Nanoplatform-Based Molecular Imaging

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NANOPLATFORM-BASED MOLECULAR IMAGING

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NANOPLATFORM-BASED MOLECULAR IMAGING

Edited by

Xiaoyuan Chen Laboratory of Molecular Imaging and Nanomedicine National Institute of Biomedical Imaging and Bioengineering National Institutes of Health Bethesda, Maryland

A JOHN WILEY & SONS, INC., PUBLICATION

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C 2011 by John Wiley & Sons, Inc. All rights reserved. Copyright


Published by John Wiley & Sons, Inc., Hoboken, New Jersey. Published simultaneously in Canada. No part of this publication may be reproduced, stored in a retrieval system, or transmitted in any form or by any means, electronic, mechanical, photocopying, recording, scanning, or otherwise, except as permitted under Section 107 or 108 of the 1976 United States Copyright Act, without either the prior written permission of the Publisher, or authorization through payment of the appropriate per-copy fee to the Copyright Clearance Center, Inc., 222 Rosewood Drive, Danvers, MA 01923, (978) 750-8400, fax (978) 750-4470, or on the web at www.copyright.com. Requests to the Publisher for permission should be addressed to the Permissions Department, John Wiley & Sons, Inc., 111 River Street, Hoboken, NJ 07030, (201) 748-6011, fax (201) 748-6008, or online at http://www.wiley.com/go/permission. Limit of Liability/Disclaimer of Warranty: While the publisher and author have used their best efforts in preparing this book, they make no representations or warranties with respect to the accuracy or completeness of the contents of this book and specifically disclaim any implied warranties of merchantability or fitness for a particular purpose. No warranty may be created or extended by sales representatives or written sales materials. The advice and strategies contained herein may not be suitable for your situation. You should consult with a professional where appropriate. Neither the publisher nor author shall be liable for any loss of profit or any other commercial damages, including but not limited to special, incidental, consequential, or other damages. For general information on our other products and services or for technical support, please contact our Customer Care Department within the United States at (800) 762-2974, outside the United States at (317) 572-3993 or fax (317) 572-4002. Wiley also publishes its books in a variety of electronic formats. Some content that appears in print may not be available in electronic formats. For more information about Wiley products, visit our web site at www.wiley.com Library of Congress Cataloging-in-Publication Data: Nanoplatform-based molecular imaging / edited by Xiaoyuan Chen. p. ; cm. Includes bibliographical references and index. ISBN 978-0-470-52115-1 1. Molecular probes. 2. Diagnostic imaging. I. Chen, Xiaoyuan. [DNLM: 1. Molecular Imaging–methods. 2. Molecular Imaging–trends. 3. Molecular Probes–diagnostic use. 4. Nanoparticles–diagnostic use. 5. Nanotechnology–trends. WN 180 N1865 2010] QP519.9.M64N36 2010 616.07 54–dc22 2010007984 Printed in the United States of America eBook ISBN: 978-0-470-76703-0 oBook ISBN: 978-0-470-76704-7 10 9 8 7 6 5 4 3 2 1

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CONTENTS

Preface

ix

Acknowledgments

xi

Contributors

PART I

xiii

BASICS OF MOLECULAR IMAGING AND NANOBIOTECHNOLOGY

1. Basic Principles of Molecular Imaging

3

Sven H. Hausner

2. Synthesis of Nanomaterials as a Platform for Molecular Imaging

25

Jinhao Gao, Jin Xie, Bing Xu, and Xiaoyuan Chen

3. Nanoparticle Surface Modification and Bioconjugation

47

Jin Xie, Jinhao Gao, Mark Michalski, and Xiaoyuan Chen

4. Biodistribution and Pharmacokinetics of Nanoprobes

75

Nagesh Kolishetti, Frank Alexis, Eric M. Pridgen, and Omid C. Farokhzad

PART II

NANOPARTICLES FOR SINGLE MODALITY MOLECULAR IMAGING

5. Computed Tomography as a Tool for Anatomical and Molecular Imaging

107

Pingyu Liu, Hu Zhou, and Lei Xing

6. Carbon Nanotube X-Ray for Dynamic Micro-CT Imaging of Small Animal Models

139

Otto Zhou, Guohua Cao, Yueh Z. Lee, and Jianping Lu

7. Quantum Dots for In Vivo Molecular Imaging

159

Yun Xing

8. Biopolymer, Dendrimer, and Liposome Nanoplatforms for Optical Molecular Imaging

183

David Pham, Ling Zhang, Bo Chen, and Ella Fung Jones v

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9. Nanoplatforms for Raman Molecular Imaging in Biological Systems

197

Zhuang Liu

10. Single-Walled Carbon Nanotube Near-Infrared Fluorescent Sensors for Biological Systems

217

Jingqing Zhang and Michael S. Strano

11. Microparticle- and Nanoparticle-Based Contrast-Enhanced Ultrasound Imaging

233

Nirupama Deshpande and J¨urgen K. Willmann

12. Ultrasound-Based Molecular Imaging Using Nanoagents

263

Srivalleesha Mallidi, Mohammad Mehrmohammadi, Kimberly Homan, Bo Wang, Min Qu, Timothy Larson, Konstantin Sokolov, and Stanislav Emelianov

13. MRI Contrast Agents Based on Inorganic Nanoparticles

279

Hyon Bin Na and Taeghwan Hyeon

14. Cellular Magnetic Labeling with Iron Oxide Nanoparticles

309

S´ebastien Boutry, Sophie Laurent, Luce Vander Elst, and Robert N. Muller

15. Nanoparticles Containing Rare Earth Ions: A Tunable Tool for MRI

333

C. Rivi`ere, S. Roux, R. Bazzi, J.-L. Bridot, C. Billotey, P. Perriat, and O. Tillement

16. Microfabricated Multispectral MRI Contrast Agents

375

Gary Zabow and Alan Koretsky

17. Radiolabeled Nanoplatforms: Imaging Hot Bullets Hitting Their Target

399

Raffaella Rossin

PART III NANOPARTICLE PLATFORMS AS MULTIMODALITY IMAGING AND THERAPY AGENTS 18. Lipoprotein-Based Nanoplatforms for Cancer Molecular Imaging

433

Ian R. Corbin, Kenneth Ng, and Gang Zheng

19. Protein Cages as Multimode Imaging Agents

463

Masaki Uchida, Lars Liepold, Mark Young, and Trevor Douglas

20. Biomedical Applications of Single-Walled Carbon Nanotubes

481

Weibo Cai, Ting Gao, and Hao Hong

21. Multifunctional Nanoparticles for Multimodal Molecular Imaging

529

Yanglong Hou and Rui Hao

22. Multifunctional Nanoparticles for Cancer Theragnosis Seulki Lee, Ick Chan Kwon, and Kwangmeyung Kim

541

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23. Nanoparticles for Combined Cancer Imaging and Therapy

vii

565

Vaishali Bagalkot, Mi Kyung Yu, and Sangyong Jon

24. Multimodal Imaging and Therapy with Magnetofluorescent Nanoparticles

593

Jason R. McCarthy and Ralph Weissleder

25. Gold Nanocages: A Multifunctional Platform for Molecular Optical Imaging and Photothermal Treatment

615

Leslie Au, Claire M. Cobley, Jingyi Chen, and Younan Xia

26. Theranostic Applications of Gold Nanoparticles in Cancer

639

Parmeswaran Diagaradjane, Pranshu Mohindra, and Sunil Krishnan

27. Gold Nanorods as Theranostic Agents

659

Alexander Wei, Qingshan Wei, and Alexei P. Leonov

28. Theranostic Applications of Gold Core–Shell Structured Nanoparticles

683

Wei Lu, Marites P. Melancon, and Chun Li

29. Magnetic Nanoparticle Carrier for Targeted Drug Delivery: Perspective, Outlook, and Design

709

R. D. K. Misra

30. Perfluorocarbon Nanoparticles: A Multidimensional Platform for Targeted Image-Guided Drug Delivery

725

Gregory M. Lanza, Shelton D. Caruthers, Anne H. Schmieder, Patrick M. Winter, Tillmann Cyrus, and Samuel A. Wickline

31. Radioimmunonanoparticles for Cancer Imaging and Therapy

755

Arutselvan Natarajan

PART IV TRANSLATIONAL NANOMEDICINE 32. Current Status and Future Prospects for Nanoparticle-Based Technology in Human Medicine

783

Nuria Sanvicens, F´atima Fern´andez, J.-Pablo Salvador, and M.-Pilar Marco

Index

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PREFACE

This book focuses on the rational design of water-soluble, biocompatible nanoparticles for the visualization of the cellular function and follow-up of the molecular processes in living organisms without perturbing them. Molecular imaging probes based on nanotechnology hold great potential in diagnosis, imaging guided intervention, and treatment response monitoring of diseases. This book is logically organized by including the basics of molecular imaging, general strategies of particle synthesis and surface chemistry, applications in computed tomography (CT), optical imaging, magnetic resonance imaging (MRI), ultrasound, multimodality imaging, and theranostics, and finally clinical perspectives of nanoimaging. This comprehensive title provides expert opinions on the latest developments in molecular imaging using nanoparticles. This book consists of 32 chapters and was contributed by nearly 100 authors worldwide, who are among the world’s prominent scientists in material science and/or molecular imaging. Part I consists of Chapters 1–4 Chapter 1 describes the basic principles of molecular imaging, how nanoparticles can be applied to different molecular imaging modalities, and challenges in developing nanoparticle-based molecular imaging probes; Chapter 2 highlights the general strategies to produce narrowly dispersed nanomaterials for molecular imaging; Chapter 3 emphasizes the importance of surface modification to render nanoparticles biocompatible and suitable for molecular imaging applications; and Chapter 4 talks about the toxicity and factors such as size, shape, coating, and surface charge that affect the biodistribution and pharmacokinetics of nanoprobes. Part II consists of Chapters 5–17 Chapter 5 illustrates the basic principles of CT, the evolution of CT imaging technology, and the rationale for nanoparticle-based CT contrast agents; Chapter 6 describes the advantages of fascinating carbon nanotube field emission X-ray technology over conventional thermionic X-ray tubes that are used in current X-ray imaging systems; Chapter 7 describes the use of unique optical properties of semiconductor quantum dots (QDs) for near-infrared fluorescence imaging in living animals; Chapter 8 introduces macromolecular nanoconstructs such as biopolymers, dendrimers, and liposomes as carriers for fluorophore conjugation and optical imaging; Chapter 9 summarizes recent progress in developing nanoplatforms for Raman imaging of biological systems; Chapter 10 summarizes the work in using single-walled carbon nanotubes (SWNTs) as near-infrared fluorescent sensors for biomolecule detection; Chapter 11 describes the use of micro- and nanoparticles as ultrasound contrast agents; Chapter 12 proposes the use of metal nanoparticles in ultrasound-based photoacoustic and magnetoacoustic imaging modalities; Chapter 13 reports the progress on magnetic resonance imaging (MRI) contrast agents based on inorganic nanoparticles; Chapter 14 emphasizes the use of iron oxide nanoparticles for cellular labeling followed by T2 - and T∗2 -weighted MRI; Chapter 15 covers the use of rare earth based nanoparticles for MR imaging as positive contrast agents; Chapter 16 reviews the top–down microfabrication technology to synthesize multispectral MRI contrast agents; ix

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PREFACE

and Chapter 17 gives an overview of the strategies to label nanoparticles with radionuclides to study in vivo distribution. Part III consists of Chapters 18–31 Chapter 18 introduces techniques to incorporate imaging agents into lipoproteins and to reroute lipoproteins to cancer specific epitopes; Chapter 19 exemplifies the use of protein cages such as virus capsids and ferritins as platforms for MRI contrast agents and fluorescent imaging agents; Chapter 20 provides a comprehensive summary of the state-of-the-art of SWNTs for multimodality biomedical imaging applications; Chapter 21 reviews the progress in the controlled synthesis, surface modification, and multimodality imaging applications of multifunctional nanoparticles in recent years; Chapter 22 argues the use of cancer theranostics as a promising new strategy in cancer management, permitting simultaneous cancer diagnosis, drug delivery, and real-time monitoring of therapeutic efficacy; Chapter 23 provides more examples of multifunctional nanoparticles for combined cancer imaging and therapy (theranostics); Chapter 24 describes the recent progress in modifying magnetic nanoparticles for multimodality imaging as well as targeted treatment of a number of diseases; Chapter 25 introduces gold nanocages as contrast agents for optical bioimaging (such as optical and spectroscopic coherence tomography amd photoacoustic tomography) and photothermal treatment; Chapter 26 describes the biological inertness, ease of manufacture and bioconjugation, and presumed lack of toxicity of gold nanoparticles for simultaneous sensing, imaging, and treatment of tumors; Chapter 27 presents the recent developments in the chemistry and photophysics of gold nanorods and their applications toward biological imaging and photothertmally activated therapies; Chapter 28 describes a number of gold core–shell nanostructures for cancer molecular optical imaging, controlled drug delivery, and photothermal ablation therapy; Chapter 29 describes a novel temperature and pH-responsive magnetic nanocarrier that combines tumor targeting and controlled drug release capabilities; Chapter 30 deals with perfluorocarbon nanoparticles as a multidimensional platform for targeted image-guided drug delivery; and Chapter 31 describes the use of radiolabeled nanoparticles and radiolabeled immunonanoparticles for imaging and therapy. Part IV is the concluding Chapter 32 that highlights some of the nanoparticle-based novel technologies for molecular imaging, diagnosis, and drug delivery formulations. The limitations and future challenges of nanoparticle-based systems are also discussed. Bethesda, Maryland

Xiaoyuan Chen

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ACKNOWLEDGMENTS

The editor thanks the nearly 100 authors throughout the world for their contributions and collaboration on this book project. The editing work of this book was accomplished using a significant amount of the editor’s spare time including family time. Therefore the editor also thanks his wife, Michelle Ji, and his daughter, Grace Chen, for their wonderful support and understanding.

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CONTRIBUTORS

Frank Alexis, Department of Bioengineering, Clemson University, Clemson, South Carolina, USA Leslie Au, Department of Biomedical Engineering, Washington University, St. Louis, Missouri, USA Vaishali Bagalkot, School of Life Sciences, Gwangju Institute of Science and Technology, Gwangju, South Korea R. Bazzi, Laboratoire Physico-Chimie des Electrolytes, Colloides et Sciences Analytiques, Universit´e Pierre et Marie Curie, Paris, France C. Billotey, Laboratoire CREATIS–Animage, Universit´e Claude Bernard, Lyon, France S´ebastien Boutry, Department of General, Organic and Biomedical Chemistry, NMR and Molecular Imaging Laboratory, University of Mons, Mons, Belgium J.-L. Bridot, Service de Chimie G´en´erale, Organique et Biom´edicale, Laboratoire de RMN et d’Imagerie Mol´eculaire, Universit´e de Mons-Hainaut, Mons, Belgium Weibo Cai, Departments of Radiology and Medical Physics, School of Medicine and Public Health, University of Wisconsin–Madison, and University of Wisconsin Carbone Cancer Center, Madison, Wisconsin, USA Guohua Cao, Department of Physics and Astronomy, University of North Carolina at Chapel Hill, Chapel Hill, North Carolina, USA Shelton D. Caruthers, Department of Medicine, Washington University Medical School, St. Louis, Missouri, and Philips Healthcare, Andover, Massachusetts, USA Bo Chen, Center for Molecular and Functional Imaging, Department of Radiology and Biomedical Imaging, University of California, San Francisco, California, USA Jingyi Chen, Department of Biomedical Engineering, Washington University, St. Louis, Missouri, USA Xiaoyuan Chen, Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, and Laboratory for Molecular Imaging and Nanomedicine, National Institute of Biomedical Imaging and Bioengineering, National Institutes of Health, Bethesda, Maryland, USA Claire M. Cobley, Department of Biomedical Engineering, Washington University, St. Louis, Missouri, USA

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Ian R. Corbin, Department of Medical Biophysics, University of Toronto, Toronto, Ontario, and Division of Biophysics and Bioimaging, Ontario Cancer Institute, Toronto, Ontario, Canada Tillmann Cyrus, Department of Medicine, Washington University Medical School, St. Louis, Missouri, USA Nirupama Deshpande, Department of Radiology and Molecular Imaging Program at Stanford, Stanford University School of Medicine, Stanford, California, USA Parmeswaran Diagaradjane, Department of Radiation Oncology, University of Texas M.D. Anderson Cancer Center, Houston, Texas, USA Trevor Douglas, Department of Chemistry and Biochemistry and Department of Plant Science, Center for Bio-inspired Nanomaterials, Montana State University, Bozeman, Montana, USA Luce Vander Elst, Department of General, Organic and Biomedical Chemistry, NMR and Molecular Imaging Laboratory, University of Mons, Mons, Belgium Stanislav Emelianov, Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA Omid C. Farokhzad, Department of Anesthesiology, Brigham and Women’s Hospital, Harvard Medical School, Boston, Massachusetts, USA F´atima Fern´andez, Applied Molecular Receptors Group, CIBER de Bioingenier´ıa, Biomateriales y Nanotecnolog´ıa, IQAC-CSIC, Barcelona, Spain Jinhao Gao, Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, USA Ting Gao, Tyco Electronics Corporation, Menlo Park, California, USA Rui Hao, Department of Advanced Materials and Nanotechnology, College of Engineering, Peking University, Beijing, China Sven H. Hausner, Department of Biomedical Engineering, University of California– Davis, Davis, California, USA Kimberly Homan, Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA Hao Hong, Departments of Radiology and Medical Physics, School of Medicine and Public Health, University of Wisconsin–Madison, Madison, Wisconsin, USA Yanglong Hou, Department of Advanced Materials and Nanotechnology, College of Engineering, Peking University, Beijing, China Taeghwan Hyeon, National Creative Research Initiative Center for Oxide Nanocrystalline Materials, and School of Chemical and Biological Engineering, Seoul National University, Seoul, South Korea Sangyong Jon, School of Life Sciences, Gwangju Institute of Science and Technology, Gwangju, South Korea

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Ella Fung Jones, Center for Molecular and Functional Imaging, Department of Radiology and Biomedical Imaging, University of California, San Francisco, California, USA Kwangmeyung Kim, Biomedical Research Center, Korea Institute of Science and Technology, Seoul, South Korea Nagesh Kolishetti, Department of Anesthesiology, Brigham and Women’s Hospital, Harvard Medical School, Boston, Massachusetts, USA Alan Koretsky, Laboratory of Functional and Molecular Imaging, National Institute of Neurological Disorders and Stroke, National Institutes of Health, Bethesda, Maryland, USA Sunil Krishnan, Department of Radiation Oncology, University of Texas M.D. Anderson Cancer Center, Houston, Texas, USA Ick Chan Kwon, Biomedical Research Center, Korea Institute of Science and Technology, Seoul, South Korea Gregory M. Lanza, Department of Medicine, Washington University Medical School, St. Louis, Missouri, USA Timothy Larson, Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA Sophie Laurent, Department of General, Organic and Biomedical Chemistry, NMR and Molecular Imaging Laboratory, University of Mons, Mons, Belgium Seulki Lee, Biomedical Research Center, Korea Institute of Science and Technology, Seoul, South Korea Yueh Z. Lee, Department of Physics and Astronomy and Department of Radiology, University of North Carolina at Chapel Hill, Chapel Hill, North Carolina, USA Alexei P. Leonov, Department of Chemistry, Purdue University, West Lafayette, Indiana, USA Chun Li, Department of Experimental Diagnostic Imaging, University of Texas M.D. Anderson Cancer, Houston, Texas, USA Lars Liepold, Department of Chemistry and Biochemistry and Department of Plant Sciences, Center for Bio-Inspired Nanomaterials, Montana State University, Bozeman, Montana, USA Pingyu Liu, Palo Alto Unified School District, Palo Alto, California, USA Zhuang Liu, Institute of Functional Nano & Soft Materials, Soochow University, Suzhou, Jiangsu, China Jianping Lu, Department of Physics and Astronomy, Curriculum in Applied Sciences and Engineering, University of North Carolina at Chapel Hill, Chapel Hill, North Carolina, USA Wei Lu, Department of Experimental Diagnostic Imaging, University of Texas M. D. Anderson Cancer Center, Houston, Texas, USA

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Srivalleesha Mallidi, Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA M.-Pilar Marco, Applied Molecular Receptors Group, CIBER de Bioingenier´ıa, Biomateriales y Nanotecnolog´ıa, IQAC-CSIC, Barcelona, Spain Jason R. McCarthy, Center for Molecular Imaging Research, Harvard Medical School and Massachusetts General Hospital, Charlestown, Massachusetts, USA Mohammad Mehrmohammadi, Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA Marites P. Melancon, Department of Experimental Diagnostic Imaging, University of Texas M. D. Anderson Cancer Center, Houston, Texas, USA Mark Michalski, Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, USA R. D. K. Misra, Center for Structural and Functional Materials, University of Louisiana at Lafayette, Lafayette, Louisiana, USA Pranshu Mohindra, Department of Radiation Oncology, University of Texas M.D. Anderson Cancer Center, Houston, Texas, USA Robert N. Muller, Department of General, Organic and Biomedical Chemistry, NMR and Molecular Imaging Laboratory, University of Mons, Mons, Belgium Hyon Bin Na, National Creative Research Initiative Center for Oxide Nanocrystalline Materials, and School of Chemical and Biological Engineering, Seoul National University, Seoul, South Korea Arutselvan Natarajan, Department of Radiology and Molecular Imaging Program at Stanford, Stanford University School of Medicine, Stanford, California, USA Kenneth Ng, Institute of Biomaterials and Biomedical Engineering, University of Toronto, Toronto, Ontario, Canada P. Perriat, Groupe d’Etudes de M´etallurgie Physique et de Physique des Mat´eriaox, Universit´e Claude Bernard, Lyon, France David Pham, Center for Molecular and Functional Imaging, Department of Radiology and Biomedical Imaging, University of California, San Francisco, California, USA Eric M. Pridgen, Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, Massachusetts, USA Min Qu, Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA C. Rivi`ere, Laboratoire de Physique de la Mati`ere Condens´ee et Nanostructures, Universit´e de Lyon, Lyon, France Raffaella Rossin, Department of Biomolecular Engineering, Philips Research Europe, Eindhoven, The Netherlands S. Roux, Laboratoire de Physico-Chimie des Mat´eriaux Luminescents, Universit´e de Lyon, Lyon, France

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J.-Pablo Salvador, Applied Molecular Receptors Group, CIBER de Bioingenier´ıa, Biomateriales y Nanotecnolog´ıa, IQAC-CSIC, Barcelona, Spain Nuria Sanvicens, Applied Molecular Receptors Group, CIBER de Bioingenier´ıa, Biomateriales y Nanotecnolog´ıa, IQAC-CSIC, Barcelona, Spain Anne H. Schmieder, Department of Medicine, Washington University Medical School, St. Louis, Missouri, USA Konstantin Sokolov, Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, and Department of Medical Physics, University of Texas M.D. Anderson Cancer Center, Houston, Texas, USA Michael Strano, Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, Massachusetts, USA O. Tillement, Laboratoire de Physico-Chimie des Mat´eriaux Luminescents, Universit´e de Lyon, Lyon, France Masaki Uchida, Department of Chemistry and Biochemistry and Department of Plant Science, Center for Bio-Inspired Nanomaterials, Montana State University, Bozeman, Montana, USA Bo Wang, Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA Alexander Wei, Department of Chemistry, Purdue University, West Lafayette, Indiana, USA Qingshan Wei, Department of Chemistry, Purdue University, West Lafayette, Indiana, USA Ralph Weissleder, Center for Molecular Imaging Research, Harvard Medical School and Massachusetts General Hospital, Charlestown, Massachusetts, USA Samuel A. Wickline, Department of Medicine, Washington University Medical School, St. Louis, Missouri, USA Jurgen K. Willmann, Department of Radiology and Molecular Imaging Program at Stanford, Stanford University School of Medicine, Stanford, California, USA Patrick M. Winter, Department of Medicine, Washington University Medical School, St. Louis, Missouri, USA Younan Xia, Department of Biomedical Engineering, Washington University, St. Louis, Missouri, USA Jin Xie, Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, and Laboratory for Molecular Imaging and Nanomedicine, National Institute of Biomedical Imaging and Bioengineering, National Institutes of Health, Bethesda, Maryland, USA Lei Xing, Department of Radiation Oncology, Stanford University School of Medicine, Stanford, California, USA

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CONTRIBUTORS

Yun Xing, Department of Material Science and Engineering, University of Dayton, Dayton, Ohio, USA Bing Xu, Department of Chemistry, Brandeis University, Waltham, Massachusetts, USA Mark Young, Department of Chemistry and Biochemistry and Department of Plant Science, Center for Bio-Inspired Nanomaterials, Montana State University, Bozeman, Montana, USA Mi Kyung Yu, School of Life Sciences, Gwangju Institute of Science and Technology, Gwangju, South Korea Gary Zabow, Laboratory of Functional and Molecular Imaging, National Institute of Neurological Disorders and Stroke, National Institutes of Health, Bethesda, Maryland, USA Jingqing Zhang, Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, Massachusetts, USA Ling Zhang, Center for Molecular and Functional Imaging, Department of Radiology and Biomedical Imaging, University of California, San Francisco, California, USA Gang Zheng, Department of Medical Biophysics and Institute of Biomaterials and Biomedical Engineering, University of Toronto, Toronto, Ontario, and Division of Biophysics and Bioimaging, Ontario Cancer Institute, Toronto, Ontario, Canada Hu Zhou, Community Cancer Center of Roseburg, Roseburg, Oregon, USA Otto Zhou, Department of Physics and Astronomy, Curriculum in Applied Sciences and Engineering, and Lineberger Comprehensive Cancer Center, University of North Carolina at Chapel Hill, Chapel Hill, North Carolina, USA

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PART I

BASICS OF MOLECULAR IMAGING AND NANOBIOTECHNOLOGY

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CHAPTER 1

Basic Principles of Molecular Imaging SVEN H. HAUSNER Department of Biomedical Engineering, University of California–Davis, Davis, California, USA

1.1 INTRODUCTION The ability to identify diseased tissue for detection and treatment remains a central goal for medical research. Several noninvasive or minimally invasive diagnostic modalities have been developed which allow one to obtain anatomical, physiological, and molecular information. “Molecular imaging” can be defined as in situ visualization, characterization, and measurement of biological processes in the living organism at the molecular or cellular level. Diagnosis and visualization at the molecular level, that is, detection of a disease in its infancy, may significantly improve treatment and patient care. By combining two or more imaging modalities, each with its different strengths, high-quality complementary (e.g., molecular and anatomical) information can be obtained and analyzed in the context of each other. This has led to the rise of dual- and multimodality imaging approaches. Depending on the modality, imaging probes or contrast agents are required or highly desirable; they can range in size from single atoms to cell-sized constructs. Nanoparticles, that is, entities with dimensions in the range of several tens of nanometers, can display desirable pharmacokinetic properties and permit the combination of different clinically relevant moieties (e.g., targeting groups, molecular beacons, and contrast agents for different modalities, surface coatings, enclosed payload) in a single unit. The inclusion of a therapeutic component yields “theranostics.” Taken together, nanotechnology-based molecular probes offer the promise for tailor-made clinical tools required for “personalized medicine.” This chapter provides an introductory overview of molecular imaging, major imaging modalities, and imaging probes, with particular focus on the promises and challenges of nanoparticle-based compounds.

1.2 IMAGING IN MEDICINE Most areas of clinical practice require identification and localization of diseased tissue for detection and treatment. Ideally, reliable, specific, and noninvasive high-contrast Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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BASIC PRINCIPLES OF MOLECULAR IMAGING

whole-body evaluations would allow physicians to detect serious abnormalities before patients present with symptoms, thus permitting early intervention, thereby increasing the chance for cure or, at a minimum, allow for better patient management and improved quality of life. Given these incentives, it is clear that practical (i.e., minimally inconvenient for the patient) and affordable (i.e., overall cost-saving to the health care system and society) diagnostic approaches are highly desirable. Ever since Wilhelm R¨ontgen’s first use in 1895 of the then newly discovered X-rays to noninvasively image the interior of the body, the keen interest in medical imaging has been met by increasingly sophisticated technologies (Fig. 1.1). While R¨ontgen’s X-ray image was a grainy two-dimensional anatomical projection, physicians nowadays have access to tomographic (three-dimensional) imaging modalities with, depending on the technique, submillimeter resolution, which allows visualization of anatomical, physiological, and, increasingly, molecular (cellular) biological information. Since diseases often arise from changes on the molecular and cellular levels, long before manifesting themselves in detectable large-scale physiological or anatomical changes, molecular imaging is gaining increasing attention. If a disease can be diagnosed and visualized at the molecular level, that is, detected in its infancy, it can be treated at a much earlier stage, the treatment’s efficacy can be determined much sooner and, if necessary, the treatment plan can be adjusted accordingly. This benefits the individual patient and society as a

FIGURE 1.1 (Left) Wilhelm R¨ontgen’s (1845–1923) first X-ray image, depicting the hand of his wife, Anna, taken on 22 December 1895. (Right) A slice of a modern whole-body multimodality positron emission tomography/computed tomography (PET/CT) scan showing glucose metabolism within the body, including a large, metabolically active tumor (arrow). (PET/CT image courtesy of Dr. Cameron Foster and Dr. Ramsey Badawi, UC Davis Medical Center, Davis, California.)

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IMAGING IN MEDICINE

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whole. Molecular biology is discovering a growing number of disease-specific cellular targets and is determining their distribution in patient populations [1]. For certain diseases this has already had significant effects on determining beforehand which patients will benefit from a certain treatment (“patient stratification”). A prime example is testing for the expression of HER2/neu in breast cancer for prognosis, as well as for selection and monitoring of treatment: expression has been linked to aggressiveness of the disease, but it also provides a target for highly effective treatment with antibodies (Trastuzumab, Herceptin® ) [2, 3]. Similarly, monitoring glucose metabolism with the imaging agent 18 F-fluorodeoxyglucose (18 F-FDG) has proved itself to be the preferred approach for staging, restaging, and evaluation of response to treatment for several cancers [4]. Concurrently with the advances in molecular biology, engineers and physicists are developing increasingly sophisticated imaging instrumentation capable of localizing imaging agents in the body at high sensitivity and high resolution in short acquisition time [5]. By bridging the clinical and engineering worlds, research in imaging agents plays a central role. To that end, the development of target-specific (and disease-specific) nanoparticle-based molecular probes draws on research in several fields including biology, molecular biology, medicine, chemistry, and biomedical engineering. 1.2.1 Molecular Imaging Rather than relying only on intrinsic large-scale differences of tissue characteristics (e.g., density) or passive accumulation of administered probes to reveal disease in vivo, molecular imaging strives to make use of disease-specific (“targeted”) interactions of imaging probes with the target tissue on a molecular and a cellular level. The goal is the real-time in situ visualization of biological processes in the living organism. This focus is also reflected in the Society of Nuclear Medicine’s definition of molecular imaging as “an array of non-invasive, diagnostic imaging technologies that can create images of both physical and functional aspects of the living body. It can provide information that would otherwise require surgery or other invasive procedures to obtain. Molecular imaging differs from microscopy, which can also produce images at the molecular level, in that microscopy is used on samples of tissue that have been removed from the body, not on tissues still within a living organism. It differs from X-rays and other radiological techniques in that molecular imaging primarily provides information about biological processes (function) while [computed tomography] CT, X-rays, [magnetic resonance imaging] MRI and ultrasound, image physical structure (anatomy)” [6]. As stated above, the information obtained is linked to which imaging modality is chosen. Individual imaging modalities can be grouped by the energy spectrum and energy type evaluated (X-ray, photons, sound; positrons), the resolution that can be achieved, and the type of information obtained (anatomical, physiological, cellular/molecular) (Table 1.1). Widely used clinical imaging modalities include magnetic resonance imaging, ultrasound (US), computed tomography, as well as positron emission tomography (PET) and single photon emission computed tomography (SPECT). All of these modalities allow for the noninvasive imaging of living subjects. Although the first three imaging modalities are primarily anatomical and not molecular, the two types of modalities can be combined for dual- or multimodality imaging. In addition, MRI, US, and CT can be used with molecular imaging probes, especially as part of nanoplatforms. In addition, a number of more specialized optical modalities are being used or are under investigation, including endoscopic methods [12].

6 M, P

M, P

M, P

Positron emission tomography (PET)

Single photon emission computed tomography (SPECT)

Typeb

Optical imaging (fluorescence and bioluminescence)

Imaging Modalitya −15

Sensitivity (Concentration of Imaging Probe/Contrast Agent)

Photon emitted by radioactive isotope of imaging probe.

∼10−11 mole/L [7]

∼10−11 –10−12 mole/L

Depth

0.5–2 mm (preclinical) 10–15 mm (clinical)

1–2 mm (preclinical) 4–8 mm (clinical)

No limit

No limit

Yes

Yes

∼1 to ∼10 mm Centimeters (Yes, limited quantification possible)

Resolution

Quantitative Modality

Minutes to tens of minutes

Minutes to tens of minutes

Seconds to minutes

Typical Scan Acquisition Time

Clinical and preclinical. High cost. Versatile imaging probe chemistry. Possibility to distinguish different radioisotopes based on photon energy.

Clinical and preclinical. High cost. Versatile imaging probe chemistry.

Preclinical; limited clinical translation (close to skin or requiring endoscopic approaches) Low cost. Depth limitation based on wavelength-dependent absorption by tissue. Resolution is depth dependent. Two-dimensional (surface) image.

Other

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511-keV Photons generated during annihilation of positron emitted by radioactive isotope of imaging probe.

Fluorescence: External As low as ∼10 mole/L excitation light absorbed by fluorochrome of imaging probe and reemitted at longer wavelength. Bioluminescence: Chemiluminescence of enzymatic reaction.

Basis for Detection

TABLE 1.1 Widely Used Imaging Modalities

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b

Less than ∼50 ␮m (preclinical) <500 ␮m (clinical) [9]

<100 ␮m (preclinical) <1 mm (clinical) [9]

Absorption of focused Low (gram amounts <10 ␮m external X-rays by of contrast agent (preclinical) tissue (or contrast required) agent).

Primary molecular imaging modalities are listed in bold. A, anatomical; M, molecular/cellular; P, physiological. Source: Adapted from Willmann [11] and Weissleder [12].

a

A, P

A, P (M) Echoes of tissue (or High (single imaging probe) microbubbles— generated by volume ∼0.004 high-frequency pL—can be (∼1–40 MHz) detected) [8] sound waves propagating through tissue.

A, P (M) Interaction of external 10−3 –10−4 mole/L magnetic field and radiofrequencies with atomic nuclear spins (of tissue or contrast agent) depending on environment of nuclei.

No limit

Up to ∼25 cm [10]

No limit

Yes

Yes

Yes

Minutes

Seconds to minutes

Minutes to hour

Clinical and preclinical. Excellent bone and lung contrast. Poor soft-tissue contrast.

Clinical and preclinical. Low cost. Frequency used determines resolution and (inversely) penetration depth. Targeted imaging limited to vasculature. High operator dependency.

Clinical and preclinical. High cost. Long scan acquisition time. Excellent soft tissue contrast.

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Computed tomography (CT)

Ultrasound imaging (US)

Magnetic resonance imaging (MRI)

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Regardless of the imaging modality chosen, quantifiable high-resolution images and reasonable acquisition times are highly desired, and if modalities are combined they should yield relevant additional (e.g., anatomical plus molecular) information. Molecular imaging per se is complementary to primarily anatomical imaging (Table 1.1). This is the motivation behind the ongoing push toward dual-/multimodality imaging where molecular imaging data are collected at the same time as anatomical imaging data (Fig. 1.1). This synergistic approach, in which superimposed tomographic images are analyzed, allows the physician interpretation of the molecular imaging data within the anatomical context. The tremendous benefits of this diagnostic approach have also been recognized by the manufacturers of clinical imaging equipment. This has led to the rapid spread of integrated hybrid PET/CT and SPECT/CT scanners in recent years. Dual-modality scanners are now becoming the norm rather than the exception in the clinic [5, 13]. Similarly, hybrid PET/MR scanners are now becoming available; they are eagerly awaited for tasks where the molecular imaging data have to be interpreted in the context of soft tissue, such as, for example, within the brain. Engineering and technical challenges are largely the reason that the availability of hybrid-MR systems has been lagging behind their CT counterparts [14, 15]. The instrumentation for the various modalities has also been adapted for preclinical applications [16]. By using mice, rats, nonhuman primates, or other animal models, specialized small animal scanners allow dedicated imaging in a preclinical research setting. Spatial resolution is generally higher because the subjects can be moved closer to the detectors and the instruments are specifically designed for the reduced dimension required. Because of their small body size, whole-body imaging is easily possible for several species with many of the imaging modalities.

1.3 MAJOR IMAGING MODALITIES 1.3.1 Optical Imaging (Fluorescence and Bioluminescence) Optical imaging is finding increasing clinical use in several specialized applications, largely using endoscopic (or similar fiberoptic intravital) methods or in regions with limited tissue thickness (e.g., the breast) [12, 17]. Still, the major application of optical imaging lies in preclinical use for small animal studies, chiefly thanks to relatively low cost and simple setup: the subject is placed in a light-tight box and imaged with a highly sensitive charge-coupled device (CCD) camera. A considerable number of optical probes and tags are commercially available, making optical imaging the most popular preclinical imaging modality [5]. For fluorescence imaging the subject is typically illuminated by an external source with excitation light that is absorbed by the fluorophore of an imaging probe. The fluorophore then emits light at lower energies (longer wavelengths) that is detected by the camera. Ideally, the light involved should be in the near-infrared range (∼650–900 nm), where absorbance by blood is minimal. For bioluminescence imaging, no external excitation is required; rather, the faint light emitted by certain biological processes is measured directly. In laboratory studies, this can be achieved by linking a “reporter gene” encoding for a luminescent protein (usually luciferase) to the gene of interest and genetically transferring them into the animal before the study. After administration of an exogenous substrate (e.g., luciferin) light is generated only at sites where the genes are expressed. A similar approach can also be used for modified fluorescence imaging. In this case, the gene for green fluorescent protein (GFP) or one of its derivatives is commonly used as the reporter. Under

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illumination, locally expressed GFP emits light that is detected by the CCD camera. An advantage of this approach is the possibility of longitudinal studies because injection of a substrate is not necessary for visualization, whereas the useful window for bioluminescence after a luciferin injection is usually only about 5–30 minutes. Owing to the fact that a carefully designed optical probe can be switched on and off in vivo as a result of chemical or physicochemical transformations, “activatable” or “smart” fluorescent probes have been developed that can respond to the presence and level of biological markers at sites within the body [18]. This has been used in preclinical tumor models to monitor treatment response using a near-infrared fluorophore (NIRF)-based imaging probe responsive to the level of matrix metalloproteinase (MMP)-2. Treatment reduced the level of MMP-2 expressed by the tumor, which was reflected in a reduced signal emitted by the imaging probe. Several challenges exist for optical imaging. For fluorescence imaging, they include high background signals caused by tissue autofluorescence [19] and limited stability (photobleaching) of many small-molecule fluorophores. Bioluminescence does not have the same problems, but researchers face the tasks of genetically engineering the animal model and detecting very faint signals. Both approaches are constrained by depth limitations due to scattering and absorbance by overlying tissue and the concomitant difficulties with exact quantification. If spatial resolution is not a major concern, whole-body optical imaging is possible for small rodents (especially mice) since scattering and absorption are limited because of the small body size [20]. Perhaps more than for other imaging modalities, a notable number of new approaches based on different technologies are being investigated for optical imaging [12]. Fluorescence lifetime imaging (FLIM), photoaccoustic imaging, multispectral imaging [21], self-illuminating fluorescent imaging probes [19], Raman microscopy techniques, and tomographic fluorescence systems are among the exciting approaches currently under development [12]. Some of them rely entirely on endogenous contrast and do not require the administration of any exogenous probes. Two examples are coherent anti-Stokes Raman scattering (CARS) and optical coherence tomography (OCT). CARS is a nonlinear Raman technique that measures the vibrational spectra of light scattered from illuminated biological specimens. Analysis of the spectra allows conclusions about the constituents of the tissue close to the surface. It has been used in vivo to map lipid compartments, protein clusters, and water distribution at subcellular resolution [22]. OCT is a technique based on light scattering that can be described as an optical version of ultrasound (see below). Despite a shallow penetration depth of only about 2–3 mm, it is attractive since it yields real-time very-high resolution (1–15 ␮m) “optical biopsy” images that are comparable to conventional histopathology. It is finding applications in ophthalmic, gastrointestinal, and intravascular imaging using noninvasive or minimally invasive instrumentation such as handheld probes, endoscopes, catheters, laparoscopes, or needles [23]. 1.3.2 Radionuclide-Based Imaging Modalities: Positron Emission Tomography (PET) and Single Photon Emission Computed Tomography (SPECT) Because of high sensitivity and absence of depth limitations, PET and SPECT are the two molecular imaging modalities that have risen to prominence in both the clinical and preclinical settings. They require the administration of a positron- or single-photonemitting radioisotope, usually attached to a larger molecule. Examples are [18 F]fluorine in

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2-[18 F]fluoro-2-deoxy-glucose (18 F-FDG), [123 I]iodine in [123 I]metaiodobenzylguanidine (123 I-MIBG), or radioactive metal isotopes captured by a chelator (e.g., [64 Cu]copper or [111 In]indium in chelator-bearing proteins and antibodies). As such, both imaging modalities rely completely on exogenous probes for imaging. For both imaging modalities, the availability, the chemistry, and the radioactive half-life of the chosen isotope have to be considered. This is illustrated by comparing the two popular PET isotopes [11 C]carbon and [18 F]fluorine. [11 C]Carbon is an attractive isotope because it can directly replace a nonradioactive carbon without changing the molecular structure of a compound. However, it has a half-life (t1/2 ) of only 20.4 min, necessitating production in an on-site cyclotron and limiting preparation of the imaging probe to a handful of very fast chemical reactions. By contrast, the nearly 2-h half-life of [18 F]fluorine allows a much wider range of chemistries and even some degree of shipment of the imaging probe from central production facilities to outlying hospitals by ground or air. Since the fluorine atom usually takes the place of another element (often a hydrogen atom), possible effects on pharmacokinetics have to be evaluated during drug development. Regardless of which radionuclide-based imaging modality is employed, it is important to use a radioisotope whose physical (radioactive) half-life is matched to the pharmacokinetics of the imaging probe to ensure a sufficiently high signal-to-noise ratio at the time of imaging [24]. Since most nanoparticles have long blood circulation times, it may take up to a few days before the level in the target tissue has risen significantly over background levels. In order to match the long biological half-life of the probe, long-lived radioisotopes are often required. Fortunately, many such radioisotopes are available. For example, the SPECT isotopes [123 I]iodine, [99m Tc]technetium, and [111 In]indium have half-lives of 13.2 h, 6.0 h, and 67.3 h, respectively, and long-lived PET isotopes include [64 Cu]copper, [124 I]iodine, and [89 Zr]zirconium (t1/2 = 12.7 h, 100.2 h, 78.4 h, respectively). PET, in particular, distinguishes itself through its high sensitivity combined with the ability to image effectively without depth limitation [25]. As mentioned earlier, especially 18 F-FDG has helped clinical PET to play a prominent role in cancer detection and monitoring of response to treatment because it allows the visualization of glucose hypermetabolism associated with many malignancies and whole-body PET scans permit the detection of distant metastases (Fig. 1.1). Delicate biological systems (e.g., the brain) can be imaged with minimal disturbance of the molecular processes investigated thanks to the extremely low amount of imaging probe required. The signal detected by the scanner originates from the radioactive decay of a positron-emitting radioisotope prepared in a cyclotron prior to incorporation into the imaging probe. The positron loses energy by scattering through the tissue until undergoing annihilation with an electron, resulting in the emission of two 511-keV photons at an angle of nearly 180◦ . The pair of photons is detected by a cylindrical array of scintillators connected to photomultiplier tubes (PMTs). Image quality is greatly improved by only accepting valid coincidences and rejecting random events stemming from background radiation: only signals obtained in opposite detectors within a narrow time window, commonly 2–5 ns, are accepted as originating from the same positron-decay event. Resolution-limiting factors are the average range positrons travel before undergoing annihilation (“positron range”), the noncollinearity of the two photons emitted, and detector geometry [26, 27]. The positron range is isotope specific as it depends on the energy with which the positrons are emitted. It can range from <1 mm to >5 mm for common PET isotopes; for [18 F]fluorine it is approximately 0.7 mm [26, 28]. Clinical PET scanners have a typical resolution on the order of several millimeters. Submillimeter resolution is possible

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with dedicated small animal scanners, allowing detailed studies of mouse, rat, and monkey models, including brain imaging [25, 29]. Contrary to PET, SPECT directly measures single gamma-ray photons emitted by the decay of the chosen radioisotope. The scanner consists of scintillator/PMT detector heads that slowly rotate around the object in a stepwise fashion, collecting data every few degrees for a certain time. Since only single photons are measured, collimators are used to obtain directional information. Only photons traveling along the axis of the collimator are captured by the detector. This allows SPECT to reach resolutions comparable to PET (Table 1.1), but only at the cost of reduced sensitivity. Depending on the collimator used, sensitivity is reduced by at least about one order of magnitude [7, 11]. On the other hand, since the isotopes commonly used for SPECT have relatively long half-lives on the order of several hours to days, longitudinal studies are greatly facilitated and shipping from distant production facilities is possible. Because of the chemistry of the isotopes involved, straightforward preparation (radiometal chelation) protocols exist for many SPECT imaging probes. In addition, different isotopes can be distinguished based on the isotope-specific energies of the emitted photons (e.g., [123 I]iodine: 159 keV; [99m Tc]technetium: 140 keV; or [111 In]indium: 171, and 245 keV). 1.3.3 Magnetic Resonance Imaging (MR Imaging or MRI) MRI is based on the principles of nuclear magnetic resonance (NMR). The nuclei of certain isotopes, most notably naturally occurring hydrogen (i.e., protons), possess nuclear spins that can be aligned along strong external magnetic fields such as those produced inside an MRI scanner. Alignment can be either parallel or antiparallel to the field. Similar to spinning top toys, the spins are not totally aligned with the field but precess around the field vector, resulting in what can be viewed as cone-shaped distributions. Following the application of transverse radiofrequency (rf ) pulses to perturb both the longitudinal and transverse alignments, the nuclear spins return to their previous (ground) state, generating a faint signal detected by a receiver coil. Longitudinal (T1 recovery) and transverse (T2 decay) relaxations are measured by the scanner and form the basis for the tomographic image. Relaxation times are dependent on the local environment of the nuclei and on their density within the tissue. Diseased tissue may therefore be detected by its effect on the local density and environment. Furthermore, the rf pulses and their sequence can be modified in order to weight images toward primarily detecting differences in T1 or T2. Even though protons are virtually ubiquitous throughout the body as part of water and fat molecules, physical limitations, most notably the fact that only a miniscule fraction of spins are aligned properly along the field to generate a signal, result in the relatively low sensitivity of MRI, coupled with long acquisition times (Table 1.1) [11]. Still, MRI is widely used as it permits imaging of soft tissue at high resolution without depth limitation. While images can be acquired by solely relying on native protons (endogenous contrast), the image quality can be greatly improved by administration of certain contrast-enhancing agents. After accumulation in the tissue, these agents change the protons’ local environment, thereby shortening their relaxation times (T1 and T2, measured as relaxivities R1 = 1/T1, and R2 = 1/T2). Gadolinium-chelate-based paramagnetic agents are frequently used to increase T1 contrast, resulting in local brightening in the tomographic image. Newly developed T1 contrast agents include long-circulating manganese oxide (MnO) nanoparticles [30] and gadolinium-containing micelles and liposomes amendable to functionalization [31].

