals facilitate skin transplantation and examination and because they have well-tolerated grafts of scaffolds for dermal regeneration. Finally, for our hind limb ischemia model, we commonly use immune-defficient strains, because they allow us to study the migration and tissue repair effects of human cells, while keeping their cellular identities distinct from the mouse background. It is important to note that in all experiments with animals, guidance and approval should be obtained from an Institutional Animal Care and Use Committee at the investigators’ institutions.

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20.3 Closed Transverse Fracture Model in Mice for Assessment of Bone Repair The average person in a developed country can expect to sustain two bone fractures over the course of his lifetime. Although many of these fractures heal by themselves, the type of fracture, location of the damaged bone, or patient’s health and age can strongly impact proper bone regeneration. The quest to find a method to create consistent, transverse, noncomminuted fractures in rodents for research purposes began in 1970. Jackson et al. developed a fracture device using three-point bending with a pneumatic punch press to create transverse fractures in the prepinned femurs of rats [35]. In 1984, Bonnarens and Einhorn improved upon the previous design by replacing the pneumatic punch press with a gravity-driven dropped weight and spring mechanism [36]. The Bonnarens and Einhorn method is currently the most well-known and most cited method of creating transverse fractures in rodents. In 2004, Manigrasso et al. modified this method for use on the prepinned femurs of mice, characterizing these fractures in the process [37]. In 2008, Marturano et  al. redesigned the fracture apparatus specifically for use on murine femurs. Their publication [38] gives a detailed diagram on the construction of the fracture apparatus and includes some major improvements, such as a notch in the anvils to aid with positioning the hind limb, uniting the striker and drop weight into one unit, and eliminating the need for the return spring. They also performed systematic impact mass and velocity experiments, mathematical modeling, and validation studies. With the parameters developed by Marturano et al., they were able to create favorable transverse fractures with a success rate over 80% [38]. An example of this method was recently published studying the role of inflammatory mediators during bone repair [39]. Although low numbers of animals were used per group, small variances in measurements such as histomorphometry and gene expression were achieved, indicating a highly reproducible injury and response. In this fracture methodology, both the prepinning surgery and creating the fracture are relatively simple (Fig. 20.1). Using a fracture apparatus to break the femur allows the creation of multiple replicable, transverse, noncomminuted fractures. The fractures created with this method are closed with minimal tissue trauma while, in contrast, open operative created fractures add increased local soft tissue trauma and increased risk of infection [35]. Internal instead of external fixation is much easier to establish, allows for natural and free movement of the mouse immediately after recovery from surgery anesthesia, and eliminates the possibility of affecting fracture healing with abnormal circulatory flow, disuse osteoporosis, muscle wasting, and joint stiffness [35]. After euthanization, the pin and needle wedge are easily removed without disturbing the fracture and callus, which allows one to use the sample for multiple uses, such as microcomputed tomography, histology, immunohistochemistry, RNA extraction, and mechanical testing. Samples can be collected immediately after fracture creation or months later. This technique can be used on mice of varying strains, sizes, weights, ages, and gender [38, 40].

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Fig. 20.1  Closed transverse femur fracture in mice. (a) Under appropriate anesthesia, the pin and needle wedge have been inserted between the femoral condyles into the bone marrow cavity of the femur. Analgesics are used to prevent pain. After bone fracture, pin placement, fracture, and fracture configuration are confirmed by X-ray (b)

20.4 Full Skin Defect Model in Mice for Assessment of Vascularization During Tissue Regeneration In cases of severe skin injury, dermal replacement is critical, since dermal tissue induces scaring instead of regeneration of normal dermis [41]. The possibility to replace skin defects using bioartificial products began with the culture of keranitnocytes [42] and has achieved great promise with the development of biodegradable SDR [43]. Albeit these important contributions, important improvements on bioengineered SDR are still required. For example, almost 100 million individuals worldwide suffer from DFU, a defect that occurs in 15–20% of all diabetics and results in limb loss in 25% of all DFU cases [44, 45]. Current treatments can heal only 50% of DFU and similar unmet needs exist in other chronic wounds. Even clinical trials using SDR containing living fibroblasts and/or keratinocytes have demonstrated only modest improvement in healing DFU, where 51% of foot ulcers heal in the SDR-treated groups, compared to 32% in controls. Most tissue engineering approaches rely on the rapid vascularization capacity of the implanted structures. In fact, in the absence of blood vessels, engineered tissue

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Fig. 20.2  Quantitative analysis of vascularization in full skin defect. Under appropriate anesthesia, a 15 mm diameter full skin defect is surgically created (a), the wound bed is covered by a mesh (b), and a scaffold for dermal regeneration is affixed (c). In order to visualize the blood vessels, skin (including the scaffolds) is removed after euthanasia and transilluminated, thus vascularization of the entire scaffold is clearly visualized (d). Afterward, digital segmentation of selected areas is performed (e) and further quantified as the percentage of white pixels

cannot be perfused with oxygen, nutrients, and other cells involved in the regenerative process and grafts undergo central necrosis. Consequently, high infection rates and low regeneration are general problems in the field of DFU tissue engineering [46]. Ex vivo, scaffold vascularization can be visualized with light, fluorescence, confocal, and electron microscopes. In general, these methods present adequate resolution, but require special labeling, are time-consuming, and expensive. Most important, staining techniques such as antibody-based identification of endothelial cells do not distinguish between functional vessels and nonperfused vessels. Other methods such as computerized tomography, magnetic resonance imaging, and positron electron tomography can overcome some of these difficulties and are commonly used in visualization of blood vessels, but they require sophisticated equipment. Moreover, these techniques do not allow the study of microvascular processes, due to their low resolution power [47]. Other techniques to visualize functional vessels are intravascular perfusion with tracers such as contrast phase liquid (followed by a radiographic analysis) and perfusion of the animal with latex or other plastic materials followed by tissue clearance or degradation [48]. However, results are strongly dependent on the levels of artificial perfusion and distribution of tracers in the animal. Furthermore, perfusion of the tissue with tracers excludes the use of that tissue for other studies such as protein or RNA analysis. We have developed a full skin defect model that presents several advantages compared to other described methods [49]. The use of skin as a model for tissue regeneration represents advantages that are intrinsic to the nature of skin, including high transparency, large surface, easy manipulation, external location and tissue homogeneity. In this model, full skin defects are surgically created in the back of the mouse (Fig. 20.2). Due to the symmetry

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of the animal, bilateral defects can be created and compared. In this model, multiple wounds allow internal comparisons and decrease the number of animals required for the study. After excision of the skin (Fig.  20.2a), a mesh is placed over the wound bed (Fig. 20.2b). Since most of scaffolds are biodegradable in vivo, this step minimizes artifacts that could be generated during tissue harvesting. Also, this procedure prevents wound shrinking, a commonly observed phenomenon in loose skinned animals that can later create bias in the evaluation of wound surface regeneration. Over the mesh, a scaffold is placed and surgically fixed (Fig. 20.2c). After the surgical implantation of the mesh, animals are kept for defined periods of time and then euthanized to evaluate the levels of skin regeneration. In order to quantify tissue vascularization, the whole skin, including the scaffold, is removed and quickly placed, upside down, over a light source. During transillumination, a digital picture is taken and further analyzed by digital segmentation (Fig. 20.2d–e). Although postmortem, this method is not invasive to the reconstituted tissue and both scaffolds and animals are intact, allowing further experiments such as histology or RNA extraction to be performed with the harvested material. Digital segmentation pictures are analyzed by the VesSeg-Tool program developed by our group [49]. This program generates semiautomatic segmentations which can be used to quantify vascularization with respect to area or length. With this model, different scaffolds or biomaterials can be easily compared. Moreover, the regenerative potential of different cell populations (e.g., stem or gene-modified cells) can be evaluated by seeding the cells in the scaffold before transplantation [50]. This model can also be used to evaluate pro- or antiangiogenic substances that are applied directly in the scaffold or systemically in the animal.

20.5 Hind Limb Ischemia for Quantification of Blood Flow Restoration Peripheral arterial disease affects more than ten million people in the USA, particularly in older patients, diabetics, and smokers [51]. Regeneration of blood vessels involves two mechanisms: angiogenesis, which is the growth and extension of preexisting capillaries, and arteriogenesis, the growth of functional arteries [52]. Both angiogenesis and arteriogenesis are commonly studied using the hind limb ischemia model [53]. It consists of ligation of the femoral artery, obstructing the flow of blood to the hind limb. This extremely traumatic injury is naturally repaired in young mice within a few weeks (typically four), through the formation or enlargement of collateral blood vessels [34]. Since originally described [54], the hindlimb ischemia protocol has undergone several minor modifications. A common surgical approach is ligation of the femoral artery at distal and proximal sites and removal of the comprised arterial fragment. In our hands, better results are obtained when in addition to ligation and excision, two major collateral arteries, the deep femoral artery near the proximal ligation site and the superficial epigastric artery near the distal site, are also ligated with sutures.

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The reason for these additional ligations is to prevent these preexisting collateral arteries that support blood flow to the hind limb when the femoral artery has been blocked from favoring angiogenesis rather than arteriogenesis originating from small arteries. Blood flow restoration that is too rapid would hamper the study of novel treatments. On the other hand, severe ischemia may result in necrosis and loss of the extremity. It is important to have enough functional tissue at the site of the injury to recruit stem cells and allow repair. A common method to quantify blood flow restoration is laser scanning Doppler imaging. An advantage of this method is that it is not invasive and can be performed in the same animal for many time points, from the day of surgery to the experimental end point. Typically, the study ends when blood flow in the ischemic limb is comparable to the contralateral, uninjured hind limb (control). Then, animals can be euthanized for further studies including histochemistry and RNA extraction. A negative aspect of scanning Doppler imaging is the sensitivity of the method. Only robust differences will be noticed and we recommend scanning at the highest possible resolution. This may be time-consuming, and a possible solution is to scan a smaller area of the hind limb, such as only the thigh or the foot. In addition, we suggest placing the animal on a heat pad with ventral side down for a few minutes prior to scanning, to avoid variation due to temperature changes. The optimal setup must be established by each investigator and kept constant during the entire experiment. The hind limb ischemia model has allowed us to demonstrate specific migration of ALDH+ cells to sites of injury following tail vein injection [31]. It also has allowed us to study the effect of hypoxia on MSC-mediated blood flow restoration [32].