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To obtain negative contrast enhancement (darkening) in T2-weighted images, superparamagnetic particles such as iron oxide nanopraticles are typically used. Biocompatibility and ease of synthesis have made colloidal superparamagnetic iron oxide (SPIO) and ultrasmall superparamagnetic iron oxide (USPIO) particles particularly popular. In addition, the continued pursuit of improved magnetic properties has yielded metal-doped (MnFe2 O4 , FeFe2 O4 , CoFe2 O4 , NiFe2 O4 ), metallic (coated iron, cobalt, nickel), and bimetallic (FePt, FeCo) nanoparticles [32]. Composition, size, surface characteristics, and potential aggregation of the particles, as well as external experimental parameters, all can have an effect on relaxivity and thus the MRI image [32]. Because of the abundance of hydrogen atom within the body, clinical MRI still to a large extent relies on the detection of protons, using contrast agents or high-field MRI scanners to improve contrast [11]. Another way to increase signal strength is to employ hyperpolarized agents. Optical pumping can greatly enrich the excited spin state of certain isotopes (e.g., [3 He]helium, [129 Xe]xenon), resulting in signal enhancements of up to 106 [11]. However, because the hyperpolarized state lasts only a few seconds, the agents need to be prepared immediately prior to imaging and, even then, the time window available for imaging is extremely limited. In addition to the isotopes mentioned so far, certain isotopes of other elements can be detected as well, provided their natural abundance is high enough (e.g., [23 Na]sodium, [31 P]phosphorus) or they can be administered in sufficient quantities (e.g., [17 O]oxygen as water, [13 C]carbon as glucose, [19 F]fluorine as fluorocarbons). Because MRI is based on NMR detection, it is possible to tune special scanners to specifically detect spectra of different nuclei. This forms the basis for magnetic resonance spectroscopy (MRS), which is beginning to open the door for quantitative or semiquantitative physiological and molecular MRI by tracing the distribution of individual compounds (e.g., cancer-related metabolites) [33]. To map brain activity, functional MRI (fMRI) detects blood-oxygenlevel-dependent changes in the environment of the iron contained in hemoglobin, a measure for the hemodynamic response linked to neural activity. Other novel functional imaging techniques include diffusion-weighted imaging (DWI) and diffusion tensor imaging (DTI) as well as perfusion MRI [33]. DWI and DTI assess the movement of free water molecules in tissue. Perfusion MRI measures blood-transport dynamics by following signal changes with special imaging protocols (dynamic susceptibility contrast MRI, DSC-MRI; dynamic contrast enhanced MRI, DCE-MRI). 1.3.4 Ultrasound (US) Imaging Ultrasound imaging records the reflections of high-frequency sound waves by internal body structures. Focused sound waves, originating from a transducer placed on the skin, propagate through the body and generate echoes as they encounter tissues of different density. The echoes, in turn, are captured again by the transducer and the data are processed to generate digital images. By performing Doppler sonography, velocity can be measured; this approach is commonly employed for blood-flow examinations. Thus physiology as well as anatomy can be imaged with US [8]. Resolution can be increased by choosing higher frequencies, which, however, lead to a concomitant decrease in penetration depth. Since US imaging does not involve the use of ionizing radiation, it is considered a particularly “safe” imaging modality. Therefore, and because of relatively low cost, ease of use, and wide availability, US is the most popular clinical imaging modality [9]. The image quality depends on sound frequency, sound speed, sound attenuation, sound backscatter, data processing, and

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transducer handling [11]. Because of the latter, US imaging is more dependent on the experience and skill of the operator than the other imaging modalities. Although not necessary for US imaging, contrast agents can be used. Because of their ability to significantly increase the signal intensity, gas-containing microbubbles, along with echogenic liposomes and perofluoro-emulsion nanoparticles, ranging in size from about 100 nm to 8 ␮m, are most commonly used [8]. They allow for greatly improved imaging of small blood vessels that are otherwise difficult to distinguish from surrounding tissues. Their use, however, is in fact largely limited to imaging the vascular compartment because their particle size precludes escape from blood vessels (“extravasation”). Despite this considerable limitation, clinically approved US contrast agents have been developed and targeted US imaging has been demonstrated with surface-modified microbubbles for diseases that result in molecular changes to the vascular compartment. For example, RGDpeptide modified microbubbles have been used to target the integrin ␣v ␤3 , a cell-surface receptor linked to tumor angiogenesis [8]. US contrast agents such as microbubbles can also be used for drug delivery. It has been shown that drug-carrying bubbles can be fragmented by high-intensity ultrasound, leading to the local release of a therapeutic agent [34]. 1.3.5 Computed Tomography (CT) For CT a series of X-ray images is acquired and reconstructed into a tomographic image. Focused X-rays emerge from a source rotating stepwise around the subject and, after traveling through the body, are captured by a detector situated opposite the source. Contrast is based on differences in attenuation of the X-rays as they pass through the body. Dense tissue (bones) absorbs significantly more energy than soft tissue, while air (lung) absorbs less. As a result, CT is primarily an anatomical imaging modality that is especially useful for bone and lung imaging but does not show particularly good soft tissue contrast. For contrast-enhanced CT X-ray absorbing agents are administered. Typically, large amounts of iodine-based compounds have to be administered intravenously to achieve an appreciable soft tissue contrast. The compounds have a largely nonspecific distribution and fast clearance kinetics, although work on improving these characteristics is ongoing [9]. For contrast-enhanced CT of the gastrointestinal tract, insoluble contrast agents based on barium sulfate are given orally or administered as an enema. Recently, lanthanide-bearing compounds, bismuth sulfide nanoparticles, and targeted gold nanorods have been studied as potential future CT contrast agents with increased image contrast and longer circulation times than the standard iodine agents [35–37]. Although CT is just at the verge of also being explored as a molecular imaging modality, it is already extremely useful in providing a high-resolution anatomical framework in which to interpret molecular imaging (e.g., PET or SPECT) data (Fig. 1.1) [13]. The use of PET/CT and SPECT/CT hybrid scanners has improved diagnostic accuracy in the clinic and facilitated preclinical drug evaluation [5].

1.4 MOLECULAR IMAGING PROBES Even when ex vivo diagnostic tests (e.g., blood, urine, Pap smear) exist for a disease [1], the ability to obtain data on location and local amount of a disease-specific indicator, and thereby the disease itself, within the body is what makes molecular imaging such a powerful tool. Nowhere is this perhaps more evident than in oncology where tumor heterogeneity, possible involvement of lymph nodes, and the potential of metastases create an extremely

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challenging diagnostic target. Detection based on disease-related changes in the local environment provides one possible avenue. This includes imaging of metabolic rate (e.g., glucose consumption), cell proliferation (i.e., DNA synthesis), or a change in the microenvironment (e.g., local pH, concentration of oxygen or proteases) typically associated with cancers. Indeed, a number of small-molecule compounds have been developed into highly successful probes for these conditions, as exemplified by the PET-probes 18 F-FDG (for glucose consumption), 18 F-FLT (for cell proliferation), and 18 F-FMISO or 64 Cu-ATSM (for hypoxia) [38]. Even though these probes for local conditions provide very useful information, they still tend to be relatively nonspecific and may miss slow-growing neoplasms [38] or tumors in certain locations such as excretory organs that have a high nonspecific background signal. Since upregulation of intracellular markers and cell-surface receptors is a hallmark of disease states, aiming for these disease-specific targets is highly attractive, provided that suitable ligands are available. For imaging, cell-surface receptors may be favored over intracellular targets because they do not require internalization into the cell and therefore can be accessed more readily, particularly by nanoparticle-based imaging probes. If the nanoparticle is designed for concomitant drug or gene delivery, subsequent internalization into the cells becomes important. Giving examples from oncology, Table 1.2 illustrates the range of receptor targets and targeting ligands presently available. The examples are taken from MRI imaging [39], but are equally representative for other molecular imaging modalities [12, 18, 38, 40]. Most imaging probes are administered intravenously and allowed to distribute through the body; Notable exceptions include experimental implantable microprobes containing nanoparticles [41], intrinsically produced optical imaging probes (i.e., fluorescent or bioluminescent proteins [42]) expressed by reporter genes, and locally injected imaging probes for sentinel lymph node mapping during cancer surgery [43]. Typically, as the individual probes circulate through the body they can accumulate in tissues, be excreted from the body, and/or be metabolized. To be clinically useful, molecular imaging probes should be based on high-affinity high-specificity ligands and show a relatively low background signal [44]. The ratio of imaging probe in the target tissue versus in the surrounding tissues (i.e., the signal-to-noise ratio) determines how much image contrast can be achieved. In other words, the image depends on how well the imaging probe has been taken up in the target tissue and how well it has cleared from other regions of the body at the time the image is acquired. Therefore some uptake period is frequently required between injection and imaging. Depending on the desired information and the pharmacokinetics of the imaging probe, it can range from a very short time (e.g., seconds to minutes for [15 O]oxygen-water in blood flow studies [45] to over tens of minutes (e.g., for 18 F-FDG) to days (e.g., for antibodies). Size has a significant impact on the time imaging probes circulate in the body. Generally, small molecules are eliminated fast, resulting in a rapidly decreasing background signal, while larger compounds, including nanoparticles [46], have longer circulation times. Thus, as outlined earlier, for radionuclide-based imaging the radioactive half-life of the radioisotope has to be matched to the pharmacokinetic profile of the imaging probe to ensure that a sufficient radioactive signal remains by the time an adequate amount of the imaging probe has accumulated in the target tissue in order to obtain statistically significant contrast. For nanoparticles this usually requires radioisotopes with half-lives on the order of several hours to days. Algorithms help to choose radioisotopes with an appropriate half-life: by calculating a suitable “imaging figure of merit” (IFOM) that takes into account parameters such as the size of the target organ (or tumor), the measured decay corrected organ uptake (% injected

15

Source: Peng [39] and references therein.

Magnetism-engineered iron oxide (MEIO) nanoparticles PEG-IO

SPIO USPIO PEG-SPIO Streptavidin-SPIO

Chlorotoxin

RGD4C Luteinizing hormone releasing hormone (LHRH) CREKA peptide Arg-Gly-Asp (RGD) Folic acid Antibody to prostate-specific membrane antigen (PSMA) Herceptin

Folic acid F3 peptide

Monoclonal antibody A7 Herceptin Methotrexate Chlorotoxin peptide

Folic acid

Monoclonal antibody-610 Antibody to carcinoembryonic antigen (CEA) Monoclonal antibody L6 Transferrin Monoclonal antibody-Her-2/neu EPPT peptide

Targeting Ligands

Membrane-bound matrix metalloproteinase-2 (MMP-2)

Her-2/neu receptors

Clotted plasma proteins ␣v ␤3 integrins Folate receptor PSMA

Rat glioma

NIH3T6.7

In vivo

In vivo

In vivo In vivo In vivo In vitro

In vitro In vivo

Melanoma cells Breast cancer Breast cancer Human epidermoid carcinoma Human epithelial mouth carcinoma Human prostate cancer cells

In vitro In vivo

In vivo In vivo In vitro In vitro

In vitro

In vivo

In vivo In vivo In vitro

In vitro In vivo

Experimental Conditions

Human cervical cancer cells Rat glioma

Colorectal carcinoma NIH3T6.7 Human cervical cancer cells Rat glioma

Breast, colon, pancreas, and lung cancer cell lines Human epithelial mouth carcinoma

Underglycosylated mucin-1 antigen (uMUC-1) Folate receptor

Colorectal tumor antigen Her-2/neu receptors Folate receptor Membrane-bound matrix metalloproteinase-2 (MMP-2) Folate receptor Surface-localized tumor vasculature ␣v ␤3 integrins LHRH receptor

Intracranial tumor LX-1 Rat mammary carcinoma Breast cancer

Colon carcinoma cell lines Colon tumor

Tumor

Surface antigen Transferrin receptor Her-2/neu receptors

Surface antigen CEA

Targets

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Biofunctional PEG-SPIO SPIO encapsulated with photodynamic agent HFn-IO SPIO

Dextran-coated superparamagnetic maghemite (␥ -Fe2 O3 ) nanocrystals Ferumoxides (SPIO) Iron oxide nanocrystals (Fe3 O4 ) SPIO SPIO

CLIO-NH2

MINO USPIO Streptavidin-conjugated SPIO

USPIO SPIO

Iron Oxide Nanoparticles

TABLE 1.2 Targeted Iron Oxide Nanoparticles for Tumor Imaging with MRI

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dose/gram, % ID/g) or organ activity (%ID/organ), the area under the curve (AUC) and their respective changes over time, predictions for an optimal imaging time, and thus radioactive half-life are possible [24, 47]. Similar calculations can be performed to obtain a “therapeutic figure of merit” (TFOM). Allometric analysis is used to obtain predictions for clinical trials of a novel probe based on preclinical data from animal models. Most imaging modalities can yield quantitative data after careful calibration (Table 1.1). The previous paragraph illustrates the importance of quantification during early probe development and for clinical translation. Once an imaging probe has reached this step, quantitative image-based data analysis becomes even more relevant. Newly developed imaging probes need to be evaluated in patients, and sensitivity (ideally 100% = no falsenegative examinations) and specificity (ideally 100% = no false-positive examinations) need to be established and compared to standard test methods (i.e., a “gold standard” such as biopsy/histopathology). Since molecular imaging is founded on the ability to quantify specific disease-related molecular profiles and target densities, the clinical goal is to go beyond mere detection of disease and to permit staging, restaging, and evaluation of response to treatment. To that end, image-derived quantitative parameters for evaluating the progression of a disease have been developed. For example, in oncology the “standard uptake value” (SUV) of a PET probe within a chosen “region of interest” (ROI)—the tumor—can be used as prognostic indicator [4, 48] and response to treatment can be assessed based on the change in SUV, often soon after the treatment since molecular changes precede changes detectable by anatomical or physiological imaging [49]. Regardless of the imaging modality, important parameters of an imaging protocol include (1) the imaging probe and its formulation, (2) the type and model of the scanner/detector, (3) patient selection and preparation, (4) data acquisition, (5) data processing, and (6) data analysis. All parameters need to be carefully controlled, recorded, and analyzed to yield results that are comparable over time and between different hospitals [50]. This is a prerequisite for meaningful results from clinical trials evaluating new imaging probes or new therapy regimens, and for subsequent clinical use based on reliable image acquisition and data handling protocols approved by regulatory agencies [49]. Ultimately, to be successful in clinical use, a new imaging probe needs to pass several tests: Is it better than competing diagnostic approaches? Does it have a cost-effective role in clinical management? Is the interpretation of the imaging data easy [4]? Once established, a well-studied imaging probe such as 18 F-FDG can then also play an important role in future drug development [51]. 1.4.1 Nanoparticle-Based Probes Molecular imaging can be performed with imaging probes of vastly different sizes. Entities can be the size of single atoms (e.g., a 18 F fluoride ion), small molecules (e.g., the modified glucose 18 F-FDG, the radiometal complex 99m Tc-sestamibi), peptides, antibodies, or constructs approaching the dimension of small cells (e.g., micrometer-sized microbubbles [34]). Typical “nanoparticles” fall somewhere in the middle of this range. As outlined below, to be clinically useful the preferred size of nanoparticles is roughly between 10 and 100 nm [52]. Thus this classification covers compounds being similar in size to hemoglobin (6.5 nm), an antibody (12 nm), a hepatitis virus (45 nm), or an influenza virus (130 nm), although the constructs approaching the micrometer range are also commonly included in the definition. The nanoparticles themselves can extend in one to three dimensions, as depicted schematically in Figure 1.2, and show unique size- and shape-dependent properties that can be beneficial for imaging.

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Quantum Dot

Nanoshell

Gold

Liposome

Iron Oxide

FeCo

Perfluorocarbon

Nanotube

Dendrimer

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FIGURE 1.2 Schematic depictions of representative nanoparticles for targeted molecular imaging in living subjects. (Reproduced with permission from Cai [9].)

Nanoparticle-based probes are receiving considerable attention because the particles form a convenient platform for nanodiagnostics and nanomedicines that can combine different clinically relevant properties in a single unit. Each particle can carry highly specific targeting ligands, antibodies (for targeting or treatment), different imaging probes, drugs, or various combinations there of (Fig. 1.3). Depending on the nature of the particle, it itself can assume some of these roles. Examples include fluorescent quantum dots detectable by optical imaging, iron-based MRI contrast agents, carbon nanotubes used for targeted thermal ablation [53], and drug-carrying bubbles for ultrasound imaging. When an imaging

FIGURE 1.3 Schematic depiction of a targeted multifunctional nanoparticle carrying imaging probes, drugs, and antibodies. (Reproduced with permission from Rao [52].)

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probe, that is, a diagnostic entity, is combined with a drug, the resulting compound is also called a “theranostic” or “theragnostic” [40, 54, 55]. Ideally, this integrated theranostic approach will permit diagnosis, targeted delivery of therapy, and monitoring of response to treatment Sumer 2008 [56]. By choosing the right combination of targeting moieties and therapeutic components, the theranostic can home in on multiple markers and target cell subpopulations, thereby providing a tool to address disease heterogeneity and adaptive resistance. The treatment can, if necessary, quickly be adjusted based on the rapid feedback obtained. Although still in its early stages, the hope is that this approach will allow true “personalized medicine” based on individualized, molecular diagnosis within a heterogeneous disease and patient population, something that cannot be accomplished by taking small biopsy samples or following standardized “one-dose-fits-all” protocols. As already alluded to in the previous paragraphs, nanoparticle-based probes offer several advantages over standard probes Sanvicens 2008 [57, 58]. First, the size of nanoparticles allows them to carry a large number of targeting ligands. The binding behavior is influenced by the ligands’ affinities, their density, and by the combination of the chosen ligands. Because of the resulting multivalent binding (“avidity”), a nanoparticle can exhibit high-affinity binding even though low-affinity ligands were employed [59] and different receptors may be targeted simultaneously. Second, moieties for multiple imaging modalities can be combined onto a single particle. This makes the nanoparticles amendable to dual-/multimodality imaging and allows one to draw on the strengths of different imaging modalities (e.g., high resolution and high sensitivity, Table 1.1). For a comparable small-molecule probe, achieving desirable signal strengths for each of the different imaging modalities can be challenging because the ratio of the different imaging moieties (e.g., an optical dye molecule and a radioisotope-bearing prosthetic group) cannot be easily changed [60] and the introduction of additional groups can significantly affect pharmacokinetics. By contrast, the ratio of the different moieties can be tailored more easily on nanoparticles, be it to achieve simultaneous detection in all modalities, or to use a high-sensitivity “beacon” (e.g., a PET-isotope) to trace low-sensitivity particles (e.g., MRI particles) at very low concentrations [61, 62]. Third, nanoparticles used as theranostics can carry large numbers of therapeutic entities like small molecules or peptides. Depending on the type of nanoparticle used, these payloads can be stowed away in the core, eliminating deleterious interference with pharmacokinetics and biodistribution. In addition, different types of drug molecules can be combined in a single particle. Fourth, drug-delivery kinetics can be tuned for a particular target and monitored using molecular imaging. A “controlled release” after delivery to the target helps to protect the rest of the body from harmful side effects, while at the same time increasing the “effective dose” in the target tissue. The drugs may be released by a microenvironmental stimulus such as pH or enzymes [63] or by external stimuli. An example for the latter are drugcarrying microbubbles that can be tracked by conventional ultrasound and fragmented using destructive high-intensity US pulses, releasing the drug in the target region [34]. When drug delivery directly into the cells is desired [64], multimodality nanoparticles can be modified for efficient endocytosis by decoration with cell-penetrating groups such as Tat-peptides [65, 66] and tracked by molecular imaging. Furthermore, there is evidence that nanoparticles may be able to sidestep multidrug resistance involving protein efflux pumps [57]. Fifth, nanoparticle pharmacokinetics can be modified with surface modifications. To improve the pharmacokinetic properties and in vivo stability, the particles are oftentimes

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coated with biocompatible polymers such as polyethylene glycol (PEG) or the polysaccharide dextran. This coating effectively “hides” the particles from detection by the reticuloendothelial system (RES), which would otherwise rapidly eliminate them by opsonization and phagocytosis [63]. All these modifications have to be accomplished while keeping the hydrodynamic (HD) size roughly between 10 and 100 nm. If particles fall below the lower limit, renal filtration provides a rapid way of elimination, while, on the other hand, hepatic and splenic elimination becomes increasingly efficient for particles greater than 100 nm [46]. For solid tumors, effective delivery to the target tissue beyond the vascular compartment is another reason for the upper size limit: since the gaps in leaky vasculature are about 400 nm, extravasation is most efficient for particles well below this limit [67]. This fact also explains why microbubble-based US imaging is largely confined to the vascular compartment; for other—smaller—nanoparticles, this passive targeting phenomenon, termed “enhanced permeation and retention” (EPR) effect, together with a lack of efficient lymphatic drainage, represents an important way of tumor-directed delivery, particularly for smaller tumors [68]. Besides the hydrodynamic size, the charge of the nanoparticle can also affect pharmacokinetics and biodistribution. Ideally, a nanoparticle should exhibit a near neutral charge since positively charged particles can form aggregates with negatively charged plasma proteins, while negatively charged particles demonstrate increased liver uptake [69]. The positive effects of nanoparticle coatings on pharmacokinetics include shielding of surface charges and surface chemistry. The commercial availability of PEG in a wide range of lengths and modifications has made it particularly popular for fine-tuning of these pharmacokinetic properties. For example, one study showed that after relatively small, nontargeted quantum dots with identical 3.2-nm diameter InAs(ZnS) core were coated with PEG-chains of various lengths (PEG2 to PEG22), resulting in HD sizes ranging from 5 to 16 nm, the biodistribution and clearance characteristics varied widely [70]. Expectedly, PEGylation improved hydrophilicity, protected from opsonization, and circulation times in the blood increased with increasing chain length. In addition, by simply increasing the length of the PEG units, major organ uptake shifted from liver over kidneys and bladder to pancreas, intestine, and lymph nodes. However, it has to be underscored that in the study cited above, the quantum dots did not bear any specific targeting moiety. Fortunately for targeted nanoparticles, recent data indicate that the presence of targeting moieties may minimize the influences of other nanoparticle properties in vivo [71]: a preclinical study compared different carriers (linear polymer, dendrimer, and liposome) carrying an anticancer drug (paclitaxel) and/or an imaging probe (Cy5.5) in conjunction with a targeting moiety (a LHRH peptide). Notably, the authors conclude “that the architecture, composition, size and molecular mass of the receptor-targeted drug nanocarriers can be selected based on other than anticancer efficacy considerations (cost, type of active ingredients, difficulties in production, stability, patient compliance, etc.) ensuring that the high efficacy and low adverse side effects could be achieved automatically by tumor targeting” [emphasis added] [71]. If similar results are found in other studies, the focus might shift toward careful selection of the targeting ligand(s) while allowing a high degree of flexibility for the underlying platform [72]. 1.4.2 Challenges for Nanoparticle-Based Imaging Probes The body of work produced to date has demonstrated the promising potential of nanoparticle-based molecular imaging agents [9, 11, 12]. At the same time, the studies

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have shown some of the accompanying challenges Sanvicens 2008 [57]. They chiefly fall within three basic areas: compound preparation, toxicity, and pharmacokinetics/ pharmacodynamics. Understanding and addressing these challenges will also be highly important to gain the regulatory approval necessary for rapid and successful translation into the clinic. Some of the challenges are described next. 1. Size Control and Batch-to-Batch Reproducibility. In contrast to small-molecule probes, achieving good control over the size and precise composition of nanoparticles is more challenging. But since pharmacokinetics and, for theranostics, pharmacodynamics can be significantly influenced by these factors, gaining a thorough understanding of these manufacturing parameters and their in vivo consequences is crucial [46, 69, 71]. Taking magnetic nanoparticles used for MRI as an example, it is well known that the (superpara)magnetic properties are size dependent as surface effects play an increasing role for smaller particles and that magnetic properties can vary significantly depending on production chemistry [32, 73]. 2. Particle Stability During Storage and Formulation. Because of their size and surface properties, nanoparticles may aggregate over time. This and other deleterious changes such as decomposition and oxidation have to be evaluated. 3. Toxicity and Immune Response. The composition and shape of the nanoparticles can potentially have toxic effects. For example, many quantum dots popular for optical fluorescence imaging contain cadmium, lead, or selenium [74]. Magnetic nanoparticles used for MRI may contain gadolinium, cobalt, or nickel [31, 75]. Even if these toxic elements are shielded from direct contact with the in vivo environment by protective coatings, they may eventually leach out if residence times in the body are long. For carbon nanotubes, their shape rather than their composition could be cause for concern. Similar to quartz and asbestos, they have been shown to be capable of causing chronic inflammation and fibrosis [76]. In addition, while coatings are useful for shielding the core of the nanoparticles from the immune system, if antibodies are used for targeting, they necessarily need to be exposed to the in vivo environment, thus posing a risk for an immune response. Fortunately, less immunogenic antibody-derived targeting moieties may be available and biodegradable low-toxicity nanoparticles are being developed [72]. 4. Control Over Pharmacokinetics/Pharmacodynamics. Nanoparticles demonstrate long blood circulation times and, because of their size, show a tendency to undergo opsonization and clear the body via the liver and spleen. Clearance through the reticuloendothelial system is slower than the excretion in the urine encountered for many small molecules [69]. The longer residence time in excretory organs may not be a significant concern for imaging probes administered in trace amount, but could become pharmacodynamically relevant for theranostics carrying highly potent drug molecules or therapeutic radioisotopes. In the end, solving some of these challenges could be easier than it appears at present. As the recent study by Saad and co-workers has indicated [71], depending on which information is desired about the target, choosing the imaging modality (or modalities) and targeting moiety may be the most important factors to consider when designing the imaging agent, while leaving considerable flexibility for the underlying platform. Specifically, the authors found “that tumor-specific targeting minimized the differences between nanocarriers of distinct architecture, size, mass, and composition” [71]. If these findings represent a

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general principle, they could have wide-ranging implications for the rapidly growing field of nanoparticle-based diagnostic and treatment tools. Thus despite the seemingly overwhelming number of potential combinations of sizes, shapes, compositions, surface coatings, targeting moieties, imaging modalities, functional groups, and potential drug molecules, in the future the challenge may not be so much trying to discover a molecular imaging agent for a particular target, but rather to assemble the best combination of components from the choices available in the toolbox.

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CHAPTER 2

Synthesis of Nanomaterials as a Platform for Molecular Imaging JINHAO GAO Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, USA

JIN XIE Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, and Laboratory for Molecular Imaging and Nanomedicine, National Institute of Biomedical Imaging and Bioengineering, National Institutes of Health, Bethesda, Maryland, USA

BING XU Department of Chemistry, Brandeis University, Waltham, Massachusetts, USA

XIAOYUAN CHEN Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, and Laboratory for Molecular Imaging and Nanomedicine, National Institute of Biomedical Imaging and Bioengineering, National Institutes of Health, Bethesda, Maryland, USA

Recent successful syntheses of narrowly dispersed nanomaterials have offered a unique opportunity to probe biological interactions and visualize biological structures. This chapter highlights the general strategies to produce novel nanomaterials for molecular imaging, such as quantum dots as fluorescent markers, metal oxide nanoparticles as magnetic resonance imaging (MRI) contrast enhancement agents, metallic nanoparticles as optical contrast agents, and multifunctional nanomaterials for multimodality imaging. Well-established and thoroughly characterized nanomaterials as a general platform may offer many prominent biomedical applications with high sensitivity and high affinity via molecular imaging. 2.1 INTRODUCTION Nanotechnology refers to the control, manipulation, and study of structures and devices with length scales in the range of 1–100 nanometers, which usually falls at atomic, molecular, and macromolecular scales. Recently, the successful syntheses of narrowly dispersed Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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nanomaterials have offered a unique opportunity to probe biological interactions and visualize biological structures. A new emerging field that combines nanotechnology and biotechnology, so called nanobiotechnology, is expanding new frontiers of knowledge [1, 2]. One of the most attractive topics in the nanobiotechnology field is molecular imaging based on the nanomaterials, that is, using nanomaterials as a platform to perform molecular imaging [3]. Molecular imaging has the ability to detect the biological and biochemical processes in living subjects in a noninvasive manner, which would assist many applications in biomedicine, such as drug development, diagnosis, and therapy [4, 5]. The properties and functions conferred by nanomaterials, for example, their high volume/surface ratio, surface tailorability, improved solubility, and multifunctionality, are expanding many new possibilities for molecular imaging. The intrinsic optical, magnetic, and biological properties of the nanomaterials offer remarkable opportunities to detect and study the complex biological processes for biomedical applications in an unprecedented manner. Multifunctionality is the key advantage of nanoplatform over traditional approaches. One can integrate targeting ligands, therapeutic drugs, imaging labels, and other agents into the nanoplatform to allow targeted molecular imaging and molecular therapy. One of the most advanced and exciting forefronts of nanobiotechnology is the use of quantum dots (QDs) as fluorescent dyes in biology and medicine [6]. In the past two decades, the advances in the synthesis of QDs significantly promoted fluorescent imaging for in vitro and in vivo applications. Another rapidly evolving nanoplatform-based molecular imaging is magnetic resonance imaging (MRI), that is, using nanomaterials (e.g., iron oxide nanoparticles, MnO nanoparticles) or the nanomaterials functionalized by paramagnetic entities (e.g., Gd3+ ) as MRI contrast agents. Many research groups have intensively pursued the development of controllable uniform metal oxide nanoparticles as MRI contrast enhancement agents. Indeed, chemists developed many synthetic strategies and provided assays of nanomaterials for the potential applications in molecular imaging. Based on these well-established nanomaterials, biologists and doctors can extend the potential biomedical applications of the nanomaterials and further improve them for clinical applications. After all, the interdisciplinary collaboration in the fields of material science, chemistry, biology, and clinical medicine is very important for the development of the nanoplatform-based molecular imaging. In this chapter, we give an overview on the synthetic strategies of nanomaterials used for molecular imaging in living subjects. First, we summarize the synthesis of QDs and metal oxide nanoparticles. We then give a brief introduction to the synthesis of metallic nanomaterials (e.g., Ag, Au). Finally, we discuss the design and synthesis of multifunctional nanomaterials for multimodality molecular imaging.

2.2 SYNTHESIS OF QUANTUM DOTS AS FLUORESCENT DYES QDs find many applications as fluorescent labels in biology because the mere change of their diameters significantly shifts their emission maxima. As now well recognized, besides their narrow, tunable, and symmetric emission spectra, QDs have much greater temporal stability and resistance to photobleaching than fluorescent dyes do. Over the past three decades, the syntheses of QDs with different materials have been well developed. In the early 1980s, researchers started to focus on the synthesis of semiconductor nanostructures because they realized the photophysical properties of semiconductor structures were important for a broad range of computer and electronic applications. Henglein [7] and Rossetti et al [8].

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reported the first method for synthesizing colloidal QDs. They mixed the cadmium and sulfide salts in an aqueous buffer to produce CdS QDs. Researchers in the late 1980s and early 1990s attempted to improve the overall optical qualities of these QDs (e.g., improving particle monodispersity and size tunability) by investigating the effect of different reaction conditions (e.g., solvents, salts, precursor ratios, pH, and temperature) on their morphological and optical properties. These early fundamental research findings provided a guide for the current state-of-the-art in QD synthesis, an organometallic synthetic procedure developed by Murray et al [9]. in 1993 for making monodispersed QDs, such as CdSe nanocrystals. For the typical synthesis of CdSe QDs by an organometallic procedure, dimethylcadmium (Cd(CH3 )2 ) and selenium were initially dissolved in the organic solvent trin-octylphosphine (TOP) at defined ratios and injected in a hot coordinating solvent of tri-n-octylphosphine oxide (TOPO; 350 ◦ C) under argon gas protection. Nucleation and growth of the nanocrystals of CdSe happened after injection and lowering of the temperature from 350 to 300 ◦ C, indicated by a change in the color of the solution (from clear to light yellow to orange to red). According to this method, CdSe QDs that have different shapes can be produced by tuning the reaction conditions [10, 11]. However, Cd(CH3 )2 precursor is extremely toxic, pyrophoric, expensive, unstable at room temperature, and explosive at elevated temperatures by releasing a large amount of gas. In 2001, Peng and co-workers reported the synthesis of CdS, CdSe, and CdTe QDs using CdO as precursor [12, 13]. Compared with Cd(CH3 )2 , CdO is much safer and cheaper. This “greener” method quickly gained in popularity. The emission wavelength of Cd-based QDs can only reach about 700 nm in terms of their bandgap energy. Alivisatos and co-workers reported the synthesis of InAs nanocrystals with the emission wavelength from 700 to 1300 nm by tuning the size of the nanoparticles from 3 to 7 nm [14]. As shown in Figure 2.1, the emission wavelength of the colloidal QDs made of ZnS, CdSe, CdTe, PbS, PbSe, and InAs covers the UV to infrared range. Because of their wide absorption spectra, QDs with different emission wavelengths excited by a single light source can emit various color fluorescence; therefore QDs have a great potential for multicolor molecular imaging. However, these synthetic processes may need refinement for biological applications as issues of reproducible synthesis and inorganic passivation remain. The breakthrough in synthesizing high-quality colloidal QDs came with the reports of CdSe/CdS core–shell QDs with high quality and strong luminescence [15, 16]. Because of the toxicity and instability of the CdSe and InAs QDs surface, it is extremely important

FIGURE 2.1 Representative QDs with different materials scaled as a function of their emission wavelength superimposed over the spectrum.

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to passivate or cap the CdSe and InAs QDs with a layer of ZnS or ZnSe. The ZnS or ZnSe layer improves the fluorescence quantum yield of the QDs and protects them against photo-oxidation (which is important for minimizing cytotoxicity and for enhancing photostability). The ZnS shell has a larger bandgap energy than CdSe, eliminating the core’s surface defect states. Also, the ZnS shell has a similar bond length to that of CdSe, minimizing crystal–lattice strain and allowing for epitaxial growth. The strategy of using a ZnS or ZnSe shell to cap QDs has become a popular approach. In order to meet the requirements of in vivo biological imaging applications, the fluorescent emission wavelength of the QDs ideally should be in a region of the spectrum where blood and tissue absorb minimally but still be detectable by the instruments. Thus the QDs should emit at approximately 700–900 nm in the NIR region to minimize the problems of indigenous fluorescence of tissues. Because of their interest and high demand, the development of NIR QDs has progressed rapidly. Recently, Bawendi’s group developed several types of QDs with near-infrared emission, such as CdTe@CdSe [17], InAs@InP@ZnSe [18], and InAs@ZnSe [19] QDs with emission wavelengths between 700 and 850 nm. Peng and co-workers also reported several types of NIR QDs based on an InAs core [20, 21]. Due to the potential cytotoxicity of cadmium- and arsenic-containing QDs, it is imperative to replace the toxic metals by other more benign elements. CuInSe QDs, promising candidates, have recently been reported by Bawendi’s group [22]. The high-temperature organic phase synthetic strategy produces QDs that are nonpolar and insoluble in aqueous solvents. In order to render the QDs water soluble, surface modification of the nanoparticles is necessary, such as chemical exchange and hydrophobic–hydrophobic interaction, which sometimes needs complicated manipulations and yet with low quantum yield. The water-dispersed QDs can be synthesized directly via an aqueous method, which is simpler, cheaper, and less toxic, for example, thiol-capped CdTe QDs [23, 24]. Recently, Wang and co-workers reported the rapid synthesis of waterdispersed CdTe, CdTe@CdS core–shell, and CdTe@CdS@ZnS core–shell–shell QDs in aqueous phase assisted by microwave irradiation [25–27]. This method may open up a new route for the synthesis of water-dispersed QDs.

2.3 SYNTHESIS OF METAL OXIDE NANOPARTICLES AS MRI CONTRAST AGENTS Metal oxide nanoparticles, especially iron oxide nanoparticles, have shown many potential applications in biology and medicine. The most prominent biomedical application of iron oxide nanoparticles is their usage as MRI contrast enhancement agents, which have already received some clinical acceptance [28]. Among many types of iron oxides, superparamagnetic iron oxide (SPIO) nanoparticles have become the front runner for gaining regulatory approval. Recent advances in the synthesis of various metal oxide nanoparticles by using colloidal chemical approaches offer the capability to precisely control the size of iron oxide nanoparticles from 2 to 20 nm in diameter [29–31]. As shown in Table 2.1, besides iron oxide nanoparticles, there are several other types of metal oxide nanoparticles that have been developed as MRI contrast agents, such as manganese oxide [32] and gadolinium oxide (Gd2 O3 ) [33]. Sun and co-workers developed the synthesis of monodisperse magnetite (Fe3 O4 ) nanoparticles by a high-temperature solution phase reaction of Fe(acac)3 and 1,2hexadecanediol in the presence of oleic acid and oleylamine as surfactants [34, 35]. Particle

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TABLE 2.1 Several Types of Metal Oxide Nanoparticles as Potential MRI Contrast Enhancement Agents Metal Oxide Magnetite (Fe3 O4 ) ␥ -Fe2 O3 Manganese oxide Gadolinium oxide (Gd2 O3 )

Methods

Reference

Thermal decomposition Decomposition and oxidation Thermal decomposition Thermal decomposition, polyol synthesis

34, 35 36–38 32, 38 33, 39–41

diameter can be tuned from 3 to 20 nm by varying reaction conditions or by seed-mediated growth. For example, 4-nm Fe3 O4 nanoparticle seeds were synthesized by refluxing a reaction mixture composed of Fe(acac)3 , diphenyl ether, 1,2-hexadecanediol, oleic acid, and oleylamine. By controlling the quantity of the nanoparticle seeds and reaction conditions, they generated various larger sized Fe3 O4 nanoparticles. Figure 2.2 shows the TEM images of different sizes of Fe3 O4 nanoparticles synthesized by the seed-mediated growth. Magnetic measurements indicate that the nanoparticles are superparamagnetic at room temperature (no hysteresis); that is, the thermal energy can overcome the anisotropy energy barrier of a single particle to result in zero net magnetization of the nanoparticle assemblies in the absence of an external field. The mixture of metal (Co, Mn, and Ni) precursors in the reaction resulted in metal-doped iron oxide nanoparticles, which also show good MRI contrast enhancement [35], especially for Mn-doped iron oxide nanoparticles [42]. However, elements other than iron may cause cytotoxicity, thus limiting their in vivo applications. In the past ten years, Hyeon’s group has contributed several excellent strategies on the synthesis of monodisperse iron oxide nanoparticles. In 2001, they reported on the synthesis of monodisperse and highly crystalline maghemite (␥ -Fe2 O3 ) nanoparticles [36]. Monodisperse iron nanoparticles were first synthesized and then were transformed to monodisperse ␥ -Fe2 O3 nanocrystals by controlled oxidation using trimethylamine N-oxide ((CH3 )3 NO) as a mild oxidant. Particle size could be varied from 4 to 20 nm by altering the experimental conditions. The dominating size-controlling factor was the molar ratio of Fe(CO)5 to oleic acid. Figure 2.3 shows the TEM images of ␥ -Fe2 O3 nanoparticles with the sizes of 7, 11, and 13 nm in diameter. A subsequent study by Hyeon and co-workers [37] implied that the oxidation state of iron in these particles also depended on the sizes of the nanoparticles. Using this method, which is similar to the seed-mediated growth, they succeeded in producing monodisperse iron oxide nanoparticles with a continuous size spectrum of 6–13 nm in diameter without size-selection processes [37]. Based on the thermal decomposition of a metal–surfactant complex, Hyeon and coworkers further developed large-scale syntheses of monodisperse iron oxide nanocrystals [38]. Instead of using toxic and expensive organometallic compounds such as Fe(CO)5 , they prepared the iron–oleate complex by reacting inexpensive and environmentally friendly compounds, namely, metal chlorides and sodium oleate. They generated monodisperse iron oxide nanocrystals by heating 1-octadecene (ODE) solution with iron–oleate complex to 320 ◦ C and keeping that temperature for 30 minutes. The amount of the separated nanocrystals produced was up to 40 g with a yield higher than 95%. The iron oxide nanocrystals can easily redisperse in various organic solvents including hexane and toluene. Moreover, the size of the nanoparticles could be controlled with reasonable precision by simply varying the experimental conditions, with a range from 5 to 22 nm.

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(a)

20 nm

(b)

20 nm

(c)

20 nm

FIGURE 2.2 TEM images of Fe3 O4 nanoparticles with a diameter of (a) 6 nm, (b) 10 nm, and (c) 12 nm. (Reproduced with permission from Sun et al [35]. Copyright © 2004 American Chemical Society.)

Recently, Sun’s group reported the synthesis of iron nanoparticles with high chemical stability and good dispersion by controlled oxidation of the iron surface to crystalline Fe3 O4 shell [43]. The crystalline Fe3 O4 shell offered more robust protection to the metallic

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(a)

20 nm

(b)

20 nm

(c)

20 nm

FIGURE 2.3 TEM images of ␥ -Fe2 O3 nanoparticles with a diameter of (a) 7 nm, (b) 11 nm, and (c) 13 nm. (Reproduced with permission from Hyeon et al. [36]. Copyright © 2001 American Chemical Society.)

core of iron, and the core/shell-structured Fe@Fe3 O4 nanoparticles were stable in hexane or water dispersion for a certain period of time. So it is possible to produce small Fe nanoparticles (<10 nm in radius) with the desired stability for highly efficient biological

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separation [44], drug delivery, and high sensitivity in biological detection [45]. Sun and coworkers generated the crystalline Fe3 O4 shells by controlled oxidation of the as-synthesized nanoparticles using the oxygen transferring agent (CH3 )3 NO. The thickness of the shell was tunable by controlling the amount of (CH3 )3 NO added into the reaction mixture. For example, they produced the 2.5-nm/5-nm Fe@Fe3 O4 core/shell nanoparticles (Fig. 2.4a) by oxidizing as-synthesized Fe nanoparticles with (CH3 )3 NO. Magnetic measurements of the 2.5-nm/5-nm Fe@Fe3 O4 nanoparticles show that they are superparamagnetic with the magnetic moment reaching 61.6 emu/g. Moreover, the gap between the core and shell shown in Figure 2.4a suggests the Kirkendall effect [46–48]. The further oxidization of Fe@Fe3 O4 nanoparticles results in the hollow Fe3 O4 nanoparticles (Fig. 2.4b) [49]. Hollow Fe3 O4 magnetic nanoparticles may exhibit advantages in MR relaxation enhancement due to their unique geometry. The interface between the iron oxide hollow nanoparticles and

(a)

10 nm

(b)

20 nm

FIGURE 2.4 TEM images of (a) Fe@Fe3 O4 core–shell nanoparticles and (b) Fe3 O4 hollow nanoparticles synthesized due to the Kirkendall effect.

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the water phase is much larger than that of solid nanoparticles for unit weight or mass, which may contribute to the stronger MR contrast enhancement per unit weight [50, 51]. The thermal decomposition method developed recently has become the most effective approach for preparing high-quality iron oxide nanoparticles. However, most of the products made by the thermal decomposition approach only disperse well in organic solvents, and postpreparative procedures and modifications are necessary to transfer the hydrophobic iron oxide nanoparticles to the water phase before further biomedical applications. It is thus worthwhile to develop a direct synthesis (i.e., one-pot synthesis) method of water-dispersible iron oxide nanoparticles. Recently, Gao and co-workers reported the one-pot synthesis of water-dispersible and biocompatible Fe3 O4 nanoparticles in strong polar 2-pyrrolidone solvent using Fe(acac)3 or hydrated ferric salts as iron precursors [52–54]. The method provides a new way to prepare water-soluble Fe3 O4 nanoparticles, although the size distribution of the resulting nanomaterials has yet to be improved. Besides iron oxide nanoparticles, there are many other metal oxide nanoparticles that can act as MRI contrast enhancement agents. Hyeon and co-workers reported that the waterdispersible and biocompatible MnO nanoparticles can serve as T1 MRI contrast agent [32]. The thermolysis of Mn–oleate complex in organic solvent with high boiling point can generate MnO nanoparticles with a very narrow size distribution (Fig. 2.5a) [38]. Cao [33] reported the solution synthesis of plate-shaped Gd2 O3 nanocrystals (Fig. 2.5b) by thermal decomposition of gadolinium acetate in the presence of oleylamine and oleic acid. The high-quality Gd2 O3 nanocrystals have potential as a T1 MRI contrast agent if the toxicity of Gd(III) could be managed.

2.4 SYNTHESIS OF METALLIC NANOMATERIALS Metallic (e.g., Ag, Au, and Pt) nanomaterials have fascinated scientists for several centuries, partly because of the amazing colorful colloidal solution of metallic nanomaterials, in which resonant electron oscillations on the surface of noble metal nanomaterials create the surface plasmon resonance (SPR) that greatly enhances the absorption and Rayleigh (Mie) scattering of light by these nanomaterials. Because of their inherent ultrahigh density, enhanced resonance absorption and scattering properties, and strong Raman scattering, the metallic nanoparticles are useful for some biomedical applications, such as optical contrast agents, multimodal sensors combining optical imaging and scattering imaging, photothermal therapy, and Raman probes [55–57]. The chemical, optical, and thermal properties of the metallic nanomaterials heavily depend on the shape of the nanomaterials [58, 59]. For example, adjusting the size and shape of the Au particles from spheres to rods can tune the SPR absorption from the visible to the NIR region. In this section, we summarize the synthesis of Ag and Au nanomaterials with different shapes from nanoparticles to nanorods and nanowires, from nanocubes to nanocages. The Ag nanoparticles can easily grow in aqueous solution containing ethylene glycol and poly(vinyl pyrrolidone) (PVP) polymer to prevent aggregation, the so-called polyol synthesis. In most solution-phase syntheses of Ag nanoparticles, the results are the mixture of single-crystalline and twinned structures. In order to produce single-crystalline nanoparticles at high yields without the cumbersome separation steps, it is necessary to eliminate the twinned nanoparticles by selective etching [60]. Because PVP interacts more strongly with <100> than <111> facets and passivates silver atoms on the <100> facets, the silver

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(a)

50 nm

(b)

10 nm

FIGURE 2.5 TEM images of (a) MnO nanoparticles and (b) self-assembly of Gd2 O3 nanoplates.

nanowires along the <100> direction can be formed by reducing silver nitrate with hot ethylene glycol in the presence of PVP [61]. There are many methods to produce Au nanoparticles with uniform size and structure in either the aqueous phase or organic phase by using different surfactants. For example, started by Faraday in the 1850s, using tetrachloroauric acid (HAuCl4 ) as precursor and sodium citrate as surfactant, Au nanoparticles with different sizes (>2 nm) were formed in the aqueous phase [62]. Using HAuCl4 as precursor, sodium borohydride (NaBH4 ) as the reductant, and quaternary ammonium bromide salts (R4 N+ Br− ) as phase-transfer agents, Au nanoparticles about 8 nm in diameter can be formed in toluene by using a two-phase system [63]. In single organic phase (e.g., toluene), the large-scale synthesis of Au nanoparticles can also be easily achieved using oleylamine as reductant and surfactant [64]. Moreover, the well-established surface chemistry of Au nanomaterials with organic ligands containing the thiol group [65] generated several strategies for the synthesis of

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Au nanomaterials. In particular, thiol-derivatized Au nanomaterials can be prepared in one-step, two-phase reactions [66–68]. The stability of Au nanomaterials conferred by the attachment of thiols in self-assembled structures is remarkable; that is, the solutions show no sign of decomposition or loss of solubility even after several months of storage at room temperature [66]. One-dimensional metallic nanostructures such as nanorods and nanowires are of tremendous interest for sensing and imaging applications [69–71]. Shape anisotropy introduces new optical properties in gold and silver nanomaterials, such as longitudinal plasmon resonance bands in the visible and NIR portions of the spectrum. The synthesis of nanorods using surfactants can be achieved by electrochemical, photochemical, and seed-mediated methods [72–74]. Typically, a metal salt such as AgNO3 or HAuCl4 is reduced in the presence of a surfactant to yield one dimensional (1D) nanostructures. The aqueous metal salt solutions are reduced by the passage of an applied current, by irradiation with photons, or by the use of a chemical reductant, such as hydrazine, citrate, or ascorbic acid. The surfactant is critical to the generation of 1D nanostructures. One simple chemical reduction technique often used for the preparation of Au and Ag nanorods and nanowires is the seed-mediated method, which has been well developed by Murphy and co-workers [74–77]. First, a metal salt of gold or silver is initially reduced by a strong reducing agent such as NaBH4 to form spherical nanoparticles in water. Then, these spherical nanoparticles can be used as seeds to grow nanorods. In the presence of more metal salt, surfactants, and additional but milder reducing agents, such as ascorbic acid, seeds grow into nanorods or nanowires. Using this strategy, Au nanorods with various aspect ratios have been prepared, as shown in Figure 2.6. Among a large variety of metallic nanostructures, Ag nanocubes and Au nanocages are two impressive and interesting structures. Sun and Xia [78] reported the large-scale synthesis of Ag nanocubes with high quality (Fig. 2.7a) by reducing AgNO3 with ethylene glycol (polyol synthesis) in the presence of PVP as the surfactant. The presence of PVP and its molar ratio relative to AgNO3 both play important roles in determining the geometric shape and size of the product mostly due to the selective adsorption of PVP on the various crystallographic planes of the silver nanocrystalline seeds. Using Ag nanocubes with truncated corners as the sacrificial template, they further synthesized Au nanocages with well-controlled pores at the corners (Fig. 2.7b) by galvanic replacement reaction [79]. The resulting Au nanocages have large absorption cross sections because of the discrete dipole approximations, and they also display a large photothermal effect with absorbed photons being converted into phonons (i.e., lattice vibrations) to produce a localized temperature jump. Moreover, the Au nanocages show strong SPR absorption in the NIR region, together with potential computed tomography (CT) contrast enhancement of Au nanomaterials; therefore they are particularly attractive for biomedical applications that require the selective absorption of light at great depths [80, 81].

2.5 DESIGN AND SYNTHESIS OF MULTIFUNCTIONAL NANOMATERIALS FOR MULTIMODALITY IMAGING Nanostructures provide the platform to integrate different functional nanocomponents into one single nano-entity to exhibit multifunctional properties. To assemble different nanoparticles into a single entity offers a unique building block of materials. Based on the QDs, metal oxide nanoparticles, and metallic nanoparticles, the combination of two of them

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(a)

(b)

(c)

100 nm

FIGURE 2.6 TEM images of Au nanorods with aspect ratios of (a) ∼4.6, (b) ∼13, and (c) ∼18. (Reproduced with permission from Jana et al. [75]. Copyright © 2001 American Chemical Society.)

will produce several kinds of multifunctional nanomaterials as dual-mode imaging agents. There are already several review papers that summarize and discuss the synthesis of this type of hybrid nanostructures [82–87]. Several useful applications, in the study of subcellular processes of fundamental importance in biology, have highlighted the potential of QDs in nanobiotechnology [88–90]. Such exciting and attractive properties of the QDs have inspired the fabrication of hybrid nanostructures based on QDs, such as Co@CdSe core–shell nanocomposites [91], Fe2 O3 –CdS heterojunctions [92], FePt–CdX (X = S or Se) heterodimer or core–shell nanoparticles [93, 94], and CdSe–Au heteronanorods [95].

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(a)

(b)

200 nm

FIGURE 2.7 SEM images of (a) Ag nanocubes and (b) Au nanocages (insets, TEM images). (Reproduced with permission from Chen et al. [79]. Copyright © 2006 American Chemical Society.)

Based on the FePt magnetic nanoparticles and semiconducting chalcogenide nanocomponents, the systematic studies show that the reaction condition variables control the formation of different hybrid nanostructures. In the one-pot reaction, the sequential growth of CdX (X = S or Se) onto the FePt nanoparticles under the lower reaction temperature will result in the formation of FePt@CdX core–shell nanoparticles (Fig. 2.8a,b). However, use of a higher boiling point solvent will result in FePt-CdX heterodimer nanoparticles (Fig. 2.8c,d). The formation of heterodimer nanoparticles at the higher temperature is probably due to the difference in phase transition temperatures between the FePt and CdX components. The CdX components may melt at the higher temperature and induce dewetting from the FePt cores, resulting in the formation of heterodimeric nanostructures [93]. The synthesis of these core–shell nanoparticles and heterodimer nanoparticles is highly reproducible and general. Although the optical properties of these nanostructures must be improved for higher quantum yields, this rather simple approach indeed opens an avenue to design and synthesis of sophisticated and multifunctional nanostructures. The combination of superparamagnetism and fluorescence at the nanometer scale should help expand the biological applications of multifunctional nanomaterials. Although the fluorescence of QDs can be partially quenched by metallic nanoparticles in FePt–CdX hybrid nanostructures, the replacement of the metallic nanoparticle part with a metal oxide nanoparticle has resulted in a higher quantum yield. For example, the Fe3 O4 –CdSe heterodimer nanoparticles show an emission wavelength peak at 610 nm with quantum yield of about 38% [96]. The resulting

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Core-Shell

Heterodimer (c)

(a)

5 nm

20 nm (b)

5 nm (d)

5 nm

20 nm

5 nm

FIGURE 2.8 The formation of core–shell or heterodimer nanostructures by FePt and CdX (X = S or Se) nanocomponents. The core–shell structures of (a) FePt@CdS nanoparticles and (b) FePt@CdSe nanoparticles (insets, HRTEM images). The heterodimer structures of (c) FePt@CdS nanoparticles and (d) FePt@CdSe nanoparticles.

fluorescent magnetic nanoparticles bear two attractive features, superparamagnetism and fluorescence, which allow their intracellular movements to be controlled using magnetic force and to be monitored using a fluorescent microscope. Using these fluorescent magnetic nanoparticles, it is straightforward to assess the intracellular manipulation of nanoparticles by a magnetic force. It also promises a good model for dual-functional molecular imaging (i.e., combining MRI and fluorescence imaging) [97]. Based on CdSe dots and rods, Banin and co-workers reported on the growth process of the formation of symmetric or asymmetric heterostructures [95, 98]. They found that one-sided growth of a gold tip on the CdSe dots or rods is preceded by the two-sided growth. As shown in Figure 2.9, at lower gold precursor (AuCl3 ) concentrations, two-sided growth is evident in the simulation (Fig. 2.9a) and was confirmed by experimental result (Fig. 2.9b). As the gold concentration is increased, a marked change happens, instead of two-sided growth, the one-sided growth process becomes dominant and results in heterodimer nanostructures (Fig. 2.9c,d). Experimental analysis and theoretical modeling show that a ripening process drives gold from one end to the other. Ripening is therefore occurring effectively on the nanostructure itself, leading to a phase-segregated structure and thus extending the realm of ripening phenomena and their importance in nanostructures. The combination of metallic nanoparticles and magnetic nanoparticles certainly promises many applications in nanoplatform multimodality molecular imaging. Recently,

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(c)

(a)

(b)

(d)

100 nm

100 nm

FIGURE 2.9 The formation of CdSe–Au heterostructures at (a, b) low gold concentrations and (c, d) high gold concentrations.

there have been several strategies developed for the design and synthesis of hybrid nanostructures containing magnetic nanoparticles and metallic nanoparticles. The simple and efficient method to form magnetic nanoparticle-based heterodimer structures is the sequential growth of metallic components (e.g., Ag or Au) onto the “colloidosome,” that is, the self-assembly of nanoparticles at a liquid–liquid interface [99]. As shown in Figure 2.10a, the metallic nucleation reaction takes place on the exposed surface of the magnetic nanoparticles and produces heterodimers of two distinct nanospheres. TEM and HRTEM images (Fig. 2.10b,c) shows the Fe3 O4 –Ag heterodimer nanoparticles with uniform size and structure, indicating that the method of liquid–liquid interface heterogeneous growth is a general way to make heterodimer nanoparticles, such as FePt–Ag and Au–Ag heterodimer nanoparticles. Recently, Sun and co-workers reported another way to fabricate Fe3 O4 –Au heterodimer nanoparticles with controllable size of each part of the component [100]. As shown in Figure 2.10d, in the homogeneous organic solvent (e.g., 1-octadecene), the thermal decomposition of Fe(CO)5 on the surface of the Au nanoparticles and the following oxidation of the intermediate resulted in the uniform Fe3 O4 –Au heterodimer nanoparticles (Fig. 2.10e,f). Based on the optical properties of metallic nanoparticles and magnetic nanoparticles as potential MRI contrast agents, these types of heterodimer nanoparticles show great potential in multimodal biomedical applications, especially for molecular imaging [101–103]. Recently, Xu and co-workers developed FePt@CoS2 yolk–shell and FePt@Fe2 O3 yolk–shell nanoparticles as a potential nanodevice in controlled drug release [104]. The unique structure and properties of FePt@CoS2 yolk–shell nanoparticles give a good example for the investigation of a novel drug delivery system that uses the magnetic nanomaterials to directly encapsulate the potential anticancer drug. The premise is that

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(a)

(b)

Ag+

(c)

Ag

organic droplet

Fe3O4 10 nm

aqueous phase (d)

(e)

Au Fe

Ag

1 nm (f)

Fe(CO)5

Fe3O4(111) 0.485 nm

organic phase oxidation

Au(111) 0.24 nm

Au Fe3O4

24 nm

2 nm

FIGURE 2.10 The two approaches for the synthesis of Fe3 O4 –M (M = Ag or Au) heterodimer nanoparticles. (a) The formation of Fe3 O4 –Ag heterodimer nanoparticles at a liquid–liquid interface. (b) TEM and (c) HRTEM images of Fe3 O4 –Ag heterodimer nanoparticles. (d) The formation of Fe3 O4 –Au heterodimer nanoparticles by the decomposition of Fe(CO)5 on the surface of the Au nanoparticles followed by oxidation in organic solvent. (e) TEM and (f) HRTEM images of Fe3 O4 –Au heterodimer nanoparticles.

platinum-containing nanoparticles (e.g., FePt) without any surface coating may be used as a potential anticancer drug. The sequential growth of CoS2 porous nanoshells by the Kirkendall effect [46–48] provides the feasibility to produce the “naked” FePt nanoyolks and the special interface between the outside and inside of the shells (Fig. 2.11a). The bifunctional FePt@Fe2 O3 yolk–shell nanoparticles exhibit high cytotoxicity and strong MR contrast enhancement [50]. Recent development on the synthesis of iron oxide hollow nanoparticles [49, 105] facilitates the design of FePt@Fe2 O3 yolk–shell nanoparticles. The direct injection of Fe(CO)5 into the refluxing solution of 1-octadecene containing oleylamine and FePt nanoparticles forms FePt@Fe core–shell nanoparticles as the intermediate. The further oxidation of as-prepared FePt@Fe core–shell nanoparticles in 1-octadecene under ambient atmosphere in the presence of O2 produces FePt@Fe2 O3 yolk–shell nanoparticles with uniform size and structure (Fig. 2.11b). Given the capability of surface functionalization of these yolk–shell nanoparticles using disease-specific molecules (e.g., antibodies), one could develop the multifunctional nanoparticles that target a specific tissue for delivering the therapeutic agents to tumors and monitoring the transformation of the tumor by MRI. Based on the multimodal features, multifunctional nanomaterials provide many new tools and opportunities for the exploration of life science because of their integrated functions. For example, based on the magnetic nanoparticles as MRI contrast agents, one of

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(a)

5 nm

(b)

5 nm

FIGURE 2.11 TEM images of (a) FePt@CoS2 yolk–shell nanoparticles and (b) FePt@Fe2 O3 yolk–shell nanoparticles.

the most exciting applications is the multimodal molecular imaging that integrates optical imaging (e.g., dyes, QDs, and Au nanoparticles) or positron emission tomography (PET) imaging (e.g., isotopes) with MRI [106–110]. Multimodality molecular imaging will likely become a dominant imaging tool for future biomedical research because it has the potential to provide more accurate information that is critical for clinical diagnostics and therapeutics [111–113].

2.6 SUMMARY AND PERSPECTIVE In this chapter, we outlined the syntheses of the nanomaterials that can serve as a platform for molecular imaging. Based on QDs, magnetic nanoparticles, and metallic nanomaterials, multifunctional nanomaterials combine at least two of these components for multimodality imaging. There are several synthetic strategies for different kinds of nanomaterials depending on different purposes. Most of the synthetic strategies are organic solvent-based reactions, which result in nonpolar soluble nanomaterials, so further surface modification

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and functionalization are needed for biological molecular imaging applications. The purity, dispersibility, and stability of the nanomaterials in the physiological environment are highly important. Recent successful syntheses of narrowly dispersed nanomaterials have offered abundant opportunities for molecular imaging applications as nanoprobes. The designs and syntheses of sophisticated and multifunctional nanomaterials open up a new avenue for multimodality molecular imaging. Future nanomedicines mostly lie in the multifunctional nanoplatformbased materials, which combine therapeutic components and nanoprobes. Therefore it is necessary and important to further study and explore multifunctional nanomaterials for creating successful nanobiotechnology to meet increasing demands in therapeutic and diagnostic applications.

ACKNOWLEDGMENT We gratefully acknowledge Stanford University and Brandeis University for financial support.

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CHAPTER 3

Nanoparticle Surface Modification and Bioconjugation JIN XIE Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, and Laboratory for Molecular Imaging and Nanomedicine, National Institute of Biomedical Imaging and Bioengineering, National Institutes of Health, Bethesda, Maryland, USA

JINHAO GAO and MARK MICHALSKI Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, USA

XIAOYUAN CHEN Molecular Imaging Program at Stanford and Bio-X Program, Department of Radiology, Stanford University School of Medicine, Stanford, California, and Laboratory of Molecular Imaging and Nanomedicine, National Institute of Biomedical Imaging and Bioengineering, National Institutes of Health, Bethesda, Maryland, USA

The last decade has witnessed fast progress in nanotechnology. Especially, the marriage of nanotechnology and conventional imaging facilities has brought forth a new research realm that holds great potential in bringing revolutionary changes to the current diagnostic technologies. The interface that connects nanomaterials and functional biomolecules plays a critical role in determining the performance of nanoconjugates in living subjects, and a lot of effort has been directed to this area. This chapter attempts to review the current status of biomolecule–nanoparticle conjugate preparation and use, with an emphasis on introducing the methodologies that have been developed to achieve a physiologically stable and functioning extendable nanoparticle coating.

3.1 INTRODUCTION Nanoparticle-based contrast agents, with their unique size (several to hundreds of nanometers in diameter) and physical properties (magnetic, optical, acoustic, etc.), hold great promise for revolutionizing current radiologic methodologies by offering more accurate, specific, and detailed imaging at both the cellular and subcellular levels. Recent Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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breakthroughs in nanoparticle preparation have permitted researchers unparalleled control over nanoparticle size, shape, and composition, all of which have implications for this technology’s translatability to clinical practice [1–3]. Despite these advances in nanoparticle synthesis, progression of nanoparticle technology to the clinical arena remains challenging, in part because of the poor native biocompatibility of these particles. Many of these particles, as-synthesized, bear an unfavorable layer, which causes them to aggregate in the physiology environment and severely compromises their in vivo applications. Thus particle modification, through the alteration of the particle surface coating layer to achieve sufficient hydrophilicity, has been a critical focus area within nanoparticle research. Beyond providing biocompatibility, postsynthesis modifications are important for achieving the potential nanoparticles integrated with disease-targeting agents; much of the impetus behind developing the nanoparticles discussed here is their ability to effectively transduce the presence of disease in a targeted fashion, to serve as an in vivo biosensor or delivery agent. The conventional nanoparticle technology very often relies on the enhanced permeability and retention (EPR) effect, a nonspecific pathophysiologic property of neoplastic tissue to naturally retain and concentrate particles of a certain size, to passively target and enhance tumors [4, 5]. While this has been a useful approach in the past, highly sensitive, specific pathology targeting ultimately requires the addition of specific targeting biomolecules to illuminate biomarkers present in various disease states. Such targeting biomolecule–particle conjugates can be formed through various means, including electrostatic adsorption, biological interaction, or, more commonly, through covalent linkage. Forming robust conjugates can be a challenging process, largely because standard chemical linkage techniques often cannot be used due to the fragility of the candidate biomolecules, most of which can only withstand mild reaction conditions (pH 5–10, 4 ◦ C to room temperature). Whatever the functional agent, successful conjugation usually depends on effective nanoparticle surface modification, adding a degree of complexity when attempting to render a nanoparticle biocompatible. In this chapter, we describe standard practices and some recent developments in nanoparticle surface modification and conjugation. Surface modification is highly dependent on the nature of the native nanoparticle surface and is therefore highly variable among different types of particles. In contrast, bioconjugation techniques are more broadly applicable: amine-terminal molecules, for example, can be linked through forming an amide bond with any nanoparticle with a carboxyl-containing coating layer, independent of the nanoparticle’s core composition. Because of their interchangeability, we first summarize applicable conjugation techniques that are routinely utilized to link functional biomolecules with nanomaterials. Surface modification methods for the most common nanoparticle varieties are then discussed in the second section.

3.2 CONJUGATION TECHNIQUES In this section, we briefly introduce the most commonly utilized methods for conjugation. For a complete list of the conjugation methods and additional details, readers are referred to Bioconjugate Techniques written by Greg Hermanson [6] and Recent Advances of Bioconjugation Chemistry in Molecular Imaging by Xiaoyuan Chen [7]. Here we separate existing techniques into three broad categories: zero-length crosslinking, crosslinker modulated linking, and conjugation through streptavidin–biotin interaction.

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3.2.1 Zero-Length Crosslinking Zero-length crosslinks are chemical bonds that are formed directly between chemical groups on particles surfaces and biomolecules, without the use of a mediating linker. Typically, some chemical reagents are added to catalyze the linking between these two entities. One commonly utilized method within this class is 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC) mediated carboxylate-amine coupling. EDC is a watersoluble derivative of N,N  -dicyclohexylcarbodiimide (DCC), which converts carboxylate into an O-acylisourea intermediate that is reactive toward amine groups and results in forming a amide bond. However, this intermediate is not stable and tends to hydrolyze in water. Therefore N-hydroxysuccinimide (NHS), or its more water-soluble derivative N-hydroxysulfosuccinimide (sulfo-NHS), is usually employed in combination with EDC, for resulting in much more stable NHS–ester intermediates while retaining coupling capacity in water (Fig. 3.1). A very common strategy is to preactivate the carboxyl with NHS and then store the reagent in a solid state or in anhydrous solution. Linking can be achieved later by simply incubating it with the desired aminated species. Another example of zerolength crosslinking is the conjugation between aldehyde and amine groups. Aldehydes can form Schiff bases with amines, although these bonds are labile and easily decompose. Elevating the environmental pH can enhance but cannot completely stabilize these bonds; however, the addition of sodium cyanoborohydride can reduce the intermediate species to a secondary amine, and result in a stable linkage (Fig. 3.1). “Click chemistry” has arrived as a recent alternative to these strategies and has been widely applied in forming biomolecule–nanoparticle couplings, due to its dependability, speed, and efficiency. The most common reaction within the click chemistry paradigm, the Azide–Alkyne Huisgen Cycloaddition, is a 1,3-dipolar cycloaddition between azide and alkyne groups, resulting in a stable 1,2,3-triazole product (Fig. 3.1). This reaction is typically catalyzed by Cu(I) or, in some cases, Cu(II) and a reducing agent (e.g., sodium ascorbate). The reaction is fast, spontaneous, and irreversible, with perfect transformation

1. N

COOH

C

N

+

– Cl NH

O 2.

O

O S

– O

N HO O sulfo-NHS

EDC

O C N O O

O S O – O

H2N

C N H O

O H NH2

NaCNBH3

N

N3 N Cu(I) or Cu(II) & sodium ascorbate

N

N

FIGURE 3.1 Examples of zero-length crosslinking in bioconjugation.

HN

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yield. However, since most biomolecules and particle coating materials contain no azide and alkyne groups, both elements need to be preactivated, which can be a nontrivial task and has been the main disadvantage of click chemistry-based bioconjugation. Nonetheless, several reports describing click chemistry-mediated particle conjugation have been documented [8–10]. 3.2.2 Conjugation with Crosslinkers The natural chemical compositions of nanoparticles and biomolecules are not necessarily sufficient for direct conjugation; in these cases, a crosslinker intermediate can be used. Crosslinkers are composed of a spacer with two terminal chemical groups that bind complementary groups on the nanoparticle and the biomolecule to be conjugated. One example of such a crosslinker is BS3 (bis(sulfosuccinimidyl) suberate) with sulfo-NHS groups at each end, which functions to link two amine-possessing species (Fig. 3.2). When two complementary groups are different, a heterobifunctional crosslinker can be used. Figure 3.2 gives examples, where amine-containing particles and thiol-containing molecules (and vice versa) are coupled by applying a crosslinker with NHS ester and maleimide on opposing ends (such as sulfo-succinimidyl-(4-N-maleimidomethyl) cyclohexane-1-carboxylate (sulfo-SMCC) and 4-maleimidobutyric acid N-hydroxysuccinimide ester) [11, 12]. Notably, since forming conjugations with homobifunctional crosslinkers may inevitably cause crosslinking among the same species, it is very often necessary to preconvert the active group of one of the to-be-coupled moieties and thereafter use heterobifunctional crosslinkers to achieve linking. For instance, instead of using BS3 to achieve coupling between two amine-containing species, it is advisable to thiolate one component and modulate the coupling with sulfo-SMCC. 3.2.3 Streptavidin–Biotin Interaction Besides using chemical coupling methods, specific biological interactions can also be utilized to form particle–biomolecule conjugates, the most popular of which is the streptavidin–biotin interaction. The streptavidin–biotin interaction is one of the strongest known non–covalent bonds, and is stable in a wide variety of conditions. Because streptavidin has four biotin binding sites, it can work as a bridge to link two or more biotinylated species (Fig. 3.3). As compared with chemical reactions, the streptavidin–biotin interaction O O NaO S O

N

O O

O

N

O

O O

O BS3

O S O ONa H N

NH

O

O O

O O

N

H N

O S O ONa

O O

O O

O S OH O

O O

HN

O O

O

N O

N O NH

sulfo-SMCC

HS

H N O O

O N

H N O O

O

FIGURE 3.2 Examples of crosslinker-mediated bioconjugation.

O N O

S

N H

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Streptavidin

FIGURE 3.3 Forming conjugation via streptavidin–biotin interaction.

is highly specific and less likely to alter biomolecules in a way that might result in a functional loss. Despite the considerable differences in primary structure, avidin, which was found in egg whites, is able to bind to biotin with a similar affinity as that of streptavidin. One major reason that makes it less applied in bioconjugation, however, is the severe nonspecific binding issue of avidin. With a highly basic pI of 10 (as compared to that of 5–6 for streptavidin), avidin can interact with molecules via ionic interaction; moreover, as a glycoprotein, it has the potential concern of binding to carbohydrate receptors that are expressed on many types of cell surfaces. Addressing that, a deglycosylated version of avidin, NeutrAvidin, was later developed and employed. By removing a carbonhydrate section, not only is the lectin binding risk of avidin avoided, but the protein pI also has been decreased to a nearly neutral level (pI 6.3), making it a popular alternative to streptavidin. In this section, we briefly introduce the most common techniques for nanoparticle– biomolecule conjugation. For each nanoparticle–biomolecule pair there are usually several options available to achieve conjugation; the choice of which specific method to use is based on the expense, effectiveness, particle stability, and bioactivity loss associated with each approach. Next, we introduce the most common methods used for nanoparticle surface modification. Specific considerations for surface modification of several types of nanoparticles, including quantum dots, iron oxide nanoparticles, gold nanoparticles, carbon nanotubes, silica particles, and nanocomposites, are discussed.

3.3 QUANTUM DOTS Quantum dots (QDs) refer to a large, varied class of nanostructures made from semiconductor materials, which, at nanoscale, demonstrate unique size-dependent fluorescent emissions. Much brighter and more photostable than traditional organic dyes, QDs are

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expected to serve as the next generation optical tags for molecular imaging. Accordingly, intensive studies have been conducted on QDs, from both the synthetic and surface modification perspectives [13]. Commonly used QDs have cores made of materials like CdSe, CdTe, PbS, and InAs, with tunable emissions spanning most of the visible and near-infrared regions. Although there are reports of aqueous phase-based QD synthesis, most of the QDs utilized for bioapplications are made through pyrolysis in organic solution, which generally produces higher quality crystals. Described briefly, this synthesis is performed by heating an appropriate organometallic precursor in a high boiling point organic solvent in the presence of certain surfactants (such as trioctylphosphine (TOP) and trioctylphosphine oxide (TOPO)), which passivate the particle surface and limit their growth. Typically, a coating step is subsequently applied, in which a layer of semiconductor material with a wider bandgap than the core, such as ZnS, is introduced to the particle surface. Such a coating is essential for increasing the particles’ photoluminescent quantum efficiencies and reducing the particle toxicity by suppressing the dissolution of cadmium ions [14–16]. As synthesized, the QDs are intrinsically water insoluble due to the hydrophobic TOPO/TOP coating. In order to convert the particles to a water-soluble form and to allow subsequent conjugation, an extra surface modification step is needed. Numerous approaches have been developed for this purpose, but despite their variability, most of them can generally be divided into two categories: ligand exchange and ligand addition. The former approach refers to the method of entirely replacing the original organic surface coating, while the latter refers to introducing a layer of new materials on top of the existing coating to render the particles hydrophilic. 3.3.1 Ligand Exchange For ligand exchange to be successful, the new hydrophilic ligand needs to have comparable or higher affinity toward the particle surface than the original surfactant, so that when incubated in high concentration, the nanoparticles’ original coating will be substituted by the new ligand. Thanks to the ZnS protection layer, thiolated molecues can easily be introduced to the QD suface, which form dative thiol bonds with the exterior S atoms. A host of small biofunctional molecules, including mercaptoacetic acid, mercaptopropionic acid, mercaptosuccinic acid, dithiothreitol, glutathione, histidine, DTT, and cysteine are capable of rendering QDs water-soluble [14] (Fig. 3.4). For example, Bawendi and co-workers used cysteine to coat CdSe/ZnS nanoparticles, resulting in water-soluble conjugates with a hydrodynamic diameter less than 5.5 nm [17]. Owing to their zwitterionic nature and ultrasmall size, such particles have proved stable in blood circulation and are inhibited to serum adsorption. When intravenously injected, those particles were found to be rapidly excreted via renal clearance, which is a size-exclusive excretion route with a size-cut around 7 nm. More sophisticated molecules, such as thiolated DNA oligos [18], peptides [19], or even proteins [20], can be anchored onto QDs in the same manner (Fig. 3.4). For example, thiolated biotinylated bovine serum albumin (bBSA), which is developed by treating bBSA with 2-iminothiolane, can be linked to QD surfaces, rendering them water-soluble. However, there are concerns about the longevity of dative thiol-based surface modifications in in vivo environments [14], considering the fragility of the disulfide bond between R-SH and ZnS. To protect against this, a crosslinking step can be used to reduce the risk of dissolution of the protective cap. Peng and co-workers anchored thiolated dendrimers to

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DNA oligos

Proteins S H N

S S

S

N Poly-His tagged proteins/peptides

O O HN

Zn

NH

QD

N N H

S

S

S

S

R

Polydendate ligands S S R Thiolated compounds

FIGURE 3.4 Commonly used molecues for QD surface modification.

the surface of QDs [21], producing multiple hydroxyl groups. Those hydroxyl groups were further crosslinked with second-generation dendrons, which improved particle stability while providing multiple amines for use as binding sites [22]. Alternatively, polydentate ligands, such as DHLA [23], oligomeric phosphines [24], or peptides with multiple cysteinerich domains [19], are also widedly utilized, which are expected to have higher affinity to the particle surface than monothiolated species (Fig. 3.4). Although it is most common to target S from the ZnS shell, it is also possible to utilize Zn for surface modification. For instance, proteins or peptides containing polyhistidine tag have been found to easily anchor onto QDs through Zn(II)–His interaction [25, 26] (Fig. 3.4), with a binding strength that is much stronger than most antibody–antigen interactions [16]. 3.3.2 Ligand Addition Ligand addition is mostly performed by encapsulating the original particles with a layer of amphiphilic ligand. When mixing in water, the new ligand self-assembles onto the particle surface by having its hydrophobic section intercalating into the TPOP/TOP layer, while leaving its hydrophilic tail protruding into water. Compared with those modified by ligand exchange, QDs rendered through ligand addition are generally more resistant to agglomeration that may otherwise occur during conjugation, long-term storage, or exposure to complex physiologic environments. This added stability, however, comes with a cost. The ligand addition usually leads to a significant hydrodynamic size increase, which could potentially compromise the particle’s in vivo mobility and penetrability. Nevertheless, a variety of amphiphilic compounds have been utilized for this purpose, including phospholipids [27, 28], amphiphilic saccharides [29], acrylic acid polymers [30, 31], and others. For example, Ballou [32] utilized polyacrylic acid and its PEGylated derivatives to coat QDs; the resulting adducts showed potential as sentinel lymph node imaging agents [32]. In another effort, Nie and colleagues utilized a triblock copolymer (consisting of a polybutylacrylate segment, a polyethylacrylate segment, a polymethacrylic acid segment,

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(a)

(b)

(c)

FIGURE 3.5 (a) Schematic illustration of triblock copolymer coated QDs. (b) The chemical composition of the triblock copolymer (c) The postulated mechanism of the QD tumor retention, which is believed to involve both passive targeting (left) and active targeting (right). (Adapted from Gao et al. [33] with permission.)

and a hydrophobic hydrocarbon side chain) for QD modification (Fig. 3.5). Such polymercoated QDs were conjugated with a prostate-specific membrane antigen (PSMA) targeting antibody through EDC/NHS chemistry and were administrated into prostate cancer bearing mice [33]. Fluorescence optical imaging confirmed successful particle accumulation in the tumor area, attributed to both the EPR effect and specific antibody–antigen interaction.

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3.3.3 Protein Encapsulation Proteins can be utilized as biomaterials to be adsorbed onto QDs to render the particle watersoluble. For instance, dihydrolipoic acid-modified QDs are highly negatively charged but are not sufficiently stable in water. In order to improve particle solubility, a layer of avidin, which is positively charged at neutral pH, can be adsorbed onto the particles directly without performing chemical conjugations. The resultant particles are highly stable in water and can easily be linked with other biomolecules through biotin–avidin interaction [34]. Another example of protein-based surface modification is the use of GroEL. GroEL is a cylinder-like chaperonin protein, which has a hydrophobic cavity with an inner diameter of about 4.5 nm. The as-synthesized QDs, being hydrophobic, can be accommodated in the GroEL cavity, resulting in a protein–QD hybrid. An interesting feature of this hybrid is that the release of QD particles can be governed by the addition of ATP, which has potential implications for biosensor applications [35]. Expected to be the next-generation fluorescent labels for bioapplications, the biocompatibility of quantum dots has been a great concern. Previously utilized QD formulas are mainly constituted of CdSe, CdTe, and PbS, which contain extremely toxic metals such as Cd and Pb, and therefore have limited prospects in clinics. Addressing such issue, recently, several non-Cd-based QDs have successfully been developed by Peng and several other groups. Especially, InAs/InP/ZnSe core/shell/shell QDs have been synthesized by a one-pot approach and were converted to be water soluble with mercaptopropionic acid (MPA) [36]. The success of commercially available options, like Qdots from Invitrogen Inc., has removed some of the impetus for developing improved coating strategies, since much research is now conducted with these particles. Nevertheless, due to the large hydrodynamic size (50 nm) and unclear surface chemical composition of such particles, individual surface modification is still preferred in many cases.