20.6 Translation from Rodent Models to Humans Most in vivo models to study tissue damage and regeneration are developed in mice. However, the translation of possible therapies to humans commonly requires larger preclinical animal models that better resemble human anatomy and the appropriately sized tissues’ intrinsic abilities to heal. While this topic is treated extensively in the next chapter, it bears mention here as well. For example, to study cartilage damage and repair, it is difficult to create a surgical defect in mice that can be compared to the human situation, because the joints are very small and the cartilage extremely thin [55]. In contrast, the goat (Capra aegagrus) is much better suited to the study of cartilage repair, because the subchondral bone, thick cartilage, and joint anatomy are much closer to the human size and anatomy. The thick cartilage also allows creation of partial defects, which are not possible in small animal models [55]. Dogs (Canis lupus familiaris) are a common large animal model in particular for gene therapy studies of naturally occurring genetic diseases, because a high percentage of genetic diseases present in the dog are true orthologues of human diseases caused by mutations in the same genes [56]. For example, since the discovery that a canine X-linked muscular dystrophy faithfully mimics the phenotype of human Duchenne

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muscular dystrophy [57], the model has been commonly used to understand the pathophysiology of the disease, as well as for testing gene and cell therapies [58]. Pigs have similar anatomy and physiology to humans. Therefore, pigs are used to study cardiovascular damage and repair [59, 60]. In addition adult pigs present slow osteochondral intrinsic healing potential, allowing a critical assessment of new treatment strategies [61]. Other studies, however, are restricted by the very fast growth of pigs, limiting the time available to assess tissue-engineered constructs [62]. Horses (Equus ferus caballus) are the largest animal model available, typically used in orthopedic research. However, most applications are tested in naturally occuring instances of tissue damage, rather than in induced damage models, limiting the use of randomized controls in equine studies. Sheep models are available to test bone, tendon, and cartilage healing using the appropriate control groups [63].

20.7 Conclusions During the 1990s, experimentation on animals has increasingly emphasized the “three Rs”: reduction (i.e., minimize the number of animals used), refinement (i.e., optimize generation of the maximum amount of data), and replacement (i.e., substitution by in  vitro studies wherever possible) [64]. Consequently, novel methods have been described to create highly reproducible damage models and accurately quantifiable assesment of tissue repair. In this chapter, we have described three practical examples in detail: a system to generate consistent transverse femoral fractures in mice, a precise method to quantify vascularization during scaffold-mediated skin regeneration, and a model for vascular ischemia. The development of robust small animals models like these is essential to moving the field of regenerative medicine forward.

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8. Racki WJ, Covassin L, Brehm M et al (2010) NOD-scid IL2rgamma(null) mouse model of human skin transplantation and allograft rejection. Transplantation 89:527–536 9. Sanchez PV, Perry RL, Sarry JE et al (2009) A robust xenotransplantation model for acute myeloid leukemia. Leukemia 23:2109–2117 10. Meyerrose T, Olson S, Pontow S et al (2010) Mesenchymal stem cells for the sustained in vivo delivery of bioactive factors. Adv Drug Deliv Rev 62(12):1167–1174 11. Bonde J, Wirthlin L, Kohn DB, Nolta JA (2008) Human hematopoietic cell culture, transduction, and analyses. Curr Protoc Hum Genet Chapter 13:Unit 13 17 12. Zhou P, Hohm S, Capoccia B et al (2008) Immunodeficient mouse models to study human stem cell-mediated tissue repair. Methods Mol Biol 430:213–225 13. Wirthlin L, Hess D, Zhou P, Nolta J (2008) Human stem cells for tissue repair. Biol Blood Marrow Transplant 14:151–153 14. Hess DA, Craft TP, Wirthlin L et al (2008) Widespread nonhematopoietic tissue distribution by transplanted human progenitor cells with high aldehyde dehydrogenase activity. Stem Cells 26:611–620 15. Hofling AA, Vogler C, Creer MH, Sands MS (2003) Engraftment of human CD34+ cells leads to widespread distribution of donor-derived cells and correction of tissue pathology in a novel murine xenotransplantation model of lysosomal storage disease. Blood 101: 2054–2063 16. Meyerrose T, De Ugarte D, Hofling A et al (2007) In vivo distribution of human adipose-derived mesenchymal stem cells in novel xenotransplantation models. Stem Cells 25:220–227 17. Meyerrose TE, Herrbrich P, Hess DA, Nolta JA (2003) Immune-deficient mouse models for analysis of human stem cells. Biotechniques 35:1262–1272 18. Williams JT, Southerland SS, Souza J, Calcutt AF, Cartledge RG (1999) Cells isolated from adult human skeletal muscle capable of differentiating into multiple mesodermal phenotypes. Am Surg 65:22–26 19. Galmiche MC, Koteliansky VE, Briere J, Herve P, Charbord P (1993) Stromal cells from human long-term marrow cultures are mesenchymal cells that differentiate following a vascular smooth muscle differentiation pathway. Blood 82:66–76 20. Theise ND, Nimmakayalu M, Gardner R et  al (2000) Liver from bone marrow in humans. Hepatology 32:11–16 21. Theise ND, Badve S, Saxena R et al (2000) Derivation of hepatocytes from bone marrow cells in mice after radiation-induced myeloablation. Hepatology 31:235–240 22. Petersen BE, Bowen WC, Patrene KD et  al (1999) Bone marrow as a potential source of hepatic oval cells. Science 284:1168–1170 23. Lagasse E, Connors H, Al-Dhalimy M et al (2000) Purified hematopoietic stem cells can differentiate into hepatocytes in vivo. Nat Med 6:1229–1234 24. Orlic D, Kajstura J, Chimenti S et al (2001) Bone marrow cells regenerate infarcted myocardium. Nature 410:701–705 25. Wu GD, Nolta JA, Jin YS et al (2003) Migration of mesenchymal stem cells to heart allografts during chronic rejection. Transplantation 75:679–685 26. Jackson KA, Majka SM, Wang H et al (2001) Regeneration of ischemic cardiac muscle and vascular endothelium by adult stem cells. J Clin Invest 107:1395–1402 27. Zimmet JM, Hare JM (2005) Emerging role for bone marrow derived mesenchymal stem cells in myocardial regenerative therapy. Basic Res Cardiol 100:471–481 28. Davidoff AM, Ng CY, Brown P et al (2001) Bone marrow-derived cells contribute to tumor neovasculature and, when modified to express an angiogenesis inhibitor, can restrict tumor growth in mice. Clin Cancer Res 7:2870–2879 29. Crisa L, Cirulli V, Smith KA, Ellisman MH, Torbett BE, Salomon DR (1999) Human cord blood progenitors sustain thymic T-cell development and a novel form of angiogenesis. Blood 94:3928–3940 30. Peichev M, Naiyer AJ, Pereira D et al (2000) Expression of VEGFR-2 and AC133 by circulating human CD34(+) cells identifies a population of functional endothelial precursors. Blood 95:952–958

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31. Capoccia BJ, Robson DL, Levac KD et al (2009) Revascularization of ischemic limbs after transplantation of human bone marrow cells with high aldehyde dehydrogenase activity. Blood 113:5340–5351 32. Rosova I, Dao M, Capoccia B, Link D, Nolta JA (2008) Hypoxic preconditioning results in increased motility and improved therapeutic potential of human mesenchymal stem cells. Stem Cells 26:2173–2182 33. Rosová I, Link D, Nolta J (2010) shRNA-mediated decreases in c-Met levels affect the differentiation potential of human mesenchymal stem cells and reduce their capacity for tissue repair. Tissue Eng 16:2627–2639 34. Sondergaard CS, Hess DA, Maxwell DJ et al (2010) Human cord blood progenitors with high aldehyde dehydrogenase activity improve vascular density in a model of acute myocardial infarction. J Transl Med 8:24–30 35. Jackson RW, Reed CA, Israel JA, Abou-Keer FK, Garside H (1970) Production of a standard experimental fracture. Can J Surg 13:415–420 36. Bonnarens F, Einhorn TA (1984) Production of a standard closed fracture in laboratory animal bone. J Orthop Res 2:97–101 37. Manigrasso MB, O’Connor JP (2004) Characterization of a closed femur fracture model in mice. J Orthop Trauma 18:687–695 38. Marturano JE, Cleveland BC, Byrne MA, O’Connell SL, Wixted JJ, Billiar KL (2008) An improved murine femur fracture device for bone healing studies. J Biomech 41:1222–1228 39. Wixted JJ, Fanning PJ, Gaur T et al (2009) Enhanced fracture repair by leukotriene antagonism is characterized by increased chondrocyte proliferation and early bone formation: a novel role of the cysteinyl LT-1 receptor. J Cell Physiol 221:31–39 40. Manigrasso MB, O’Connor JP (2008) Comparison of fracture healing among different inbred mouse strains. Calcif Tissue Int 82:465–474 41. Machens HG, Berger AC, Mailaender P (2000) Bioartificial skin. Cells Tissues Organs 167:88–94 42. Rheinwald JG, Green H (1975) Serial cultivation of strains of human epidermal keratinocytes: the formation of keratinizing colonies from single cells. Cell 6:331–343 43. Burke JF, Yannas IV, Quinby WC Jr, Bondoc CC, Jung WK (1981) Successful use of a physiologically acceptable artificial skin in the treatment of extensive burn injury. Ann Surg 194:413–428 44. Dorresteijn JA, Kriegsman DM, Valk GD (2010) Complex interventions for preventing diabetic foot ulceration. Cochrane Database Syst Rev CD007610 45. Boulton AJ (2008) The diabetic foot: grand overview, epidemiology and pathogenesis. Diabetes Metab Res Rev 24(Suppl 1):S3–S6 46. Laschke MW, Harder Y, Amon M et al (2006) Angiogenesis in tissue engineering: breathing life into constructed tissue substitutes. Tissue Eng 12:2093–2104 47. McDonald DM, Choyke PL (2003) Imaging of angiogenesis: from microscope to clinic. Nat Med 9:713–725 48. Bergeron L, Tang M, Morris SF (2006) A review of vascular injection techniques for the study of perforator flaps. Plast Reconstr Surg 117:2050–2057 49. Egana JT, Condurache A, Lohmeyer JA et al (2009) Ex vivo method to visualize and quantify vascular networks in native and tissue engineered skin. Langenbecks Arch Surg 394:349–356 50. Egana JT, Fierro FA, Kruger S et al (2009) Use of human mesenchymal cells to improve vascularization in a mouse model for scaffold-based dermal regeneration. Tissue Eng Part A 15:1191–1200 51. Niiyama H, Huang NF, Rollins MD, Cooke JP (2009) Murine model of hindlimb ischemia. J Vis Exp 23:1035 52. Carmeliet P (2000) Mechanisms of angiogenesis and arteriogenesis. Nat Med 6:389–395 53. Limbourg A, Korff T, Napp LC, Schaper W, Drexler H, Limbourg FP (2009) Evaluation of postnatal arteriogenesis and angiogenesis in a mouse model of hind-limb ischemia. Nat Protoc 4:1737–1746

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54. Couffinhal T, Silver M, Zheng LP, Kearney M, Witzenbichler B, Isner JM (1998) Mouse model of angiogenesis. Am J Pathol 152:1667–1679 55. Ahern BJ, Parvizi J, Boston R, Schaer TP (2009) Preclinical animal models in single site cartilage defect testing: a systematic review. Osteoarthritis Cartilage 17:705–713 56. Casal M, Haskins M (2006) Large animal models and gene therapy. Eur J Hum Genet 14:266–272 57. Cooper BJ, Winand NJ, Stedman H et  al (1988) The homologue of the Duchenne locus is defective in X-linked muscular dystrophy of dogs. Nature 334:154–156 58. Wang Z, Chamberlain JS, Tapscott SJ, Storb R (2009) Gene therapy in large animal models of muscular dystrophy. ILAR J 50:187–198 59. Lowe HC, Schwartz RS, Mac Neill BD et al (2003) The porcine coronary model of in-stent restenosis: current status in the era of drug-eluting stents. Catheter Cardiovasc Interv 60:515–523 60. Schwartz RS, Murphy JG, Edwards WD, Camrud AR, Vliestra RE, Holmes DR (1990) Restenosis after balloon angioplasty. A practical proliferative model in porcine coronary arteries. Circulation 82:2190–2200 61. Gotterbarm T, Breusch SJ, Schneider U, Jung M (2008) The minipig model for experimental chondral and osteochondral defect repair in tissue engineering: retrospective analysis of 180 defects. Lab Anim 42:71–82 62. Rashid ST, Salacinski HJ, Hamilton G, Seifalian AM (2004) The use of animal models in developing the discipline of cardiovascular tissue engineering: a review. Biomaterials 25:1627–1637 63. Pearce AI, Richards RG, Milz S, Schneider E, Pearce SG (2007) Animal models for implant biomaterial research in bone: a review. Eur Cell Mater 13:1–10 64. Mayer FL, Whalen EA, Rheins LA (1994) A regulatory overview of alternatives to animal testing: United States, Europe, and Japan. J Toxicol Cutaneous Ocul Toxicol 13:3–22

Chapter 21

Use of Large Animal and Nonhuman Primate Models for Cell Therapy and Tissue Engineering Alice F. Tarantal and Karina H. Nakayama

Abstract  The intent of this chapter is to highlight the contributions of large animal models including pigs, sheep, goats, dogs, and nonhuman primates. These species are crucial for clinical translation of new cell and tissue engineering approaches for the treatment of human diseases. Choice of species, age, and model validation are necessary to ensure outcomes are predictive, and recapitulate human development, anatomy, physiology, and disease. The overriding objective is to select a model that can reliably assess the safety and efficacy of new therapies beyond the discovery phase and to obtain results that can be translated to clinical trials in children and adults.

Abbreviations ACL A-V CT CTSA DOA ECFC ECM ESRD HA

Anterior cruciate ligament Arteriovenous Computed tomography Clinical and Translational Science Award Deoxycholic acid Endothelial colony forming cell Extracellular matrix End-stage renal disease Hydroxyapatite

A.F. Tarantal (*) Departments of Pediatrics, Cell Biology and Human Anatomy, School of Medicine, UC Davis Clinical and Translational Science Center, California National Primate Research Center, University of California, Davis, CA, USA e-mail: [email protected] H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6_21, © Springer Science+Business Media, LLC 2011

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MI MSC NIH PAD PCL PDGF PGA PLGA PLLA PTFE RPE SIS TCP VEGF

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Myocardial infarction Mesenchymal stem cells National Institutes of Health Peripheral arterial disease Poly(e-caprolactone) Platelet-derived growth factor Polyglycolic acid Poly(lactide-co-glycolic acid) Poly-l-lactide acid Polytetrafluroethylene Retinal pigmented epithelium Small intestinal submucosa Tricalcium phosphate Vascular endothelial growth factor

21.1 Introduction A variety of species used in biomedical research serve as models for human health and disease, and represent a necessary step in the development of new diagnostics and therapies. Both small and large animal models provide substantial contributions and aid in overcoming the many roadblocks to clinical translation, and fill a gap in developing safe and effective treatments for humans across the lifespan. The intent of this chapter is to highlight the contributions of large animal models including pigs, sheep, goats, dogs, and nonhuman primates, and the need for teams of investigators with a range of expertise to participate in clinical translation. It is important to note that this chapter does not address the spectrum of studies conducted to assess stem and progenitor cells for regenerative purposes, which are covered elsewhere in this book, or related topics such as immunosuppression and tolerance induction to avoid rejection of donor cells. Indeed, the induction of donor-specific tolerance remains a challenge, and readers are referred to reviews and other chapters in this book that discuss the importance of large animal models in developing translational strategies for this purpose [1]. Studies have shown that tolerance is much more difficult to achieve in human and nonhuman primates when compared to mice due to differences in the immune system, and that findings in mice do not necessarily translate well to the human clinical setting [2–6]. This chapter presents tissue engineering studies for regenerative purposes that have been conducted in large animal species, and areas where further investigation and multidisciplinary approaches are needed to move closer to clinical trials. Topics are presented anatomically and also highlight current clinical needs. The authors acknowledge that many excellent publications have not been included due to space constraints, and hope that the publications referenced throughout the chapter adequately reflect the current literature.

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21.2 Cardiovascular System Heart and blood vessel diseases remain the greatest cause of morbidity and mortality in the USA today and include disorders of the valves and carotid and coronary arteries, PAD, as well as congenital defects, cardiomyopathy, and MI. Approximately 25% of adults (roughly 47 million in the USA alone) have also been identified with metabolic syndrome thus raising concerns that the quantity of individuals with cardiovascular disease will continue to rise, and with greater frequency in younger patients [7, 8]. Coronary artery bypass grafts, stents, ventricular assist devices, and heart transplant are procedures that are currently used clinically [9]. However, new regenerative strategies are needed to overcome current limitations including durability, infection, thrombosis, and the challenges of limited donor organ availability. Although rodents are useful for understanding molecular and related mechanisms, large animal models are important for cardiovascular research because of closer similarities when compared to humans [10, 11]. Promising results in rodent models for PAD, for example, have not predicted clinical success [12] and, similar to hematopoietic stem cell transplantation [6], have raised questions regarding relevance beyond early discovery. Differences between mice and humans include heart rate, oxygen consumption, response to loss of regulatory proteins, and contractile proteins that warrant needs for larger species [10]. Some anatomical and functional differences have also been identified such as the coronary circulation of pigs, which have no anastomoses between vascular branches, and the dog, where coronary arteries can be extensively collateralized [13]. The rhesus monkey model has also shown relevance for allograft transplantation, and examining the effects of pretransplant systemic inflammation on the posttransplant arterial wall [14]. Pigs, sheep, goats, and nonhuman primates have been used to explore vascular grafts and valve replacements, and to recapitulate MI, ischemia, ventricular pressure and volume overload, and pacing-induced dilated cardiomyopathy [10, 11, 15]. New imaging and interventional approaches have also benefited from large animal model research [16, 17], and imaging remains an essential component of in vivo studies focused on new regenerative techniques [18, 19].

21.2.1 Valves Replacement valves are used routinely and, as stated in a recent review, choice is dependent upon “…which complication one wants to avoid or reduce to a minimum” [9]. New tissue-engineered valves are needed for improved durability and biocompatibility, and that will allow growth and remodeling as well as eliminate the potential for thrombus formation [20]. Current techniques in aortic valve replacement utilize bioprosthetic valves that eventually fail because of calcification and degeneration [21]. Decellularized valves provide a platform for host cells, and acellular pig and sheep valves have been implanted into sheep and demonstrated in  vivo recellularization and adaptive remodeling [22, 23]. Valves were initially

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repopulated with interstitial cells and, at 11 months, the valve leaflets were approximately 80% repopulated and contained smooth muscle cells with no evidence of inflammation. In more recent studies, valves treated with DOA exhibited complete recellularization and an absence of calcification when implanted into the pulmonary position in juvenile sheep and pigs [24]. The pig model was included in these studies because of challenges due to rapid growth and greater potential for thrombus formation when compared to sheep. Replacement mitral DOA valves showed recellularization with limited calcification after 6 months, although thrombosis and calcification were noted in explants recovered when no antithrombotic therapy was given. These authors speculated that the mechanobiological environment may impact cellular repopulation, and highlighted a study by Baraki et al. [25] where decellularized aortic valves were only proximally covered by functional endothelial cells after 9 months in juvenile sheep. The authors concluded that recellularization may occur to a lesser extent in the aortic position when compared to the pulmonary position, and proposed that preseeding valves with autologous endothelial cells may provide benefit. The use of acellular valves has been a source of controversy due to the potential for thrombotic and neointimal formation compared to the use of preseeded valves [26]. While Dohmen et al. [27] found no differences in the degree of recellularization between DOA valves without cells and those that were preseeded with endothelial cells, the belly region and free edge of the valves remained acellular. These findings further supported preseeding particularly because the surface of decellularized valves makes them highly susceptible to platelet and fibrin deposition. In a recent study, trileaflet valves fabricated from biodegradable synthetic scaffolds and seeded with autologous myofibroblasts and endothelial cells were integrated in self-expanding stents and implanted as pulmonary valve replacements in sheep [28]. In vivo imaging showed thickened leaflets after 8 weeks with abundant amounts of collagen, although no significant differences were found in vitro or in vivo regarding ECM composition, morphology, or mechanical properties, and there was no evidence of inflammation. While promising, the authors noted that the long-term fate and safety of this approach requires further investigation. Anti-nonGal xenoantibodies are a major barrier to the survival of genetically modified porcine xenografts. Commercially available bioprosthetic heart valves from pigs or cows are known to retain the Gal antigen responsible for xenoantibodymediated rejection [29]. Genetic manipulation of porcine donors has significantly prolonged the survival of grafts placed into nonhuman primate recipients, but antinonGal xenoantibodies and thrombosis still limits the ability of these grafts to function on a long-term basis. Extensive studies have been conducted in an African Green monkey abdominal wall resection model to assess the effect of the Gal epitope on host response to porcine-derived ECM [30]. Gal-deficient pigs have since been proposed as a possible source for new, potentially calcium-resistant bioprosthetic heart valves [31].