3.4 IRON OXIDE NANOPARTICLES Iron oxide nanoparticles (IONPs) are well known for their superior magnetic properties. This, together with their native biocompatibility and inexpensiveness, make IONPs important actors in molecular imaging, specifically as contrast agents for MRI [37–42]. As compared with other particles, the additional magnetic interaction among IONPs makes the stabilization and functionalization of them especially challenging. 3.4.1 Surface Modification of IONPs Synthesized from Aqueous Solution IONPs are conventionally made in aqueous solution by coprecipitating Fe(II) and Fe(III) precursors in the presence of a base, such as ammonium hydroxide and sodium hydroxide [39, 43, 44]. Due to their high surface energy and magnetic interactions, bare IONPs tend to form aggregations, so additives, in most cases hydrophilic polymers, are usually applied in the synthesis, which helps modulate particle growth and provide electrostatic and steric forces to counteract the IONP attractive forces. Conventionally applied polymers include polyvinylpyrrolidone (PVP), dendrimer, polyaniline, and dextran [39, 45, 46]. Among them, Feridex particles (AMAG Pharmaceuticals), which are essentially dextran-coated IONPs, are an FDA approved agent for use in MR imaging for the detection of lesions

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in liver and spleen. With an overall size above 50 nm, Feridex particles are dominantly removed from circulation through the reticuloendothelial system (RES) uptake and are selectively concentrated in liver and spleen [47, 48]. Combidex particles, another AMAG product currently in clinical trials, have a similar iron oxide core but a shorter dextran for coating. With a small hydrodynamic size, Combidex particles possess a longer circulation time and aim for lymph node imaging [49]. Also, Schering AG is promoting Resovist as a liver-specific contrast agent in MRI, which basically is carbon carboxydextran-coated iron oxide nanoparticles [50, 51]. Biomolecule immobilization on the polymer-coated IONPs is achieved by making use of the functional groups on the surface polymers. For instance, dendrimer-coated IONPs, having terminal amine groups, can form conjugation with bovine serum albumin using glutaraldehyde as the homobifunctional crosslinkers [52]. In cases when more than one kind of chemical group is available, it is possible to load multiple functional entities onto the particles. One such example is polyaspartic acid (PASP)-coated nanoparticles, which have both carboxyl and amine groups available on the particle surfaces. Lee and co-workers covalently linked both a 64 Cu-DOTA chelate and the ␣v ␤3 integrin targeting peptide c(RGDyK) onto the IONPs. The chelate was covalently bound to the carboxyl group through EDC/NHS mediated chemistry, while the peptide, in order to avoid the self-crosslinking issue, was prethiolated and subsequently coupled with the amines from PASP-IONPs using 4-maleimidobutyric acid N-hydroxysuccinimide ester as the bridge linker [53]. Notably, this was one of the first examples of an integrated MRI/PET probe with specific targeting capacity, which is significant given the ongoing clinical efforts of building integrated MRI/PET detectors. For easy conjugation, sometimes the nanoparticle coatings need to be pretreated before linking with functional biomolecules. For instance, bioconjugation with dextran-coated IONPs is usually preceded by treatment with epichlorohydrin and ammonia, which converts the dextran hydroxyls to amines. The resulting crosslinked iron oxide (CLIO) nanoparticles have been a very effective platform and have been integrated with various kinds of biomolecules, including transferrin, annexin V, anti-VCAM monoclonal antibody (mAb), anti-E-Selectin mAb, oligonucleotides, and TAT peptides [54–58]. While surface modifications are usually introduced during particle synthesis, stabilizing agents can be grafted onto IONPs postsynthesis as well. For example, Zhang and co-workers grafted PEGylated silanes onto the IONPs to generate stable conjugates. The silane moiety can anchor onto the IONP surface by forming Si–O–Fe bonds with the surface Fe atoms, and the long PEG chains help to avoid particle aggregation. Further conjugation can be derived from the amine groups on the ends of the PEG chain (Fig. 3.6) [59–61]. Similarly, Jon and co-workers utilized a PEG-silane copolymer poly(TMSMA-r-PEGMA) to coat IONPs. Multiple silanes make the polymer adhere strongly to the particle surface, and the resultant conjugates are very stable in vivo, with long circulation half-lives [62, 63].

3.4.2 Surface Modification of IONPs Synthesized from Pyrolysis Although widely utilized, there are several drawbacks for the IONPs produced in waterbased synthesis, such as poor crystallinity, suboptimal size and shape control, and unfavorable magnetization [64, 65]. These concerns have facilitated the progress of pyrolysis-based IONP synthesis, as commonly used in the QD synthesis described previously. Fe(acac)3 , Fe(CO)5 , and, more recently, Fe-oleateare are the most utilized of the iron precursors, which

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FIGURE 3.6 Using PEGylated silane to modify and stabilize IONPs. The resulting particles can be covalently linked with functional entities, such as methotrexate (MTX) in this case. (Adapted from Kohler et al. [60] with permission.)

decompose, nucleate, and grow to desired size in high boiling solvent with the presence of oleic acid/oleylaimne [66, 67].

Ligand Addition As is the case with TOPO/TPO coating for QD passivation, oleic acid/oleylamine forms a hydrophobic layer on the IONP surface during the pyrolysis-based synthesis. The particles can be made water soluble with some single-chained surfactants, such as dioctyl sodium sulfosuccinate, but due to the low stability, surfactants with multiple hydrophobic tails are preferred [68]. Phospholipid analogs, particularly their PEGylated derivatives, are widely used as effective phase transfer ligands. A variety of phospholipids with a wide selection of functional terminal groups are now commercially available from outside vendors, like Avantilipids Inc. Hutman and co-workers used maleimide-terminated PEG 2000 phospholipids to coat IONPs, which were subsequently conjugated with RT1 anti-MHC II antibodies. These conjugates showed targeting specificity for renal medulla of the rat, a section of the kidney in which MHC Class II is specifically expressed [69]. Larger polymers possessing multiple hydrocarbon chains have also been reported for IONP modification. In one such example, poly(maleic anhydride alt-1-tetradecene) was first coated onto IONPs with the multiple alkyl chains intercalating into the oleic acid/oleylamine coating layer (Fig. 3.7). Next, bis(6-aminohexyl) amine was added, which induced crosslinking among the polymer anhydride groups. In an aqueous environment, the unreacted anhydride

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FIGURE 3.7 Poly(maleic anhydride alt-1-tetradecene)-based nanoparticle surface modification. After coating and crosslinking, biomolecules can be linked to the particles by forming amide bonds with the surface terminal carboxyls. (Adapted from Pellegrino et al. [70] with permission.)

groups were subjected to hydrolysis and were converted to carboxyls, ensuring the particle water solubility while creating new bonding sites for future bioconjugation [70].

Ligand Exchange While many of the previous ligand addition strategies are interchangeable between QDs and IONPs, ligand exchange approaches are generally exclusive. For most IONPs, it is the iron–carboxyl interaction that needs to be overcome by the newly added ligands. It is possible to achieve this by simply utilizing highly concentrated carboxyl-terminated compounds [71], but the efficacy of this approach is unsurprisingly low so that ligands with higher surface affinity are still highly desirable. Recent research has demonstrated that bidentate compounds, which can bind the surface Fe atom in chelating fashion, are far more effective as IONP surface-exchange ligands. One such example is meso-2,3-dimercaptosuccinic acid (DMSA), which is a small compound with two carboxyls at adjacent positions [12, 65]. By incubating it with IONPs overnight, such a ligand

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Fe3O4

COOH

Fe3O4

COOH

O

COOH

X

O

O

OH

X

n

O

Oleylamine & oleic acid OH

OH

NH

FIGURE 3.8 Surface modification of IONPs with PEGylated dopamine. Further function docking can be done with the terminal carboxyl. (Adapted from Xie et al. [73] with permission.)

can bind to the particle surface, while leaving the thiols exposed on the outer surface. A portion of these surface thiols can form disulfide bonds among each other, which helps to stabilize the particle coating. The remaining unreacted thiols, on the other hand, can be utilized for further bioconjugation purposes. It is worth noting that such chelation is formed with two carboxylates binding to one Fe atom, which is different from the case of using dithiol ligands for QD modification, where thiols bind with S from the ZnS coating. Similarly, dopamine and its derivatives, which contain two adjacent hydroxyl groups from the catechol moiety, can also serve to form a chelate complex with the Fe on the particle surface [72]. Particle stability can then be further established by PEGylating the dopamine conjugates with PEG diacids. The resulting particle conjugates are very stable in aqueous solution and have terminal carboxyl groups that can be used for further functionalization (Fig. 3.8) [73].

3.4.3 Polymer and Protein Adsorption Multidendate ligands such as PEI have also been applied as IONP capping ligands. Containing multiple primary, secondary, and tertiary amines, PEI has strong affinity toward the IONP surface and can easily form conjugates with IONPs. The resulting adducts are highly positively charged in aqueous solution and have been exploited as cell labeling agents [74] and as transfection agents [46, 75]. Compared with PEI alone, these PEI–IONP conjugates may have enhanced cellular permeability, if used in conjunction with a vertically oriented magnetic field [76]. Being highly positively charged, the PEI coating may also work as the foundation layer to induce a layer-by-layer coating, where materials

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FIGURE 3.9 BSA-coated manganese ferrite nanoparticles. The surface amines from BSA were modified with SPDP, followed by coupling with HS-PEG-RGD and HS-siRNA-Cy5. The resulting nanoconjugates proved to have both MR imaging and gene regulation capacities. (Adapted from Lee et al. [78] permission.)

with alternating charge natures are deposited onto the particles in succession. For instance, a copolymer, poly(ethylene oxide)-block-poly(glutamic acid) (PEO-PGA), was adsorbed onto the PEI–IONPs through the electrostatic interaction between PEI and PGA [77]. More recently, Cheon and his colleagues used bovine serum albumin (BSA) to coat manganese-doped IONPs, which yielded stable nanoconjugates. With N-succinimidyl-3(2-pyridyldithio) propionate (SPDP) as the crosslinker, they successfully coupled thiolated RGD and thiol-siRNA onto the nanoparticles, resulting in an all-in-one target-specific nanoplatform with both MR imaging and siRNA delivery capacities (Fig. 3.9) [78].

3.5 GOLD NANOPARTICLES Gold nanoparticles have the potential to become effective probes in an array of imaging modalities, such as CT, photoacoustic, and surface-enhanced Raman spectroscopy (SERS). From the synthesis perspective, modulating the shape of gold nanoparticles is well established, with those in the form of spheres [2, 79, 80], cubes [81], rods [82, 83], cages [81, 84], and wires [85] already developed. Such shape control is important for manipulating the nanostructure’s spectral absorption profile, which affects their roles as imaging probes. Common 10-nm spherical gold particles have characteristic surface plasmonic absorption at around 520 nm. Tuning the particle size may alter the spectra, but to a limited degree. For example, the maximum absorption of gold nanoparticles with diameters of 48.3 and 99.3 nm are 533 nm and 575 nm, respectively [86], both in a spectral region where light has limited tissue penetrating depth. Changing the nanoparticle configuration to a nanorod structure, on the other hand, can dramatically shift the absorption spectrum to the nearinfrared region (650–900 nm), which potentiates their role as probes in photoacoustics and mediators in photothermal nanotherapeutics (Fig. 3.10) [81, 82, 84, 87]. Due to the strong interaction between thiol and gold, the surface modification of gold nanostructures is overwhelmingly conducted with thiolated species. Typically, a bifunctional thiolated compound is utilized, such as carboxyl-ended alkanethiol with various lengths of hydrocarbon chains [79]. When applied, the thiol group will anchor onto the particle surface, and the terminal carboxyl group can be utilized for forming conjugates. Alternatively, biomolecules can be prethiolated, after which they can be directly loaded onto the gold particle surface [88–90]. Thiolated DNA oligos have consistently been utilized for

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A

B

C

D

500

500

E

100

100

100

1.2 1

Abs.

0.8

e

0.6

b

c

d

0.4 0.2

Seed

a

a

b

c

d

e

Increasing Aspect Ratio

0 400

500

600 700 Wavelength (nm)

800

900

FIGURE 3.10 Transmission electron micrographs (top), optical spectra (bottom left), and photographs (bottom right) of gold nanorods of various aspect ratios. (Adapted from Murphy et al. [87] with permission.)

Au nanoparticle functionalization, and the resultant particle conjugates have been found to be useful in many aspects of biodiagnostics [91]. Unlike QDs, which may prefer multidendate compounds for the sake of stable binding, in most scenarios, a monothiol compound is sufficient to result in stable gold nanoconjugates. In fact, monothiol compounds can be advantageous over multidentatethiol compounds for affording better loading capacity. In one such demonstration, two oligonucleotides were loaded onto 13-nm gold particles: one was tetrathiol modified while the other was monothiol modified. The two oligos share the same sequence, which aims at inhibiting EGFR expression. It was found that, for each 13-nm Au nanoparticle, 45–50 tetrathiol strands can be loaded, whereas the number for the monothiol oligo is 110–120. In vitro experiments showed that, while both particles were able to penetrate cell membrane and suppress the EGFR expression, the monothiol group was 35 times more efficient in inducing knockdown, presumably due to an improved oligo loading efficacy [92]. Tuning the anisotropy of Au nanostructures, aside from altering the particle absorption spectrum, also changes the particle’s surface chemistry. To date, most gold nanorod synthesis is conducted in water with cetrimonium bromide (CTAB) as the capping ligand, as compared to citric acid or alkyl acid/alkyl amine, which are commonly used in spherical gold nanoparticle production. The CTAB forms a thick double-layer coating over the growing gold rods during synthesis, which constrains the particle radial growth and makes the particle sidewall hard to access. The ends of the rods are relatively less protected and can serve as the binding sites for thiolated compounds. Figure 3.11 is one good demonstration of this model: CTAB-coated gold nanorods were first modified with thioctic acid and were subsequently conjugated with anti-mouse IgG through EDC-mediated coupling. Afterwards, mouse IgG, with two binding sites, was added, which was recognized by the anti-mouse IgG and resulted in crosslinking among the nanorods. The fact that the particle aggregation was formed in a head-to-tail configuration confirms that only at the

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FIGURE 3.11 Head-to-tail assembly of Au nanorods, which is caused by the interaction between IgG and the anti-mouse IgG that was immobilized at the ends of the Au nanorods. (Adapted from Chang et al. [94] with permission.)

ends of the nanorods did the ligand exchange and the subsequent antibody immobilization occur [93]. Besides ligand exchange, gold nanorod surface modification can also be done through a layer-by-layer deposition approach. This takes advantage of the highly positively charged CTAB coating, which can absorb one layer of negatively charged species, thereafter making it ready for another round of absorption but with positively species, and so on. For example, poly(sodium-4-styrenesulfonate) PSS(−) and poly(diallyldimethyl-ammoniumchloride) PDADMAC(+) were alternatively applied to generate a multilayer coating on the gold nanorod surface [95, 96]. One good feature of this approach is that functional molecules such as antibodies can easily be immobilized onto the outer layer via electrostatic interaction rather than through chemical conjugation [97]. Recently, a green process was developed by Katti and colleagues to prepare gold nanoparticles. In brief, without applying conventional chemicals, soybean extract was utilized in the particle preparation. Multiple components of the extract are believed to play a role in particle formation, either as reductants (sucrose, stachyose, glycinin, ␤-conglycinin, trypsin inhibitors, soybean vacuolar protein p34) or stabilizers (raffinose, isoflavones, daidzein, glycitin, and various proteins), or both. The resulting gold nanoparticles are nontoxic and highly stable and can be coupled with other species by taking advantage of the multiple amines from the surface proteins [98]. Using a similar strategy, the same group also utilized gum arabic, a plant-derived biocompatible protein, to synthesize gold nanoparticles [99, 100].

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3.6 CARBON NANOTUBES Unlike the aforementioned particles, carbon nanotubes (CNTs) are not yielded from wet chemistry routes; therefore no ligands passivate the surface. However, the missing of a surface coating does not necessarily facilitate the surface modification of CNTs. Rather, the hydrophobic nature, the propensity to form bundles, and the poor solubility even in organic solvents make the modification and bioconjugation of CNTs a nontrival one. 3.6.1 Direct Surface Modification The intact CNTs have sidewalls structurally resembling graphite, which is inert and is unfavorable for chemical modifications. In order to generate binding sites, some extreme oxidative conditions are applied to the CNTs, which rusults in defects and generates carboxyl groups on the CNT surfaces [101]. For example, after treating single wall nanotubes (SWNTs) in refluxing 2.5 M HNO3 for two 36-h periods with an interval of 30-minute sonication, the SWNTs can be made water-soluble. At pH 7, these modified CNTs showed a surface potential of −75 mV, attributable to multiple surface carboxyl groups [102]. Alternatively, cooking CNTs with succinic or glutaric acid acyl peroxide at 80–90 ◦ C leads to the immobilization of 2-carboxyethyl or 3-carboxypropyl on the nanotube surface [103]. Subsequently, conjugation chemistry can be applied to covalently link molecules such as fluorescein, biotin, sugar moieties, oligonucleotides, and proteins onto the nanotubes [102–109]. Besides strong acid treatment, another common approach of CNT modification is via 1,3-dipolar cycloaddition with azomethine ylide and its derivatives, which anchor onto the CNT surface by forming a pyrrolidine ring. Reported functional molecules that have been conjugated to CNT via this means include organic dyes, peptides, and even antibiotics [110]. The two surface modification strategies can be applied simultaneously, and by doing so, more than one type of binding site can be generated on the CNT surface, permitting the loading of multiple entities. In one such demonstration, amphotericin B, a commonly used antifungal agent, and an organic dye, FITC, were both covalently linked to the CNTs. The former was coupled to the end of the aminotriethylene glycol chain that was preanchored onto the CNTs through a 1,3-dipolar cycloaddition, while the later was coupled to the short PEG chain that was derived from the carboxyl groups generated from acid treatment (Fig. 3.12) [111]. 3.6.2 Ligand Addition Due to the hydrophobic nature of the surface, CNTs can be modified as well with amphiphilic compounds. Many types of surfactants have been documented for doing this job, including sodium dodecyl sulfate (SDS), triton-X-100, CTAB, sodium dodecylbenzene sulfonate (SDBS), and sodium dodecane sulphonic acid (SDSA) [101, 113–115]. With sonication, these surfactants can have their hydrophobic hydrocarbon chain interact with the graphite surface and achieve an effective CNT dispersion. One interesting finding is that SDBS is much more efficient at making CNT suspension than its analog SDS. It could be due to the longer hydrocarbon chain of SDBS compared to SDS. But more importantly, it is attributed to the possession of an aromatic ring in SDBS, which allows for the additional ␲-␲ interaction with the CNT surface that further stabilizes the nanostructure. It

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FIGURE 3.12 Both amphotericin B and FITC were covalently immobilized onto CNTs but via different routes. (Adapted from Wu et al. [112] with permission.)

was supported by a separate study, in which trimethyl-(2-oxo-2-phenylethyl)-ammonium bromide, trimethyl-(2-naphthalene-2-yl-2-oxo-ethyl)-ammonium bromide, trimethyl-(2oxo-2-phenan-threne-9-yl-ethyl)-ammonium bromide, and trimethyl-(2-oxo-2-pyrene-1yl-ethyl)-ammonium bromide were utilized to solubilize CNTs. Although the later two compounds are suboptimal in forming micelles themselves, they were found far more effective in achieving CNT water-suspension than the former two, suggesting that their association with CNTs is mainly a ␲-␲ interaction. To achieve even stronger ligand–CNT interaction, polycyclic aromatics, such as 1pyrenebutanoic acid succinimidyl ester, have been applied to modify CNTs [116]. Also, a diblock polymer, PS-b-PAA, which has multiple aromatic styrene groups, was immobilized on the CNTs. After crosslinking with diamine linkers, the resulting CNTs were reported to have good solubility in water (> 0.5 mg/L) [117]. The layer-by-layer deposition can be applied on the foundation laid by the aromatic compounds. For instance, the CNTs could be rendered positively charged first with 1-pyrenepropylamine hydrochloride, followed by

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FIGURE 3.13 One example of using a layer-by-layer approach to modify CNTs. (Adapted from Artyukhin et al. [119] with permission.)

alternative depositions of negatively charged polystyrene sulfonate (PSS) and positively charged poly (diallyldimethylammonium chloride) (PDDA) to achieve water-solubility (Fig. 3.13) [118]. Among the many ligands that are suitable for ligand addition, PEGlyated phospholipids might be one of the best options, given their effectiveness, ease of functionalization, and biocompatibility. In a recent study conducted by Gambhir and co-workers, a group of mice were intravenously administrated phospholipid–CNT conjugates and were carefully monitored for 4 months [120]. Even though some CNTs were found trapped in the liver and spleen without degradation, evidence of toxicity was found neither from the survival, clinical, and laboratory parameters nor from the following necropsy and tissue histology studies. This result, while confirming the biosafety of phospholipid–CNT hybrids, suggests that the toxicity of CNTs that was previously observed was likely due to the coating materials, rather than the CNTs themselves [121]. The Dai group has led efforts to utilize phospholipid–CNT conjugates for several applications, including tumor imaging, drug delivery, and gene delivery [122–126]. Specifically, phospholipid-branched-PEG was used to modify CNTs and to load paclitexal (PTX); the resultant conjugates showed improved therapeutic effects compared to PTX alone [125]. 3.6.3 Direct Adsorption Biomolecules can also be loaded onto CNTs through direct physical interaction. For example, single strand DNA (ssDNA) can be adsorbed spontaneously onto SWNT surfaces, with its sugar–phosphate backbones exposed to the aqueous environment to render water solubility of the CNTs. The driving force underlying the conjugate formation is attributed to the ␲-␲ stacking between the bases and the graphite surface. Due to the electrostatic and torsional interaction within the sugar–phosphate backbones, the adsorbed DNAs

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encompass the CNT in a helical manner [127–129]. Since chemical manipulation and polymerization of DNA is well understood, DNA–CNT hybrids can easily be modified for function addition. Erie and co-workers utilized thiolated DNA to modify CNTs, followed by the addition of CdSe/ZnS QDs. Atomic force microscopy (AFM) confirmed the formation of a CNT–DNA–QD hybrid and showed a distance of 14 nm between the adjacent QDs on the CNTs, implying an end-to-end arrangement of the DNAs on the CNTs [130]. Proteins can also be adsorbed directly onto CNTs. For instance, streptavidin was found to be easily adsorbed onto the CNT surface in a well-ordered self-assembly [131]. Although the mechanism is not yet fully understood, it is believed to involve a baseline hydrophobic interaction and aromatic stacking and may also result from the amino–carbon interaction between proteins and CTNs [132]. The adsorbed streptavidin retains its functionality and still actively binds to biotin and its derivatives. Double stranded DNA fragments, when biotinylated, can be immobilized onto such a streptavidin–CNT hybrid through biotin–streptavidin interaction. It is worth mentioning that both the DNA–CNT adducts and the protein–CNT adducts are highly stable and can sustain dialysis without forming aggregations, which is rarely observed for other CNT formulas [3].

3.7 SILICA PARTICLES Silica nanoparticles are another type of biomaterial that has been extensively studied. Interesting features of silica particles include (1) their biocompatibility, (2) the sophisticated size control (from 20 nm to several ␮m) [133], (3) good water-solubility, (4) mesoporosity [134], and (5) the ready incorporation of imaging modalities into the matrices. The main precursor of the silica particle is tetraethyl orthosilicate (TEOS), which hydrolyzes and coagulates during synthesis to form the main body of the particles. However, lacking conjugation-friendly chemical groups on the surface, the silica particles are not readily functionalizable. To facilitate bioconjugation, aminopropyltrimethoxysilane (APS) and mercaptopropylmethoxysilane (MPS) are usually co-coagulated with TEOS, which introduces amines or thiols to the particle surface. If some functional molecules are precoupled with APS/MPS and co-coagulated with TEOS, then the functions can be imported into the silica particles. For example, organic dyes and Gd–DTPA complex were preconjugated with APS and mixed with TEOS as particle precursors. The resulting silica particles may possess thousands of dyes or Gd chelates in the matrices and are optically or magnetically active [135–137]. Aside from accommodating small molecules, various nanoparticles can also be incorporated into silica matrices as well. In other words, silica coagulation can serve as a coating strategy, which imports water solubility and facile functionalizability to the particles. Reports on employing silica to coat IONPs, gold nanoparticles, and QDs have been well documented [133, 138–142]. A great virtue of utilizing such a coating strategy is that it allows the encapsulation of more than one component in a fast and even quantitative way [143–148]. Nie and co-workers utilized silica to encapsulate both IONPs and QDs, creating a hybrid that retained both magnetic and optical properties [146]. Koole et al. [149] reported a dual-function core–shell–shell (CSS) nanoparticle, which were basically silica nanoparticles loaded with QDs and Gd complexes. While the QDs were encapsulated within the silica matrices via the above-mentioned method, the Gd chelates were loaded on the surface of the particles, which was achieved by making the particle surface hydrophobic

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FIGURE 3.14 Preparation of multifunctional silica nanoparticles embedded with Gd chelates and QDs. (Adapted from Koole et al. [149] with permission.)

with otcadecanol followed by the adsorption of a hybrid phospholipid layer constituted of PEG-DSPE, Gd-DTPA-DSA, and Mal-PEG-DSPE. (Fig. 3.14) [149].

3.8 NANOCOMPOSITES A current effort in molecular imaging is the development of complex systems with more than one imaging entity, which encourages multimodality imaging. One way to achieve this goal is to encapsulate different nanostructures into one unit, as is the case when utilizing a silica coating. Alternatively, integration can be achieved at the chemical synthesis stage, which is advantageous in affording better ratiometric control and size limitations. Many efforts have recently been dedicated to the synthesis of composite nanoparticles, which resulted in a burst of reports on novel heterodimer particles [150]. Such particles typically possess two side-by-side substructures and therefore two different surfaces. In some cases, the two surfaces can be modified with the same ligand. For example, FePt–Au nanoparticles can be modified with PEGylated DHLA, which can bind with both Pt and Au atoms [151]. More commonly, however, two different ligands are needed, each targeting one moiety of the heterodimeric particles. For instance, Sun and co-workers utilized thiolated PEG and PEGylated dopamine to modify Fe3 O4 –Au nanoparticles, where the former ligand was anchored on the Au side and the latter one resided on the Fe3 O4 side (Fig. 3.15) [152]. For some particles, the two composites are assembled in a core–shell instead of a sideby-side manner. In these cases, the imbedded moiety has no contribution to the surface and

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FIGURE 3.15 (a) Schmatic illustration of the Au–Fe3 O4 dumbell nanoparticle surface modification. (b, c) TME images of Au–Fe3 O4 nanoparticle in hexane and in water, respectivelly. (Adapted from Xu et al. [152] with permission.)

the surface modification is just about the moiety that is exposed. For instance, Fe3 O4 @Au nanocomposites can be made water-soluble by sodium citrate, which is known to have a high affinity toward the Au surface but not the Fe3 O4 surface [153]. For a more detailed introduction to recent progress on nanocomposites, readers are referred to several other excellent review papers [150, 154, 155].

3.9 CONCLUSION In summary, surface modification is essential for the development of nanoparticle-based imaging probes. It makes the nanoparticles water-soluble and affords them an interface for biomolecule conjugation. With the increased availability of animal facilities worldwide, more and more experiments are conducted at the in vivo level. In this setting, better control over the particle pharmacokinetics is of critical importance, which is highly dependent on the nanoparticle surface modification and bioconjugation.

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CHAPTER 4

Biodistribution and Pharmacokinetics of Nanoprobes NAGESH KOLISHETTI Department of Anesthesiology, Brigham and Women’s Hospital, Harvard Medical School, Boston, Massachusetts, USA

FRANK ALEXIS Department of Bioengineering, Clemson University, Clemson, South Carolina, USA

ERIC M. PRIDGEN Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, Massachusetts, USA

OMID C. FAROKHZAD Department of Anesthesiology, Brigham and Women’s Hospital, Harvard Medical School, Boston, Massachusetts, USA

4.1 INTRODUCTION The field of medicine has made significant progress in the areas of diagnosis and therapy due to recent advances in nanotechnology. Early disease diagnosis substantially increases patient survival. Therefore, significant research efforts have been aimed at the development of new imaging agents and imaging systems. Nanoprobes are a class of imaging agents that consist of metal and inorganic materials such as quantum dots (QDs) [1–6], magnetic nanoparticle (MNPs) [7–11], carbon nanotubes (CNTs) [12, 13], and gold nanoparticles (GNPs) [14–19]. These agents are synthesized using a “bottom–up” approach. . . for precise control over physicochemical properties, allowing particles to be engineered with small sizes, long blood circulation times, and a functionalized surface for targeting capabilities. The ability to tailor the properties of nanoprobes combined with the discovery of new disease markers or pathological defects have led to increases in the specificity and sensitivity of these agents. Clinical translation of nanoprobes follows a path similar to drugs, with the evaluation of efficacy, stability, toxicity, purity, and pharmacokinetics. However, because of their larger size and slower excretion rate, national agencies have had to create specific guidelines

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to test short- and long-term interactions of nanoprobes with the biological environment. Research into the synthesis and screening of nanoprobes to improve their safety profiles is a very active area due to their clinical potential. More recently, progress has been made in the development of theragnostics [20–29], a term denoting the fusion of therapeutics and diagnostics. Nanoprobes are being combined with drug delivery systems to enable the simultaneous imaging and treatment of diseases, further enhancing their potential use in biomedical applications. Theragnostic agents are receiving increasing attention as pharmacogenomics increases in clinical use [30]. In this chapter, nanoprobe technologies, specific toxicological characteristics, clinical developments, pharmacokinetics, and multifunctional nanoprobes for theranostic applications will be reviewed.

4.2 NANOPROBE TECHNOLOGIES Various imaging techniques developed over the past few decades have contributed to the advancement of cancer and cardiovascular research and their translation to present medical practice [31–33]. Imaging techniques have aided the understanding of cancer biology as well as the detection and monitoring of tumors. Recent reviews have covered various advances in imaging technologies and their applications for cancer diagnosis [31, 32, 34]. These technologies are clinically used based on their sensitivity, spatial resolution, and tissue penetration. For example, X-rays are used for imaging hard tissues such as bones while ultrasound, magnetic resonance imaging (MRI), and computed tomography (CT) are used to image soft tissues such as the heart, lungs, and tumors. [31]. These techniques provide information about tissue defects or pathologies. Their spatial resolution varies from microns to millimeters. Recently, various microscopic and other intravital optical techniques like tomography or macroscopic reflectance have been developed to study genetic, molecular, and cellular events in vivo [31]. Numerous microversions of imaging modalities such as microCT, microMRI, and microPET (positron emission tomography), as well as microscopy techniques such as fluorescence, multiphoton, and electron microscopic techniques, are primarily used in cellular studies [31, 32]. Depending on the imaging technique, the contrast agents/molecular probes could be proteins, dyes, iodinated molecules, microbubbles, or radiolabeled molecules. Several nanomaterials are promising candidates (Fig. 4.1) as imaging agents (nanoprobes) such as iron oxide-based magnetic nanoparticles, quantum dots, carbon nanotubes, and gold nanoparticles. Unlike conventional imaging agents and therapeutics, nanoprobes are stable in the biological environment, have enhanced bioavailability, display absorption at various wavelengths, and provide better contrast or magnetic moment. In general, nanomaterials provide higher sensitivity and specificity than molecular imaging agents [31, 35]. 4.2.1 Quantum Dots A unique property of quantum dots (QDs) is their increased absorbance with increasing separation between excitation and emission wavelengths. Attenuation of signal during imaging is a major concern due to tissue thickness [36]. Absorbance, scattering, and waterto-hemoglobin ratio can affect the performance of QDs [37, 38]. One of the advantages of QDs is their ability to emit in the visible and near-infrared (NIR) emission wavelengths, where the biological tissues can transmit the light [39]. QDs are >10 times brighter than

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FIGURE 4.1 Examples of promising nanoprobes: quantum dots (QDs), gold nanoparticles (GNPs), magnetic nanoparticles (MNPs), and carbon nanotubes (CNTs).

fluorescent dyes and significantly less sensitive to photobleaching. QDs are used to image coronary vasculature, blood vessels, and the lymphatic system due to their superior fluorescence properties [40]. For example, NIR QDs have been used to image the coronary vasculature of a rat [37]. Coronary vasculature of thickness 282 ␮m in a rat was imaged with a signal-to-background ratio of 5 using an inexpensive white light of 2 mW/cm2 after intravenous (IV) injection of 2.5 nmol QDs, while conventional fluorophores like IRDye78CA needed 12.5 mW/cm2 to get similar resolution. QDs with higher emission wavelengths (>655 nm) were used to image blood vessels of chick chorioallantoic membrane (CAM) [41], a model for studying angiogenesis. These QDs eliminated all chick-derived autofluorescence and improved depth of field imaging. In another example, green light-emitting QDs remained fluorescent and detectable in the capillaries of adipose tissue and skin of a living mouse following IV injection [42]. Conventional FITC-dextran was unable to provide equivalent resolution even after using 5 times more imaging power, indicating that QDs have higher resolution when imaging was performed at a depth of 250 ␮m in a live mouse. Similarly, Gao et al. [38] reported targeted QDs for cancer imaging in prostate cancer tumor xenograft mice and estimated that tumor imaging sensitivity was increased to a detection limit of 10 to 100 cancer cells. In addition, cadmium selinide (CdSe)–zinc sulfide (ZnS) core–shell QDs were shown to be present in the liver and spleen for 4 months post-injection suggesting high in vivo stability over a long period of time. All of these reports demonstrate that QDs are excellent fluorescent probes for imaging various blood vessel tissues. They are extremely bright and photostable, have low background interference, and an improved depth-of-imaging over conventional fluorescent dye conjugates. Recently, major research and development efforts have focused on large-scale synthesis and size-controlling synthetic schemes to produce nanoparticles with high monodispersity

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for precise control of emission characteristics. QDs are an emerging class of nanoprobes for in vivo optical imaging [1, 43, 44] but have not reached the clinical phase of development mainly due to toxicological concerns.

4.2.2 Carbon Nanotubes Carbon nanotubes (CNTs) are rolls of graphene sheets, with unparalleled physical, mechanical, and chemical properties. CNTs can easily be synthesized in large quantities via chemical vapor deposition, arc discharge, and laser ablation processes. The two main types of CNTs are single-walled carbon nanotubes, SWNTs (0.4–2 nm diameter and 20–1000 nm length), and multiwalled carbon nanotubes, MWNTs (1–100 nm diameter and 1000–5000 nm length). MWNTs are distinguished by having multiple layers of graphene rolled on top of each other, resulting in strnger mechanical properties and chemical resistance. Due to the sharp densities of electronic states, SWNTs have distinctive optical properties [45]. A recent report provided detailed information about various methods used for both covalent and noncovalent functionalization of CNTs [46]. CNTs offer new opportunities in biomedical research because of their large surface area, which could allow the efficient loading of multiple molecules. CNTs have been used for in vivo animal imaging and have potential applications in various optical imaging techniques such as Raman imaging, photoablation, and photoacoustic molecular imaging of cancer [46–48]. The intrinsic near-infrared fluorescence makes SWNTs a promising optical agent [49]. For example, gadolinium-loaded SWNTs have been shown to have greater relaxivity than conventional MRI agents and therefore might become a promising contrast agent for MRI in the near future [50]. However, the clinical translation of CNTs will depend on how well-controllled the synthesis processes are, toxicological studies in various animal models, and interactions between the administered CNTs and the immune complement system.

4.2.3 Gold Nanoparticles Gold nanoparticles (GNPs) are interesting candidates for imaging/diagnostic and theragnostic applications due to their biocompatibility, inert nature, resistance to oxidation, and strong absorption of NIR radiation. Recent studies on colloidal gold nanoparticles have focused on CT applications in place of traditional iodinated molecules dissolved in liquids [51], or as X-ray and MRI contrast agents [51]. Their tunable plasmon resonance enables them to be specifically engineered to a desired wavelength for a particular application [52]. One distinct advantage of GNPs is the ease of surface functionalization using either thiol/disulfide sulfydryl or click chemistry approaches [52]. GNPs systems are currently being developed for high-sensitivty assays that require functionalization of the surface with biological molecules. For example, Hirsch et al. [52] developed a gold nanoparticle-based immunoassay to detect immunoglobulins in whole blood with a detection limit in the range of ng/mL. Mirkin’s group [15, 53–55] and others [14, 56] have developed quantitative multiple nucleic acid detection technologies, including a colorimetric sensing system based on the formation of nanoparticle networks and changes in color due to different distances between each nanoparticle. In addition to detection applications, GNPs have great potential to become a future theragnostic agent by combining the imaging capabilities with drug delivery and/pr photothermal therapy.

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4.2.4 Magnetic Nanoparticles Iron oxide magnetic nanoparticles (MNPs) are the most commonly used metallic nanoparticles due to their superior biocompatibility. Magnetic properties of iron oxide MNPs are characterized by the shortening of the T2 relaxation times of water, allowing imaging and diagnosis based on MRI [57]. MNPs do not possess magnetic properties outside an external magnetic field in contrast to other magnetic particles [58]. Unlike large magnetic particles with different domains of spins, MNPs have only a single domain of mutually aligned spins with a very large magnetic moment, making them especially effective contrast agents [59]. Because of the large magnetic moment, iron oxide MNPs are also known as superparamagnetic iron oxide (SPIO). MNPs are hydrophobic and require surface modification to disperse them in aqueous solutions. Multiple approaches are used to solubilize MNPs including polymer, surfactant, lipid-based, protein, and metal or oxide coatings [21, 22, 31, 34, 60–64]. Coating improves the nanoparticles’ stability, prevents immunogenic response, reduces opsonization, and has a limited effect on the nanoparticles’ magnetic properties. Polymeric coatings of magnetic nanoparticle have been shown to be the most promising approach and formulations used in the clinic or in clinical development are based on polymer coatings with affinity to iron oxide crystals [11]. SPIO-based MNPs are already used in medical imaging, mostly for imaging liver tissues due to their high accumulation in Kupffer cells, while ultrasmall paramagnetic iron oxide (USPIO) nanoparticles are in clinical development for lymph node imaging due to their longer circulation time [59, 65]. 4.2.5 Summary Nanoprobe imaging technologies have been developed based on the optical properties of materials, although the materials have not been optimized for biomedical applications. Most of the efforts have been in the synthesis of monodisperse nanoparticles with reduced impurities and surface modifications to improve their stability and biocompatibility. Thus, in the past 20 years, numerous imaging approaches using nanoprobes have been developed to improve resolution, sensitivity, and therapeutic applications but their possible clinical translation is limited by their immunogenicity and cellular and tissue toxicity.

4.3 NANOPROBE TOXICITY Nanoprobe toxicity is a concern for biomedical and clinical applications [66]. Different toxicity profiles are addressed depending on the material used for imaging or residual impurities from synthesis. Multiple factors such as aggregation, formation of reactive oxygen species, and nonspecific binding to proteins and cellular membranes influence the mechanism of toxicity in vivo. In general, surface coating, inert nanomaterials and/or biodegradability properties, and small size have been shown to decrease cell and tissue toxicity [67]. However, a universal toxicity assay to assess nanoprobes is difficult due to difference in cell types and assays used. Toxicities associated with different nanomaterials include stability, oxidative stress, genotoxicity, platelet coagulation, tissue-specific toxicity, hepatocyte toxicity, and immunogenicity [68, 69]. Therefore, in order to determine the potential toxicity of new nanomaterial formulations, the National Cancer Institute, NCI established the Nanotechnology Characterization Laboratory (NCL) to test and accelerate the transfer of technologies to the clinic through physical, in vitro, and in vivo characterization.

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Specifically, nanoprobes are tested for accuracy, sensitivity, and specificity. The NCL serves as a national resource of knowledge to facilitate the regulatory process and study preclinical toxicology parameters. The NCL has developed a series of characterization tests, including toxicity, that have been incorporated into the American Society for Testing and Materials (ASTM). The three new ASTM standards are based on the evaluation of nanomaterial toxicity in kidney and liver cells (E2526), hemolysis (E2524), and stimulation or inhibition of bone marrow macrophages (E2525). In addition, Ralph Weissleder’s group [70] developed an in vitro screening platform based on multiple cellular assays and cell types to analyze nanoprobe toxicity profiles in vitro and predict their in vivo behavior. A hierarchical clustering method has been used to create an identification map of nanomaterials with various surface and core properties that correlates to biological activity. However, it is accepted that three-dimensional (3D) toxicity assays of drugs represent a better structural framework to reproduce the complex interactions in tissues [71, 72]. Recently, Kotov and co-workers tested nanoparticle liver toxicity using a 3D cell culture system based on polyacrylamide hydrogels and demonstrated significant differences when compared to two-dimensional (2D) assays [72]. Because the 3D toxicity assays are emerging technologies and not yet ready for industrial applications and standardization, only toxicity data based on 2D assays and in vivo studies will be discussed. 4.3.1 Quantum Dots Toxic heavy metals like cadmium, mercury, lead, and arsernic are part of the nanocrystalline core of QDs covered by biologically inert zinc sulfide. Two major factors contributing to QD toxicity include their colloidal instability and heavy metal constituents leaching out, leading to metal poisoning [67, 68, 73]. In general, divalent cations such as cadmium or gadolinium have a long half-life in humans and are toxic in the body due to their accumulation in the placenta and kidneys, possibly leading to nephrotoxicity and kidney failure. Due to the large variety of QDs and disparity in experimental conditions to assay their toxicity, such as varied concentration and exposure times, it is only possible to draw conclusions based on their stability and surface coatings. In vitro studies with QDs were used to relate cytotoxic events to the release of potentially toxic elements and the size, shape, surface, and cellular uptake of QDs. Several groups investigated the release of cadmium ions from QDs into culture media by ICP-MS or fluorimetric assays. All these studies concluded that the release of cadmium ions correlates with cytotoxicity due to oxidative stress [74–76]. Yang et al. [77] studied the concentration of cadmium in the liver and kidneys after intravenous administration of 40 pmol of cadmium-based QDs containing ZnS shell and a hydrophilic surface coating (PEG-5000, Molecular weight, MW = 5 kDa). It was observed from the ICP-MS analysis that cadmium concentration increased over the course of 28 days and reached 40% in the liver and 40% in the kidneys with respect to the injected dose [77]. These results were confirmed using fluorescence microscopic analysis showing the presence of intact QDs in both the liver and kidneys. Free cadmium may be redistributed to the kidneys via hepatic production of metallothionein, which can lead to malignant transformation of cells and ultimately induce cancer [78]. In addition, oxidative degradation of QDs by exposure to air and UV irradiation released cadmium ions, leading to in vitro and in vivo toxicity [76, 79, 80]. To protect the body from metal poisoning, QD surfaces must be capped with a passivating agent, typically zinc sulfide (ZnS), cadmium sulfide (CdS), or bovine serum albumin (BSA). These core/shell QDs were prepared in nonpolar solvents and transferred to water using

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thiolated ligands. CdSe is used over cadmium telluride (CdTe) due to its higher stability toward oxidation. Hence these QDs were more stable and reported to be less cytotoxic. However, concentrations in the range of 10 ␮g/mL and long-term exposure (∼24 h) to cells were found to decrease cellular metabolic activity. The critical factors contributing to the toxicity of core/shell QDs are not yet clear, but it suggests a possible role of the Cd2+ leakage and QD aggregation [81]. In addition, carboxylated QDs were synthesized in aqueous solution with core CdTe and stabilized by mercapto acids (mercaptoacetic acid or mercaptopropio acid) due to their ease of synthesis, low cost, and immediate utility in biological buffers. The cytotoxicity of these thiolate-stabilized QDs was reported to be due to aggregation and degradation [74, 82]. Oxidative release of cadmium ions from the CdSe QDs by exposure to air or ultraviolet irradiation was investigated by Bhatia and co-workers [76]. Under the same conditions, CdSe QD cores coated with small thiolated ligands were also found to be toxic. Another report showed that cadmium ions release through oxidative degradation of the QDs and bind to sulfhydryl groups on a variety of intracellular proteins, causing decreased functionality in many subcellular organelles [83]. QDs were found to induce reactive oxygen species (ROS), leading to oxidative stress, inhibition of cell proliferation and induction of cell differentiation. The ROS toxicity was cell-dependent and oxidation of intracellular DNA and protein led to apoptosis. In the presence of ultraviolet radiation, QDs were found to catalyze the formation of ROS [82, 84]. Coating QDs with a thick ZnS shell was found to reduce the ROS production [85, 86]. Aldana et al. [87] showed photo-instability led to the degradation of thiolated coatings and aggregation of 3.2-nm size QDs. More stable colloidal QDs with core–shell structures (CdSe–ZnS) were prepared by encapsulation of the QDs in amphiphilic polymers and crosslinked silica [88]. These core–shell QDs with polymer coatings were found to be optically stable and less toxic than their counterparts coated in small ligands like trioctylphosphine oxide ligand [38, 39, 88, 89]. However, high concentrations (>5 × 109 QDs/cell or >2.3 ␮M) of these QDs were shown to cause toxic or inflammatory responses in Xenopus embryos for reasons that are not fully understood [89, 90]. Another report showed that colloidal instability causing precipitation of nanoparticles on cells led to the detachment of cells from cell culture substrates [83]. Other studies suggest that the toxicity is controlled by the size of the QDs. For example, cystamine-coated QDs (2 nm) were shown to accumulate in the nucleus. Small QDs (below 3 nm) were found to adsorb nonspecifically to intracellular proteins and impair cellular functions by invading the nucleus and binding to histones or nucleosomes, causing DNA damage [43, 91–93]. Biotinylated CdSe/ZnS QDs incubation with supercoiled DNA for 60 min showed 56% and 29% of DNA damage in the presence of UV exposure and dark, respectively [91, 93]. DNA damage was due to the free radicals generated by light and oxidation. In addition, Geys et al. [94] reported higher toxicity of negatively charged nanoparticles change to period and end sentence. They observed that negatively-charged nanoparticles caused pulmonary vascular thrombosis and suggested that the negative charge contributed to the activation of the coagulation through aggregation with plasma proteins. A current problem for in vivo applications is that a major fraction of the QDs remain at the site of injection for a very long time [95] and the long-term fate of these QDs in the body is yet to be clearly understood. Recently, Yang et al. [77] reported a long-term study (over 7 weeks) in mice showing that signal was found in the intestinal tract, suggesting fecal as the primary excretion pathway. Tissue histology and neurological evaluations did not show apparent changes 100 days post-injection. The long-term toxicity of QDs coated with crosslinked polymers does not show harmful effects at dosages of ∼10 mg/kg, but

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distress was reported at 42 mg/kg. This dose-dependent toxicity was previously reported for QDs containing mercury, which is known to have major safety concerns. These results suggest that the potential release of cytotoxic elements from QDs and their distribution in specific organs, tissues, cell types, and subcellular locations must be well understood before they can be used in humans. Due to their enhanced optical properties, QDs might be injected at doses below their cytotoxicty threshold, although the minimum dose for precise diagnosis remains to be determined. QDs with reduced or no heavy metal toxicity and colloidal stability may find an important role as clinical contrast agents in future optical imaging techniques. 4.3.2 Gold Nanoparticles Gold is known to be an inert material and it is reported that gold nanoparticles (GNPs) (size range 10–250 nm) can be taken up by various human cells without any cytotoxic effects at up to 4000 particles per cell [96, 97]. Murphy and co-workers did not observe any in vitro cytotoxic effect using gold nanoparticles of various sizes (4, 12, and 18 nm) and surface coatings (anionic, neutral, and cationic) on leukemia cells [98, 99]. The purification process for gold nanoparticles was a key step because residual chemical impurities such as surfactants were shown to be toxic to the cells at the nanomolar level. Ofek et al. [100] showed that zebrafish embryos treated with gold nanoparticles at sizes ranging from 3 to 100 nm did not show significant increase in mortality with increasing concentrations up to 250 ␮mol of gold in contrast to silver nanoparticles. The difference in toxicity between silver and gold nanoparticles was thought to be due to the chemical composition and residual impurities of silver nitrates [100]. Leonov et al. [101] showed that polystyrenesulfonate sodium salt (PSS, 70 kDa) could efficiently remove residual surfactant from gold nanorods. However, cytotoxicity profiles using three different cell lines (porcine kidney, human liver carcinomas, and human nasopharyngeal carcinomas) [101] showed unexpected results indicating that purified gold nanorods had higher toxicity than the surfactant-coated nanorods. These results indicated that PSS adsorption on the gold nanorods is not stable enough, leading to complexes with residual surfactants [101]. Several other studies have examined the effect of gold nanoparticle size on in vitro and in vivo toxicity [67, 102]. Hainfeld et al. [51] observed no in vivo toxicity after 30 days for 1.9-nm sized gold nanoparticles using mice as an animal model (dose = 2.7 g/kg) and the LD50 was found to be 3.2 g/kg. The nanoparticles were quickly cleared through kidneys without significant accumulation in the liver and spleen. Others have shown that gold nanoparticles with hydrodynamic diameter smaller than 2 nm are more toxic than larger particles. Maria et al. [103] showed that small gold nanoclusters (Au55 ) with a size of 1.4 nm are toxic to various human cancer and healthy cell lines in contrast to larger gold nanoparticles. Toxicity studies with 18-nm and 1.4-nm GNPs showed that larger NPs were less toxic [104], with the results attributed to the intercalation of GNPs in the DNA major groove. The smaller GNPs were also found to easily translocate in significant amounts through the blood barrier of the respiratory tract. Positively-charged, small nanoparticles were shown to increase cytotoxicity due to electrostatic interactions with cell membranes and/or DNA. No significant cytotoxicity differences have been reported due to shape when comparing spherical and nanorod gold nanoparticles. However, most of the surfactants used for stabilizing gold nanoparticles have been found to be cytotoxic. Therefore significant efforts are being devoted to using biocompatible materials. These results were consistent with data showing that gold nanorods

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coated with polyethylene glycol were less toxic than surfactant-coated nanorods based on cellular micromotility assays. No general conclusion can be drawn at present, although most of the research has clearly shown the effect of the size on the increased cytotoxicity profile of GNPs. Thus, these results suggest that more in vivo studies need to be carried out to understand the size, charge, and shape-dependent toxicity of GNPs. 4.3.3 Carbon Nanotubes Carbon nanotubes are known to cause significant cytotoxic responses such as induction of oxidative stress, inhibition of cell proliferation, and apoptosis in different cell lines [105]. Carbon nanotubes have also demonstrated major cytotoxicity in the lungs upon inhalation [106–109]. Lam et al. [106] showed pulmonary toxicity of SWNTs in mice after intratracheal instillation. All the mice showed epithelioid granulomas and interstitial inflammation after 7 and 90 days. Similar pulmonary toxicities were observed in rats by another group [110]. The gene expression of macrophages that had taken up SWNTs showed activation of oxidative stress and an inflammatory response [111]. For example, interleukin-6 (IL-6) was overexpressed (25-fold increase) upon incubation of macrophage with SWNTs [109]. The potential hazard of SWNTs strongly depends on the metal content and the size of the agglomerates [112]. CNT products contain impurities such as amorphous carbon and metals such as, Co, Fe, Ni, and Mo. The amount and type of impurities depends on the manufacturer and synthetic methods [112, 113]. Elemental analysis of raw SWNTs (uncoated) using ICP-MS showed substantial content of iron (10%), sodium (0.03%), and nickel (0.02%). A large number of studies over the past several years reported varied toxicity levels suggesting a dependence on the type of nanotube materials used and functionalization strategies. In vitro and in vivo toxicological studies have suggested reduced toxicity due to surface functionalization of carbon nanotubes [108, 114–116]. For example, glycodendrimercoated SWNTs were found to be nontoxic to HEK292 cells at a concentration of 100 ␮g/mL, while the non-functionalized SWNTs greatly inhibited cell growth [116]. In another report, biotinylated SWNTs were found to be nontoxic up to concentrations of 0.05 mg/mL in HL60 cells [114]. In mice, 151 and 47 mg of PEG-functionalized SWNTs was found to be nontoxic even after a 4-month period [108]. SWNTs suspended in Tween-80 exhibited toxicities due to accumulation in the liver and lungs after 3 months at higher dose (40 mg/kg), while no detectable toxicity was observed at 2 mg/kg [117]. However, SWNTs functionalized with phenyl-carboxylic groups were shown to be less cytotoxic than surfactant-coated (Pluronic F108) SWNTs. Functionalized CNTs with biocompatible surface coatings have been shown to be nontoxic with greater renal clearance and insignificant reticuloendothelial system (RES) uptake and are promising candidates for future studies [46, 50, 108, 117–119]. For example, PEGylated SWNTs administered through IV injection in mice at a concentration of 3 mg/kg showed normal blood chemistry and histological observations after 4 months [108]. This study did not show significant systemic toxicity of SWNTs after a single dose injection but white blood cell concentrations decreased by 50%. In another example, 20 ␮g of linear and branched PEG functionalized SWNTs showed no toxic side effects, low RES uptake, and near complete clearance two months after IV injection [119]. Although no acute toxicity is reported using histological assessment of tissues, multiple reports suggested increased levels of oxidative stress. Overall, SWNT have been shown to be more toxic than MWNTs, possibly due to their geometry [67]. Poland et al. [120] showed that non-functionalized long MWNTs pose a carcinogenic risk in mice.

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The report systematically studied various commercially available MWNTs with different lengths. Upon administration at a concentration of 100 mg/mL via peritoneal injection in mice, MWNTs demonstrated asbestos-like, length-dependent, pathogenic behavior. The longer length of MWNTs is believed to be the reason for the inflammatory response and granuloma formation within a week post-injection due to less aggregation and increased macrophage uptake. 4.3.4 Magnetic Nanoparticles The chemical toxicity of iron and its derivatives is well studied [121, 122]. Iron oxides have been shown to degrade in vivo according to natural iron pathways. Normal human tissues contain iron or iron oxides in the form of hemosiderin, ferritin, and transferrin [123]. Therefore, numerous iron oxide-based magnetic nanoparticles (MNPs) have been reported and studied for MRI applications over the last two decades [60–62]. SPIOs must be functionalized during the synthesis process for biocompatibility before medical use [59]. A variety of methods have been developed to encapsulate SPIO particles within a sheath of benign polymers such as dextran [124], polysaccharides [125], PEG, and polyethylene oxide [126]. Weissleder and co-workers reported the proof of concept profiling 50 different iron oxide based nanoprobes. The in vitro assays addressed the effect of the core composition, coating, and surface functionalization on the interactions of nanoparticles with the biological environment [70]. All the nanoparticle formulations were organized into clusters based on their biological activity and compared to Feridex as an FDA approved control [70]. The data from the four cell toxicity assays, five different cell types, and four different doses and incubation times showed that the core composition had a strong contribution to the biological effects. For example, nanoparticle formulations with carboxylic or ethylenediamine surface coatings had similar biological acitivty. The analysis of the data showed that crosslinked iron oxide (CLIO)-amine was in the same cluster as dextran-coated Feridex even though their surface coatings were significantly different. More importantly, results showed that these two formulations did not affect monocytes in the blood and spleen in contrast to the control sample that was clustered in a different group based on in vitro assays. In addition, Jain et al. [127] did not observe significant differences in the liver enzyme activity in vivo compared to the group of mice injected with saline. Overall, the histology and enzymatic activity data suggested that the oxidative stress was minor and did not affect cellular and tissue integrity. Iron oxide concentration in liver did not exceed 300 ␮g/g of tissue during the experimental period of 3 weeks, which included the redistribution of iron through protein binding. Iron oxide concentrations were 10 times lower than the toxic concentration reported to develop cirrhosis and hepatocellular carcinoma and 4 times lower than the dose used for Feridex in humans (2.6 vs. 10 mg/kg, respectively). Many iron oxide-based MNPs are thought to have suitable toxicity profiles and are now being evaluated in clinical trials. The promising results of the Phase I clinical trial using ultrasmall superparamagnetic iron oxide colloids (USPIOs) BMS180549 showed mild to moderate side effects for 45% of the volunteers [128, 129] and the most common adverse event was urticaria. This is possibly due to the dosage of iron oxide NPs used for diagnostic imaging (1–2 mg Fe/kg bodyweight), which was less than the normal iron store and dose required to develop chronic iron toxicity. The reported iron concentration for cirrhosis and hepatocellular carcinoma is over 4 mg of Fe per gram of wet liver [63]. These SPIO formulations passed the standard toxicological and pharmacological requirements, most

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likely due to the low doses used in humans and their use as a single-dose contrast agent. However, patients injected with SPIO generated anti-dextran antibodies and developed allergic reactions. 4.4 NANOPROBE CLINICAL TRANSLATION Two SPIOs and one USPIO are currently approved by the FDA for clinical use. In 1996, Feridex I.V. became the first FDA approved nanoprobe for imaging liver lesions based on the differential amount of Kupffer cells in metastatic tumors in the liver, allowing detection of small metastases using magnetic resonance imaging (MRI). Feridex is a SPIO nanoparticle (120–180 nm) coated with dextran (MW = 10 kDa) given at a dose of 0.56 mg Fe/kg and has a circulation half-life of ∼1 hour in humans. In addition, Resovist, which is an SPIO formulation coated with carboxydextran (Molecular weight = 20 kDa) dosed at ∼2.6 mg/kg, was approved in 2001 for the European market. The advantage of Resovist is safety because it allows rapid injection without cardiovascular side effects and lumbar pain. This is due to its smaller size (∼60 nm) and increased cell internalization. In June 2009, the FDA approved the first USPIO nanoprobe, Ferumoxytol, which is used for the treatment of iron-deficiency anemia in adult patients with chronic kidney disease. Clinical data showed that patients treated with Ferumoxytol demonstrated greater hemoglobin enhancement compared to iron given orally. In addition, many USPIO magnetic nanoparticles (Table 4.1) are currently in different clinical trials as contrast agents for several different applications. In TABLE 4.1 Various Nanoprobes Under Clinical or Preclinical Investigations Trade Name

Composition

Status

Indication/Application

Size (nm)

Circulation Time (h)

Feridex I.V. (Ferumoxides AMI-25) Combidex (Ferumoxtran10, AMI-227, BMS-180549) Gastromark (Ferumoxsil AMI-121) Resovist (Ferucarbotran SHU-555A) Supravist (Ferucarbotran SHU-555C) Ferumoxytol code 7228

Dextran-coated SPIO

Approved

Liver imaging, cellular labeling

120–180

2

Dextran-coated USPIO

Phase IV

15–30

24–36

Siloxane-coated SPIO

Approved

Lymph nodes, RES directed liver diseases, macrophage imaging, cellular labeling Bowel marking, oral GI imaging

300

Oral

Carboxydextrancoated SPIO

Approved

Liver imaging, cellular labeling

60

2.4–3.6

Carboxydextrancoated SPIO

Phase III

Blood pool agent, cellular labeling

21

6

Cyclomethyl dextran-coated USPIO Citrate-coated VSSPIO Pegylated colloidal gold nanoparticles

Phase II

Macrophage imaging, blood pool agent

30

10–14

Phase II

Blood pool agent, cellular labeling Tumor therapy and tumor vasculature imaging

7

0.5–1.5

30

2–6

VSOP-C184 Aurimune (CYT-6091)

Phase I

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general, USPIOs have longer circulation times and are mostly used for cancer metastases diagnosis. For example, Ferumoxtran-10 was shown to be a very sensitive and specific agent in human cancer trials allowing accurate diagnosis of metastatic lymph nodes with a reported circulation half-life of ∼24 h [130]. Ferumoxtran-10 is currently being tested in 8 clinical trials for tissue neoplasm (brain and bladder), aortic aneurysm, and prostate metastases diagnosis (www.clinicaltrials.gov) in Europe. Aurimune (CYT-6091) is the only system based on materials other than iron oxide that has reached clinical trials. Aurimune delivers TNF-␣ bound to PEG-coated gold nanoparticles (∼27 nm) for solid tumor therapy [131] and possible imaging abilities. TNF-␣ is a potent cytokine with antitumor cytotoxicity which requires incorporation into a nanocarrier formulation to reduce systemic toxicity. The toxicity data from the Phase I trial showed that gold nanoparticles could deliver 3 times more TNF-alpha systemically than a free dose of TNF-alpha. The TNF-alpha dose was reported to be 1.2 mg and considered to be in the therapeutic range used previously in isolated limb perfusion (ILB).

4.5 FACTORS AFFECTING THE BIODISTRIBUTION AND PHARMACOKINETICS Nanoparticle-based technology is being increasingly employed in drug delivery and imaging applications. For most applications, intravenous administration of nanoprobes is required. Compared to in vitro studies, different challenges arise in vivo due to the increased complexity of the organism. The unique physiochemical properties of nanoprobes include size, shape, chemical composition, surface chemistry, porosity, and agglomeration state. These properties control their interactions with the biological environment. In addition to toxicity, understanding the fate of nanoprobes in a complex biological environment is very important for potential clinical translation. Biodistribution and pharmacokinetic studies give a detailed knowledge about the fate of nanoprobes in the body, including circulation time, excretion pathway, degradation time, and plasma protein binding. We discuss the main factors affecting nanoprobe biodistribution and pharmacokinetics such as size, shape, surface functionality/coating, and charge. 4.5.1 Effect of Size In general, nanoparticles enter the body through the vein, skin, lungs, or gastrointestinal adsorption. Circulating nanoparticles in the body will differentially distribute/accumulate in major organs according to cell interaction properties and pathological defects [132].For example, filtration in the spleen and liver is dependent on the opsonization effect and related to the size of the nanoparticles. The endothelial barrier controls and prevents the diffusion of circulating nanoparticles out of the bloodstream. Blood vessel defects can form under pathological conditions, leading to enhanced permeability and differential accumulation of nanoparticles at specfic site in the body. This phenomenon can be taken advantage of to facilitate nanoparticle (10-500 nm) accumulation in cancer tissue due to the relatively leaky tumor vasculature and absence of a lymphatic network, resulting in the enhanced permeability and retention effect (EPR). Hence the effect of particle size on interactions with plasma components, blood cells, endothelium, and distribution in various tissues are of great interest.

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There are only a few reports of QD biodistribution and pharmacokinetics parameters. However but it is well established that QDs accumulate significantly in the reticuloendothelial system (RES), including the liver, spleen, and the lymphatic system [135–137]. Kim et al. [39] reported a critical size of ∼15–19 nm for the retention of QDs in the sentinel lymph node during imaging of the lymphatic system. The injection of QDs of different colors at different intradermal locations allowed imaging of their drain to common nodes [138, 139]. Similarly, it was reported that QDs smaller than 9 nm could migrate into the lymphatic system with up to 5 nodes showing fluorescence [140]. QDs with size ∼9 nm could be entirely eliminated from the kidneys and directly extravasate out of blood vessels into interstitial fluid [140]. Another study found that the renal clearance of QDs was closely related to hydrodynamic diameter and the renal size threshold was defined as ∼5–6 nm [141]. The blood half-life of QDs varied from 48 min to 20 h for sizes ranging from 4.4 to 8.7 nm, respectively, demonstrating rapid urinary excretion [141]. Smaller QDs (∼5 nm) were found in detectable amounts only in the liver (4.5%) and kidneys (2.6%). Larger QDs (∼8 nm) showed higher uptake in the liver (26.5%), lungs (9.1%), and spleen (6.3%) 4 h post-intravenous injection. This was possibly due to longer circulation times allowing higher accumulation in tissues [141]. However, more studies need to be carried in vivo to determine the effects of QD size. Many reports on the biodistribution of MNPs examined the effect of particle size [142–144]. In vivo biodistribution data for five neutral particles with different sizes ranging from 30 to 90 nm was reported. [142] and the results showed that smaller nanoparticles (<50 nm) have longer blood circulation time (>10 min for 50% clearance) and less uptake by the liver (∼20% for small particles and >50% for larger particles). After 20 min, it was found that the liver uptake increased as the size of the NPs increased. Pharmacokinetic profiles were studied with nanoparticles of three different sizes ranging from 46 to 75 nm to determine the clearance time. It was found that increasing the particle size decreased the plasma half-life, with 46-nm particles showing 50% clearance from the bloodstream in the first 10 min in contrast to only 5 min for 90-nm particles [142]. Neuberger et al. [143] reported that 200-nm MNPs were cleared faster than MNPs of sizes 30–100 nm. It was also demonstrated that for iron oxide particles smaller than 40 nm in diameter, both biodistribution and blood half-life were mostly controlled by the coating material rather than the mean particle size [11, 145]. For example, circulation half-life of two magnetic nanoparticles with the same polymeric surface coating, Ferumoxtran-10, had a longer circulation time compared to Feridex due to its smaller size (∼15–30 nm vs. ∼120 nm, respectively). The slower clearance of Ferumoxtran-10 was believed to be due to larger particle size and/or surface charge affecting degradation and biodistribution. Blood half-lives of magnetic nanoparticles in human are reported to be up to 2 days and USPIOs were found to have the longest circulation time partially due to their smaller size and lower accumulation in the liver. Functionalizing the surface of MNPs with targeting ligands can significantly affect the circulation half-life and biodistribution. Reddy et al. [146] found that MNPs of 40 nm showed increased tumor contrast half-life in mice from 39 min to 123 min by attaching a targeting F3 ligand. Montet et al. [147] showed that the MNP–RGD conjugates with sizes 28 and 36 nm were found to have blood half-lives of 180 and 207 min in rats. Biodistribution studies 24 h after IV administration showed accumulations in the liver, spleen, skin, kidneys heart, and tumor [147]. According to the reported data, the overall size of MNPs should be smaller than 200 nm to evade rapid splenic filtration [148], but larger than ∼5 nm to avoid renal clearance [141].

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Carbon nanotubes have been shown to penetrate into the lungs and cross the blood–brain barrier, leading to neurotoxicity. The first study of CNTs in animals to compare tissue distribution characteristics after subcutaneous, oral, and intravenous administration was performed using iodine-labeled hydroxyl-functionalized carbon nanotubes (125 I-CNT) with a length of 300 nm. It was found that CNT accumulated in the stomach, kidneys, and bone similar to small molecules. The results showed no tissue damage and excretion through urine [149]. Later, this same group reported that the functionalized CNTs showed persistent liver accumulation after IV injection [150]. Additional biodistribution data using indiumlabeled carbon nanotubes (111 I-CNT) also showed accumulation in kidneys, bone, muscle, and skin. The circulation half-life was found to be ∼3 h with rapid clearance through urine. In a recent report, yttrium-labeled carbon nanotubes (86 Y-CNT) with 42 ± 17 nm length were shown to be cleared from the bloodstream within 3 h when injected intravenously in mice [151]. The PET images indicated that the major sites of accumulation of 86 Y-CNT were the liver (18%), spleen (14%), kidneys (8%), and to a less extent bone (2%). The uptake of CNTs in the kidneys, liver, and spleen was slightly lower when the CNTs were administered via intraperitoneal (IP) injection, suggesting a relatively slow egress from the intraperitoneal compartment into the vascular compartment. It was observed that the clearance of CNTs from the kidneys was much faster than from the spleen and liver. Elgrabli et al. [152] used commercial uncoated MWNTs with a diameter of 20–50 nm and length of 500–2000 nm to study the biodistribution in rats. The results showed CNT bio-persistence and clearance 6 months after respiratory administration. The MWNTs were intratracheally instillated and 53% of the initial dose accumulated in the lungs after 24 h and 16% of MWCNTs were still found in the lungs after 6 months. However, most of the biodistribution studies were carried out using a radiolabel tracking method. Therefore, the dissociation of the label could affect the measurements. There are few reports describing the size-dependent cellular uptake and biodistribution of gold NPs over the last decade. Chithrani et al. [96] reported in vitro cellular uptake of GNPs with different sizes in HeLa cells. It was found from the uptake kinetic studies that the uptake half-lives of 14-, 50-, and 74-nm colloidal gold nanoparticles were 2.1, 1.9, and 2.24 h, respectively. De Jong and co-workers studied size-dependent (10-, 50-, 100-, and 250-nm) biodistribution of spherical gold nanoparticles after intravenous administration in rats [97]. A clear difference was observed between the biodistribution of the 10-nm particles and the larger NPs. Small gold nanoparticles (10 nm) were found in various organs including the blood, kidneys, testes, thymus, heart, lungs, and brain, whereas larger particles (>10 nm) were only detected in the blood, liver, and spleen after 24 h [97]. However, Wolfgang and co-workers studied the biodistribution of gold NPs with 1.4- and 18-nm sizes in rats after intravenous injection and found no clear differences between the nanoparticle sizes tested [104]. Similar accumulation and biodistribution of the NPs was observed in the liver, blood, carcass, spleen, skin, kidneys, urine, and feces 24 h post-injection. Intratracheal administration of the same nanoparticles using rats as an animal model showed 99.8% and 91.5% accumulation in the lungs after 24 h for 18-nm and 1.4-nm GNPs, respectively. Small amounts of the instilled 1.4-nm NPs (0.6–3.3% I.D.) were detected in the blood, urine, skin, and carcass, whereas no detectable amounts of larger nanoparticles (18 nm) were found [104]. Hillyer and Albrecht [153] showed in mice that metallic colloidal gold nanoparticles of different sizes (4, 10, 28, and 58 nm) distributed to other organs after oral administration in mice. The 4-nm GNPs showed differential accumulation in the kidneys, liver, spleen, lungs, and even in the brain in contrast to the 58-nm particles detected solely inside the gastrointestinal tract. Hainfeld et al. [51] subcutaneously injected 1.9-nm GNPs

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into mice and examined their long-term distribution (11 and 30 days) in various organs. The biodistribution of the gold nanoparticles in an EMT-6 syngeneic mammary carcinoma xenograft mouse model showed very low retention of GNPs in the liver and spleen due to elimination by the kidneys after 24 h. Pharmacokinetic studies revealed a biphasic clearance, with a drop of 50% between 2 min and 10 min post-injection followed by a slower decrease of another 50% between 15 min and 1.4 h. Accumulation of the NPs in the kidneys (10.6% I.D.), tumor (4.2% I.D.), liver (3.6% I.D.), and muscle (1.2% I.D.) was evaluated 15 min post-injection. In another report, Vijaya Kattumuri et al. [154] used 15–20-nm size gum arabic stabilized GNPs to study biodistribution in larger animals such as pigs. Following intravenous administration of the GNPs (0.8 mg/kg), identical amounts (∼50 ppm) of Au were observed in the lungs and liver after 24 h, while 40 ppm of gold was measured in the spleen. After 72 h, 43–69% of the administered dose was retained in the liver while only 1–2% was observed in the kidneys, indicating slow clearance of GNPs. 4.5.2 Effect of Shape Most NPs exhibit a spherical shape as a result of surface energy minimization during their synthesis. The advent of mimicking nature with nanoparticles of different shapes such as a virus or bacterium is now possible due to the development of new fabrication processes. The ability to make particles with shapes other than spheres is opening a path to new design solutions for systemically administered particulates. Decuzzi et al. [155] reported that the intravascular journey of the particle can be broken down into three events: (1) margination dynamics, (2) firm adhesion, and (3) control of internalization. This report compared the predictions of mathematical models and showed that the particle geometry plays an important role in all three events [155]. In addition, recent studies performed using polystyrene nanoparticles of different shapes showed that the phagocytosis pathway by macrophages exhibits a strong dependence on shape [156, 157]. A recent report used a physiologically based pharmacokinetic (PBPK) model along with experimental data of biodistribution of QDs with various shapes, sizes, and exposure routes, to predict the biodistribution of other QDs [158]. The model estimated partition coefficients for various tissue concentrations of QDs in the blood, kidneys, liver, muscle, and skin using the experimental data from commercial QD705. The ellipsoid particles showed increases in the steady-state concentration of QDs in the liver and kidneys, similar to experimental observations. This model considered that the QDs initially distributed to the kidneys, then redistributed to other compartments with higher partition coefficients but much slower flow rates per mass of tissue as these compartments approach steady-state concentrations. Gold nanoparticles can also be made in different shapes such as rods and pyramids with an aspect ratio of 1:1 to 1:5 [96]. The cellular uptake in HeLa cells was studied with various shapes and found that spherical particles have a higher probability of uptake than rod-shaped nanoparticles with aspect ratios 3 and 5. Niidome et al. [159] studied the biodistribution of PEG(5000)-functionalized gold nanorods (65-nm length and 11-nm width) and surfactant-stabilized gold nanorods in mice. The results showed that the pegylated gold nanorods had a blood circulation time of 30 min (∼50% I.D.), while the residual CTABfunctionalized nanorods in blood were significantly lower (∼5%) [159]. No detectable amounts of gold were present in the blood, while 35% was accumulated in the liver and very small amounts were in the lungs, spleen, and kidneys, showing the clearance of gold nanorods after 72 h. In contrast, CNTs, which are known to have high aspect ratios, were shown to have a rapid blood clearance (circulation half-life = 1 h) in rabbits due to their

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rapid renal excretion and accumulation in the liver [49]. Another study showed a blood circulation half-life of 3 h for the water-soluble DTPA-functionalized SWNTs in mice after IV administration [160]. 4.5.3 Effect of Surface Coating and Charge Other than size and shape, the nanoparticle surface coating and charge can modulate the in vivo fate of nanoprobes. NP coatings affect surface properties such as surface charge (zeta potential), surface functionalities, and degree of hydrophilicity. Surface modification of QDs is necessary due to their low solubility in polar solvents. There are many reports addressing the surface modification of QDs with a variety of coatings such as small molecules or encapsulation in amphiphilic polymers to increase surface hydrophilicity, reduce nonspecific binding, and prolong blood circulation [38, 89, 161, 88, 162]. QD structure and surface properties have been found to strongly impact plasma halflife. The half-life of anionic, carboxylated QDs in the bloodstream of mice was significantly increased from 4.6 min to 71 min by coating QDs with PEG polymer chains of MW 5000 [135]. Schipper et al. [163] evaluated the biodistribution of commercially available PEGcoated CdSe QD525 and QD800 in mice. Both types of QDs rapidly accumulated in the liver within 6 min postinjection. In mice, PEG-coated CdSe/ZnS QDs were rapidly removed from the bloodstream into organs of the RES with detection possible for up to 4 months post-injection using fluorescence imaging [135]. TEM studies on the tissues revealed that these QDs retained their morphology, suggesting QDs are stable in vivo given the proper coating. The blood circulation half-life of these QDs was significantly increased by coating QDs with poly(acrylic) or a short PEG(750), while the half-life increased to 72 min with an increase in the length of PEG(5000) as determined from venipuncture studies [135]. Multiple approaches have been used to attach ligands to the surface. For example, heterofunctional ligands have been attached to the surface of QDs with the hydrophilic group exposed [164, 165]. In addition, biological molecules such as monoclonal antibodies (J591) or EGF receptor (erbB1) have been attached to the surface, serving the dual purpose of increasing surface hydrophilicity and acting as a targeting ligand [161, 166, 167]. Fc fragments of antibodies conjugated to QDs were reported to have longer circulation times due to the reduction of nonspecific interactions [168]. Another report demonstrated slower removal of albumin-coated QDs from the bloodstream before accumulation in the liver compared to QDs without albumin [137]. Surface charge will lead to different mechanisms of interaction with macrophages and may affect the intracellular fate of internalized NPs. Even though cationic NPs show increased cell internalization, supramagnetic particles with negative surface charge were found to exhibit a high but nonspecific affinity for the plasma membrane, favoring adsorption and endocytosis in endosomal compartments of HeLa cells [169]. In another report, rpositively-charged MNPs were prepared by the coating of monocrystalline iron oxide nanoparticles(MIONs) with cationic poly-l-lysine. The blood half-life was found to be only 1–2 min in comparison to 2–3 h for the uncharged variant [170]. Small neutral MION particles distributed to lymph nodes, whereas the positively-charged particles of similar size were found to be rapidly taken up by the liver [63]. Jallet and co-workers carried out a systematic study to analyze the effect of charge on MNPs as contrast agents for MRI [142]. When the surface charge of these nanoparticles was varied from neutral to negative, there was a slight decrease in the particle size while the change to a positive charge led to an increase in the size of the MNPs. The biodistribution results showed that charged

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particles had greater accumulation in the liver relative to neutral particles. Twenty minutes after injection of neutral MNPs into the bloodstream via intravenous injection, 40% of the NPs remained in the bloodstream. In contrast, only 10% of the negatively charged particles remained in the blood. Uptake of positively charged particles (∼60%) was higher than neutral particles (40%), but lower than the negatively charged MNPs (∼75%) in the liver. The uptake of MNPs in the liver was almost 3 times higher (∼75%) for negatively charged MNPs compared to neutral MNPs [142]. MNPs of 100 nm showed 20% more uptake when the surface charge was positive compared to neutral [142]. Citrate-coated iron oxide NPs (8.7 nm) had a blood half-life of 1 h in humans due to their highly anionic surface charge, while a commercially available Ferumoxtran-10 had a blood half-life of 24–36 h in humans and 2–3 h in rats [11]. Similarly detectable amounts of charged particles were observed in the carcass (bones, skin, muscles, and the whole head). Overall, this study concluded that the negative charge on the particle surface resulted in enhanced liver uptake [142], while the MNPs with neutral surfaces were taken up the least by the liver. More importantly, coated magnetic nanoparticles remained in the liver for a few months with sustained contrast properties and were shown to be mainly eliminated through feces in human studies. CNTs were functionalized with various groups on the surface to improve water solubility. Yuliang and co-workers used hydroxyl-functionalized, iodine-labeled carbon nanotubes (125 I-CNT), which showed blood clearance and accumulation in the stomach, kidneys, and bones [149]. Amine-functionalized CNTs with indium labeling cleared the blood compartment [160] and all tissues within 3 h except for the kidneys, liver, and spleen. Rapid clearance of indium-labeled CNTs and significant accumulation in the kidneys, liver, spleen, and, to a small extent, bone was reported within 1 h after administration. Wenxin and co-workers reported water-soluble MWNTs by surface functionalization with glucosamine (MWNT-G) [171]. MWNT-G as injected into mice by intraperitoneal injection quickly spread to measurable levels in the blood, heart, lungs, liver, spleen, kidneys, stomach, intestines, coat, muscle, enterogastric area, and feces, behaving like a small molecule. A significant amount of total activity was retained throughout the 24-h study, particularly in the stomach. Samples collected from urine and feces showed >70% activity after 24 h, indicating MWNT-G was excreted predominantly via urine and feces. The pharmacokinetic profiles indicated a blood circulation half-life of about 5.5 h. Another study by Liu et al. [172] showed PEG-wrapped SWNTs can escape the RES for a blood circulation half-life up to 2 h. A more recent study revealed the increase in blood circulation time from 2 to 15 h when branched (MW = 7 kDa) PEG is used instead of linear (MW ranging from 2 Da to 12 kDa, methoxy end group) [119]. When the branched PEG is covalently bound to SWNTs, the blood circulation time is prolonged to 15.3 h and a low hepatic uptake was observed in the biodistribution study [118]. Most of the SWNTs were shown to accumulate in the spleen and liver in contrast to other studies using a radiolabeled tracking method. PEG-SWNTs were found to be distributed throughout most of the organs within 1 h and there were still considerable amounts of PEG-SWNTs present in the liver and spleen after 50 days. However, this study confirmed earlier reports of feces as the major excretion path. McDevitt et al. [173] showed significant difference in biodistribution and pharmacokinetics for targeted and nontargeted CNTs. The targeted CNTs showed more accumulation in the spleen and liver while the nontargeted CNTs showed more accumulation in the kidneys. Gold nanoparticles tend to agglomerate in a few hours without proper stabilizers or coatings. Typical stabilizers are surfactants, organic carboxylates, long-chain amines or thiols, sugars, carbohydrates, and proteins [154]. Coating gold nanoparticles with gum

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arabic, a hydroxyproline-rich arabinogalactan, has been found to have enhanced kinetic inertness and high in vivo stability. These nanoparticles were found in the lungs, liver, and spleen within 30 min [154] and significant amounts of NPs were present in the liver 72 h after intravenous administration. Another report showed negatively charged GNPs obtained by coating with water-soluble sulfonated phosphane accumulated in the lungs (>90%) 24 h after intratracheal instillation, while 47% were found in the liver when administered via intravenous injection [104]. However, more systematic studies of biodistribution of gold nanoparticles with variations in surface charge are needed before they are tested in the clinic.

4.6 THERAGNOSTIC NANOPROBES Nanoparticle delivery systems are becoming increasingly recognized as a potential therapeutic vehicle with multiple examples approved by the FDA for therapy. The similarities to nanoprobes have led to the design of multifunctional nanoparticle systems able to perform imaging and therapy simultaneously. In general, the therapeutic agent adsorbed on the surface of the nanoprobe is directly exposed to the biological environment and delivered at a dose controlled by the surface loading, which is reported to be in the range of 10–100 molecules in most cases. There are two promising therapeutic approaches (Fig. 4.2) that have been combined with imaging. The first approach is based on the combination of imaging and hyperthermia. The second approach is based on the combination of imaging and delivery of cytotoxic drugs. We discuss both strategies for the combination of therapy and imaging. Currently, Aurimune (CYT-6091) is the only multifunctional nanoparticle with imaging and delivery properties that has reached the clinic. CytImmune Sciences Inc. developed a nanoparticle system with the ability to bind TNF-␣ on the surface of gold nanoparticles through electrostatic interactions. Preliminary SEM micrographs of nanoparticles accumulated in breast tumor tissue sections in contrast to healthy tissues showed targeting of the nanoparticles by the EPR effect. Other formulations are still in the discovery stage using combinations of drugs such as TNF-␣ with paclitaxel, doxorubicin, or interleukin–12. Although the load of therapeutic agent is reported to be several hundreds of molecules,

FIGURE 4.2 Various promising nanoprobes used for theragnostics (see Refs. 175–183).

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recent data from a Phase I clinical trial reported that gold nanoparticles could deliver a therapeutic dose (www.cytimmune.com). Most other theragnostic nanoprobes are in early discovery/development stages, but the early clinical results are clearly demonstrating the potential opportunities for the development of theragnostic agents. Gold nanoprobes have been used mostly as a hyperthermia therapeutic agent by heating the tumor through the application of a magnetic field. The local increase in temperature kills the surrounding cells. In vivo results [174] of nanoshell-mediated NIR (near-infrared) thermal therapy showed that the nanoparticles induced irreversible cancer tissue damage at a temperature ∼40 ◦ C in a human breast cancer xenograft. However, the distribution of nanoparticle in the tumor might significantly affect the control of the local temperature. Similarly, Schwartz et al. [175] reported efficient therapy in large animals for brain cancer. The results showed that the temperature reached ∼70 ◦ C in tumor tissues and ∼50 ◦ C in normal white and grey matter, which is expected to significantly damage nondiseased areas of the brain. In addition, Ross and co-workers reported targeted iron oxide based nanoparticles as a delivery system (∼40 nm) of photofrin, which can be activated. . . with wavelength light (∼630 nm) to treat brain cancer [146]. Light irradiation activates the photosensitizer to produce a single oxygen leading to apoptotic and necrotic cytotoxicity. Photofrin loading on the nanoparticles was ∼4% w/w and was administered intravenously at a dose of 7 mg/kg using a rat 9L tumor xenograft model implanted into the animal brain. Tumor size was significantly reduced and the 50% survival time was extended to at least double that observed in the nontargeted nanoparticle control group [146]. MRI was used to monitor changes in the tumor and the nanoparticles showed magnetic resonance contrast of the targeted tumor and therapeutic efficacy of photofrin [146]. Similarly, Weissleder and coworkers have reported a macrophage-targeted theragnostic system based on photodynamic therapy for cardiovascular diseases [176, 177]. A potent sensitizer, 5-(4-carboxyphenyl)10,15,20-tryphenyl-2,3-dihydroxychlorin (TPC), was covalently attached to the surface of the iron oxide nanoparticle. The load was estimated to be three photosensitizer molecules per nanoparticle and the results showed efficient killing of macrophages [177]. Others have combined imaging and therapeutic functions using drug–material interactions as a loading strategy. For example, Farokhzad and co-workers reported the loading of doxorubicin, a potent chemotherapeutic agent, through intercalation between base pairs of aptamers conjugated to the surface of quantum dots [178]. In this case, aptamers were used as a specific targeting ligand to cancer cells and molecular carrier of the chemotherapeutic drug [178]. This system was precisely designed to have doxorubicin loaded on the surface quenching the quantum dot fluorescence. Therefore, the quantum dot lit up only when doxorubicin was released, allowing simultaneous delivery of the chemotherapeutic drug into specific cells and monitoring of the delivery using fluorescent imaging [178]. Adair’s group [179, 180] reported intracellular imaging and delivery using biocompatible calcium phosphate nanoparticles (∼27 nm). This system provided controlled drug delivery through pH dissolution of calcium phosphate in the acidic tumor environment. In vitro studies showed high uptake of the nanoparticles in bovine aortic endothelial cells and efficient inhibition of human vascular smooth cells using hexanoyl-ceramide (Cer-6) as a model drug. This technology is now being developed by Keystone Nano for imaging and delivery of therapeutic agents. Recently, Perez and co-workers used poly(acrylic acid) (PAA)-coated iron oxide nanoparticles loaded with paclitaxel, a potent chemotherapeutic agent, and an NIR fluorescent dye to combine optical imaging, magnetic resonance imaging, and drug delivery [181]. The nanoparticle was functionalized with a targeting

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ligand to specifically bind cancer cells. Nanoparticle size was characterized to be ∼90 nm and taxol loading was 11 molecules per nanoparticle. Taxol release kinetics were accelerated under esterase and acidic pH -mediated degradation. In addition, the results showed multimodal imaging capabilities using NIR and MRI in vitro. Other groups have reported quantum dots delivering nucleic base therapeutics, although endosomal escape and unpacking still remain a challenge. Recently, Gao and co-workers reported the surface modification of quantum dots with tertiary amine groups for the intracellular delivery of siRNA [182]. The results showed uptake and distribution of quantum dots in the cell cytoplasm. Gene silencing efficiency using formulations suspended with and without serum was better than lipofectamine and no significant toxicity of the vehicle was reported. The quantum dot gene silencing efficiency is reported to be dependent on the ratio of carboxylic group/tertiary amine groups, which control endosomal escape. Quantum dots have also been developed to monitor RNAi delivery and improve gene silencing using cationic liposome [183].