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21.2.2 Patches Approximately 50% of postacute MI patients develop heart failure due, in part, to significant fibrosis and loss of contractile cells throughout the myocardium. While studies have focused on a variety of cell populations for direct injection into the myocardium, these investigations have not resulted in long-term improvements [32]. Thus, studies have focused on seeding artificial scaffolds (e.g., HA, collagen, fibrin) with the intent of using a cardiac patch to lessen scar expansion and progression to heart failure. Studies in dogs evaluated ECM derived from porcine urinary bladder to serve as an inductive scaffold for myocardial repair, compared to a Dacron patch [33]. The matrix-repaired region showed an increase in systolic contraction over the 8-week implantation period, and with an approximate thickness equivalent to the native right ventricular wall. Histological analysis showed a fibrotic reaction surrounding the Dacron patch and no evidence of myocardial regeneration, whereas the scaffold site contained cardiomyocytes accounting for approximately 30% of the remodeled tissue. A MI pig model was used to assess a patch implant (ECM or expanded PTFE) [34] and the success rate was noted to be associated with infarct size and location. Miyagawa et al. [35] used autologous skeletal sheets and showed improved cardiac function by attenuating cardiac remodeling in a pig ischemic myocardium model. Studies with pig hearts have also shown that natural scaffolds can be obtained by decellularization using detergent-based perfusion, with gross structural and biochemical properties retained [32]. While cardiac patches may provide a method for surgical implants, injectable materials may be less invasive and provide a way to replace damaged ECM or a scaffold for cell delivery [36]. Porcine myocardium has been decellularized and processed to form a myocardial matrix with the ability to gel in vitro [37]. The resulting myocardial matrix maintained a complex composition, including glycosaminoglycan content, and was able to self-assemble to form a nanofibrous structure. Endothelial and smooth muscle cells were also shown to migrate toward the matrix in vitro; this matrix is under assessment in a pig model via transendocardial catheter delivery [36].

21.2.3 Grafts Current treatment for PAD includes surgical or endovascular revascularization to improve blood flow, but outcome is typically poor, and the end result is frequently amputation [38]. The use of autologous or synthetic grafts has been shown to improve blood flow around occluded arteries, but this approach presents challenges. Patients in need may not be able to provide a suitable vessel, thus tissueengineered constructs that retain patency have been proposed as a promising

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alternative for long-term use. Studies have focused on tubular constructs made from decellularized tissue, biodegradable polymer or biopolymer scaffolds, cell sheets, and biomimetic materials [39]. Porcine models were initially used to develop vascular grafts with PGA seeded with smooth muscle and endothelial cells that have shown patency postimplantation [40]. Small caliber vascular grafts have been engineered using decellularized canine ureters seeded with autologous peripheral vein endothelial cells and myofibroblasts [41]. In vitro studies demonstrated reasonable levels of burst strength and compliance of the grafts similar to native carotid arteries. Grafts have been implanted in a canine carotid arterial replacement model and shown to remain patent after 6 months compared to nonseeded decellularized ureters and a PTFE tube, which became occluded after 1 week in vivo. The seeded grafts showed no signs of aneurysm or thrombus formation; however, the arterial wall was not completely regenerated. Baboons have also been studied to assess in vitro confluent endothelialized PTFE vascular prostheses and showed continued confluent endothelialization over time with no evidence of fibrin deposition [42]. These studies enabled translation of this approach to human clinical trials, where it was shown that expanded PTFE prostheses lined with autologous endothelial cells were successful in improving the patency of small diameter vascular grafts [43]. In recent studies, engineered vascular grafts using human allogeneic or canine smooth muscle cells grown on tubular PGA scaffolds have been tested in a baboon model of A-V access for hemodialysis, and in a dog model of peripheral and coronary artery bypass [44]. The grafts demonstrated excellent patency and resisted dilatation, calcification, and intimal hyperplasia. The authors proposed that the grafts could provide readily available options for patients without suitable autologous tissue or those that are not candidates for synthetic grafts. Recent studies in sheep have focused on microsurgical creation of an A-V loop embedded in an isolation chamber filled with fibrin matrix [18]. Constructs were implanted for up to 6 weeks and vascularization was shown using CT angiography and magnetic resonance angiography, followed by posttransplant microCT and histology, and with clinically relevant dimensions. Cho et al. [45] developed tissue-engineered vascular grafts using autologous bone marrow-derived cells and decellularized pig abdominal aortas to assess growth potential and vascular remodeling in the abdominal aortas of young pigs. Eighteen weeks postimplant, all grafts were patent with no sign of thrombus formation, dilatation, or stenosis when assessed by CT, and histological and immunohistochemical analyses revealed regeneration of endothelium and smooth muscle and the presence of collagen and elastin. The outer diameter of the engineered grafts was shown, in most cases, to increase in proportion to body weight and the native aorta. The growth potential of engineered grafts was also assessed by Brennan et al. [46] in a juvenile lamb model where PGA nonwoven mesh tubes coated with a 10% copolymer solution (l-lactide and e-caprolactone) and bone marrow mononuclear cells were implanted in the inferior vena cava. All grafts explanted at 6 months were patent and demonstrated evidence of growth over time, similar to other studies where grafts were used in the pulmonary circulation [47].

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Studies suggest that chronic ischemia can also stimulate collateral growth although collateral formation can be influenced by a variety of factors including age [48]. Vasculogenesis is the process of blood vessel development de  novo from endothelial precursors, and the isolation of circulating endothelial progenitors suggests that under some circumstances (e.g., tissue or vascular injury) new vessels can form with recruitment [49]. Endothelial progenitors play a role in vessel formation (angiogenesis or vasculogenesis) and the regenerative potential of endothelial and other stem and progenitor cells is currently being evaluated in clinical trials [38]. Human ECFC frequency and blood vessel forming potential has been shown to decrease in adult peripheral blood when compared to umbilical cord blood, and investigations across the lifespan (fetal to aged) in rhesus monkeys have demonstrated that the proliferative potential and vessel forming ability of ECFC progeny declines with age [50]. Age-related differences have also been shown in the structural, mechanical, and compositional properties of SIS-ECM harvested from pigs from different age groups [51]. These studies highlight the importance of age as a factor when designing studies for clinical translation.

21.3 Respiratory System Lung diseases such as chronic obstructive pulmonary disease, pulmonary fibrosis, and cystic fibrosis have limited treatments available that can halt disease progression, thus new therapies are needed. Significant advances have been made in understanding the role of endogenous lung progenitor cells in development and disease, and how they can be used for regenerative purposes [52, 53]. Reviews have also addressed the need for studies in large animal models with enhanced clinical relevance [54]. Similar to the kidney, the three-dimensional architecture of the lung presents substantial challenges, and while engineered lung transplants have been initiated in smaller species they have not yet been reported in large animal models. In contrast to the lung, preclinical studies in pigs enabled the translation of a decellularized cadaveric human trachea seeded with autologous epithelial and MSC-derived chondrocytes to form a functional airway that was successfully transplanted into an adult with bronchial stenosis and a child with long segment tracheal stenosis (both compassionate use) [55–57]. Prior studies have shown that autologous MSC-derived chondrocytes and epithelial cells could be seeded on a nonimmunogenic tracheal matrix and successfully grafted into pig tracheas. Other studies by Remlinger et  al. [58] have indicated that hydrated pig decellularized tracheal matrix could also serve as a functional scaffold for tracheal reconstruction. Further preclinical studies in these and other large animal models are needed to continue to develop and refine these clinically relevant techniques.

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21.4 Digestive System Tissue engineering of the small intestine remains experimental despite attempts to develop a functional substitute for the treatment of diseases such as short bowel syndrome [59]. Studies have used PLLA/PGA copolymer for a scaffold as well as a nanocomposite of polyhedral oligomeric silsesquioxane and poly(caprolactoneurea)urethane to develop porous scaffolds using a solvent casting/particulate leaching technique shown to support epithelial cell proliferation and growth in vitro. Porcine-derived, xenogeneic ECM obtained from either SIS or urinary bladder submucosa has been used as a tissue scaffold for esophageal repair in a dog model [60]. The xenogeneic scaffolds were resorbed within 30–60  days and showed replacement with skeletal muscle oriented appropriately and contiguous with adjacent esophageal skeletal muscle. Further studies showed that ECM scaffolds with autologous muscle, but not ECM scaffolds or muscle alone, facilitated the in situ reconstitution of structurally and functionally acceptable esophageal tissue [61]. In other studies, fabricated scaffolds formed from PGA nonwoven fabric with or without oral keratinocytes were rolled around a polypropylene tube and placed within the dog omentum, then removed for esophageal replacement 3 weeks postabdominal surgery [62]. The tissue-engineered esophagus with autologous cells showed good distensibility, and the dogs remained problem-free for the duration of the study (420 days). Orthotopic liver transplantation for liver failure could be avoided if new tissue engineering strategies are developed [63]. A study with transplanted genetically modified pig livers into baboons resulted in survival for up to 8 days [64], although complications associated with acute rejection were noted. Until a functional and safe tissue-engineered alternative is developed, extracorporeal artificial liver devices will remain the standard of care to bridge patients until donor liver, or alternative therapies, become available.

21.5 Urinary System The most common causes of ESRD in the adult population include diabetic nephropathy, hypertension, and glomerular disease. Data accumulated by the US Renal Data System (http://www.usrds.org/) also indicates that the rate of pediatric ESRD has tripled since 1980. Currently, nearly 80% of individuals awaiting organ donation are on the waitlist for a kidney. New regenerative therapies are clearly needed, and it is likely that different diseases will require targeted strategies to ultimately be effective [65, 66]. Hemodialysis is the current standard of care but, while life sustaining, has limitations as it does not mimic kidney function, and morbidity and mortality remains high [67]. A bioartificial implantable kidney, or Renal Assist Device, has been proposed as an alternative to dialysis and whole organ transplantation [68, 69]. Further understanding of the response of the kidney to acute and chronic injury is needed to guide effective regenerative strategies [65].

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21.5.1 Kidney The causes of renal failure and the determinants of progression to ESRD in children differ from those in the adult population; in children the major cause and need for dialysis and transplantation is developmental anomalies of the kidneys. Although congenital obstructive nephropathy, for example, is a major cause of chronic renal failure in children, little progress has been made toward developing effective methods for treatment [70, 71]. Studies have been initiated to develop new ways to address kidney regeneration in the rhesus monkey model of obstructive renal disease, a model that closely recapitulates findings in humans [70, 72]. Although other animal models have provided important information on the pathogenesis of this disease, the monkey model has distinct developmental similarities when compared to humans, such as the period of nephrogenesis, which begins in both species in the late first trimester and continues throughout the mid-third trimester. Monkeys and humans share many characteristic features because of their close phylogenetic relationship, and similarities in development, physiology, immunology, and anatomy provides the opportunity to explore translational strategies with high clinical relevance. Acellular kidney scaffolds have been developed from rhesus monkey kidneys across the lifespan (fetal to adult) and shown to retain critical ECM proteins and well-preserved morphology [72]. The scaffolds have also been repopulated using a renal explant/scaffold model, as well as isolated glomeruli, tubular fragments, or single dissociated cells from different aged donors, and compared to recellularization with human embryonic stem cells [72, 73]. These studies showed that the explants and cell populations recellularized the scaffolds in an age-dependent manner.

21.5.2 Bladder and Urethra Pediatric diseases such as urethral obstruction and congenital bladder anomalies can lead to glomerular injury and renal failure if surgical augmentation cystoplasty is not performed. Other congenital abnormalities, such as spina bifida and hypospadias, frequently require surgical reconstruction; nonurologic specimens are generally used including gastrointestinal segments, which are not without complications [74]. Xenogeneic ECM has been explored for approximately two decades, typically using pig organs and bladder for urogenital surgery [75]. Studies in pigs, for example, demonstrated success when allogeneic SIS replaced a ureteral segment [76, 77] and further investigations indicated that preseeding grafts enhanced patency [78]. Oberpenning et  al. [79] were the first to develop autologous urinary bladders from PGA/PLGA-molded polymer scaffolds which were seeded with smooth muscle and urothelial cells. These constructs were implanted into a dog model where the engineered bladders demonstrated capacity to retain urine and developed a trilayer structure of smooth muscle bundles covered by a sheath of submucosa and uroepithelial lining by 3 months, and ingrowth of neural tissue by 6 months. Further studies demonstrated no evidence of toxicity or immune rejection using this approach [80].

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Reddy et  al. [81] reconstructed the bladder wall in a pig model using acellular matrix allografts. These studies showed calcification and approximately 30% tissue shrinkage. Further studies in pigs have focused on the use of bovine collagen I scaffolds seeded with autologous fibroblasts and keratinocytes, compared to bladder reconstruction using complete or mucosa-free ileum patches [82]. The tissue-engineered bladders persisted for 3  months and showed less distensibility when compared to native tissue, and with approximately 6% shrinkage. More recently, Yang et al. [83] showed that preservation of bioactive factors such as PDGF and VEGF were associated with migration and proliferation of human bladder smooth muscle cells as well as umbilical cord endothelial cells within a porcine acellular matrix. Loai et al. [84] further showed neovascularization and increased urothelium and smooth muscle generation of porcine bladder acellular-HA and VEGF incorporated scaffolds after transplant into pigs. A new technique for ureteral reconstruction in pigs has been reported by Wolters et al. [85] where the external jugular vein was used for an autologous graft to replace a segment of the ureter with and without an endoluminal biodegradable PLLA stent. The biodegradable stent served as a scaffold for urothelial ingrowth into the vein. After 6 months, the graft was relined with urothelium and morphologically resembled the native ureter, and showed similar function. The authors noted that further studies are needed to ensure that strictures do not develop over time.

21.6 Musculoskeletal System Tissue engineering has emerged as a promising approach to overcome significant clinical challenges in the treatment of musculoskeletal disorders, using a variety of natural, synthetic, and mixed composites [86] and animal models [87, 88]. Tissuespecific cell delivery has been considered for articular cartilage and meniscus repair, osteoarthritis and rheumatoid arthritis, segmental bone defects and nonunions, ligament and tendon repair, osteonecrosis, intervertebral disc repair, and spinal fusion. Platelet-rich plasma is an autologous blood product which has been used in conjunction with engineering approaches because it is a rich source of growth factors and cytokines (e.g., PDGF and VEGF) [89]. Clinical trials in progress use matrixbased delivery of cells such as MSC for bone and cartilage repair, and in some cases include bone morphogenetic proteins, many of which were initially tested in large animal models (http://clinicaltrials.gov/).

21.6.1 Muscle, Tendons, and Ligaments With a focus on accelerating tendon-bone healing, Derwin et al. [90] used a dog model to demonstrate augmentation of acute repair of rotator cuff tendons with a PLLA repair device that provided a tendon-bone scaffold for host tissue deposition and ingrowth, improved functional and biomechanical outcomes, and enhanced healing.

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ACL tear is one of the most common knee injuries, and recent studies have focused on combining scaffold/cytokine constructs to replace traditional grafts. Collagenplatelet composites have been implanted into pig ACL to simulate the wound healing of a fibrin clot [91]. After 3  months, improvements in yield load and linear stiffness were observed as well as increased cell density. ECM from pig SIS has also been used for musculotendinous repair and has been shown to degrade 3 months after implantation in a dog model of Achilles tendon repair, with remodeling into collagen-rich connective tissues similar to the native tendon [92].

21.6.2 Bone and Cartilage The use of bioceramic scaffolds in dog, sheep, and goat models has been reviewed by Cancedda et al. [93], and improved healing noted when scaffolds were seeded with bone marrow cells. Initial bone formation has also typically been shown at the periphery likely reflecting limited vascularization of the inner portions of the constructs. To address this concern, Beier et al. [94] used the A-V loop sheep model to show axial vascularization of a clinically approved bone substitute, HA/ß-TCP mixed with a fibrin gel. Studies by Torigoe et al. [95] also addressed concerns that porous ceramics such as HA and TCP typically are insufficient to repair large bone defects, and showed that incorporation of bone marrow stromal cells into a TCP construct with autologous plasma promoted osteogenesis in a monkey heterotopic bone formation model. Platelet-rich plasma has also been used for bone healing of tibial defects in a goat model [96] where an injectable scaffold (TCP/chitosan) led to complete healing by 16 weeks. Improved mechanical strength, biocompatibility, and osteoinductive properties were shown when compared to controls. However, the combination of mineralized collagen and platelet-rich plasma had no significant effect when assessed in a sheep model of cervical discectomy and fusion [97]. A composite injectable scaffold consisting of cartilage microparticle acellular tissue matrix and fibrin glue combined with chondrocytes was delivered to loadbearing defects of autologous pig articular cartilage [98]. The repaired cartilage was found to be similar in hardness and appearance to that of native cartilage, and chondrocyte proliferation and angiogenesis were observed. Canine chondrocytes have also been encapsulated in agarose gels to form three-dimensional chondral constructs or combined with porous trabecular tantalum metal cylinders to form osteochondral constructs, and demonstrated mechanical properties and glycosaminoglycan content similar to native cartilage [99]. A study in pigs using a biodegradable PCL nanofibrous scaffold seeded with allogeneic chondrocytes or human MSC, or acellular PCL scaffolds, were used to repair iatrogenic, full-thickness cartilage defects [100]. Six months postimplant, the MSC-seeded constructs showed the most complete repair, with a smooth hyaline-like cartilage surface and mechanical properties similar to native hyaline cartilage. These studies also showed that the use of human MSC, which have been proposed to be immunosuppressive, did not result in graft rejection, which has been noted in studies with other species.

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21.6.3 Periodontal Bioengineering efforts focused on tooth and bone constructs have included the use of stem and progenitor cells from dental pulp, periodontium, and alveolar bone [101]. The minipig tooth/bone implant model was developed to generate functional, biological tooth substitutes. Dental stem cells from tooth buds have also been used to seed fabricated PGA/PLLA scaffolds with good results. Periodontal tissue engineering and the induction of cementogenesis have also been extensively studied in baboons [102].

21.7 Skin Substantial efforts over the past decades have focused on new engineering constructs to replicate human skin to promote and improve the quality of wound healing in acute and chronic conditions [103]. Tissue-engineered substitutes for autologous skin grafts have been proposed to provide “off-the-shelf” products, and include epidermal, dermal, and mixed substitutes, several of which are commercially available, while others remain in development. Generally, both use natural or synthetic scaffolds, sometimes in combination with cells [104]. While some of the newer constructs under development are first evaluated in vitro, in vivo studies are necessary to address effectiveness particularly for full-thickness skin wounds, and have included pigs and sheep. As noted in a recent review, the model chosen has a significant impact on translation of results to humans because wounds in rodents heal primarily by contraction, whereas pig skin is more comparable to human skin and heals by granulation [104]. Readers are referred to a comprehensive table in this publication [104] that highlights 79 studies performed in animal models, of which 21 included pigs and one was conducted in sheep. Two additional studies reported on an engineered skin substitute that was developed from epidermal keratinocytes and dermal fibroblasts obtained from pig skin biopsies grown in collagen/fibrin hydrogel scaffolds [105], and tested in a pig model with full-thickness skin defects [106]. These investigations showed that large-scale autologous grafts were feasible in this model and could potentially address the needs of both children and adults. The one sheep study noted in the above review focused on fetal tissue engineering for postnatal skin replacement [107]. Since this publication, other studies have addressed protective measures to alleviate the extent of neural damage that can arise with neural tube defects, and conditions where the developing spinal cord is chronically exposed to amniotic fluid. Such defects have been induced in lambs then repaired with either a collagen-based biomaterial or SIS and demonstrated minimal loss of function [108]. A similar study conducted in a full-thickness fetal sheep model with a biodegradable collagen scaffold loaded with VEGF and fibroblast growth factor-2 also showed increased vascularization at the wound site and enhanced healing [109].

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21.8 Neural and Visual Systems Animal models of neurodegenerative disorders, traumatic brain injury, and spinal cord injury have been developed, and several reviews have addressed the studies conducted on the regenerative potential of stem and progenitor cells [110, 111]. It has been proposed that because rodents show high rates of spontaneous recovery from induced injuries they may not accurately predict outcomes in humans emphasizing the need for large animal models for neural repair [112]. Substantial information has been gained in understanding posttraumatic ischemia and the role of inflammation as a contributing factor to secondary damage in spinal cord injury. However, replacing cells and lost tissue by filling the physical gap remains a challenge [113]. From a cellular perspective, one target has been demyelinated axons that traverse the injury site, and the use of cell substrates such as Schwann cells and oligodendrocyte progenitors, including those derived from human embryonic stem cells. Hydrogels have been used as three-dimensional tissue engineering scaffolds in the central nervous system and shown to fill an injury site and provide a growth-promoting environment; when combined with dextran and chitosan such hydrogels have been permissive for neurite extension [114]. Other materials studied include PGA, collagen, alginate, and synthetic gels such as poly hydroxypropyl methacrylamide. While these engineering approaches show promise, more work is needed to determine the best combination(s) for future human applications.

21.8.1 Spinal Cord Injury Old World nonhuman primate species such as macaques are well suited for spinal cord injury and repair studies compared to New World species because they better model humans when investigating recovery of fine motor skills [115]. A surgical model of acute spinal cord injury to parallel Brown-Séquard syndrome has been developed in the African Green monkey to evaluate biomaterials for repair [116]. PLGA scaffolds with and without human neural stem cells were shown to persist for at least 40 days, with degradation and clearance noted within 82 days postimplantation. A 20-point observational behavioral scoring system was used to assess functional recovery; all subjects studied exhibited a return to baseline neuromotor scores in the ipsilateral hindlimb. The authors noted that more definitive preclinical studies were needed to evaluate the efficacy of the PLGA scaffolds in the promotion of recovery and repair. Studies have also been conducted with cynomolgus macaques to evaluate a guidance channel seeded with autologous Schwann cells to promote regeneration of spinal nerve root axons [117]. Cells were initially obtained to seed polyacrylonitrile/polyvinyl chloride channels and placed within a laminectomy site. Modest regeneration and muscle reinnervation were found; the investigators noted that the use of new absorbable polymers and neurotropins might be more effective and improve outcome.