4.7 CONCLUSION Early detection is a critical component of improving patient survival in many diseases. The development of new imaging systems and imaging agents has significantly improved the ability to detect disease earlier in its progression. Nanoprobes are an emerging technology that has the potential to increase the sensitivity and specificity of imaging systems. Early designs of nanoprobes have shown great promise, but there are significant concerns over the toxicity of these probes since they are nanomaterials composed of metallic or inorganic components. Despite toxicity concerns, nanoprobes have already reached the clinic for the diagnosis of liver lesions and cancer metastases in lymph nodes. As synthesis and modification techniques improve and a greater understanding of interactions between these nanomaterials and tissues in the body is reached, the capability to design safer nanoprobes should be possible. The ability to target nanoprobes should allow their clinical applications to expand from liver lesions to other metastatic sites, such as the brain and bones. Another promising application of nanoprobes is as theragnostic agents, combining imaging and therapeutic modalities. These systems have the potential to simultaneously image and treat disease, as well as exploit synergistic cytotoxicity by combining approaches such as hyperthermia therapy with chemotherapeutic agents. Advanced designs of nanoprobes and theragnostic agents are anticipated to significantly improve patient care and compliance. The multidisciplinary approach of combining chemistry and bioengineering with toxicology and clinical research has the potential to lead to novel agents with advanced, multifunctional properties.

ACKNOWLEDGEMENT The authors acknowledge funding from National Institutes of Health Grants CA119349 and EB003647, and the David-Koch-Prostate Cancer Foundation Award in Nanotherapeutics. EMP is supported by a National Defense Science and Engineering Graduate Fellowship (NDSEG). OCF has a financial interest in BIND Biosciences and Selecta Biosciences.

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the stabilization of gold nanoparticles: in vivo pharmacokinetics and X-ray-contrast-imaging studies. Small 2007, 3, 333–341. Decuzzi, P.; Pasqualini, R.; Arap, W.; Ferrari, M. Intravascular delivery of particulate systems: does geometry really matter? Pharm. Res. 2009, 26, 235–243. Champion, J.; Mitragotri, S. Shape induced inhibition of phagocytosis of polymer particles. Pharm. Res. 2009, 26, 244–249. Champion, J. A.; Mitragotri, S. Role of target geometry in phagocytosis. Proc. Natl. Acad. Sci. U.S.A. 2006, 103, 4930–4934. Lee, H. A.; Leavens, T. L.; Mason, S. E.; Monteiro-Riviere, N. A.; Riviere, J. E. Comparison of quantum dot biodistribution with a blood-flow-limited physiologically based pharmacokinetic model. Nano Lett. 2009, 9, 794–799. Niidome, T.; Yamagata, M.; Okamoto, Y.; Akiyama, Y.; Takahashi, H.; Kawano, T.; Katayama, Y.; Niidome, Y. PEG-modified gold nanorods with a stealth character for in vivo applications. J. Control. Release 2006, 114, 343–347. Singh, R.; Pantarotto, D.; Lacerda, L.; Pastorin, G.; Klumpp, C. d.; Prato, M.; Bianco, A.; Kostarelos, K. Tissue biodistribution and blood clearance rates of intravenously administered carbon nanotube radiotracers. Proc. Natl. Acad. Sci. U.S.A. 2006, 103, 3357– 3362. Wu, X.; Liu, H.; Liu, J.; Haley, K. N.; Treadway, J. A.; Larson, J. P.; Ge, N.; Peale, F.; Bruchez, M. P. Immunofluorescent labeling of cancer marker Her2 and other cellular targets with semiconductor quantum dots. Nat. Biotechnol. 2003, 21, 41–46. Xing, Y.; Smith, A. M.; Agrawal, A.; Ruan, G.; Nie, S. Molecular profiling of single cancer cells and clinical tissue specimens with semiconductor quantum dots. Int. J. Nanomed. 2006, 1, 473–481. Schipper, M. L.; Cheng, Z.; Lee, S.-W.; Bentolila, L. A.; Iyer, G.; Rao, J.; Chen, X.; Wu, A. M.; Weiss, S.; Gambhir, S. S. MicroPET-based biodistribution of quantum dots in living mice. J. Nucl. Med. 2007, 48, 1511–1518. Bruchez, M., Jr.; Moronne, M.; Gin, P.; Weiss, S.; Alivisatos, A. P. Semiconductor nanocrystals as fluorescent biological labels. Science (N.Y.) 1998, 281, 2013–2016. Chan, W. C. W.; Nie, S. Quantum dot bioconjugates for ultrasensitive nonisotopic detection. Science (N.Y.) 1998, 281, 2016–2018. Oliver, L.; Mnica L.-G.; Christine, S.; Jorg, A.; Horst, K.; Andreas R. B. Specific integrin labeling in living cells using functionalized nanocrystals. Small 2007, 3, 1560–1565. Pathak, S.; Davidson, M. C.; Silva, G. A. Characterization of the functional binding properties of antibody conjugated quantum dots. Nano Lett. 2007, 7, 1839–1845. Jayagopal, A.; Russ, P. K.; Haselton, F. R. Surface engineering of quantum dots for in vivo vascular imaging. Bioconjug. Chem. 2007, 18, 1424–1433. Wilhelm, C.; Gazeau, F.; Roger, J.; Pons, J. N.; Bacri, J. C. Interaction of anionic superparamagnetic nanoparticles with cells: kinetic analyses of membrane adsorption and subsequent internalization. Langmuir 2002, 18, 8148–8155. Papisov, M. I.; Bogdanov, A., Jr.; Schaffer, B.; Nossiff, N.; Shen, T.; Weissleder, R.; Brady, T. J. Colloidal magnetic resonance contrast agents: effect of particle surface on biodistribution. J. Magn. Magn. Mater. 1993, 122, 383–386. Jinxue, G.; Xiao, Z.; Qingnuan, L.; Wenxin, L. Biodistribution of functionalized multiwall carbon nanotubes in mice. Nucl. Med. Biol. 2007, 34, 579–583. Liu, Z.; Cai, W.; He, L.; Nakayama, N.; Chen, K.; Sun, X.; Chen, X.; Dai, H. In vivo biodistribution and highly efficient tumour targeting of carbon nanotubes in mice. Nat. Nano. 2007, 2, 47–52.

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173. McDevitt, M. R.; Chattopadhyay, D.; Kappel, B. J.; Jaggi, J. S.; Schiffman, S. R.; Antczak, C.; Njardarson, J. T.; Brentjens, R.; Scheinberg, D. A. Tumor targeting with antibodyfunctionalized, radiolabeled carbon nanotubes. J. Nucl. Med. 2007, 48, 1180–1189. 174. Hirsch, L. R.; Stafford, R. J.; Bankson, J. A.; Sershen, S. R.; Rivera, B.; Price, R. E.; Hazle, J. D.; Halas, N. J.; West, J. L. Nanoshell-mediated near-infrared thermal therapy of tumors under magnetic resonance guidance. Proc. Natl. Acad. Sci. U.S.A. 2003, 100, 13549–13554. 175. Schwartz, J. A.; Shetty, A. M.; Price, R. E.; Stafford, R. J.; Wang, J. C.; Uthamanthil, R. K.; Pham, K.; McNichols, R. J.; Coleman, C. L.; Payne, J. D. Feasibility study of particle-assisted laser ablation of brain tumors in orthotopic canine model. Cancer Res. 2009, 69, 1659–1667. 176. McCarthy, J. R.; Jaffer, F. A.; Weissleder, R. A macrophage-targeted theranostic nanoparticle for biomedical applications. Small (Germany) 2006, 2, 983–987. 177. Jaffer, F. A.; Libby, P.; Weissleder, R. Optical and multimodality molecular imaging: insights into atherosclerosis. Arterioscler. Thromb. Vasc. Biol. 2009, 29, 1017–1024. 178. Bagalkot, V.; Zhang, L.; Levy-Nissenbaum, E.; Jon, S.; Kantoff, P. W.; Langer, R.; Farokhzad, O. C. Quantum dot-aptamer conjugates for synchronous cancer imaging, therapy, and sensing of drug delivery based on bi-fluorescence resonance energy transfer. Nano Lett. 2007, 7, 3065–3070. 179. Kester, M.; Heakal, Y.; Fox, T.; Sharma, A.; Robertson, G. P.; Morgan, T. T.; Altinoglu, E. I.; Tabakovic, A.; Parette, M. R.; Rouse, S.; Ruiz-Velasco, V.; Adair, J. H. Calcium phosphate nanocomposite particles for in vitro imaging and encapsulated chemotherapeutic drug delivery to cancer cells. Nano Lett. 2008, 8, 4116–4121. 180. Morgan, T. T.; Muddana, H. S.; Altinoglu, E. I.; Rouse, S. M.; Tabakovic, A.; Tabouillot, T.; Russin, T. J.; Shanmugavelandy, S. S.; Butler, P. J.; Eklund, P. C.; Yun, J. K.; Kester, M.; Adair, J. H. Encapsulation of organic molecules in calcium phosphate nanocomposite particles for intracellular imaging and drug delivery. Nano Lett. 2008, 8, 4108–4115. 181. Santimukul, S.; Charalambos, K.; Jan, G.; Perez, J. M. Drug/dye-loaded, multifunctional iron oxide nanoparticles for combined targeted cancer therapy and dual optical/magnetic resonance imaging. Small 2009, 5, 1862–1868. 182. Yezhelyev, M. V.; Qi, L.; O’Regan, R. M.; Nie, S.; Gao, X. Proton-sponge coated quantum dots for siRNA delivery and intracellular imaging. J. Am. Chem. Soc. 2008, 130, 9006–9012. 183. Chen, A. A.; Derfus, A. M.; Khetani, S. R.; Bhatia, S. N. Quantum dots to monitor RNAi delivery and improve gene silencing. Nucleic Acids Res. 2005, 33, e190.

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PART II

NANOPARTICLES FOR SINGLE MODALITY MOLECULAR IMAGING

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CHAPTER 5

Computed Tomography as a Tool for Anatomical and Molecular Imaging PINGYU LIU Palo Alto Unified School District, Palo Alto, California, USA

HU ZHOU Community Cancer Center of Roseburg, Roseburg, Oregon, USA

LEI XING Department of Radiation Oncology, Stanford University School of Medicine, Stanford, California, USA

5.1 INTRODUCTION X-ray and computed tomography (CT) imaging play a pivotal role in the diagnosis, staging, treatment planning, and image-guided intervention of various diseases. They are at the foundation of contemporary medical imaging and small animal imaging for biomedical research. It is estimated that 70–80% of all imaging procedures in medical applications entail the use of X-ray or CT imaging. The simplest implementation of X-ray imaging is the “plain film imaging,” and the next most common X-ray imaging approach is CT, which allows cross-sectional imaging of the body with exquisite depiction of anatomic detail. There are about 35,000 X-ray CT instruments installed worldwide for clinical applications. One of the primary advantages of X-ray imaging is the inherently simple basis of image contrast, which is the absorption of X-rays. For this reason, it has become the modality of choice in most radiology clinics. X-ray imaging can be used to image virtually every part of the body and is used for diagnosing orthopedic cases, cancer, heart disease, circulatory disease, and respiratory disease. In addition, it is also used as an aid for therapeutic interventions and as a means to follow the effects of treatment. During the last few years, innovations in X-ray detector technology have provided the capability to perform subsecond imaging of millimeter thin slices, which has opened up a host of new applications, such as blood vessel imaging, cardiac imaging, imaging of calcifications, large field imaging, and the ability to separately image the different vascular phases of medical diagnostic products. Finally, CT also plays an important role in molecular imaging of human diseases and animal model studies by providing unique and accurate spatial anatomy information. Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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In this chapter, the physical and mathematical foundations of CT imaging are first reviewed. Recent advancements in CT technologies, such as multislice CT, cone-beam CT, respiration-gated CT, and four-dimensional (4D) CT, are highlighted. Novel strategies of scatter removal, noise suppression, and dose reduction are also discussed. In the second next of the chapter, we summarize the characteristics and performance of various contrast agents for enhanced CT imaging. The success of CT imaging techniques in molecular imaging is intimately dependent on the use of X-ray contrast agents to better differentiate soft tissues and to reveal the diseased regions. In reality, while contrast media find important applications in both micro-CT small animal imaging research and clinical settings, most commercially available CT agents possess blood clearance properties that are not suitable to the time scale of small animal CT protocols. The sensitivity of the agents is also problematic. Development of novel CT contrast agents represents a forefront of micro-CT and molecular imaging research. After a brief review of conventional iodine-based contrast media, we discuss some emerging agents for disease-specific applications. Clinically, the development of disease-specific contrast agents will allow us to realize the enormous promise of molecular biology research in routine practice.

5.2 PRINCIPLE OF COMPUTED TOMOGRAPHY Although the words of “computed tomography” (CT) could have much wider meaning, the terminology is accepted to specially refer to X-ray transmission computed tomography. In this technology the projections from the X-ray transmitted through the object under investigation are mathematically processed to construct a two-dimensional image. The word tomography originates from the Greek word tomos, which means “a section” or “a cutting.” Tomographic techniques are not restricted to transmitted X-rays. In positron emission tomography (PET) the information projections are from the emitted photons instead of the transmitted ones. In optical projection tomography (OPT) the source is visible light instead of X-rays. In traditional imaging technologies such as camera or X-ray radiography the recording media are used to store the image projected onto them. In tomography the image does not exist but is mathematically constructed from indirect information by computation based on the Radon transform theory. 5.2.1 Physics of X-rays X-rays are a form of electromagnetic radiation emitted when electrically charged particles release their potential or kinetic energies in the form of photons of energy from kilo electron volts (keV) to mega electron volts (MeV), or wavelength from picometers (10−9 meters) to nanometers (10−6 meters). In CT systems the X-ray is generated in an X-ray tube. The tube contains a thermal filament and a rotational anode target in a vacuum chamber. The target is set to a positive voltage relative to the filament, usually from 30 to 150 kV. The electrons emitted from the filament are accelerated by the electric field toward the target. When these electrons strike the target, they experience electrostatic forces from the electrons and nuclei in the target and are decelerated in a short distance. According to electrodynamics, a charged particle will emit electromagnetic radiation when it is accelerated or decelerated. The kinetic energies suddenly lost from the incident electrons are released in the form of electromagnetic

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waves called bremsstrahlung radiation. The electrons which lose most of their kinetic energies in the first collision emit high-energy photons, and those electrons which lose their energies gradually emit low-energy photons. Thus bremsstrahlung radiations form a continuous energy spectrum, whose maximum energy equals the acceleration potential times the electron charge. As in most of the X-ray tubes used for diagnostics, the acceleration potential is high enough so that electrons in the inner orbits of the target atoms may be kicked out by the incident electrons, or by the photons generated by the incident electrons, and the vacancies are filled up by higher orbit electrons. In this process a photon will be emitted whose energy equals the energy difference between the two orbits. Since this process has a resonance property, the photons emitted form the characteristic energy spikes on top of the bremsstrahlung spectrum. The energies of these spikes are specified by the target material, independent of the acceleration potential. The X-ray photons then pass through a window that separates the vacuum in the tube from the ambience, and form the X-ray beam. The window usually consists of a layer of millimeter-thick low-Z metal, which also plays a role in filtration, absorbing most of the low-energy photons. The spectrum from a kilovolt X-ray tube is specified by its acceleration potential, the target material, and the filtration. Figure 5.1 shows the output spectra from a tungsten target with acceleration potentials of 80 kV, 100 kV, and 120 kV with a 2.5-mm Al filter. The surface normal of the tungsten target is 10◦ with respect to the incident electron beam. The data of a spectrum specified in this way can be generated by Monte Carlo simulations or by empirical formulas [1]. When an X-ray beam passes through a material, the photons in the beam are either scattered or absorbed. Therefore the intensity of the outgoing beam is attenuated in comparison with the incident beam. If the incident photons are not monochromic, usually the lower-energy photons are attenuated more than the high-energy photons, so the transmitted 10

Photon Number (arbitrary)

9 8 7 6 5 4 3 2 1 0

0

0.02

0.04

0.06

0.08

0.1

0.12

0.14

Photon Energy (MeV) FIGURE 5.1 Energy spectrums from an X-ray tube with 80 kV (dotted), 100 kV (dashed), and 120 kV (solid) acceleration potentials, with 2.5 mm Al filtration. The spectrums are generated with program Spektr [2]. The continuous component is from bremsstrahlung radiation and the spikes are the tungsten characteristic peaks. Note that the maximum photon energy equals the acceleration voltage.

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beam will contain a higher fraction of high-energy photons than low-energy photons than the incident beam. This effect is called “beam hardening.” Since the X-ray emitted from the target contains a large portion of low-energy photons, and these photons do not make significant contributions to the image contrast, as we will discuss shortly, a filter of a few millimeters of aluminum is usually placed in front of the X-ray tube window to reduce the low-energy component. Also, beryllium filtration is used for low-energy and copper in high-energy systems. The filtration thus changes the shape of the output spectrum. A photon in the energy range between 1 and 200 kV may experience one of three types of interactions when it passes through a material: Rayleigh scattering, photoelectric absorption, and Compton scattering. In Rayleigh scattering, or coherent scattering, the photon is bounced by an electron or an atom without loss of its energy. It is this process that causes the sky to be blue. Rayleigh scattering mainly happens with low-energy photons. In X-ray radiography and CT it can cause image blurring. Photoelectric absorption happens when an electron is knocked out from an inner orbit of an atom. The energy of the photon is completely absorbed and causes ionization. In Compton scattering the photon energy is partially transferred to an electron. The outgoing photon changes its direction and has lower energy compared with the incident photon. The probabilities of all three interactions increase with electron density and are energy dependent. As the energy increases, the probabilities of Rayleigh scattering and the photoelectric absorption reduce. The probability of Compton scattering increases at low energy, then stays nearly constant over the energy range for X-ray imaging. When a monochromic X-ray beam passes through a material, the attenuation can be described by the Beer–Lambert law: I = I0 exp(−␮x)

(5.1)

where I and I 0 are the transmitted intensity and incident intensity, respectively, ␮ is the linear attenuation coefficient (in cm−1 ), which is energy dependent, and x is the thickness (in cm) of the body through which the beam penetrated. The Beer–Lambert law can easily be understood by looking at Figure 5.2. In this figure the X-ray beam of intensity I 0 is incident on a uniform material from the left side and transmits to the right side. The beam passes through the first block with some attenuation, and we denote the beam intensity outgoing from the block as I  , and that from the second block as I  , and so on. Because every block generates the same attenuation, I  /I0 = I  /I  = · · · we have I/I 0 = (I  /I 0 )n = exp[−n ln(I 0 /I  )]. That is, the attenuation is an exponential function of the material thickness. The logarithm of the attenuation by a unit thickness of the material is the linear attenuation coefficient ␮.

I0

I'

I"

...

I

FIGURE 5.2 Attenuation of X-ray through material. The material is made by n identical blocks.

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The linear attenuation coefficient ␮ is often expressed as the product of mass attenuation coefficient (␮/␳ ) and density ␳ : ␮ = (␮/␳ )␳

(5.2)

The significance of this expression comes from the fact that for all condensed matter the mass attenuation coefficient (␮/␳ ) is an intrinsic property of the material, depending only on its elemental compositions. Therefore water and ice have the same ␮/␳ , and so do diamond and graphite. Since the mass attenuation coefficient of all the elements have been tabulated over a wide range of photon energies (NIST), once the elemental compositions of a material are known, its (␮/␳ )mat can be calculated from the mass attenuation coefficients of the component elements (␮/␳ )elem and their mass weights welem (␮/␳ )mat =




welem (␮/␳ )elem

(5.3)

elem

Mass Attenuation Coefficients (cm²/g)

The mass attenuation coefficients of water, soft tissue, aluminum, copper, lead, and tungsten are shown in Figure 5.3. The mass attenuation coefficients of all the materials are strongly energy dependent in the energy range up to 100 kV. From Figure 5.3 it can be seen that the low-energy component of X-rays, especially at photon energies below 10 kV, all the materials have very large attenuation coefficients. Therefore the component of an X-ray beam in this energy range is not useful for imaging of a body because it will be entirely absorbed after a layer of a few millimeters. In other words, nearly all the materials are opaque. On the high-energy end of the spectrum, a photon of energy higher than 250–300 kV would have long penetration depth comparable with typical body size, making all the tissues nearly equally transparent so the X-ray would not create significant contrast. These phenomena dictate that the spectrum ranges of X-ray imaging systems

10

10

10

10

10

10

4

water soft tissue Al Cu Pb W

3

2

1

0

-1

10

0

10

1

10

2

10

3

Photon Energy (keV) FIGURE 5.3 Mass attenuation coefficients ␮/␳ of water, soft tissue, aluminum (Al), copper (Cu), lead (Pb), and tungsten (W). Note that mass attenuation coefficients of water and soft tissue are almost overlapped in this figure.

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must lie between 10 and 250 kV. The upper limit of the spectrum is set by the acceleration potential of the X-ray tube and the lower limit is set by filtration. If the photon energies in the incident X-ray beam expand to a spectrum S0 (E), after passing through an object of thickness x, the transmitted beam intensity I(x) is expressed by extending Eq. (5.1) as  I (x) =

 d E S(E, x) =

d E S0 (E) exp(−␮(E)x)

(5.4)

where S(E, x) is the transmitted beam spectrum at thickness x. If the material in the object is inhomogeneous, the mass attenuation coefficient and the density vary with the location, and Eq. (5.4) is further extended as  I (x) =

 d E S(E, x) =

  S0 (E) exp −

x 0

 ␮ (E, ␰)␳ (␰)d␰ d E ␳

(5.5)

If a wide X-ray beam passes through an object, different parts of the beam pass through different parts of the body and will experience different attenuations. The spatial intensity variations then form an image projected to the record medium. This is how an X-ray radiographic image is formed. In cases where the spectrum is concentrated within a relatively narrow range of energy, the beam can be approximated to be monochromatic, and Eq. (5.5) is reduced to   I (x) ≈ I0 exp −

x 0

␮ (␰)␳ (␰)d␰ ␳



  = I0 exp −

x

 ␮(␰)d␰

(5.6)

0

or − ln

I ≈ I0



x

␮(␰)d␰

(5.6a)

0

From Eq. (5.5) and (5.6a) it is seen that each pixel in a projection is an overlap of the linear attenuation coefficients of many layers of the object. If the object is divided into many small voxels, the logarithm of the projection pixel value would be proportional to the sum of the attenuations of the voxels along the path of the ray that projects to the pixel. If the X-ray is incident on the object from a different direction, in the new projection the logarithm of the pixel values would be proportional to the sum of another group of the voxel attenuations. Therefore if the X-ray beam can be arranged to pass through the object from many different directions, the three-dimensional distribution of the attenuations of the voxels could be resolved from the projections. In 1917 an Austrian mathematician, Johann Radon, proved that it is in principle possible to reconstruct a cross-sectional image of an unknown object using an infinite number of projections through the object [3], which set the theoretical foundation of computed tomography. Recently, Liu et al. [4] proposed a higher level image reconstruction theory, the P-transform, which unified X-ray and photoacoustic CT reconstruction theories into one theory. The Radon transform becomes a special case of the P-transform. Further description of the P-transform is beyond the scope of this chapter. Readers who are interested in the P-transform theory may refer to the original P-transfer theory [5].

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5.2.2 Mathematics of Computed Tomography

Radon Transformation

The Radon transform of a function f (x, y) can be expressed as 

R(r, ␣) = R[ f (x, y)] =

f (x, y)␦(r − x cos ␣ − y sin ␣)d x d y

(5.7)

where r and ␣ are the polar coordinates of the projection, as shown in Figure 5.4a. The integration runs over the whole x-y plane. For a given value of ␣, a series of line integrals are computed along the lines r = x cos ␣ + y sin ␣, perpendicular to the vector r = (r, ␣), where each r corresponds to a line integral. Figure 5.4b demonstrates the result of the line integrals of the object in Figure 5.4a at angle ␣. Solving the function f (x, y) in Eq. (5.7) from known projections R(r, ␣) is called image reconstruction. There are many algorithms performing the function of image reconstruction.

Principle of Image Reconstruction mation takes the form

When applied to X-ray CT, the Radon transfor 

− ln[I (r, ␣)/I0 ] = R(r, ␣) = R[␮(x, y)] =

␮(x, y)␦(r − x cos ␣ − y sin ␣)d x d y

(5.7a) where ␮(x, y) is the linear attenuation coefficient at point (x, y), as in Eqs. (5.6) and (5.6a). Rotating the coordinate system (x, y) for an angle ␣ about its origin, so that x = x cos ␣ + y sin ␣ and y = y cos ␣ − x sin ␣ become the new coordinate system, Eq. (5.6) is then rewritten as    (5.8) − ln[I (x , ␣)/I0 ] = R(x , ␣) = ␮(x, y) dy 

FIGURE 5.4 The Radon transformation of a cross section of an object. In this illustration f (x, y) is uniformly distributed inside the square in (a) so that in (b) the values of the Radon transformation are proportional to the lengths of the line segments across the object.

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Taking the Fourier transformation of Eq. (5.8) with respect to x , we have  FR (k, ␣) =

∞

−∞



R(x  , ␣)e−i2␲kx d x  =



∞



−∞

␮(x, y)e−i2␲kx d x  dy 

(5.9)

The linear attenuation coefficient ␮(x, y) can be solved by inverse Fourier transformation as 



␲

␮(x,y) =

∞

d␣ 0

−∞

FR (k, ␣)e

i2␲kx 

 |k|dk =



␲

∞

d␣ 0

−∞

FR (k, ␣)ei2␲k(x cos ␣+y sin ␣) |k|dk (5.10)

Equation (5.10) is the back projection formula. With this equation the linear attenuation coefficient can be calculated from the projection data to form a cross-sectional image. A series of such cross-sectional images along a longitudinal axis are then stacked together to form a three-dimensional (3D) image. Therefore the essential step is the two-dimensional (2D) reconstruction of the cross-sectional image from the projected images. Although in principle Eq. (5.10) can be used to solve the problem of reconstruction, Eq. (5.7a) that relates ␮(x, y) to the projected image I(r, ␣) is much simplified. This equation assumes the X-ray is monochromic or can be approximated to be monochromic, and the projection is free of noise. In modern CT systems more sophisticated mathematics are used in the reconstruction, which will be discussed in more detail as a specific topic.

5.3 EVOLUTION OF CT IMAGING TECHNOLOGY The basic components of a CT system includes an X-ray source, a detector or an array of detectors, a collimator that shapes the X-ray beam, a gantry carrying the source and detectors, a patient couch, and one or more computers that control the system, acquire the data, and reconstruct the image. In medical CT systems the computers are also responsible for administration, data archiving, data transferring, and image display.

5.3.1 The First to Fourth Generations of CT Scanners The first CT scanner was designed by Godfrey Hounsfield in the early 1970s by exactly realizing the Radon transformation shown in Figure 5.4 and Eq. (5.7). In the system an X-ray source and a single detector were set at the opposite sides of the object. The X-ray was confined to a narrow beam (pencil beam) aiming at the detector. The source–detector pair was moved together so that the object was scanned by parallel beams. The pair was then rotated a small angle and the procedure was repeated to scan the object at a different angle. This arrangement is shown in Figure 5.5a. The data acquisition ran through the angles over 360◦ . Hounsfield and A. M. Cormack shared the 1979 Nobel Prize in Medicine and Physiology for this invention. In a first generation CT scanner, the X-ray is confined to a narrow pencil beam, and the single detector senses only the primary photons. Because the scattered photons are going in different directions, they are ignored. The scattering reduces the intensity of the signal reaching the detector but this amount can be compensated by calibration.

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(a)

(c)

115

(b)

(d)

FIGURE 5.5 Diagrams of four generations of CT: (a) first, (b) second, (c) third, and (d) fourth generations. Grey spheres represent X-ray sources.

The second generation of CT is shown in Figure 5.5b, in which the single detector is replaced by a linear array of detector elements. The X-ray was now confined in a fan beam so that all the detector elements can acquire data simultaneously. The source and the detector array translate to cover the subject being imaged, then rotate to the next angle. The data acquisition speed is multiplied because several detector elements work simultaneously, and the X-ray source is used more efficiently. But there is a price for these advantages: the noise due to scattering caused by using a wider X-ray beam. It is possible to add a grid in front of the detector array to reduce the scattering noise [6]. However, this solution has limitations because the grid not only adds extra weight to the detector but also takes away the detector area, thus reducing the efficiency. Because the wall between adjacent grid openings must be thick enough to create sufficient attenuation to the oblique photons, these walls will occupy detector area. In high spatial resolution detectors, the area taken by the grid wall becomes significant so that the benefit of noise reduction brought from the grid could be outweighed by the reduction of the element areas. In the third generation of CT, a longer array of detector elements is used and the X-ray from the single source is confined to a wider fan beam so that the whole width of the object can be covered. Therefore the translation is no longer necessary, and only the gantry that holds the source and detector array needs to rotate. The detectors are arranged on an arc, as shown in Figure 5.5c. Most modern clinical CT systems use this arrangement. The noise level of the third-generation CT scanner is higher than the second-generation scanner if all the components are the same, because of the wider X-ray beam angles. But

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the data acquisition speed is much faster, which brings tremendous benefit in medical applications, especially since it makes use of higher spatial resolution with smaller detector element area. Since the first CT scanner appeared in the 1970s, major efforts in CT engineering have been focused on increasing the data acquisition speeds and the spatial resolutions. This is discussed in the next sections. The fourth-generation CT has a set of detectors that are arranged in a ring over 360◦ around the object, and the source is mounted on a rotational gantry, as shown in Figure 5.5d. The detectors are on a fixed ring. Therefore the wiring can be greatly simplified compared to third-generation CT. The fixed relative positions of source and detectors in the third generation are no longer here. It is possible in principle to read out all the detector elements even if most of them are not irradiated, but such a procedure would slow down the acquisition speed and add work load to the reconstruction software. Therefore extra electronics and mechanical components are needed to synchronize the data acquisition—a complication that restricts the application of this generation. 5.3.2 The Fifth Generation of CT Scanner In CT history the fifth generation of CT is usually referred to as the electron beam CT, or the EBCT. The EBCT was the result of applying CT in a cardiology study, which requires a CT with higher temporal resolution. Research showed that in order to get a clear tomogram of the fastest moving part of a living heart, a complete data acquisition time should be no greater than 19.1 ms [7]. It’s obvious that there is no CT with mechanically rotating gantry that can reach this speed. An EBCT scanner can reach the goal by replacing the heavy X-ray tube with an electronically steered X-ray source. In an EBCT scanner, the X-ray source consists of a semicircular tungsten anode ring housed in a huge vacuum chamber with an electron gun, as shown in Figure 5.6. An electronically steered high-intensity electron beam generates a fast moving X-ray source along the tungsten target. Among various EBCT designs, GE’s eSpeed scanner is the fastest CT scanner in the world. Compared to other CT systems, EBCT technology has its own shortcomings, such as higher capital investment and lower signal-to-noise ratio (SNR). Currently, about 130 EBCT systems are in service worldwide. 5.3.3 CT Scanner with Nanotube X-ray Sources With the emergence of an X-ray source based on the field emission effect from the tip of a carbon nanotube [8], a new type of CT system was proposed by Alexei Ramotar in 2006 [9]. In this design the fixed detectors are arranged in a ring over 360◦ around the object as in the fourth generation, but the rotational single source is replaced by a fixed ring of many sources as shown in Figure 5.7. The sequence that turns the sources on and off is controlled by electronics. If the X-ray beam intensity of the nanotube source is sufficient, a scanner designed from this concept will completely omit the mechanical complication of the rotational gantry, and the data acquisition can be of very high speed and in more flexible patterns than just a rotational sequence. An example of possible applications of the technique is breast CT for mammograms. Compared with the current widely used breast-compression mammogram technique, breast CT offers patients more comfort and is highly sensitive in detecting early stage tumors [10]. There are restrictions to the application of traditional CT to mammograms, one of which is the consideration of radiation dose to the lungs and the heart, because of the geometry of

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Focus, bend & steering coils

Electron gun

X-rays Shielding

Detector rings

Patient Electron beams Vacuum chamber

Tungsten anode rings

Fan x-rays

FIGURE 5.6 Side view (top) and cross view (bottom) of diagram of an EBCT scanner. (From Pingyu Liu’s 1997 proposal with author’s permission.)

the breast and the limited space between two breasts. Using the new technique the detector elements could be arranged on a spherical surface around the breast in layers parallel with breast base, for left and right breasts, respectively. The sources could be arranged on the same spherical surface with the beams collimated in the corresponding layers, with limited dose to the chest wall underneath.

FIGURE 5.7 Diagram of CT scanner with nanotube X-ray sources. The little dits represent the detectors and the black lines represent the X-ray sources. In this arrangement both the sources and detectors are fixed so that the rotation gantry can be completely omitted. The emissions from the sources are controlled by electronics so the data acquisition speed can be very high and the emission can be in flexible patterns other than a rotational sequence. In the figure one of the sources is collimated so that the X-rays are confined to certain detectors.

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5.3.4 Power and Data Transmission in a CT System Power and data transmission between the stationary components and the rotational gantry of a CT system can be realized in two possible ways. Conventionally, transmission is through cables wrapped on a rack. When the gantry rotates in one direction, the cables wind up on the rack. When the gantry rotates in another direction the cables unwind. Because the cable cannot be very long, the gantry needs to rotate alternately back and forth. This technique is used in early CT systems and most modern cone-beam CT systems, where a slow gantry speed can be tolerated. In most modern CT scanners a slip ring technique is used to transmit the power and data through contact brushes and slip rings. In some systems the data are transmitted through RF couplings, or fiberoptic rotary joints. In other CT systems a more flexible optical data coupling method is utilized. An important specification of a CT system is its gantry rotation time, which determines system’s temporal resolution. In a modern CT system, gantry rotation speed can reach 0.35 second per cycle. On the rim of a gantry 2 meters in diameter, the centrifugal acceleration can reach 24g. Therefore on the gantry the components must be carefully arranged to balance the momentum. 5.3.5 Helical CT If the patient couch moves simultaneously when the gantry is rotating, the patient will see the source moving along a helical path. This arrangement is called the helical or spiral CT. With slip rings the gantry can rotate continuously so the helix can extend over the length of the whole body. In third-generation CT with an array of detectors, the distance the couch moves during one gantry rotation period defines one slice thickness, which is the spatial resolution of the image along the axial direction. This technique is used in most modern medical CT systems, because it allows fast and high-resolution imaging. Fullbody CT imaging using such a system with 0.625-mm slice thickness can be completed in 1.5 minutes, and the image reconstruction can be completed on the fly. Fast imaging speed is important for medical applications. A shorter-time scan reduces the artifacts caused by patient motion. Many CT contrast agents stay in the body only a short period of time (a few minutes) and then are washed out by the urinary system. An imaging system must complete the data acquisition within that period. Most importantly, although all the CT systems are able to provide stationary anatomical information, a fast-imaging system can also provide dynamical and functional information. In cases where motion is periodic, as the respiration or heartbeat, the acquisition and reconstruction of time-resolved or 4D CT images can be accomplished with gating technology. During gated acquisition, the CT runs in scene mode, in which the couch moves step by step instead of continuously. In each step the gantry rotates several revolutions; each corresponds to a phase. In this mode the patient sees the source moving along circles. As the couch steps through the body of the patient, these circles stack to form a scan. The projections are then sorted according to the phases of the motion, and those belonging to a phase are picked up for reconstruction. When displayed, the reconstructed images are shown phase by phase as a 3D movie. Most helical CT systems can be operated in scene mode. 5.3.6 Multidetector CT The acquisition speed and spatial resolution of helical CT can be further improved by adding one or more rows of detectors to the array, so that during one gantry revolution more than

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one slice of data can be collected at the same time. Multiple-slice CT, also referred to as multidetector CT (MDCT), makes a narrower slice thickness, more slices, and larger volume coverage simultaneously. MDCT scanners with 320 slices are now available in the market. There are many technical concerns in the design of MDCT systems. Because of the scattering effect, the MDCT data contains higher noise compared with single-slice CT. The increase in the number of slices is accomplished by reducing the row width of the detectors, and the reduction of the row width is limited by the detector sensitivity, which is proportional to the detector element area. Most manufacturers restrict the total detector width to 20 or 24 mm when the number of slices is 2, 4, 16, 20, and 24, and to 40 mm when the number of slices is 40 and 64. MDCT delivers the same X-ray dose to the object as single-slice CT for the same spatial resolution, but to get the same signal-to-noise ratio MDCT needs a higher dose because of the higher noise level. Especially, to take advantage of the high spatial resolution of MDCT, a significantly higher dose would be necessary to create a sufficient signal level from the smaller area of the detector elements. Many manufacturers have developed dose control mechanisms to automatically optimize the dose for each slice of a scan. The image reconstruction time of MDCT is expected to take longer than that of singleslice CT due to the larger amount of data. However, because of improved computer technology the reconstruction time does not significantly delay the total scan. The algorithm reconstructing on-the-fly can complete a slice reconstruction within half a second after acquisition. The fast speed and high spatial resolution brought by MDCT makes dynamic imaging of moving organs possible. An example is functional vascular imaging. In CT coronary angiography, cardiac data can be acquired continuously during 5 or 6 heartbeats. The dynamic images of the moving heart can be reconstructed corresponding to electrocardiographic gating signals. In polytrauma applications, the whole-body scan over 2 meters along the patient body can observe an upper extremity angiogram followed by pelvic and lower extremity angiography with one bolus of contrast. 5.3.7 Cone-Beam CT (CBCT) Replacement of a pencil beam and a single detector by a fan beam and a 1D detector array evolves CT from the first generation to the second generation, and extending the array size further evolves CT to the third generation [11, 12]. The array can also be extended in another direction to form a two-dimensional detector array. To irradiate the two-dimensional array the beam needs also to form a cone shape. In many cone-beam CT systems the detector areas are large enough that the beam can cover the whole region of interest in the body, so that during the acquisition of data the object does not have to move. These systems usually connect the stationary components and the gantry through cables, and rotate only one revolution for a scan. Typical scan time over 360◦ is about 1 minute, comparable to the acquisition time with helical CT. Compared to helical CT, cone-beam CT has a much simpler structure. The cable connection does not require high precision machining as the slip ring does; therefore the cost is significantly reduced. Also, cone-beam CT can be integrated into special systems such as a radiation treatment machine without interfering with the primary function of the system. The noise level of cone-beam CT is higher from 2D scattering. To obtain the same signal-to-noise level, cone-beam CT needs to deliver a several times higher dose to the object than helical CT. Also, a large-area 2D detector introduces another challenge—the

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data transmission time. A detector with 4000 × 3000 pixel size needs about 1 second to read out; therefore the frame rate cannot be faster unless the image pixels are combined in 2 × 2 binning or 4 × 4 binning. 5.3.8 Scattering Removal and Noise Reduction in CBCT In single-slice CT the X-ray beam and the detector span a narrow plane. Most of the scattered photons fall outside the plane. Only a small portion of the photons scattered in the plane are recorded. In cone-beam or multidetector CT the X-ray is collimated to a wide beam. The possibility of a scattered photon mixing with the primary photons is therefore much higher. The projection images from cone-beam or multidetector CT are therefore much noisier than those from single-slice CT, and the uncertainty in the reconstruction is correspondingly higher, which brings higher noise to the reconstructed image. Research in scattering removal techniques has shown that correction of the scattering could be carried out in several ways. If the CT systems allows object–detector distance variation, setting the distance larger helps to reduce the scattering effect, because the primary beam goes along the path defined by the desired geometry and the scattered photons move in large angles. The scattering effect can be measured by setting a small opaque shield between the source and the object, to block the primary beam from reaching the detector, so that the signal in the shielded area is all from the scattering from the object. A deconvolution algorithm in image processing based on a scattering model can be used to separate the primary beam and the scattered components in projections or sinographs. Monte Carlo simulations can fit the images as the summation of the primary beam and the scattered component. The detailed descriptions of these techniques can be found in the work by Zhu et al. [13] and the references quoted. A typical method to reduce the noise is to average a pixel in an image with its neighbors. This method also reduces the spatial resolution because it smoothes out the edges. In the work reported by Wang et al. [14], an algorithm was proposed that uses an anisotropic filter to reduce the noise. In this algorithm a pixel in a sinograph is replaced by the weighted average of the neighboring pixels. The weight is proportional to the exponential of difference between the pixel under consideration and its neighbor, so that the larger the difference the less the weight, to conserve the edges. After several iterations of such averaging the random noise can be significantly suppressed while the features remain. Numerical experiments using phantom images showed that faint features and a small spot originally flooded in noise in the original images can clearly be seen in the processed images. Advances in scattering removal and noise suppression techniques make it possible to use a lower dose to obtain sufficient image qualities in cone-beam CT and multidetector CT. The cost of using the techniques is a longer after-processing time.