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21.8.2 Macular Degeneration Degeneration of retinal neural cells leads to age-related macular degeneration where loss of photoreceptors is the primary cause of vision loss. Scaffolds have been used for the treatment of retinal degeneration in animal models, using natural and synthetic polymers (e.g., collagen, alginate, HA, fibrin, PLGA, polycaprolactone) [118]. In vivo studies in pigs with pig and cow lens capsule or bovine corneal ECM as the biomaterial with pig RPE cells showed that positioning in the subretinal space was crucial for success [119]. Studies by Del Piore et al. [120] focused on the transplant of organized primary porcine RPE sheets with 50% gelatin placed into the subretinal space, and showed that the RPE survived with no evidence of inflammation for 3  months. However, the investigators concluded that further study was needed to produce uniform repopulation of a sizable portion of Bruch’s membrane with a monolayer of transplanted RPE. Monkey corneal endothelial cells have also been cultured in sheets on a collagen type I gel membrane and inserted into the anterior chamber in monkeys with induced corneal endothelial dysfunction [121]. Animals that received the cell sheets recovered corneal clarity that was maintained for up to 2 years in contrast to those that received only cells or collagen sheets alone.

21.9 Conclusions Animal models are clearly essential to understand biological functions, study complex human diseases, and address therapeutic efficacy and safety of new tissue engineering strategies proposed for use in humans of all age groups. As is evident by the current state of the field, translational studies conducted by multidisciplinary teams with relevant large animals and nonhuman primates are crucial to ensure that new ideas are appropriately tested and for a sufficient duration. A full understanding of the animal model chosen, including anatomical and physiological similarities and differences when compared to humans, remains an essential component in study design. The choice and age of the model is also an important consideration from a lifespan health perspective, and if reliable predictions of outcome in children and adults are to be achieved. Validation of animal models and standardization of technologies used across species is essential if newly developed engineering approaches are to be effectively translated for a host of human congenital and acquired diseases. One of the constraints that can impact the choice of species remains the reagents available. For example, for some species, there are limited cross-reactive antibodies and a paucity of relevant assays. Resources are readily available for nonhuman primates that describe available reagents to facilitate research with these species (Nonhuman Primate Reagent Resource; http://nhpreagents.bidmc.harvard.edu/ nhp/). The USDA also has an online Porcine Immunology and Nutrition database that addresses cross-reactive antibodies and assays; other similar resources for large animal species can be found online at several academic institutions.

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Of primary importance in translational research is the formation of research teams that can take a novel engineering idea from the discovery phase through preclinical investigation and ultimately to human clinical trials. Expertise in multiple areas such as engineering, the chosen animal model, and in the conduct of clinical trials provides the synergy needed to ensure good study design and a greater likelihood for success. NIH programs, such as the National Primate Research Centers and other nonhuman primate colonies and large animal resources (e.g., National Swine Resource and Research Center), provide opportunities for such multidisciplinary collaborations. Once initial studies have been completed and a promising approach developed, early discussions with the FDA and enlisting the aid of those with expertise in Good Manufacturing Practice, as well as industry alliances, can help to effectively move the research to the next level. The NIH CTSA program (http://www.ctsaweb.org/) was envisioned as a national mechanism to transform research discovery to efficient clinical practice, and aid in providing the environment necessary to support synergistic teams, train the next generation of translational and clinical investigators, and help overcome roadblocks and bottlenecks to clinical translation. The CTSA program facilitates research, encourages the development of national resource databases, and is engaged in consortium-wide activities that provide many of the tools and intellectual infrastructure necessary to bring novel tissue engineering ideas to human clinical practice. Acknowledgments  Dr. A.F. Tarantal is Staff Scientist and Unit Leader at the California National Primate Research Center (NIH RR00169), and directs the NHLBI Center for Fetal Monkey Gene Transfer for Heart, Lung, and Blood Diseases (NIH HL85794), and the Translational and Pilot Programs in the UC Davis Clinical and Translational Science Center (CTSC) (NIH RR024146). Ms. K.H. Nakayama is a predoctoral scholar in the UC Davis Stem Cell Training Program (CIRM T1-00006 and TG2-01163) and a former CTSC T32 trainee.

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About the Editor

Harold S. Bernstein, M.D., Ph.D., is Professor of Pediatrics (Cardiology) at the University of California San Francisco, where he is also a senior member of the UCSF Cardiovascular Research Institute and the Eli and Edythe Broad Center of Regeneration Medicine and Stem Cell Research at UCSF. Dr. Bernstein’s research focuses on stem cell biology, cardiac and skeletal muscle development, and tissue regeneration. He has received numerous awards recognizing his work, and is an elected member of the Society for Pediatric Research and the American Pediatric Society. He has also been recognized as a Fellow and an Established Investigator of the American Heart Association, and a Fellow of the American Academy of Pediatrics Section on Cardiology and Cardiac Surgery.

H.S. Bernstein (ed.), Tissue Engineering in Regenerative Medicine, Stem Cell Biology and Regenerative Medicine, DOI 10.1007/978-1-61779-322-6, © Springer Science+Business Media, LLC 2011

415

Index

A Acellular tissue matrices, 282 Acetylated low-density lipoprotein (Ac-LDL), 212, 213 Actin cytoskeletal remodeling, 86–87 Adipose tissue animal models large, 255, 259 mouse, 255–257 rat, 255, 258 sheep, 262, 264 autografts, 252 BAT vs. WAT, 253 biomaterials drug and growth factor delivery, 255 native scaffolds, 255 synthetic scaffolds, 254–255 cell sources, 254 critical factors, 259 in vivo soft tissue engineering approaches arteriovenous pedicles, 261, 263 hybrid implants, 260–261 microchannels and/angiogenic cues, 261, 262 MSC sources, 55–56 Adipose tissue derived stem cell (ASC), 254, 256–259 Adult stem cells, hollow organ engineering, 278–279. See also Adipose tissue AFPS cells. See Amniotic-fluid and placentalderived stem (AFPS) cells Alkaline phosphate (ALP), 105, 106, 108, 112, 113 Alternatively activated macrophage (M2), 367 American Association of Blood Banks, 12, 13 American Society of Blood and Marrow Transplant, 12 Amniotic-fluid and placental-derived stem (AFPS) cells, 277–278

Amyotrophic lateral sclerosis (ALS), 45, 59–60 Amyotrophic lateral sclerosis, MSC therapy, 59–60 Anterior cruciate ligament (ACL) tear, 403 Arginine-glycine-aspartic acid (RGD) adipose tissue regeneration, 262 bioactive signals, biomaterials, 130, 131 bone healing, 112 laser-based layer-by-layer (LbL) stereolithography technique, 134 mechanical conduit, gene expression regulation, 88 Arteriovenous (A-V), 261–262 ATP-binding cassette sub-family G member 2 (ABCG2), 167 Atrial myosin light chain 2 (MLC2a), 27, 171, 184 Atrial natriuretic factor (ANF), 170 Autoimmune encephalomyelitis, 57 B Basic fibroblast growth factor (bFGF) adipogenesis and/angiogenesis, 255 adipose tissue regeneration, 258, 261–263 growth factor, skeletal muscle niche, 200 hESCs, myocardial repair, 26, 27, 169–171 mesenchymal stromal cells, 60, 62 PEG–heparin complex, 132 secretion factors, MSC, 60 skeletal muscle niche, 200 BAT. See Brown adipose tissue (BAT) Bioartificial liver (BAL), 307, 308, 310 Biomaterials adipose tissue, 254–255 design rules, stem cell phenotype bioactive signals incorporation, 129–132 417

418 Biomaterials (cont.) biomaterials translation, 127–128 epidermal growth factor (EGF), 131 gastrulation, 126 mechanics cues, developmental biology, 126 mechano-homeostasis, 125 mechanotransduction, 125–126 mesenchymal stem cells (MSC), 128, 134–136 poly-ethylene glycol (PEG), 134 scaffolding proteins, 125 sonic hedgehog (Shh), 136 stem cell differentiation mechanics, 128 stem cell fate, 134–135 temporal control, 135–136 vascular endothelial growth factor (VEGF), 130–132 x, y (2D) and x, y, z (3D),coordinate control, 133–134 hollow organ engineering acellular tissue matrices, 282 alginate, 282 collagen, 281–282 composite scaffolds, 283 design and selection, 280–281 function, 280 polyesters, 282–283 host immune response, constructive remodeling acute inflammation, 358–359 anatomical site, 370–371 bioactive factors, 364 chronic inflammation, 359 innate immunity and macrophages, 366–370 processes, 358 surgical implantation, 358 T cells, 365–366 Th1 and Th2 polarized immune responses, 365 urinary bladder matrix, 361–362 Vicryl (Polyglactin 910) mesh, 360–361 regenerative medicine, 92–93 Bioreactor hepatocyte culture techniques aggregate culture, 303–305 encapsulation, 303 flat membrane system, 302 hollow fiber system, 302–303 Blood flow restoration quantification. See Hind limb ischemia model Bone and marrow transplantation (BMT), 4–10, 12, 13

Index Bone defect model fracture methodology, 383 schematic diagram, 384 Bone-marrow derived stromal cells (BMSCs), 106, 108, 112, 243 Bone morphogenetic protein 4 (BMP4), 24–26, 28, 170, 171 Bone regeneration, synthetic matrices bone healing, 103–104 carbon nanotubes, 108 core-shell electrospinning, 108–109 designing, 102–104 healing, 103–104 remodelling, 103 structure, 102–103 dual porosity, 109–111 electrospun fibers composite nanofibers, 106–109 inorganic nanofibers, 106 polymeric nanofibers, 104–106 matrix sophistication, 104–114 nanofibers composite, 106–109 inorganic, 106 natural and synthetic, 104–106 surface functionalisation calcium phosphate coatings, 112–113 cellular-based matrices, 114 electrospin/electrospray systems, 113 hydrophilicity, 111 layer-by-layer surface modification, 113–114 signalling factors, 112 Bone sialoprotein–2 (BSP), 103, 107 Brain-derived neurotrophic factor (BDNF), 25, 27, 59, 60, 64 Brown adipose tissue (BAT), 253 C Carbodiimide, 131, 281 Cardiomyocyte (CM) embryonic, 81 LEOPARD syndrome iPSC-derived, 46 mesodermal hESC derivatives, 26–27 morphology changes, 91 N-cadherin and connexin 43, 82 pluripotent stem cells, 168–176 postmitotic cell types, 23 Cell sheets biodegradable scaffolds method, 144 clinical applications corneal regeneration, 146–147 esophageal surgery, 147–148

Index pleural defect, 149–151 free thyroxine (fT4), 154 free triiodothyronine (fT3), 154 islet cell tissue, 154–157 lower critical solution temperature (LCST), 145 LSCD, 146 poly(N-isopropylacrylamide) (PIPAAm), 145, 146, 156 primary hypothyroidism, 152 single cell suspension injection, 144 Stevens-Johnson syndrome, 146 temperature responsive polymer, 145–146 thyroid tissue, 152–154 Chronic obstructive pulmonary disease (COPD), 304 Closed transverse fracture model. See Bone defect model Crohn’s disease activity index (CDAI), 58 D Dacron, vascular grafts, 212 Deoxycholic acid (DOA) Diabetes mellitus, MSC therapy, 58–59 Diabetic foot ulcer (DFU), 384, 385 Dickkopf homolog–1 (Dkk1), 24, 27, 59, 169–171 Dimethyl sulfoxide (DMSO), 65 DMD. See Duchenne muscular dystrophy (DMD) Duchenne muscular dystrophy (DMD), 197, 198 E Ectodermal hESCs derivatives, 27–28 Embryonic stem cells (ESCs) hollow organ engineering, 275 human (See Human embryonic stem cells (hESCs)) vs. iPSCs, 42–43 SCNT, 40 Endodermal hESCs derivatives, 23–26 Endothelial colony forming cell (ECFC), 399 Endothelial progenitor cell (EPC), 214 End-stage renal disease (ESRD), 299, 400, 401 Engineered muscle tissue gel matrix supports, 203–204 properties, support material, 203 scaffold supports, 204–205 Engineering complex synthetic organs esophagus

419 intrathoracic replacement, 301–302 layers, 301 synthetic and natural scaffolds, 301 kidney artificial membrane development, 299–300 decellularized natural kidney matrices, 300 design considerations, 300 ESRD, 299 lung acellular lung scaffolds, 307–309 bioreactor-based decellularization, Sprague–Dawley rats, 307–309 COPD, 304 degradation, scaffold material, 305–306 natural and synthetic polymers, 304–305 perfusion decellularization, 306–307 scaffold selection, 304 SLPC/PGA constructs, 305, 306 production approaches, 298–299 trachea airway prostheses, 302–303 airway stenosis, 302 design approaches, 303 Epidermal growth factor (EGF), 25, 27, 47, 131, 200 ESCs. See Embryonic stem cells (ESCs) Esophagus intrathoracic replacement, 301–302 layers, 301 synthetic and natural scaffolds, 301 European Group for Blood and Marrow Transplantation, 12, 62 Extracellular matrix (ECM) autografts, 224 hESCs, 25 mechanotransduction, 78–81 vascular grafts, 213, 216 F Federation for Accreditation for Cellular Therapy, 12, 13 Fibrin, 204 Fluorescence resonance energy transfer (FRET), 174 Fluoroapatite (FAp), 106, 107 Food and Drug Administration (FDA), 13, 29, 30, 60, 61, 65, 105, 219, 252, 281, 282

420 Full skin defect model advantages, 385–386 biodegradable SDR, 384 diabetic foot ulcer (DFU), 384, 385 schematic diagram, 385 techniques, scaffold vascularization, 385 VesSeg-Tool program, 386 Functional bone graft engineering allografts, 225 autografts, 224–225 biology and structure development, 223 healing, 223–224 remodeling, 223 challenges integration into host tissues, 230 long-term mechanical function, 231 vascularization, 231 clinical translation requirements anatomical shape, 229–230 cell sources and functions, 229 preparation time, 230 graft substitutes, 225 ideal characteristics, 225 in vitro culture systems, 228 in vivo models, 228–229 medical complications, 224 process of cells, 226–227 environmental factors, 227 scaffolds, 227 Functional cartilage grafts engineering agarose hydrogel system, 240–241 articular cartilage biology and structure, 238–239 injury, 239 repair, 239–240 challenges anatomical shape, 244 integration with surrounding tissue, 244–245 size and scale up, 244 chemical stimulation, 242 clinical translation requirements cell expansion, 243 cell sources, 242–243 long term graft storage, 243 mechanical preconditioning, 241 G Genetically-modified mesenchymal stromal cells, 63–64 Glycosaminoglycan (GAG), 129, 238, 241, 244

Index Good Manufacturing Practice (GMP), 62, 63 Graft-versus-host disease (GVHD), 7–8 Graft-versus-leukemia (GVL), 7, 9, 11 Granulocyte macrophage colony-stimulating factor (GM-CSF), 10, 61, 223 Green fluorescent protein (GFP), 173, 183, 184 H Hematopoietic stem cell transplantation advantage, 13 allogeneic marrow cells, 6 autologous, 11–12 dog model system, 7 graft-versus-host disease (GVHD), 7–8 graft-versus-leukemia (GVL), 9 HLA matching, 8, 10–11 human cytokines, 10 leukemia vs. aplastic anemia, 6–7 peripheral blood stem cells (PBSCs), 10 radiation, 4–5 regulatory agencies, 12–13 skin graft, 5 syngeneic marrow cells, 6 umbilical cord blood (UCB), 11 Heparan sulfate, 200 Hepatocyte growth factor (HGF), 24, 26, 56, 57, 61, 200 hESCs. See Human embryonic stem cells (hESCs) Hind limb ischemia model, 386–387 Hollow organ engineering biomaterials design and selection, 280–281 function, 280 types, 281–283 native targeted progenitor cells, 279–280 regenerative medicine strategy categories, 274 urethra, 286–288 urinary bladder, 283–286 uterus, 288 vagina, 289 stem cells adult, 278–279 AFPS cells, 277–278 embryonic, 275 induced pluripotent, 277 properties, 275 therapeutic cloning, 276 Host immune response, constructive remodeling biomaterials

Index acute inflammation, 358–359 anatomical site, 370–371 bioactive factors, 364 chronic inflammation, 359 innate immunity and macrophages, 366–370 processes, 358 surgical implantation, 358 T cells, 365–366 Th1 and Th2 polarized immune responses, 365 urinary bladder matrix, 361–362 Vicryl (Polyglactin 910) mesh, 360–361 tissue injury hemostasis, 355–356 inflammation and proliferative phase, 356 remodeling phase, 356–358 HUCPVCs. See Human umbilical cord perivascular cells (HUCPVCs) Human embryoid body (hEB) chondrocyte formation, 29 hESC-derived CM engraftment, 183, 184 myocardial repair, hESCs, 169–174 Human embryonic stem cells (hESCs) cell derivatives ectodermal, 27–28 endodermal, 23–26 mesodermal, 26–27 characteristics, pluripotent (see Pluripotent human embryonic stem cells) clinical use requirements, 29, 30 pluripotency test, 21–22 safety concerns directed differentiation, and purification protocols, 31–32 enrichment, 31–32 genetic abnormalities, hESC lines, 31 molecular beacons, 32 transplanted hESC-derived cells, immune rejection, 32–33 xenobiotic-free conditions, 30–31 sources and derivation inner cell mass, blastocyst-stage pre-implantation embryo, 18, 19 parthenogenetic embryos, 18 parthenote-blastocysts, 19 single blastomere, 18 tissue engineering, 28–29 types, 18 Human lymphocyte antigen (HLA), 8–11, 19, 20, 30, 32, 33, 55–57

421 Human umbilical cord perivascular cells (HUCPVCs), 56–57 Hydroxyapatite (HA), 225, 227 I Immune-deficient mouse models, tissue damage bone defect model fracture methodology, 383 schematic diagram, 384 full skin defect model advantages, 385–386 biodegradable SDR, 384 schematic diagram, 385 techniques, scaffold vascularization, 385 VesSeg-Tool program, 386 hind limb ischemia model, 386–387 mouse strains BALB/c, 380 C17/SCID mice, 381 mesenchymal stem cell (MSC), 382 NOD/SCID mice, 381 translation to humans, 387–388 Immune modulation, stem cell therapy antithymocyte globulins, 328 CD28 cells, 320–321 CD8+ T cells, 321 central tolerance, 320 complement system, 319 cytokines neutralization, 329–330 dendritic cells, 319 embryonic stem cell (ESCs), 330 graft injury, 327 humoral and cellular module, 318 immunological tolerance, 327 islet transplantation, 327–328 NK cells, 319–320 rituximab, 328 stem cells embryonic stem cell (ESCs), 325 induced pluripotent stem cell (iPSCs), 326 mesenchymal stem cell (MSCs), 325–326 T and B cells, 320 transplantation antigens ABO blood group antigens, 322 allogeneic transplants, 323 human leukocyte antigens, 322 minor histocompatibility antigens, 323–324 passenger leukocytes, 322–323 Tregs, 320

422 Immunoglobulin (Ig), 326, 338, 339, 358, 371 Induced pluripotent stem cell alloimmune response, 342, 344 vs. amniotic fluid stem cells, 278 bone grafts, 229 cartilage grafts, 243 liver stem cells, 325 myocardial repair, 175–176 Induced pluripotent stem cells (iPSCs) disease models, 44–47 DNA methylation, 43 vs. ESCs, 42–43 generation, schematic overview, 40, 41 hollow organ engineering, 277 molecular profiling, 42–43 reprogramming efficiency, 43–44 Oct4, 44 viral-mediated transduction, 43 SCNT vs. ESC, 40 transdifferentiated cell potentials, 47 Inducible nitric oxide synthetase (iNOS), 368 Insulin-like growth factor–1 (IGF1) adipose tissue regeneration, 258 AT-MSCs, 56 skeletal muscle niche, 200 soluble factors, skeletal muscle niche, 200, 201 stroke, 60, 61 Intercellular contact-based (juxtacrine) mechanotransduction, 81–83 Interleukin (IL), 242, 323, 356 International Society for Cellular Therapy, 12, 55 International Society for Hematotherapy and Graft Engineering, 12 K Kidney artificial membrane development, 299–300 decellularized natural kidney matrices, 300 design considerations, 300 ESRD, 299 Kinase insert domain receptor (KDR), 168, 184 L Large animal models cardiovascular system grafts, 397–399 patches, 397

Index replacement valves, 395–396 digestive system, 400 musculoskeletal system bone and cartilage, 403 muscle, tendons, and ligaments, 402–403 periodontal, 404 neural and visual systems macular degeneration, 406 spinal cord injury, 405 respiratory system, 399 skin, 404 urinary system bladder and urethra, 401–402 kidney, 401 Laser-based layer-by-layer (LbL) stereolithography technique, 134 Limbal stem cell deficiency (LSCD), 146 Liver regeneration aggregate culture, 303–305 animal models, liver disease, 311 bioartificial liver (BAL), 307, 308, 310 challenges, 312 ex vivo hepatocyte culture techniques aggregate culture, 303–305 flat membrane system, 302 hollow fiber system, 302–303 large scale, bioreactor, 302–305 microencapsulation, 303 small scale, 301–302 function, 299–300 hepatocytes, sources hepatocyte lines, 306 human donor livers, 305 primary pig hepatocytes, 306 stem cells, 306–307 implantable constructs scaffolding, 310 vascularization, 310 xenografts, 310–311 incubators,in vivo, 305 in vitro hepatic culture techniques, 301 cryopreservation, 301 isolation, 301 opportunities, 312 primary pig hepatocytes, 306 spheroid reservoir bioartificial liver (SRBAL), 308, 309 structure, 299 synthetic liver, application diagnostic, 307 implantable constructs, 309 therapeutic, 307–309