5.4 SPECIAL CT SYSTEMS FOR DIFFERENT CLINICAL AND RESEARCH APPLICATIONS 5.4.1 Cardiac CT Since the advent of the first CT system for head imaging, scientists and engineers have been working on a more challenging task, cardiac CT. The challenge here is tremendous, and it stems from the restless, fast, and complex three-dimensional motion of the heart. To

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FIGURE 5.8 Dynamic Spatial Reconstructor at the Mayo Clinic in Rochester, Minnesota (http://dpi.radiology.uiowa.edu/gallery/dsr.html).

get a clear CT image of a living heart requires significant advancement in both hardware and software of routine CT scanners. The most significant example of endeavor in hardware improvement for cardiac CT should perhaps be the Dynamic Spatial Reconstructor (DSR), which was installed at the Mayo Clinic in 1983. In CT history the DSR is known for its speed, weight, size, and complexity. The DSR was designed mainly for cardiac imaging. It has the ability to obtain up to 240 contiguous 0.9-mm thick sections in a time period as short as 1/60 second and to repeat this acquisition rate 60 times per second. The DSR consists of a gantry weighing approximately 17 tons with a length of 20.5 feet and a diameter of 15 feet. Fourteen X-ray guns reside in a hemicylindrical configuration. Figure 5.8 shows an artistic view of the DSR. Following the DSR, the next milestone in cardiac CT development was electron beam CT. All currently running EBCT systems are mainly for cardiac imaging. All main CT system manufacturers pushed out their fastest CT systems for cardiac applications, including a 64-slice detector with 0.33 second of rotation time. These newer CT scanners generate cardiac images with higher temporal and spatial resolution. Smarter data acquisition, preprocessing, and image reconstruction techniques have become increasingly available for cardiac CT images. A recent development of these techniques is the Matched Cardiac X-ray CT (MCCT). Figure 5.9 shows a diagram of the MCCT. Clinical studies have showed that reasonably good cardiac CT images can be obtained with the MCCT for most medical conditions with three heartbeat in routine CT scanners with a gantry rotation time of 1 second. 5.4.2 Micro-CT for Animal Imaging Experimental animals, especially small animals such as mice and rats, are increasingly being used to model human diseases [16] in the study of specific pathways for disease, potential therapies, and safety of new pharmaceuticals. In traditional techniques,

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FIGURE 5.9 Diagram of the MCCT (U.S. Patent 6233478, Apparatus and Method for Constructing Computed Tomography Image Slices of an Object Undergoing Cyclic Motion, May 15, 2001. Pingyu Liu) [15].

observation of disease development is based on the sacrifice and dissection of the studied animals. Such approaches are labor intensive and don’t permit longitudinal studies. The development of in vivo imaging technology in small animals revolutionarily altered the situation. The special challenge in animal imaging comes from the small size of the animals. The typical size of a mouse is a few centimeters, which is about 5% of human size (Fig. 5.10). The respiration period of a mouse is about 10% of that of human beings. A CT system designed for small animals, the micro-CT system, therefore requires both spatial and temporal resolutions an order of magnitude finer compared with a human system. A micro-CT scanner is not a simple scaling of a human CT scanner. Special considerations have to be taken into account from high resolution requirements and the associated radiation dose increase to the imaged animals. A micro-CT scanner can rotate either the object or the gantry. In rotate-object animal micro-CT systems, to avoid organ movement during rotation, the animal is set in the vertical position. This position is unnatural for animals and it induces biological effects. Most animal micro-CT systems rotate gantries, in which the animal is set in the more natural horizontal position. Cone-shaped X-ray beam, two-dimensional detector, and one-revolution gantry are most commonly used in animal CT systems [17], although slip-ring CT systems are also available in the market [18, 19]. The detector sensitivity is proportional to pixel size. The high spatial resolution of a micro-CT system requires much smaller pixel size than the detector on a human system. A typical pixel size of micro-CT detector is 50 ␮m. The small area of the detector pixel requires higher flux density of the X-ray beam to obtain significant signal-to-noise ratio, which could result in a dose delivered to the experimental animal that is significant enough to alter its biology. In some systems if the accumulated imaging time is several hours the total dose delivered to the animal could reach the lethal level.

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FIGURE 5.10 A micro-CT image of a mouse head, acquired using a GE micro-CT with 2 × 2 binning. The X-ray tube acceleration voltage is 70 kV and the tube current is 25 mA.

A special technical barrier of a two-dimensional X-ray detector is its read-out time, as described for cone-beam CT. The development of high-speed electronics is therefore urgently required. One of the factors influencing spatial resolution is the X-ray source focal spot size. The spot size of animal micro CT varies from 25 to 900 ␮m. If the projection of the focal spot to the detector plane has a size comparable to or larger than the detector pixel size, a penumbral blurring of the image will occur. On the detector the penumbra width of a sharp-edged object is Penumbra width =

ODD × spot size SOD

where ODD is the object–detector distance and SOD is the source–object distance. On the other hand, for the object we can estimate its image size on the detector surface:   SDD ODD × object size Image size = × object size = 1 + SOD SOD Where SDD = SOD + ODD is the distance from the X-ray source to the detector. Choosing smaller ODD and longer SOD is helpful in obtaining a sharper image through reduction of the penumbral blurring, if the image resolution is sufficient. Smaller source spot size can produce images with higher spatial resolution. However, it restricts the electron beam current because of the limit of the heat dissipation rate on the

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anode. A lower electron beam will generate lower X-ray fluence; therefore a longer imaging time would be required in order to create enough signal above the noise floor. As in the human system, short imaging time might be crucial in special cases as in respiration-gated imaging and when some contrast agents are applied whose wash-out time is within one or a few minutes. In practice, the selection of the spot size must be a compromise between the requirements of the X-ray fluence and the spatial resolution. 5.4.3 Dual-Modality Systems A CT system can provide detailed anatomic information for a body. Functional imaging systems such as positron emission tomography (PET) and single photon emission computed tomography (SPECT) are capable of providing dynamic metabolic information for the body, but not detailed information about the body’s anatomy. To associate the spatial distribution of the metabolic activities with the anatomic structure the subject can be scanned using CT and PET or SPECT separately, and then the two images can be combined through image registration. The subject must not move in -between the two scans. With a dual-modality system combining CT and PET or SPECT, such registration becomes simple. Also, dualmodality systems used together with molecular-specific agents play important roles in molecular imaging studies and applications, as discussed in more details in Section 5.5. In a PET–CT dual-modality system the two units are placed next to one another [20]. Figure 5.11 shows a Siemens Inveon microPET/CT imaging system with a bore size of 10 cm in diameter. Its main study objects are small animals. Note that in the picture a dedicated PET imaging system is in front of a CT imaging system. The animal couch moves through one modality to another. The system is capable of imaging a variety of radionuclides such as 18 F, 15 O, 13 N, and 11 C. The spatial resolution for PET imaging is 1.4 mm, and the X-ray CT system is on the order of 50 ␮m and can achieve a resolution of 15 ␮m depending on the size of the specimen . A PET scan is about 7–14 min, and a CT scan is 10–12 min. The two images are reconstructed separately. The user can fuse them or display them individually.

FIGURE 5.11 Siemens Inveon microPET/CT. Its bore size is 10 cm in diameter and can accommodate animals up to the size of a small rabbit.

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SPECT uses radiation isotopes having half-life times of several hours to a few days and emitting ␥ -rays of about 100 keV, most commonly 99m Tc, 111 I, and 123 I. Glucose or drugs coded with these isotopes are injected into the patient and follow the blood flow to the target organ. The ␥ -rays are detected by a gamma camera, which consists of a lead or tungsten collimator and an array of photon multiplier tubes. Since a gamma camera can cover only a small solid angle, usually two to four such cameras are used in SPECT. During the CT scan, the SPECT system is turned off and shields are moved to the front of the highly sensitive gamma cameras, to avoid bombarding the photon multiplier tube with the high-intensity X-ray beam. When the SPECT system is on, the X-ray tube of the CT system is off to avoid possibly scattering photons into the gamma camera. 5.4.4 Dual Energy CT As shown in Eq. (5.5), the X-ray attenuation by a material depends on the density ␳ and the mass attenuation coefficient ␮/␳ of the material. In some applications only the geometry information is required, as in bone fracture diagnostics, in which attention is paid to increasing the contrast between different structures. In other applications, as in radiation treatment planning, both the densities and elemental components of tissues are important because the dose absorption is defined by the two. In regular conditions the density and elemental compositions are related and an empirical tissue model can be applied to find them from the CT numbers. A CT system with two energy spectra, on the other hand, can provide more direct information about the material properties [21]. In dual energy CT the same object is imaged twice with different X-ray tube acceleration potential settings, say, 80 and 140 kVp, respectively. The tube current of lower energy is usually adjusted higher than that of the higher energy so that the contrasts of the two images can be comparable. An algorithm proposed by Torikoshi et al. [22] modeled the linear attenuation coefficient ␮ of a CT voxel as ␮ = ␳ e [Z 4 F(E, Z ) + G(E, Z )]

(5.11)

where ␳ e Z 4 F(E, Z) and ␳ e G(E, Z) denote a photoelectric term and a scattering term, respectively, which can be derived from quantum mechanics; ␳ e is the electron density; Z is the atomic number; and E is the photon energy. Measuring the linear attenuation coefficients with two different energy X-rays, we obtain simultaneous equations with respect to the unknown variables of ␳ e and Z as follows: ␮(E 1 ) = ␳ e [Z 4 F(E 1 , Z ) + G(E 1 , Z )] ␮(E 2 ) = ␳ e [Z 4 F(E 2 , Z ) + G(E 2 , Z )]

(5.12)

Therefore if the X-ray beam is monoenergetic, Z4 =

␮(E 2 )G(E 1 , Z ) − ␮(E 1 )G(E 2 , Z ) ␮(E 1 )F(E 2 , Z ) − ␮(E 2 )F(E 1 , Z )

(5.13)

and the effective value of Z can be solved numerically. If the X-ray beam energy expands to a spectrum, the equation becomes [23]  [S1 (E)Z 4 F(E, Z ) + G(E, Z )]d E ␮1 =0 (5.14) − ␮2 [S2 (E)Z 4 F(E, Z ) + G(E, Z )]d E

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where S1 (E) and S2 (E) are the spectra of the two energy settings, respectively. Since all the functions in Eq. (5.14) are known, the ratio ␮1 /␮2 can be tabulated as the function of effective Z. Thus the Z image can be obtained from the CT images of two energy settings. Then by either form of Eq. (5.12) the electron density ␳ e can be mapped out. The capability of differentiating the effective Z of different tissues in a dual-energy CT image makes it possible to display clear soft tissue contrast. The challenges come from the facts that the effective Z solved from the Eq. (5.14) is highly sensitive to the uncertainty of ␮1 /␮2 , and the spectra of the two energy beams need to be known in high accuracies [40]. Dual-energy CT has been available in market, as introduced by Reference [41] for example. With these systems it is possible to “remove” out some structures, such as bone, from the reconstructed images and display the blood vessel only, which brings significant advantages in the clinical diagnostics.

5.5 CT CONTRAST MEDIA AND MOLECULAR CT The intrinsic contrasts of different types of tissues are often not sufficient for radiologists to make definitive diagnoses or to accurately determine the extent of diseased tissue or certain normal structures (e.g., blood vessels or prostate gland) in X-ray and CT imaging. Various contrast media have been developed over the years and used along with the X-ray or CT imaging. There are two types of X-ray contrast agents currently approved for human use: barium sulfate suspensions, which are used for GI tract imaging, and water-soluble aromatic iodinated contrast agents, which are used primarily for imaging of the blood vessels and/or urinary system. It is estimated that the annual uses of barium suspensions and iodinated media in the United States are 5 and 20 million, respectively. Despite the tremendous success of these iodinated contrast media, the search for better contrast agents has never ceased. The major problems with existing contrast media are that they have very low retention rate and are not tissue specific. In some rare cases, the iodinated media can be severely toxic, as manifested with cardiovascular, anaphylactic, and pain reactions. A novel contrast agent that circumvents these problems is thus highly desirable. In recent years, due to the advancement of nanotechnology, some new contrast agents have emerged and showed their promising future. The field of nanotechnology has experienced significant advance in synthesis, characterization, and novel applications for various nanoscale devices. One of the important areas of nanotechnology is the synthesis and characterization of polymer, metallic, or magnetic particles, which exhibit a variety of novel properties and functions with the possibility of molecular sensing and imaging. Interfacing nanotechnology, biotechnology, and medicine has opened new vistas for biomedical research and provided an unprecedented opportunity to achieve a fundamental understanding of biological processes and contribute to disease prevention, detection, and therapy. In the last decade, the magnetic nanoparticles for MRI T2 enhancement have been extensively studied and the results showed significant promise for clinical applications. In parallel, the potential of using metallic nanoparticles has been employed to augment the intrinsic contrasts of X-ray and CT imaging. While it is a known fact that heavy elements are efficient absorbers of X-rays, the use of metallic nanoparticles as contrast media has yet to be achieved. X-ray CT is among the most convenient and widely used imaging tools in hospitals today. However, in contrast to magnetic resonance imaging (MRI) and various nuclide imaging modalities, CT is generally not considered as a molecular imaging modality since targeted and molecularly specific contrast agents have not been fully developed. But this situation is changing. Popovtzer’s group synthesized gold nanorods (AuNRs) and conjugated them

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CT attenuation (HU)

HU (Hounsfield units)

240 200 160 120 80 40

La

ra l

O

ry

nx

ca

nc

er

+A La uN ca ry R nc nx +A er ca +A 9 nc u N O er R ra (w +A lc ith 9 an La ou c ry er tA nx (w uN ca ith R nc ou ) O er tA ra + uN lc Au an R N ) R ce +K r+ H Au R Fi N I-3 R br +K ob H la R st M I-3 G +A el ol u a d N n na om R +A no a+ 9 ro Au ds N in R +A wa 9 te rs ol ut io n W at er

0

FIGURE 5.12 CT attenuation (HU) of SCC head and neck cancer cells and positive and negative control samples. Bar graph with standard deviation of three samples: larynx and oral cancer cells that were targeted with A9 antibody-coated gold nanorods (AuNRs), larynx and oral cancer cells without gold nanorods, and larynx and oral cancer cells targeted with nanorods that are coated with nonmatching antibodies (KHRI-3); normal fibroblast and melanoma cells targeted with A9 antibodies, bare gold nanorods in water solution (2.5 mg/ mL), and water [21].

with UM-A9 antibodies, which home specifically to squamous cell carcinoma (SCC) head and neck cancer [24]. Two SCC human head and neck cancer cell lines (106 cells/mL) were used: oral cancer UM-SCC-1 and larynx cancer UM-SCC-5. Both cancerous cell lines were shown before to have a significant overexpression of the A9 antigen. CT imaging was performed on the SCC cells, which were targeted with the UM-A9 antibody-coated gold nanorods. When compared with the CT number of control samples, they found that the attenuation coefficient for the molecularly targeted cells is over 5 times higher than for identical but untargeted cancer cells or for normal cells, as shown in Figure 5.12. Kim et al. [25] also utilized gold nanoparticles (GNPs) as a contrast agent for X-ray CT imaging. They prepared uniform GNPs (similar to 30 nm in diameter) by general reduction of HAuCl4 by boiling with sodium citrate. The resulting GNPs were coated with polyethylene glycol (PEG) to impart antibiofouling properties, which extends their lifetime in the bloodstream. Measurement of the X-ray absorption coefficient in vitro revealed that the attenuation of PEG-coated GNPs is 5.7 times higher than that of the current iodinebased CT contrast agent, Ultravist. Furthermore, when injected intravenously into rats, the PEG-coated GNPs had a much longer blood circulation time (>4 h) than Ultravist (<10 min). Consequently, CT images of rats using PEG-coated GNPs showed a clear delineation of cardiac ventricles and main vessels. On the other hand, relatively high levels of GNPs accumulated in the spleen and liver, which contain phagocytic cells. Intravenous injection of PEG-coated GNPs into hepatoma-bearing rats resulted in a high contrast (similar to two-fold) between hepatoma and normal liver tissue on CT images. These results suggest that PEG-coated GNPs can be useful as a CT contrast agent for a blood pool and hepatoma imaging.

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FIGURE 5.13 An artistic view of a fullerene model of C60 with a heavy atom (center ball) inside the cage. (Reproduced from http://en.wikipedia.org/wiki/Endohedral fullerenes.)

Besides gold nanoparticles, other heavy metal nanoparticles have also found application in X-ray CT. Fullerenes are a new carbonic allotrope having a cage structure. Miyamoto et al. [26] investigated whether fullerenes containing one or two atoms of heavy metals could be used as X-ray contrast material. One or two atoms of dysprosium (Dy), erbium (Er), gadolinium (Gd), europium (Eu), and lutetium (Lu) were encapsulated into fullerene (C82 ), which was synthesized as a polyhydroxyl form (e.g., Gd@C82 (OH)n , n = 40, Gd— fullerenols). A model of a fullerene with one metal atom inside its cage is shown in Figure 5.13. It was found in their study that the concentrations and the CT numbers were relatively low for all solutions, and the correlation was not clear between the concentration and the CT number of the solutions. The group, however, was still optimistic about the approach and stated that “if nanotechnology progresses in the near future, it may prove to have a possibility as an X-ray contrast material.” Rabin et al. [27] developed long-circulated bismuth sulfide nanoparticles (BPNPs) as an X-ray CT imaging agent. Rabin’s group reported that they made use of a polymercoated Bi2 S3 nanoparticle preparation as an injectable CT imaging agent. This preparation demonstrates excellent stability at high concentrations (0.25 M Bi3+ ), high X-ray absorption (fivefold better than iodine), very long circulation times (<2 h) in vivo, and an efficacy/safety profile comparable to or better than iodinated imaging agents. They showed the utility of these polymer-coated Bi2 S3 nanoparticles for enhanced in vivo imaging of the vasculature, the liver, and lymph nodes in mice. CT values of different BPNP concentrations from their experiments are shown in Figure 5.14. A new nanoparticulated iodine contrast agent was prepared and tested in cardiovascular research with X-ray CT. A research team led by Fayad developed a new technique for early detection of high-risk plaque in coronary arteries using a contrast agent called N1177 containing iodinated nanoparticles [28]. High-risk plaque is characterized by its cellular and biologic structure. High-risk plaque rich in macrophages or cells can rupture, eventually causing a heart attack or stroke. Early identification of high-risk plaque in coronary arteries may be useful in preventing cardiac events but one major hurdle in detecting high-risk plaque

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2.36 M iodine (5,100)

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0.55 M bismuth

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2,000 PBS BPNP Iodine 1,000 Water (0)

0

Air (–1,000)

–1,000 0.001

0.01

0.1

1

Bismuth concentration (M)

FIGURE 5.14 X-ray absorption of BPNPs. Experimental CT opacity of bismuth solutions as a function of concentration. The horizontal dashed lines indicate the opacities of air, water, and 2.36 M (300 mg I/mL) iodine contrast agent, for comparison. Inset: X-ray fluoroscopy of tubes containing PBS (phosphate buffer saline), 0.5 M Bi2 S3 nanoparticle suspension (BPNPs), and 2.36 M iodine contrast agent (Iopromide). The error bars represent the standard deviation in the CT value of a set of 30 × 30 × 1 (900) boxels on a planar square aligned with the center of the sample receptacle [24].

is the lack of an imaging modality that allows physicians to see the composition of dangerous plaque. They tested the contrast agent N1177 for the detection of macrophages in an animal model with 64-slice CT. High-risk plaque in this animal model contained high levels of macrophages that are similar in size and content to human coronary plaques. Researchers compared the enhancement of macrophage-rich plaque after the injection of N1177 and a conventional CT contrast agent. The enhancement of the macrophage-rich plaque after the injection of N1177 was significantly higher and specific inside the vessel wall than after injecting the conventional CT contrast agent. Figure 5.15 shows axial views of the same atherosclerotic plaque (white arrowheads) in the aorta of a rabbit, obtained by CT before (a), during (b), and 2 h after the injection of N1177 (c) or a conventional contrast agent (d). Recently, Lee and Chen published a comprehensive review of dual-modality probes in molecular imaging [29]. This article introduced various emerging approaches enveloped by the term dual-modality probes with respect to their design strategies and applications for all medical imaging modalities, including PET, SPECT, MRI, CT, ultrasonography, optical CT, and photoacoustic CT. Besides extensive listing of current dual-modality imaging probes, they also discussed X-ray CT contrast agents containing bismuth, gold, and gadolinium. Alric and colleagues reported gadolinium chelate-coated gold nanoparticles as dual-imaging probes for CT–MRI [30]. They synthesized these particles by encapsulating gold nanoparticles within a multilayered DTPA shell, which is composed of gadolinium chelates bound to each other through disulfide bonds. The Au@DTDTPA-Gd nanoparticles showed dose-dependent X-ray absorption properties and relaxivities. The particles freely circulated in the blood pool without undesirable accumulation in the lung, liver, and spleen followed by either radiography or MRI. Given that surface modification of gold nanoparticles can be achieved by covalent grafting of targeting moieties, targeted Au@DTDTPA-Cid dual-imaging probes can be developed.

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(a)

(b)

(c)

(d)

FIGURE 5.15 Imaging of macrophages in atherosclerotic plaques. (A–D) Axial views of the same atherosclerotic plaque (white arrowheads) in the aorta of a rabbit, obtained by CT before (a), during (b), and 2 h after the injection of N1177 (c) or a conventional contrast agent (d). These images were acquired in two separate imaging sessions. Before the injection of the contrast agent, the atherosclerotic plaque could not be differentiated from the surrounding tissues, whereas a strong enhancement was detected in the atherosclerotic plaque after the injection of N1177 (but not the conventional contrast agent) [25].

It is interesting that the future molecular X-ray CT contrast agent should be a full-function entity, which combines targeting, imaging, and therapy functions in one nanoparticle. Correspondingly, the nanoparticle contains three parts in its structure: disease-specific targeting part, X-ray absorber, and disease-specific drug. A schematic diagram of the concept is shown in Figure 5.16. When treating a patient with a certain disease, first, a full-function contrast agent for the disease is selected. Then the patient is administered the agent. The disease-specific targeting part of the agent nanoparticle leads the whole nanoparticle to the diseased tissue. When most contrast agent particles are aggregated on the diseased tissue, an X-ray CT scan or radiography is taken. The images would yield

X-ray absorber

Disease-specific targeting part Disease-specific therapeutic agent

FIGURE 5.16 Schematic diagram of the future full-function molecular CT contrast agent.

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clinically valuable information about the size and severity of the disease. At the same time, the disease-specific therapy agent, which may be a drug or a radiation source, is activated to attack the diseased cells.

5.6 MEDICAL APPLICATIONS OF CT Modern X-ray CT is an excellent tool to show how powerful a joint effort of physics, mathematics, medicine, and engineering could be. Since X-ray CT is a noninvasive examining and measuring method, it finds a wide range of applications in medicine, bioengineering, pharmaceutics, heavy machinery, geology, and archeology. Here we just briefly discuss CT’s medical application. 5.6.1 Diagnosis Among all the applications, the primary function of CT is to diagnose abnormal structures in a body. As X-ray radiography, CT is a noninvasive diagnostic technique. With CT the cross sectional image can be clearly displayed without overlapping, enabling easy differentiation of the anatomic structures. CT images are digital, which makes their archiving and transferring much simpler than traditional X-ray films. The process of CT imaging is fast and safe. The spatial resolution can be adjusted for different applications A CT image is a map of X-ray attenuation. We have seen (Section 5.2.1) that the linear attenuation coefficient of a material is proportional to its density. A medical CT scan is carefully calibrated; its relation to the density is very stable and is sensitive to the change of electron densities. Therefore CT images show strong contrasts at material interfaces where the density abruptly changes, such as at the boundaries of body and ambience, bone and soft tissue, cranial bones and brain, and thorax muscles and lung. In a CT image these structures can easily be outlined, and abnormalities can be clearly located and associated to anatomic structures. The attenuation of a material in a CT image is displayed as its Hounsfield number. The Hounsfield numbers of all normal tissues are well known, so a change of tissue density or tissue material can easily be detected and quantified. Many diseases cause tissue and organ structural change. Shown in Figure 5.17 are tumors in the lung and sternum. The density of

FIGURE 5.17 Chest cross sections. (Left) A cross section of the right lung is entirely filled by a tumor. (Right) An abnormal mass is displayed in the sternum. In these images the letter I shown in the blue boxes indicates the views are from inferior, and the scale on the right-hand side of the images shows centimeters.

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FIGURE 5.18 (Left) Edema in right brain shown as volume with density close to water. (Right) A dense tumor in the left brain.

a normal lung is close to that of air. Therefore a solid lung tumor is clearly distinct from the normal lung tissue, as shown on the left of Figure 5.17. On the right, the tumor is revealed as an abnormal contour in the sternum region. On the left of Figure 5.18, a volume of edema can be observed as its density is lower than that of normal brain tissue. In the right a tumor is located because its density is higher than the normal gray matter. 5.6.2 Digitally Reconstructed Radiograph As a stack of 2D CT images is reconstructed from projections; a projection can also be reconstructed from the data contained in the stack of images. A digitally reconstructed radiograph (DRR) image looks like the image from an X-ray radiograph, but it can be created from different directions and any specified focal locations. A DRR is especially useful in radiation therapy simulation. In external X-ray beam radiation treatment the source comes from a point. It is essential to place the source at a number of directions so that the beams can focus at the tumor but avoid the organs at risk. Sometimes blocks made by heavy metals are placed in front of those organs to block the X-ray beam. A “beam eye’s view” DRR in which the focus location is set to the source point is constructed, from which the doctor can clearly see if the tumor is in the field of view and if the organs at risk can be blocked or are outside of the X-ray beam. By rotating the source around the body, DRRs from different directions can be investigated to select the best angles for radiation delivery; then the locations and the shapes of the blocks can be outlined on these DRRs. DRRs are also useful for guiding the patient setup before an actual treatment. An example of a DRR is shown in Figure 5.19. 5.6.3 Radiation Therapy In modern radiation therapy, a series of 2D images of the patient body are taken with CT, MRI, or ultrasound. The radiation dose can be delivered using an external beam or a

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FIGURE 5.19 CT image of a lung tumor shown in a radiation oncology treatment planning software: (top left) Transverse view, (bottom left) coronal view, (bottom right) sagittal view, and (top right) DRRs from different directions. The tumor is delineated by red loops in the views.

source placed inside the body. To ensure precise delivery of the prescribed radiation dose, a computer simulation is performed using the images of the patient for a specific type of treatment, and the sources are arranged in a way that optimizes the dose in the target and spares the surrounding tissues. In most cases CT images are used in the simulation. Thanks to the physics studies of radiation–tissue interaction over the past few decades, the dose deposited to the tissues along a path of the radiation beam can be predicted based on information provided by the CT images. Figure 5.19 shows a radiation treatment plan simulation, in which the target is irradiated by external X-ray beams from several directions. Mega voltage X-rays from a linear accelerator (linac) are commonly used in externalbeam radiation therapy because of their high penetration power. In a modern linac, an on-board kilovolt cone-beam CT device is usually installed for patient setup and to monitor the day-to-day anatomy change during a course of treatment [31–33].

5.6.4 CT Angiography With contrast agents injected into the bloodstream, detailed structures of artery and vein systems can clearly be displayed, as shown in Figure 5.20, which is crucial to detect vascular disorders such as malignant tumors. CT angiography requires short imaging time and has high spatial resolution. The multidetector helical CT device makes this application possible. The surface rendering in computer graphic technique makes the 3D blood vessel images appear more realistically. In some images the bone structures can be removed from the dataset to ease the evaluation of the vessels [21].

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FIGURE 5.20 Volume-rendered reconstruction of cranial vessels detected using iodine contrast agent [21]. The osseous structures are removed.

5.7 FUTURE PERSPECTIVES In this chapter various aspects of CT imaging and its applications are described. Recently, a CT scanner with flat-panel detector has become commercially available, which has the potential to push the spatial resolution of human CT imaging to a few hundred microns. Exciting advancements have also been made in image reconstruction methods. Thanks to the development of computer technologies (such as multicore PC and GPU) over the years, computationally expensive iterative reconstruction methods are entering into routine clinical image reconstruction. The major advantages of these iterative approaches include the ability of incorporating prior system knowledge, insensitivity to noise, lower radiation dose, and effectiveness in dealing with incomplete data. Additionally, some powerful signal processing techniques, such as compressed sensing [34–36] and total variation, are finding natural applications in CT image reconstruction with sparse or incomplete projection data. Indeed, with effective use of prior knowledge and the compressed sensing method, it is now possible to reconstruct a CT image with projection data that are deemed far from sufficient by the classical Nyquist–Shannon criterion. The approach also shows significant potential for CT reconstruction with noisy (or low-dose) and 4D CT image reconstruction. Phase contrast CT represents another interesting development and further enriches the armamentarium of tomographic imaging. This imaging technique reconstructs the refractive index distribution of a weakly absorbing object from a set of tomographic projection measurements. Early development of the imaging technique had relied on the use of a bright and coherent synchrotron radiation source, which greatly limits the application of the modality [37]. However, this has recently changed. By dividing a conventional X-ray beam into many line sources via an absorption grating, it has been demonstrated that phase contrast imaging is attainable [38, 39]. The coherence occurs in the direction perpendicular to the grating lines for each line source and the gradient of the phase distribution is measured

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with shear interferometry. A major advantage of the phase contrast CT is its superior capability in delineation of the boundaries of the involved structures. It is anticipated that a combination of phase contrast and absorption brain CT may yield soft tissue details comparable to MRI. Finally, there is a critical need for high temporospatial resolution molecular CT imaging. While much progress has been made toward this goal, CT imaging is still considered primarily an anatomical imaging modality because of the lack of effective disease-specific probes and, more fundamentally, the inherent low sensitivity of X-ray imaging. Currently, CT has been widely used in combination with other molecular imaging techniques, such as PET and optical imaging devices, to provide useful anatomical and biological information. With the recent advancements in nanoparticle-based contrast media, the role of CT imaging in biological research is being redefined.

ACKNOWLEDGMENT The authors thank Don Russell, Bob Bramall, and Jo Acord for their valuable input to the writing of this chapter. LX wishes to acknowledge research support from the National Cancer Institute (5R01 CA98523 and 1R01 CA133474), National Science Foundation (0854492), and Varian Medical Systems (Palo Alto, CA).

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12. Li, T.; Koong, A.; Xing, L. Enhanced 4D cone-beam computed tomography using an on-board imager. Med. Phys., 2007, 34, 3688–3695. 13. Zhu, L.; Wang, J.; Xing, L. Noise suppression in scatter correction for cone-beam CT. Med. Phys., 2009. 36(3), 741–752. 14. Wang, J., et al. Dose reduction for kilovotage cone-beam computed tomography in radiation therapy. Phys. Med. Bio, 2008, 53, 2897–2909. 15. Liu, P. Apparatus and method for constructing computed tomography image slices of an object undergoing cyclic motion, U.P.a.T. Office, Editor. 2001: United States of America. 16. Badea, C.T.; et al. In vivo small-animal imaging using micro-CT and digital subtraction angiography. Phys. Med. Biol. 2008, 53, R319–R350. 17. Holdsworth, D.W.; Thornton, M. Micro-CT in small animal and specimen imaging. Trends Biotechnol. 2002, 20, S34–S39. 18. Greschus, S. Potential applications of flat-panel volumetric CT in morphologic and functional small animal imaging. Neoplasia 2005, 7, 730–740. 19. Ford, N.L.; et al. Time-course characterization of the computed tomography contrast enhancement of an iodinated blood-pool contrast agent in mice using a volumetric flat-panel equipped computed tomography scanner. Invest. Radiol. 2006, 41, 384–390. 20. Xing, L. The value of PET/CT is being over-sold as a clinical tool in radiation oncology. For the proposition. Med. Phys. 2005, 32(6), 1457–1458. 21. Johnson, T. R. C., et al. Material differentiation by dual energy CT: initial experience. Eur. Radiol. 2007, 17, 1510–1517. 22. Torikoshi, M.; et al. Electron density measurement with dual-energy X-ray CT using synchrotron radiation. Phys. Med. Bio. 2003, 48, 673–685. 23. Bazalova, M.; et al. Tissue segmentation in Monte Carlo treatment planning: a simulation study using dual-energy CT images. Radiotherapy Oncol. 2008, 86, 93–98. 24. Popovtzer, R.; et al. Targeted gold nanoparticles enable molecular ct imaging of cancer. Nano Lett. 2008, 12(8), 4593–4596. 25. Kim, D.; et al. Antibiofouling polymer-coated gold nanoparticles as a contrast agent for in vivo X-ray computed tomography imaging. J. Am. Chem. Soc. 2007, 24(129), 7661–7665. 26. Miyamoto, A.; et al. Development of water-soluble metallofullerineces as X-ray contrast media. Eur. Radiol. 2006, 16, 1050–1053. 27. Rabin, O.; et al. An X-ray computed tomography imaging agent based on long-circulating bismuth sulphide nanoparticles. Nat. Mater. 2006, 5, 118–122. 28. Hyafil, F.; et al. Noninvasive detection of macrophages using a nanoparticulate contrast agent for computed tomography. Nat. Med. 2007, 13, 636–641. 29. Lee, S.; Chen, C. X. Dual-modality probes for in vivo molecular imaging. Mol. Imaging 2009, 8, 87–100. 30. Alric, C.; et al. Gadolinium chelate coated gold nanoparricles as contrast agents for both Xray computed tomography and magnetic resonance imaging. J. Am. Chem. Soc. 2008, 130, 5908–5915. 31. Xing, L.; et al. Overview of image-guided radiation therapy. Med. Dosim. 2006, 31(2), 91–112. 32. Wiersma, R.D.; Mao W.; Xing, L. Combined kV and MV imaging for real-time tracking of implanted fiducial markers. Med. Phys. 2008, 35(4), 1191–1198. 33. Lee, L.; Le, Q.; Xing, L. Retrospective IMRT dose reconstruction based on cone-beam computed tomography and the mlc positional log-file recorded during treatment. In Int. J. Radiat. Oncol. Biol. Phys. 2008, 634–644. 34. Sidky, E.Y.; Pan, X. Image reconstruction in circular cone-beam computed tomography by constrained, total-variation minimization. Phys. Med. Biol. 2008, 53(17), 4777–4807.

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35. Chen, G.T.; Kung, J.H.; Beaudette, K.P. Artifacts in computed tomography scanning of moving objects. Semin. Radiat. Oncol. 2004, 14(1), 19–26. 36. Zhu, L.; et al. Using total-variation regularization for segment-based dose optimization with field specific numbers of segments. Phys. Med. Biol. 2008, 53, 6653–6672. 37. McDonald, S. A.; et al. Advanced phase-contrast imaging using a grating interferometer. J. Synchrotron Radiat. 2009, 16 (Pt 4), 562–572. 38. Pfeiffer, F.; et al. Hard X-ray phase tomography with low-brilliance sources. Phys. Rev. Lett. 2007, 98(10), 108105. 39. Pfeiffer, F.; et al. Hard-X-ray dark-field imaging using a grating interferometer. Nat. Mater. 2008, 7(2), 134–137. 40. Bazalova, M.; Carrier, J.F.; Beautlien, L.; Verhaegen, F. Dual-energy CT-based material extraction for tissue segmentation in Monte Carlo dose calculations. Phys. Med. Biol. 2008, 53(9), 2439–2456. 41. http://www.siemens.com/innovation/en/news events/innovationnews/innovationnews-articles/ ct with dual energy simplifies diagnostics.htm.

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Carbon Nanotube X-Ray for Dynamic Micro-CT Imaging of Small Animal Models OTTO ZHOU Department of Physics and Astronomy, Curriculum in Applied Sciences and Engineering, and Lineberger Comprehensive Cancer Center, University of North Carolina at Chapel Hill, Chapel Hill, North Carolina, USA

GUOHUA CAO Department of Physics and Astronomy, University of North Carolina at Chapel Hill, Chapel Hill, North Carolina, USA

YUEH Z. LEE Department of Physics and Astronomy and Department of Radiology, University of North Carolina at Chapel Hill, Chapel Hill, North Carolina, USA

JIANPING LU Department of Physics and Astronomy, Curriculum in Applied Sciences and Engineering, University of North Carolina at Chapel Hill, Chapel Hill, North Carolina, USA

6.1 INTRODUCTION 6.1.1 Microcomputed Tomography Computed tomography (CT) pioneered by Hounsfield [1] and Cormack [2] is one of the most important breakthroughs for X-ray imaging. CT scanners are now widely used for diagnostic medical imaging [3]. Microcomputed tomography (micro-CT) has recently emerged as a powerful noninvasive imaging tool for preclinical cancer research [4–6]. In combination with careful selection of appropriate anesthesia, radiation energy, and contrast media, micro-CT has been applied successfully to image skeletal, heart, colorectal cancer, and lung cancer models [7–9]. Compared to clinical CT scanners, a much higher spatial resolution is needed to image small animals with research interest. A typical micro-CT scanner uses a fixed-anode microfocused X-ray source and a digital area X-ray detector and operates in the cone-beam geometry (Fig. 6.1). To collect all the projection images necessary for reconstruction, the source and detector pair rotates around a stationary and horizontally positioned mouse bed.

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FIGURE 6.1 Schematic of a cone-beam CT scanner. In a typical micro-CT setup for small animal imaging, the source–detector pair rotates around the sample stage to collect a series of projection images from which 3D structure is reconstructed.

Some scanners are also constructed, typically in experimental settings, with a stationary source and detector pair and a vertically positioned rotating object stage. The spatial resolution of the micro-CT scanner depends primarily on the X-ray focal spot size, the resolution of the detector, and the scanner geometry [4]. The temporal resolution is determined by the detector readout speed and the X-ray exposure time, which is in turn determined by the fluency rate of the X-ray source. Current commercial micro-CT scanners have several limitations that are directly due to the performance of the conventional thermionic X-ray source technology. These include low temporal resolution, high imaging dose, lack of specificity, and long scanning time. These limitations are of particular importance for high-resolution longitudinal studies of live tumor models where motion-induced artifacts and accumulation of imaging dose are major concerns. The required imaging dose strongly depends on the imaging resolution, because the signal-to-noise ratio (SNR) is primarily determined by the amount of photons received per voxel. To maintain the same SNR, the smaller the voxel size, the higher the imaging dose required. The concerns of imaging dose ultimately limit the spatial resolution that can be obtained. A dose of 2–20 cGy is generally required for a micro-CT scan. An exposure of 6 Gy is considered as lethal for a small rodent [6]. Current commercial micro-CT scanners offer the capability of imaging objects ex vivo with high spatial resolution. Studies on lung and colon cancer models have shown that micro-CT is capable of sequentially tracking large tumors with high fidelity. Small tumors are still hardly distinguishable from the normal tissue, which is partially attributed to the motion-induced artifacts created by peristalsis, respiration, and cardiac motion. The motioninduced artifacts can, in principle, be minimized by prospective gating, where the image acquisition is synchronized with the physiological signals, as is commonly done in clinical CT scanners. However, performing such measurements on small animals is challenging because their physiological motions are about ten times faster than those of humans. A comparison of the size and physiological functions between human, rat, and mouse is summarized in Table 6.1. Motion-induced artifacts blur the micro-CT images, resulting in significantly deteriorated spatial resolution than the nominal values.

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TABLE 6.1 Comparison of Size and Physiological Functions Between Human, Rat, and Mouse Parameter

Human

Rat

Mouse

Weight (grams) Breaths/minute Heartbeats/minute

70,000 12 60

250 85 300

25 120 600

Several new designs have been considered for high-resolution CT imaging of live animals. One is to use a high-power rotating anode source instead of a fixed-anode microfocus X-ray tube [8]. The high flux of the rotating anode X-ray tube enables a short exposure time, but at the expense of large focal spot size. To achieve the high resolution, the object is placed as far away from the source and as close to the detector as the photon flux allows. Rotation of the large X-ray source requires complicated gantry design, similar to clinical scanners. Recently, micro-CT scanners with dual source–detector pairs have also been developed by commercial vendors to increase the imaging speed.

6.1.2 Conventional Thermionic X-ray Although there has been a tremendous amount of innovation and improvement in the X-ray detection technology and imaging algorithms over the years, the basic mechanism of generating X-ray radiation has remained the same since the inventions by Roentgen and Coolidge. A conventional X-ray tube comprises a metal filament or a film (cathode), which emits electrons when it is resistively heated to over 1000 ◦ C and a metal target (anode) that emits X-rays when bombarded by the accelerated electrons [10]. For a thermionic tube X-ray radiation is switched on and off by either changing the filament temperature, which is very slow, using a grid to control the electron beam current, which has the problem with space charge accumulation, pulsing the anode high voltage, or using an external mechanical shutter. Grid-controlled tubes are not commonly used for microfocused X-ray tubes, presumably due to issues with electron beam focusing. A high-speed mechanical shutter is used in some of the commercial micro-CT scanners for gated imaging [11] with limited temporal resolution. The spatial resolution of an X-ray source is determined by the size of the focal spot—the area on the X-ray anode that receives the electron beam. The intensity of the X-ray radiation is proportional to the electron beam current and the square of the acceleration voltage, and is limited primarily by heat dissipation of the anode [12]. The smaller the focal spot size the less power can be achieved. While a state-of-art high-power rotating anode medical X-ray tube can operate at ∼100 kW at an effective focal spot size of ∼1 × 1 mm, the power of a typical microfocus tube is limited to 10–100 W due to the small focal spot size. It has been shown [13] that the maximum power of a fixed-anode microfocus X-ray tube running at the continuous mode follows the relation of Pmax ≈ 1.4(X f,FWHM )0.88 , where Pmax is the maximum X-ray tube power in watts and X f,FWHM is the focal spot size in microns. Because of the limited output power, polychromatic radiation is used in essentially all clinical and preclinical imaging applications, although simulation and synchrotron studies have shown that the image quality can be significantly enhanced and the imaging dose reduced using monochromatic or quasi-monochromatic radiation.

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Essentially all current commercial X-ray tubes are single-pixel devices, where X-ray radiation is emitted from a single focal spot. For tomography imaging either the X-ray/detector pair or the object needs to be rotated to collect projection images from different viewing angles for CT reconstruction, which limits the scanning speed, complicates the mechanical design of the scanner, and restricts the scanner configuration. The concept of spatially distributed X-ray sources has been proposed and some have been developed, including the Dynamic Spatial Reconstructor (DSR) [14], the electron beam CT (EBCT) [15], and the scanning beam digital X-ray (SBDX) [16, 17]. The DSR system uses ∼30 individual X-ray tubes mounted on a gantry around a patient to collect the images. Although high temporal resolution has been demonstrated for cardiac imaging, the cost, size, and maintenance-related issues make the design prohibitive. Recently, a dual-source CT scanner was introduced, which uses two individual X-ray tubes that rotate simultaneously to reduce the imaging time. Both the EBCT scanner and the SBDX tube use an electromagnetic field to steer the electron beam to different spots on the X-ray target to produce a scanning X-ray beam, very much like the design of a cathode-ray-tube. Such X-ray tubes are in general large and have a limited viewing range due to the difficulties in steering the high-energy electron beam.

6.2 CARBON NANOTUBE FIELD EMISSION X-RAY 6.2.1 Electron Field Emission and Early Works on Field Emission X-ray Electron field emission is a quantum process, where under a sufficiently high external electrical field electrons can escape from the metal surface to the vacuum level by quantum tunneling [18]. Compared to thermionic emission, this is a preferred mechanism because no heating is required and the emission current can be controlled by the external field to give instantaneous response time. The basic physics of field emission is summarized by the Fowler–Nordheim equation [I = ␣V 2 exp(-b␾3/2 /␤V)], which states that the emission current (I) increases exponentially with increasing voltage (V) [18]. For a metal with a flat surface the threshold field is typically around 104 V/␮m, which is impractically high. All the electron field emitters rely on field enhancement (␤) at the sharp tips/protrusions. One way to fabricate sharp field emitters is by a lithography process [19]. Such emitters have not been used in practical devices due to problems such as low emission current, poor stability, and high cost. X-ray tubes using field emission cathodes have been investigated in the past [20, 21]. In these early systems metal tips were used as the cathodes. Electrons were extracted by applying a pulsed high voltage between the target and cathode. X-ray radiation was generated when the field emitted electrons bombarded a target. The advantages of the field emission X-ray tubes compared to conventional thermionic X-ray tubes have been demonstrated in clinical studies in terms of resolution and exposure time [21]. However, the metal-tip emitters were inefficient. To obtain the high electrical field required to extract the electrons, a Max generator, which uses a series of discharging capacitors, was employed [22]. The X-ray tubes had a limited lifetime of 200–300 exposures and slow repetition rate. With a diode design, the acceleration voltage and the tube current could not be independently controlled. Field emission X-ray tubes using other types of emitters such as the Spindt tips and diamond emitter have also been investigated [23–25]. The highest electron current demonstrated in these X-ray tubes is only on the order of microamperes [25].