Index therapeutic cell mass, 308–309 extracorporeal devices, 307–308 membrane considerations, 308 oxygenation, 308 transplantation, 309 LSCD. See Limbal stem cell deficiency (LSCD) Lung acellular lung scaffolds, 307–309 bioreactor-based decellularization, Sprague–Dawley rats, 307–309 COPD, 304 degradation, scaffold material, 305–306 natural and synthetic polymers, 304–305 perfusion decellularization, 306–307 scaffold selection, 304 SLPC/PGA constructs, 305, 306 M Magnetic resonance imaging (MRI) amyotrophic lateral sclerosis (ALS), 59, 60 cardiac functional assessment, 165 patella fabrication, cartilage grafts, 244 postoperative graft surveillance, 217, 218 scaffold vascularization, 385 Major histocompatibility complex (MHC) immune modulation, stem cell therapy, 337–344 myocardial repair, 163, 172, 173 Matrigel, 203 Matrix metalloproteinase (MMP), 78, 92, 356 Mechano-homeostasis, 125 Mechanotransduction biomaterial design rules, stem cell phenotype, 125–126 Mesenchymal stromal cells (MSCs) biomaterial design rules, 128, 134–136 definition, 55 genetically-modified, 63–64 manufacture consideration culture expansion, 62 FBS, 62 GMP-grade, 63 platelet lysate, 62 traditional growth method, 63 and safety concerns, 64–65 sources adipose tissue, 55–56 HUCPVCs, 56–57 immune modulation, 57 umbilical cord blood, 56

423 therapy autoimmune diseases, 57–59 cardiovascular disease, 61–62 neurodegenerative diseases, 59–61 Mesodermal hESCs derivatives cardiomyocytes, 26–27 hematopoietic progenitor cells, 26 TGFb signaling pathway activation, 26 MicroRNA (miR), 20, 21, 30, 171, 172 MSCs. See Mesenchymal stromal cells (MSCs) Mucopolysaccharidosis Type VII (MPSVII), 381, 382 Multi-drug resistance-like protein 1 (MDR1), 167, 168 Multielectrode array (MEA), 165, 182 Multiple sclerosis, MSCs therapy, 59 Multiwall nanotube (MWNT), 108 Murine embryonic stem cell (mESC), 134, 171, 172, 307, 309 Myocardial infarction (MI) cardiac patches, 395, 397 c-kit and Sca–1, 167 electrophysiological assessment, myocardial repair, 182 hCM transplantation, 175 mechanical characteristics, tissue differentiation, 90 xenogeneic infusions, human MSCs, 57 Myocardial repair and restoration animal models, assessment electrophysiological, 182–183 hemodynamic, 179–182 cardiomyocyte (CM) application, 174–175 epigenetic regulation, 171–172 hESC-derived, 170–171 mechanical force, 171 purification, 172–173 cardiospheres, 167–168 cell sources engraftment, 176–177 mesenchymal stem cells, 163–166 pluripotent stem cells, 168–176 resident cardiac stem cells, 166–168 c-kit and Sca–1 markers, 167 side population cells, 167 stem cells human embryonic, 168–169 mesenchymal, 163–166 pluripotent, 168–176 resident cardiac, 166–168 tissue patches, 177–179

424 Myocyte enhancer factor–2 (Mef2), 47, 170, 171, 176 a-Myosin heavy chain (aMHC), 170–173 N National Institutes of Health (NIH), 147, 407 Native targeted progenitor cells, 279–280 Natural killer cell (NK) alloimmune response, stem cells, 342–344 antibodies, 346 hESCs, 24, 26 innate immune system, 336–338 MSCs, 57 NOD/SCID mice strain, 381 N-(3-dimethylaminopropyl)-N¢ethylcarbodiimide hydrochloride (EDC), 131 Nerve growth factor (NGF), 25, 27, 59, 60, 111 Neurodegenerative diseases, MSC therapy amyotrophic lateral sclerosis, 59–60 Parkinson’s disease, 60 spinal cord injuries, 61 stroke, 60–61 N-hydroxysulfosuccinimide (sulfo-NHS), 131 Nitric oxide (NO), 57, 215, 344 Notch signal transduction pathway, 82 O Octamer-binding transcription factor 4 (Oct4), 20, 21, 32, 44, 47, 56, 175, 277 Osteoarthritis (OA), 29, 101, 239, 402 Osteonectin (SPARC), 103, 224 Osteopontin (OP), 105, 228 Osteoprotegerin (OPG), 103, 107 P Parkinson’s disease (PD), 46, 59, 60, 336 PEG-based hydrogels, 202, 203 Peripheral arterial disease (PAD), 386, 395, 397 Peripheral blood stem cells (PBSCs), 10 Platelet-derived growth factor (PDGF), 132, 356 bone repair process, 224 cytokine release, 356 Hep attachment, 112 platelet-rich plasma, 402 preservation, bioactive factors, 402 signaling pathways, 223 Pluripotent human embryonic stem cells cell morphology and density, 19–20 epigenetic properties, 21 expression profiling, 20–21

Index Poly(DLlactide) (PDLLA), 15 Poly–caprolactone (PCL), 105–108, 110–112, 283, 403 Poly-ethylene glycol (PEG) based hydrogel, 202–204 vinylsulfone-functionalized, 130–131 Poly-glycolic acid (PGA) animal models, 255–259 collagen scaffold, 285 degradation rate, 305 naturally occurring a-hydroxy acids, 282 synthetic scaffolds, 254 Poly-lactic co-glycolic acid (PLGA) advantages and disadvantages, 227 animal models, 255–259 bioactive signal incorporation, 132 naturally occurring a-hydroxy acids, 282 synthetic scaffolds, 254 Poly-L-lactic acid (PLLA) muscle tissue scaffold technologies, 204 periodontal, 404 simultaneous electrospinning, 109, 110 surface activation techniques, 111 Poly-N-isopropylacrylamide (PIPAAm), 145, 146, 156, 179 Polytetrafluroethylene (PTFE), 212, 397, 398 R Receptor activator of NF-kB ligand (RANKL), 103 Regulatory T cell (Tregs), 57, 58, 338, 347, 348 Restoring blood vessels. See Vascular grafts Retinal pigment epithelium (RPE), 25, 28, 406 S SCNT. See Somatic cell nuclear transfer (SCNT) Serum response factor (SRF), 86 Side population cell, 167 Skeletal muscle niche basic fibroblast growth factor (bFGF), 200 biology organization, 198–199 regeneration, 199 composition, 200 definition, 198 DMD, 197 engineered muscle tissue gel matrix supports, 203–204 properties, support material, 203 scaffold supports, 204–205

Index function, fibrin, 204 gel matrix supports fibrin, 204 matrigel, 203 shortcomings, 204 insulin-like growth factor–1 (IGF1), 200 myogenic potential, 202 PEG-based hydrogels, 202, 203 satellite cells microenvironment, 201, 202 schematic representation, 201 scaffold supports degrapol, 205 PLLA, 204 shortcoming, 204 self-renewal, 202 transplantation, myogenic cells, 199–200 Small intestinal submucosa (SIS), 284, 285 Somatic cell nuclear transfer (SCNT), 32, 40, 41, 48, 276 Sonic hedgehog (Shh), 27, 136 Stem cell differentiation actin cytoskeletal remodeling, 86–87 cell geometry and cytoskeletal dynamics actin cytoskeletal remodeling, 86–87 cellular organization, 84 cellular shape and function, 84–86 gene expression,nuclear mechanism ECM vs. nucleus, mechanical continuity, 88 nuclear shape effect, 89 integrin-ECM interphase, 79 mechanical cues opportunities and challenges, 94 tissue formation, 89–94 mechanics, biomaterial design rules, 128 mechanotransduction cell-extracellular matrix interactions, 78–81 intercellular contact-based (juxtacrine), 81–83 microenvironment cellular organization, 84 cellular shape, 84–86 signaling pathway Notch, 82 Wnt/b-catenin, 83 tissue formation, mechanical cues biomaterials, regenerative medicine, 92–93 computational modeling, 90–91 ECM substrates,fabrication, 91–93 maturation and function, 93–94 Stem cell fate. See Biomaterials; design rules,

425 stem cell phenotype Stevens-Johnson syndrome, 146 T T cell receptor (TCR), 338–340, 346 Temperature responsive polymer, 145–146 Temporomandibular joint (TMJ), 230, 244 T helper cell (Th), 339, 365, 366 Therapeutic cloning, hollow organ engineering, 276 Thrombospondin (TSP), 103 Thyroid transcription factor–1 (TTF1), 307, 308 Tissue inhibitor of metalloproteinase (TIMP), 356 Trachea airway prostheses, 302–303 airway stenosis, 302 design approaches, 303 Transforming growth factor-b (TGFb) angiogenic regulators, 223 bone repair process, 224 chemical stimulation, 242 Hep attachment, 112 MSC, 57, 61, 63 platelets, cytokine release, 355–356 signaling, 23, 26 Transplantation antigens ABO blood group antigens, 322 allogeneic transplants, 323 human leukocyte antigens, 322 minor histocompatibility antigens, 323–324 passenger leukocytes, 322–323 Troponin I (cTnI), 170, 172 Troponin T (cTnT), 27, 171, 173, 184 Tumor necrosis factor alpha (TNFa), 347, 365, 368 U Umbilical cord blood (UCB), 11, 56–59, 381, 382, 399 Urethra bladder-derived collagen-based matrix, 287 neourethras, 286–287 tissue engineering, collagen matrix, 287 tubularized urethral repairs, 287–288 Urinary bladder, 283–286 Uterus, 288 V Vagina, 289 Vascular endothelial growth factor (VEGF)

426 angiogenic regulators, 223 bioactive signal incorporation, 130–132 Hep attachment, 112 hESC CM differentiation, 170, 171 stroke, 60–61 vascularization, 231 Vascular grafts acetylated low-density lipoprotein (Ac-LDL), 212, 213 angiography, 213 animal studies, autologous cells bone marrow-derived cells, 214–216 vascular cells, 212–213 clinical studies first-in human application, 216 late term results, 217–219 midterm results, 216–217 computed tomography

Index dacron, 212 endothelial progenitor cell (EPC), 214 homografts and heterografts, 212 polytetrafluoroethylene (PTFE), 212 synthetic conduits, thrombosis, 212 Western blot analysis, 216 Vascularization assessment. See Full skin defect model Ventricular myosin light chain 2 (MLC2v), 171, 173, 174 Vicryl (Polyglactin 910) mesh, 360–361 W WAT. See White adipose tissue (WAT) White adipose tissue (WAT), 253 Wnt/b-catenin signal transduction pathway, 83