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6.2.2 Carbon Nanotube Field Emitters and X-rays Carbon nanotubes (CNTs) are excellent electron field emitters [27]. Due to their atomically sharp tips and large aspect ratios (∼103 ), the CNTs have much larger field enhancement factors (␤), thus lower threshold fields are required for emission than with conventional emitters. They are stable at a high emission current: a stable emission current of >1 ␮A has been observed from an individual SWNT [28] and a CNT bundle [26]; we have achieved a stable emission current density of over 1 A/cm2 from macroscopic cathodes (Fig. 6.2). This is attributed to their unique physical properties including high thermal and electrical conductivities, high temperature and chemical stability, and reasonable resistance to oxidation. These properties make them attractive as “cold” cathode materials for vacuum electronic device applications. A CNT-based field emission X-ray source has recently been demonstrated [29–32]. It utilizes the CNT field emitters as the “cold” cathode for X-ray generation. A typical CNT X-ray source comprised of a CNT cathode, a gate electrode, an electrostatic focusing unit, and a metal anode in a vacuum housing is illustrated in Figure 6.3. Field emitted electrons are extracted from the CNT cathode by applying a bias voltage on the gate electrode and are subsequently focused by the focusing lenses before reaching the anode.

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Compared to conventional thermionic X-ray tubes it has several intrinsic advantages that make it attractive for diagnostic medical imaging. Because of the nature of the field emission mechanism, the source has a fast response time and electronic programmability, which are highly desirable for gated imaging of moving objects [33]. The ability to switch the X-ray radiation on and off by readily turning on/off the electron beam enables efficient utilization of X-ray radiation and better thermal management of anode heat load. The use of a “cold” cathode rather than a “hot” filament makes it possible to manufacture field emission cathode arrays with individually controllable electron emitting pixels, which can generate X-ray radiation from different focal points on the X-ray anode. By sequentially switching on and off the individual pixels, a scanning X-ray beam can be generated to image an object from different viewing angles without mechanical motion [32, 34, 35]. Figure 6.3 shows the design of a CNT multibeam field emission X-ray (MBFEX) source that was first demonstrated in our lab [32]. The new source technology can potentially lead to a truly stationary and gantry-free CT with high scanning speed and high resolution. In addition, the MBFEX source makes multiplexing—simultaneous collection of multiple images at the same time—a reality. This is accomplished using either the binary multiplexing algorithm—activating a subset of the X-ray beams according to a preselected multiplexing algorithm—or the frequency division multiplexing algorithm [36, 37]. This potentially can lead to an orders of magnitude increase in imaging data collection speed, although the trade-off between the imaging speed and the signal-to-noise ratio needs to be investigated further. Besides our team at UNC, the CNT field emission X-ray technology is now being actively investigated by a number of academic groups and major companies in the world [30, 38–41]. Potential preclinical and clinical applications including micro-CT [42–44], stationary digital breast tomosynthesis [45], and tomosynthesis-based imaging guided radiation therapy are being developed. In the next below, we briefly introduce our work on developing CNT X-ray-based dynamic micro-CT for in vivo imaging of small animal models.

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6.2.3 A Microfocus Field Emission X-ray Source for Dynamic Micro-CT To utilize the high temporal resolution and programming capability of the CNT X-ray we have developed a microfocus CNT field emission X-ray source using a modified Einzel type electrostatic lens [46]. As illustrated in Figure 6.4, the lens consists of three electrostatic focusing electrodes. Focusing electrodes 1 and 3 are made of planar metal diaphragms. Central focusing electrode 2 has the shape of a truncated cone. Electrode 1 is at the same potential as the gate electrode; electrodes 2 and 3 have independently controllable potentials. To obtain effective isotropic focal spot size, the CNT emission surface has an elliptical geometry. They are fabricated using a combined photolithography and electrophoretic deposition method developed in our lab [47]. An optical image of the finished CNT cathode

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FIGURE 6.4 (a) The electron beam optics modeling of CNT microfocus X-ray tube showing the details of the cathode, gate mesh, and focusing electrodes. The field emission cathode area is elliptical. The electron beam is uniformly focused on the anode. With a proper anode tilting the effective X-ray radiating focus spot is close to isotropic. (b) Simulated focus spot size as a function of the cathode size. Over a large range of cathode size an effective focusing power of 5 can be achieved

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and a cross-sectional SEM image of the CNT cathode are shown in Figure 6.5. The ratio between the elliptical long axis and the short axis is determined by the anode tilting angle. The dimensions of the focusing lenses were optimized by simulating the trajectory of the field emitted electrons. Figure 6.4 shows the relation between the cathode diameter and the focal diameter of the electron beam on the anode obtained from the simulation. The simulation results indicate that this design can give an optimum focusing factor of about 5. The effective focal spot size scales linearly with the linear dimension of the cathode. This means that for an isotropic effective focal spot size of 100 ␮m an elliptical CNT cathode of 2.35 mm × 0.5 mm is needed at 12o anode tilting angle. Our first generation microfocus tube is housed inside a large vacuum chamber under dynamic pumping. The setup is convenient for laboratory experiments but has limited mobility, which restricts the geometry of the micro-CT scanner. Recently, we developed a compact microfocus X-ray tube (Fig. 6.5), which can readily be mounted on a rotating gantry and has similar performance as the one housed in the large vacuum chamber.

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Extensive studies have been carried out to evaluate the performance of this CNT-based field emission microfocus X-ray tube. The source can be operated at up to 55 kVp anode voltage, effective focal spot size down to 50 ␮m, a tube current of 1.5 mA at 100-␮m focal spot size, and X-ray pulse width on the order of microseconds (although, in practice, the value is limited by the X-ray dose required for imaging).

6.3 DYNAMIC MICRO-CT BASED ON THE CNT X-RAY TECHNOLOGY 6.3.1 Design of the Micro-CT Scanner We have developed two prototype dynamic micro-CT scanners using the CNT microfocus X-ray technology at UNC. One uses the rotating object geometry and is called Cyclops; the other has a stationary and horizontally positioned mouse bed and a rotating source and detector pair and is called Charybdis. Below we briefly describe the designs of these two micro-CT scanners.

Cyclops System The Cyclops micro-CT scanner uses a stationary CNT X-ray source and stationary X-ray camera. The object is rotated to collect the projection images from different viewing angles, as illustrated in Figure 6.6. The object is suspended vertically in a custom built sample holder that is locked to a metal support attached to a computercontrolled rotation stage. The system is configured in the cone-beam geometry. The CNT microfocus X-ray source uses a 2.35-mm by 0.5-mm elliptical geometry cathode, which gives a ∼100-␮m focal spot size. It is capable of reliably delivering ∼1-mA peak anode current at up to 60-kV anode voltage. The camera is a high-speed CMOS flat-panel sensor with a CsI scintillator plate directly deposited on a photodiode array (Model C7940DK-02,

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FIGURE 6.6 A schematic of the Cyclops scanner. The CNT X-ray source placed inside a vacuum chamber and the camera are stationary. The mouse is placed vertically in a custom-built sample holder, which is rotated by a computer-controlled rotation stage. The respiration sensor is attached to the abdomen of the mouse from the top of the sample holder.

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Hamamatsu). It has 2400 × 2400 pixels with a nominal pixel size of 50 ␮m, giving an active area of 120 × 120 mm2 . Figure 6.7 is the timing diagram used to gate the X-ray pulse firing and the camera frame acquisition to the physiological signal from a live animal. The camera runs on a continuous, predefined frame rate. The camera frame rate, camera readout time, and the X-ray pulse width together determine the X-ray exposure window. The exposure window is defined as the time window between the end of the previous frame readout and the beginning of the next frame readout subtracted by the X-ray pulse width, tpw . This subtraction is to ensure the X-ray pulse would not overextend into the camera readout region. The physiological trigger is generated based on the respiration signal obtained from a BioVet physiological monitoring system (Spin Systems (QLD) Pty Ltd, Brisbane, Australia). The physiological trigger can be selected at any portion of the respiration cycle. The dynamic gating is achieved from the logic AND combination between the exposure window signal and the physiological trigger signal by an AND transistor-to-transistor logic (TTL) gate. The output of this AND gate is used as the trigger to generate the X-ray trigger signal that controls the X-ray pulse firing. An X-ray trigger pulse is enabled only by the rising edge of the TTL signal output from the logic AND gate. The physiological triggers from the respiratory gating of a free-breathing mouse may be nonperiodic, such as the three physiological triggers shown in Figure 6.7. The rising edges of the physiological triggers 1 and 3 come within the X-ray exposure window; thus they will each trigger an X-ray pulse. On the other hand, the physiological trigger 2 will not trigger an X-ray pulse because its rising edge comes outside the X-ray exposure window. In this gating scheme, the X-ray pulses (hence imaging sequences) are dynamically gated to the physiological signal of small animals (hence called dynamic scanner). It is worthwhile to point out that the X-ray firings are based on temporal coincidence between the X-ray exposure window and the physiological triggers. The total scan time for a dynamically gated scan critically depends on the rate of this temporal coincidence.

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The geometry parameters of the scanner are determined using a two-ball phantom following the procedure outlined by Noo et al. [48]. The source-to-object distance (SOD) is 135 mm. The source-to-detector distance (SDD) is 216 mm, which gives a magnification factor of 1.6. The cone angle is 14 ◦ and the field of view (FOV) of the scanner is 32 × 32 mm2 .

Charybdis Scanner The Charybdis system consists of a rotating source and detector pair and a stationary sample stage, as shown in Figure 6.8. In this design, the mouse is placed horizontally and stationary on the sample stage during scan. This configuration is preferred for in vivo measurement of the morphology and physiology of small animals. This configuration is made possible by the compact and portable CNT field emission microfocus X-ray tube recently developed in our lab [49]. The X-ray tube and the detector are mounted on two translational stages positioned on the opposite sides of a high-precision goniometer (Model 430, HUBER, Germany), which is driven by a digital stepping motor. This flexibility makes it relatively easy to optimize the scanner geometry, for example, for either high spatial resolution specimen imaging or in vivo animal imaging where a trade-off between the spatial resolution and the temporal resolution has to be considered. The X-ray detector is the same one that has been used in the Cyclops scanner. In both cases, the camera is configured to produce two-dimensional 1024 × 1024 pixels projection images at 50 × 50 ␮m2 pixel resolution. This detector configuration together with the Charybdis scanner geometry gives an effective FOV of 39 × 39 mm2 and an effective digital sampling of 38.5 ␮m at the object plane. The scanner is controlled by a homemade computer program written in LabVIEW. At the beginning of each CT scan, the program performs a series of initialization steps, which include homing the goniometer to its zero position. The program then instructs the X-ray tube, the detector, the goniometer, and a homemade dynamic gating electronic circuit (if gating is involved) to work sequentially so that projection images are acquired in the step-and-shoot mode. Multiple frame acquisition is also allowed for frame averaging. At the end of the scan, the LabVIEW program sends commands to rewind the goniometer to its zero position.

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6.3.2 Prospective Gated Micro-CT Imaging of Free-Breathing Mice

Respiratory Gated Micro-CT Imaging of Free-Breathing Mouse Respiratory motion is one of the major challenges in CT imaging of live mice. It has been the subject of numerous studies in recent years. It has been shown previously that the motions induced artifacts can be minimized by techniques including physically restricting the motion [7] and retrospective gating. In retrospective gating, projections from different phases of the respiratory cycles are taken but only the images belonging to a retrospectively selected phase are reconstructed. The method requires a significant increase in the imaging dose [50–52]. Prospective gating techniques, where the image acquisitions are synchronized with the physiological signals, have also been investigated. In most of these studies, animals are intubated and ventilator controlled [8, 53, 54]. Intubation requires dedicated equipment and special animal handling skills to avoid damage to the upper airways. The respiratory mechanics and lung function of the animal could be altered from paralysis. In addition, when animals are ventilator controlled, the pressure and volume settings can affect the morphology and mechanics, thereby making true measurements of function and structure difficult to obtain [55, 56]. Imaging of free-breathing animals is preferred for longitudinal studies and for disease models, where mechanical ventilation may introduce systematic errors into the physiologically relevant measurements. Imaging of free-breathing rodents using retrospective gating techniques has been reported in a few recent publications [50, 57]. It unavoidably requires redundant projections/exposures to minimize the missing view artifact if only one phase in the breathing cycle is to be imaged. Free-breathing rodents have also been imaged with prospective gating techniques using commercial scanners [57, 58]. However, the effective temporal resolution was only 170 ms due to limitations of the scanner. This is not sufficient to precisely define the peak inspiration phase in a mouse. In addition, each projection image had to be averaged from up to 10 frames in order to achieve a good image quality. We have recently demonstrated the potential of performing high-resolution prospective respiratory gated micro-CT imaging of anesthetized free-breathing mice using the Cyclops micro-CT scanner [59]. The imaging protocol is briefly summarized here. The scanning parameters for respiratory gated micro-CT of anesthetized free-breathing mice were 40 kVp, 0.7 mA anode current, 50-ms X-ray pulse width, 325 projections over a circular orbit of 195 ◦ with a step angle of 0.6 ◦ , and 50 bright and dark calibration images. The total scan time was typically 5–10 min depending on the mouse respiration rate. Reconstruction was done with the Feldkamp algorithm using the Cobra EXXIM software package (EXXIM Computing Corp., Livermore, CA). Data were reconstructed with Parker weighting as isotropic 512 × 512 × 512 arrays at a voxel size of 62 × 62 × 62 ␮m3 . The system performance characteristics including the system modulation transfer function (MTF), contrast-to-noise ratio, and linearity of the system have been measured. The system modulation transfer function (MTF) measured using a 10-␮m diameter tungsten wire and the oversampling method proposed by Fujita et al. [60] is shown in Figure 6.9. The 10% system MTF is 6.2 lp/mm, which represents a resolution of 81 ␮m. The measured CT imaging dose is ∼11 cGy. Respiratory-gated micro-CT imaging was performed on adult wild-type mice. The animals were anesthetized with 1–2% isoflurane delivered from a vaporizer. The animal breathing rates after anesthetization were typically in the range of 80–180 breaths/min. The anesthetized animals were placed over the pressure sensor in the mouse sample holder and secured with adhesive restraints. The animals were positioned such that the respiration

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sensor was approximately in the position of the abdomen to achieve maximum sensor coupling. Experimental procedures carried out in this study had been approved by the University of North Carolina Institutional Animal Care and Use Committee. Each mouse was scanned with the standard scanning parameters and with respiration gated to the mouse’s peak inspiration, followed with another scan gated to its end expiration. Figure 6.10 shows the typical results from the respiratory-gated micro-CT of freebreathing mice. The small structures such as vessels and higher order branching airways can be identified in all those images. The lung in the peak inspiration images appears to have a lower CT number than that in the end expiration images. And airways in the peak inspiration images also appear larger in diameter than those in the end expiration images.

Cardiac Gated Imaging of Free-Breathing Mouse The small size (∼0.1 g and ∼7 mm in the long axis) and fast beating rate (∼600 beats/min) of the mouse heart make in vivo study of its cardiac metabolism, function, and morphology difficult. So far, micro-CT imaging of mouse heart has been demonstrated in two systems: a custom-made micro-CT scanner using a rotating mouse configuration [61] and the top-of-line micro-CT scanner from GE (Locus Ultra, General Electric). Prospective cardiac gated micro-CT imaging at 150-␮m spatial resolution and 10-ms temporal resolution has been reported using the first scanner. Retrospective cardiac gating at 300-␮m spatial resolution and 12-ms temporal resolution is reported using the second system [52]. Both systems use rotating-anode X-ray tubes to provide the high photon flux needed for the high temporal resolution (10 ms). In order to compensate the large focal spot size, the X-ray tube is placed as far away from the object and detector as the photon flux allows. The Charybdis system is the first micro-CT scanner that provides high spatial (<100 ␮m) and temporal resolution (10–20 ms), stationary and horizontal mouse bed configuration, and prospective gating capability. The dynamical gating scheme is the same as the one used for Cyclops. The difference is the method for selecting physiological triggers. In the respiratory

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gated case, physiological triggers were generated entirely based on the respiration signal itself, while for cardiac imaging, the physiological triggers were generated based on both respiration signal and cardiac signal. A physiological trigger is generated if and only if the first QRS complex of the ECG occurred within an acquisition window defined at a specific phase of the respiration cycle. An example for the physiological trigger selection is shown in Figure 6.11. Once the physiological triggers were successfully generated, they then are subject to the same timing diagram shown in Figure 6.7 to trigger the X-ray firings and imaging sequences. The imaging protocol was similar to that was used with the Cyclops system, except that ECG electrodes were taped to the footpads and iodinated contrast agent (Fenestra VC, Advanced Research Technologies Inc., Montreal, Canada) was infused via tail vein injection

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FIGURE 6.11 Physiological trigger selection in cardiac imaging. The red bars indicate the physiological triggers that were generated based on the temporal coincidence between the selected phase of the cardiac signal (the S wave of ECG in this case) and the acquisition window defined at the end expiration plateaus of the respiration signal.

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FIGURE 6.12 (A) The physiological triggers were selected at 0-ms (systole) and 50-ms (diastole) delay after the first R wave of the ECG that occurred within the acquisition window. (B) The reconstructed slice images show clear differences between systole (a) and (c), and diastole (b) and (d) in the axial and coronal views, respectively.

(0.02 mL/gram of mouse) prior to imaging. For micro-CT imaging of mouse heart, 400 projections were acquired over a circular orbit of 199.5 ◦ with a stepping angle of 0.5 ◦ , and 50 bright and dark calibration images were collected. By running the detector at 1 frame/s (camera integration time = 500 ms), the scan time for a cardiac and respiratory gated mouse micro-CT was typically 15–30 min, depending on the mouse’s respiration and heart rates. Images were collected at 100-␮m spatial resolution and 20-ms temporal resolution. A representative result from respiratory and cardiac gated micro-CT of mouse heart is shown in Figure 6.12. As we can see, structures inside the heart are clearly delineated. The motion-induced artifacts are not present, indicating a proper gating.

6.4 CONCLUSION AND FUTURE DIRECTION The carbon nanotube field emission X-ray technology offers several intrinsic advantages over conventional thermionic X-ray tubes that are used in all current X-ray imaging systems. The programmability of the X-ray radiation, the high temporal resolution, the ability to produce a spatially distributed scanning X-ray beam, and the capability for multiplexing are highly attractive for X-ray imaging especially for X-ray tomography. After several years of intense research and development, the basic performance of the nanotube X-ray source including tube current, high voltage stability, focusing, lifetime, and reliability has reached the point where practical preclinical and clinical imaging applications can be expected. The novel X-ray source technology has also moved from a pure academic curiosity to industrial development and production (a MBFEX tube from XinRay Systems is illustrated in Figure 6.13; this particular model has 100 individually controllable X-ray pixels).

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(a)

3.2 cm diameter mouse holder

10 cm detector

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FIGURE 6.13 (a) Schematic of a CNT MBFEX source with 100 individually controlled X-ray pixels produced by XinRay Systems (www.xinraysystems.com). (b) One of the proposed configurations for the stationary micro-CT scanner with two orthogonally positioned source arrays and two area detectors. (From Quan and LaLush [62].)

Next-generation imaging systems for image-guided radiation therapy, stationary digital breast tomosynthesis, and homeland security applications based on the nanotube X-ray technology are being actively developed by leading system vendors. In this chapter we summarized the ongoing research in our lab at UNC Chapel Hill on high-resolution dynamic micro-CT for in vivo imaging of small animal models. The research so far has focused on developing the field emission microfocus X-ray source and micro-CT scanner. Using a single-pixel CNT field emission X-ray source, we have demonstrated the capability to perform prospective respiratory and cardiac gated CT imaging of freebreathing anesthetized mice at 100-␮m spatial resolution and 20-ms temporal resolution. The configuration of the Charybdis system is very similar to the regular cone-beam CT scanner with a stationary mouse bed and rotating source and detector pair. With further development of the MBFEX technology and the imaging reconstruction algorithm, a gantryfree and stationary micro-CT scanner can potentially be constructed. Such a system would comprise a MBFEX source with several hundred microfocused X-ray beams, a panel of X-ray area detectors, and mouse bed, as illustrated in Figure 6.13. All the components are stationary. The project images will be collected by selectively activating the X-ray beams from different projection angles without mechanical motion of the object or the source. The feasibility of such a system was recently demonstrated in our lab using a prototype 20-beam microfocus MBFEX source [63]. There is also an ongoing development effort in the reconstruction algorithm [62]. Further advancement in the imaging speed is also possible using multiplexing, which allows the acquisition of many projection images at the same time [36, 37]. However, issues such as the noise propagation and cross-scattering effect remain to be studied. Advancement in the MBFEX technology will also make it possible to investigate some advanced X-ray imaging modalities currently only accessible at remote synchrotron facilities. One example is quasi-monochromatic CT. An X-ray source with a narrow energy bandwidth will improve contrast resolution and permit K-edge imaging and dual-energy

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imaging. Studies have also shown that a significant reduction of dose can be achieved using a quasi-monochromatic beam [64]. Although the advantages of monochromatic and quasi-monochromatic imaging have been convincingly demonstrated in synchrotron-based studies, it has so far not been adapted for preclinical and clinical imaging. The principal limitation is the prohibitive long imaging time if an X-ray tube is used [64–66]. We expect that by combining the MBFEX source with a multiplexing method, quasi-monochromatic micro-CT imaging can be performed in comparable scanning time as polychromatic microCT in the near future.

ACKNOWLEDGMENT A large number of current and former students and postdoctoral fellows in our group at UNC have worked on this interdisciplinary project over the last few years. In particular, we acknowledge contributions from Xiomara Calderon-Colon, Shabana Sultana, Rui Peng, Laurel Burk, Ramya Rajaram, Lei An, Yuan Cheng, Zejian Liu, and Jian Zhang. Without them this work would not have been possible. We also acknowledge contributions from our collaborators including Prof. David LaLush for advises on imaging reconstruction; Prof. Yue Xiong, Prof. Kay Lund, and Prof. Barbara Grubb for providing the mouse models; Prof. Eric Hoffman for valuable discussions; and Dr. Bo Gao and Dr. HuaiZhi Geng for assistance in cathode fabrication. The research on carbon nanotube field emission X-ray at UNC is generously supported by the Carolina Center for Cancer Nanotechnology Excellence (NCI U54CA119343), NIH-NIBIB (4R33EB004204-01), Xintek, Inc., and the UNC University Cancer Research Funds.

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56. Nakamura, T.; Malloy, J.; McCaig, L.; Yao, L. J.; Joseph, M.; Lewis, J. ; Veldhuizen, R. Mechanical ventilation of isolated septic rat lungs: effects on surfactant and inflammatory cytokines. J. Appl. Physiol. 2001, 91, 811. 57. Ford, N. L.; Martin, E. L.; Lewis, J. F.; Veldhuizen, R. A. W.; Drangova, M.; Holdsworth, D. W. in vivo characterization of lung morphology and function in anesthetized free-breathing mice using micro-computed tomography. J. Appl. Physiol. 2007, 102, 2046. 58. Ford, N. L.; Nikolov, H. N.; Norley, C. J. D.; Thornton, M. M.; Foster, P. J.; Drangova, M.; Holdsworth, D. W. Prospective respiratory-gated micro-CT of free breathing rodents. Med. Phys. 2005 (32), 2888. 59. Guohua, C.; et al. Respiratory-gated micro-CT using a carbon nanotube based micro-focus field emission X-ray source, In Jiang, H.; Ehsan, S., Eds. SPIE: San Diego, CA. 2008, p. 691304. 60. Fujita, H. et al. Simple method for determining the modulation transfer function in digital radiography. IEEE Trans Med Imaging 1992, 11(1), 34–39. 61. Badea, C.; Hedlund, L. W.; Johnson, G. A. 4-D Micro-CT of the mouse heart. Mol. Imaging Biol. 2005, 4, 110. 62. Quan, E.; Lalush, D. S. A faster ordered-subset convex algorithm for iterative reconstruction in a rotation-free micro-CT system. Phys. Med. Biol. 2009, 54, 1061. 63. Peng, R.; et al. Stationary micro-CT scanner using a distributed multi-beam field emission X-ray source: a feasibility study. In SPIE Proceedings on Medical Imaging, in press. 64. Baldelli, P.; et al. A prototype of a quasi-monochromatic system for mammography applications. Phys. Med. Biol. 2005, 50, 2225–2240. 65. Mckinley, R. L.; et al. Simulation study of a quasi-monochromatic beam for X-ray computed mammotomography. Med. Phys 2004, 31(4), 800. 66. Marziani, M.; et al. Dual-energy tissue cancellation in mammography with quasi-monochromatic X-rays. Phys. Med. Biol. 2002, 47(2), 305.

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CHAPTER 7

Quantum Dots for In Vivo Molecular Imaging YUN XING Department of Material Science and Engineering, University of Dayton, Dayton, Ohio, USA

During the last decade, semiconductor quantum dots (QDs) have sparked tremendous interest in the biomedical research field. The unique optical properties of QDs, such as superior brightness,photostability, large Stokes shift, and tunable emission spectra, made them appealing fluorescent probes in a wide variety of applications including cell labeling, single particle tracking inside living cells, and in vivo molecular imaging. The most commonly used QDs consist of an organic solvent-grown CdSe/ZnS core–shell structure made water soluble by either ligand exchange or polymer encapsulation. Surface funtionlization with targeting moieties such as antibody and peptide has allowed successful targeted imaging of tumors inside living animals. More recent developments include non-cadmium-based QDs, smaller QDs that emit in the near-infrared range as well as a QD–luciferase hybrid system (“self-illuminating quantum dots”), which permits more sensitive detection (higher signal-to-noise ratio) and a longer tissue penetration depth. In spite of the many successes, several critical issues remain to be solved for the clinical translation of QDs as a molecular imaging agent for the human body.

7.1 INTRODUCTION Semiconductor quantum dots (QDs) are tiny light-emitting nanoparticles that have captivated scientists and engineers over the past two decades owing to their fascinating optical and electronic properties, which are not available from either individual molecules or bulk solids. Recent research has stimulated considerable interest in developing these quantumconfined nanocrystals as optical probes for biomedical applications. Compared with organic dyes and fluorescent proteins, semiconductor QDs offer several unique advantages, such as size- and composition-tunable emission from visible to infrared wavelengths, large absorption coefficients across a wide spectral range, and very high levels of brightness and photostability [1]. The long-term photostability and superior brightness of QDs make them ideal candidates for live animal targeting and imaging [2–4]. Additionally, recent Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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studies have shown that QDs have two-photon excitation cross sections of magnitude larger than those of conventional fluorescent probes now in use, a further plus for deep tissue imaging [5]. In this chapter, we cover QD core synthesis, water solubilization, biological functionalization, and animal imaging applications and finish with a discussion on the issues/limitations and future perspectives.

7.2 QUANTUM DOT STRUCTURE QDs are somewhat spherical nanocrystals in the size range of 1–10 nm in diameter (Fig. 7.1a) [6, 7]. Semiconductor nanocrystals can also be produced with other shapes such as rods and tetrapods [8], but spherical QDs are often widely used for biological applications. These particles are generally made from hundreds to thousands of atoms of group II and VI elements (e.g., cadmium selenide (CdSe) and cadmium telluride (CdTe) or group III and V elements (e.g., indium phosphide (InP) and indium arsenide (InAs)).

(a)

(c)

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FIGURE 7.1 Quantum dot structure and novel optical properties. (a) Structure of a multifunctional QD probe, showing the capping ligand TOPO, an encapsulating copolymer layer, tumor-targeting ligands (such as peptides, antibodies, or small molecule inhibitors), and polyethylene glycol (PEG) [18]. (b) Size-tunable emission spectra of QDs. This image shows ten distinguishable emission colors of ZnS-capped CdSe quantum dots excited with a near-UV lamp. From left to right (blue to red), the emission maxima are located at 443, 473, 481, 500, 518, 543, 565, 587, 610, and 655 nm [16]. (c) Excitation (dotted line) and fluorescence (solid line) spectra of fluorescein (top) and a typical water-soluble QD (bottom) [37]. (d) Superior photostability of QDs as compared to organic dyes [36].

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Recent advances have allowed the precise control of particle size, shape (dots, rods, or tetrapods) [9–11] and internal structure (core–shell, gradient alloy, or homogeneous alloy) [12, 13]. The novel optical property of quantum dots arises from the “quantum confinement” effect of the semiconductor materials. This refers to the size and composition dependence of the semiconductor bandgap energy. This effect is readily observed when one or more dimensions of a semiconductor are reduced to the nanometer regime. Absorption of a photon with energy above the semiconductor bandgap energy results in the creation of an electron–hole pair (or excitation). The absorption has an increased probability at higher energies (i.e., shorter wavelengths) and results in a broadband absorption spectrum, in marked contrast to standard fluorophores. For nanocrystals smaller than the Bohr excitation radius (a few nanometers), energy levels are quantized, with values directly related to the size of the nanoparticle. The radiative recombination of an excitation (characterized by a long lifetime, >10 ns [14]) leads to the emission of a photon in a narrow, symmetric energy band (Fig. 7.1c). This dependence of light emission on particle size allows the development of new fluorescence emitters with precisely tuned emission wavelengths (Fig. 7.1b). For example, the semiconductor cadmium selenide has a bulk bandgap of 1.7 eV (corresponding to 730-nm light emission). QDs of this material can be tuned to emit between 450 and 650 nm by changing the nanocrystal diameter from 2 to 7 nm. The composition of the material may also be used as a parameter to alter the bandgap of a semiconductor. A QD with a diameter of 5 nm can be tuned to emit between 610 and 800 nm by changing the composition of the alloy CdSex Te1 − x [15]. The classic and most commonly used QDs consist of a CdSe core and a shell layer made of ZnS or CdS. Fluorescent properties are determined by the core materials and the shell layer removes surface defects and prevents nonradiative decay (the disappearance of an excited species due to a radiationless transition), leading to a significant improvement in particle stability and fluorescence quantum yields. For biological imaging applications, these hydrophobic dots can be made water soluble by exchange with bifunctional ligands (mostly thiol and phosphine mono and multidentate ligands) or using amphiphilic polymers that contain both a hydrophobic segment or side chain (mostly hydrocarbons) and a hydrophilic segment or group (such as polyethylene glycol (PEG) or multiple carboxylate groups). Additionally, biomolecules such as antibodies and peptides can be attached to the QDs to achieve specific labeling and targeting (Fig. 7.1a).

7.3 NOVEL OPTICAL PROPERTIES As briefly noted above, QDs are made from inorganic semiconductors and have novel optical properties that can be used to optimize the signal-to-background ratio in fluorescence imaging. In comparison with organic dyes and fluorescent proteins, QDs have several advantages and unique applications. First, QDs have very large molar extinction coefficients on the order of 0.5–5 × 106 M-1 cm-1 [15], about 10–50 times larger than that (5–10 × 104 M-1 cm-1 ) of organic dyes. Therefore QDs are able to absorb 10–50 times more photons than organic dyes at the same excitation photon flux (i.e., the number of incident photons per unit area), leading to a significant improvement in probe brightness; this allows for brighter probes under photon-limited in vivo conditions (where light intensities are severely attenuated due to scattering and absorption). In theory, the lifetime-limited emission rates for single QDs are 5–10 times slower than those of single organic dyes, because of their longer excited

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state lifetimes (20–50 ns vs. less than 10 ns). In practice, however, fluorescence imaging usually operates under absorption-limited conditions, in which the rate of absorption is the main limiting factor of fluorescence emission (versus the emission rate of the fluorephore). As a result, individual QDs have been found to be 10–20 times brighter than organic dyes [16]. Second, QDs are several thousand times more stable against photobleaching (the loss of fluorescence due to photoinduced chemical damages) than organic dyes (Fig. 7.1d) and are thus well suited for continuous tracking studies over a long period of time. In addition, the relatively longer excited state lifetimes of QDs can be used to separate the QD fluorescence from background fluorescence, in a technique known as time-domain imaging [17]; since QDs emit light slowly enough that most of the background autofluorescence emission is over by the time QD emission occurs. Third, the large Stokes shifts of QDs (measured by the distance between the excitation and emission peaks) can be used to further improve detection sensitivity. This factor becomes especially important for in vivo molecular imaging due to the high autofluorescence background often seen in complex biomedical specimens. The Stokes shifts of semiconductor QDs can be as large as 300–400 nm, depending on the wavelength of the excitation light. Organic dye signals with a small Stokes shift are often buried by strong tissue autofluorescence, whereas QD signals with a large Stokes shift are clearly recognizable above the background. A further advantage of QDs is that multicolor QD probes can be used to image and track multiple molecular targets simultaneously. This is a very desirable feature, because most complex human diseases such as cancer and atherosclerosis involve a large number of genes and proteins. The ability to track a panel of molecular markers at the same time will allow scientists to better understand, classify, and differentiate complex human diseases than using a single biomarker each time. Multiple parameter imaging, however, represents a significant challenge for magnetic resonance imaging, positron emission tomography, computed X-ray tomography, and related imaging modalities. By contrast, fluorescence optical imaging provides both signal intensity and wavelength information, and multiple wavelengths or colors can be resolved and imaged simultaneously (color imaging). Therefore different molecular or cellular targets can be tagged with different colors. In this regard, QD probes are particularly attractive, because their broad absorption profiles allow simultaneous excitation of multiple colors and their emission wavelengths can be continuously tuned by varying particle size and chemical composition. For organ and vascular imaging in which micrometer-sized particles could be used, optically encoded beads (polymer beads embedded with multicolor QDs at controlled ratios) could allow multiplexed molecular profiling in vivo at high sensitivities [18, 19]. The most common scheme of quantum dot production consists of four major steps: (1) synthesis of the QD core, most often CdSe, in a high-temperature organic solvent (aqueous-based QD synthesis methods have also emerged [20]; (2) growth of an inorganic shell (usually ZnS) on the core to protect the optical properties of the core; (3) water solubilization of QDs by ligand exchange or polymeric encapsulation; and (4) functionalization of the QD surface with biologically active molecules. Each of the four steps is discussed in the following section.

7.4 CORE SYNTHESIS The most effective and reproducible synthesis procedure for monodisperse quantum dot synthesis involves the addition of semiconductor precursors to a liquid coordinating solvent at high temperature, first described by Murray and co-workers [21]. In a typical synthesis

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of CdSe quantum dots, room-temperature QD precursors, dimethylcadmium and elemental selenium dissolved in liquid trioctylphosphine (TOP), are swiftly injected into hot (290–350 ◦ C) trioctylphosphine oxide (TOPO), immediately initiating nucleation of QD crystals. The presence of coordinating solvents (i.e., TOPO or TOP) prevents the formation of bulk semiconductors by binding to the QD surfaces through their basic functional groups (phosphines, phosphine oxides). The alkyl chains from the coordinating ligands extend away from the QD surface, rendering them sterically stable as colloids and dispersible in many nonpolar solvents such as chloroform and hexane. By tuning the different parameters (precursors, coordinating ligands, etc.), QDs with diameters between 2 and 8 nm and emission spanning the entire visible spectrum can easily be synthesized with just one composition (CdSe). By also adjusting the composition (ZnS, CdS, CdSe, CdTe, PbS, PbSe, and their alloys), it is possible to generate QDs that emit all the way from 400 to 4000 nm [12, 22–24]. The quantum yield (ratio of the number of photons emitted to the number of photons absorbed, essentially the emission efficiency of a fluorophore) of the nanocrystal core synthesized as above is relatively low (less than 10%) due to the presence of gap surface states arising from surface nonstoichiometry, unsaturated bonds [21, 25]. Usually, a thin (a few atoms thick) shell of higher bandgap semiconductor material, such as ZnS or CdS, is epitaxially grown around the core to achieve better photostability and higher quantum yield. A higher bandgap material eliminates the surface defects. This layer improves quantum yield and protects the CdSe core against photo-oxidation (which is extremely important for minimizing cytotoxicity as well as for enhancing photostability). To produce a ZnS shell, a solution of dimethylzinc and hexamethyldisilathiane (in tri-noctylphosphine) is slowly added into the reaction solution of CdSe. Thickness of the shell is mediated by the amount of dimethylzinc and hexamethyldisilathiane injected in. By carefully designing the shell, quantum yields of QDs can be significantly improved (from 5% up to 90%) [25, 26]. The presence of a shell also improves QD photostability by several orders of magnitude relative to conventional fluorophores.

7.5 WATER SOLUBILIZATION As mentioned earlier, QDs made in organic solvents are coated with alkyl chains that render solubility only in nonpolar organic solvents. In order to make QDs useful for biological purposes, it is essential to modify the surface of QDs to make them water soluble. An ideal modification must meet the following criteria: (1) has good shelf stability, no flocculating over long term of storage; (2) efficiently coverts the insoluble QDs to water soluble; (3) maintains the brightness (i.e., quantum yield of the dots); and (4) keeps the overall size of the QDs small, best if below 10 nm. Unfortunately, none of the current coating strategies meet all four criteria, each with their advantages and shortcomings. Different QD solubilization strategies have been devised over the last decade: these strategies can be grouped into two major categories. The first uses “ligand exchange” and involves the substitution of the native TOP/TOPO with bifunctional ligands, each presenting a surface anchoring moiety to bind to the inorganic QD surface and an opposing hydrophilic end group that renders water compatibility. An array of thiol and phosphine mono and multidentate ligands have been tested and will be discussed in detail in the next section. The second method preserves the native TOP/TOPO on the QDs and uses variants of amphiphilic “diblock” and “triblock” copolymers and phospholipids to tightly interleave/interdigitate the alkylphosphine ligands through hydrophobic attraction, whereas the hydrophilic outer

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block permits aqueous dispersion and further derivitization. Additionally, combinations of layers of different molecules conferring the required colloidal stability to QDs have also been exploited [27, 28]. In the ligand exchange scheme, TOPO-coated QDs are mixed with a solution containing a bifunctional ligand, which competes with TOPO for binding to a metal atom on the QD surface. With excess bifunctional ligands (e.g., mercapto acetic acid) in the solution, the thiol functional groups outcompete the phosphonic oxides (from the TOPO) for binding to the QDs and the QDs become hydrophilic. Various types of ligands have been used, including simple thiol-containing molecules such as mercaptoacetic acid [29], cysteine [30], dithiothreitol [31] and dihydrolipoic acid [32], oligomeric phosphine [22], dendrons [32], and peptides [33]. This method results in smaller QDs, generally not too much bigger than the organic core/shell structure. Drawbacks of the ligand exchange method are relatively low quantum yield, pH sensitivity, and poor stability in biological buffers or insufficient robustness for further chemical modification to introduce biorecognition molecules. In the second method, coordinating ligands on the QD surface are retained and used to interact with amphiphilic block copolymers [17, 34] in silica shells [35], phospholipid micelles [36], amphiphilic polysaccharides [37], or polyanhydrides [38]. The hydrophobic domains of these chemicals strongly interact with TOPO on the QD surface, whereas the hydrophilic groups face outwards and render the QDs water soluble. Note that the coordinating organic ligands (TOP or TOPO) are retained on the inner surface of QDs, a feature that is important for maintaining the optical properties of QDs and for shielding the core from the outside environment. This approach is more effective than ligand exchange at maintaining QD optical properties (higher quantum yield) and storage stability in aqueous buffer; the drawback is the increase in the overall size of the QDs. For example, phospholipids and block copolymer coatings tend to increase the diameter of CdSe–ZnS QDs from ∼4–8 nm before encapsulation to ∼20–30 nm, a size that although smaller than most mammalian cells can still limit intracellular mobility and may preclude distance-sensitive applications such as fluorescence resonance energy transfer [39]. In animal imaging applications, increase in size may present a hurdle by impeding the extravasations of QDs to the targeted sites and by causing rapid uptake by the reticuloendothelial system (RES) and thus decreasing the bioavailability of the nanoparticles.

7.6 BIOFUNCTIONIZATION To make QDs more useful for molecular imaging and other biological applications, QDs need to be conjugated to biological molecules without disturbing the biological function of these molecules. Several successful approaches have been used to link biological molecules to QDs, including nonspecific adsorption, electrostatic interaction, mercapto (–SH) exchange, and covalent linkage (Fig. 7.2) [40]. It has been reported that simple small molecules, such as oligonucleotides [41, 42] and various serum albumins [43], are readily adsorbed to the surface of water-soluble QDs. This adsorption is nonspecific and depends on ionic strength, pH, temperature, and the surface charge of the molecule. Mattousi and coworkers [34] presented a method of conjugating proteins to QD surfaces using electrostatic interactions. The protein of interest was engineered with a positively charged domain (polyhistidine), which in turn interacted electrostatically with the negatively charged surface of DHLA-capped QD. The protein–QD conjugates prepared in this way were very stable and the fluorescence quantum yield was even higher than that from the nonconjugated dots [32].

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FIGURE 7.2 Common methods used for QD biofunctionization, including direct linkage to the TOPO-coated QDs (biomolecules can be linked include thiolated DNA (ligand exchange) or peptides with adhesive domains electrostatic interaction, and covalent linking (see text for references).

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Biological molecules containing thiol groups can be conjugated to the QD surface through a mercapto exchange process [44–48]. Unfortunately, the same problem of using thiol as an anchoring group on a ZnS surface occurs because the bond between Zn and thiol is not very strong and is dynamic. As a result, biomolecules can readily dissociate from the nanoparticle surface, causing QDs to precipitate from the solution. A more stable linkage is obtained by covalently linking biomolecules to the functional groups on the QD surface using crosslinker molecules [1, 29, 31, 34, 49]. Most water solubilization methods result in QDs covered with carboxylic acid, amino or thiol groups. Under these situations, it is easy to link QDs to biological molecules that also have these reaction groups. For example, the crosslinker 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) is commonly used to link –NH2 and –COOH groups, whereas 4-(N-maleimidomethyl)-cyclohexanecarboxylic acid N-hydroxysuccinimide ester (SMCC) can be used to crosslink –SH and –NH2 groups. Using the above methods, there have been numerous reports of conjugating QDs with various biological molecules, including biotin [35], oligonucleotides [50], peptides [32], and proteins, including avidin/streptavidin [34], albumin [51], adaptor proteins (e.g., protein A, protein G), and antibodies [34, 49]. In addition, the native functional groups (–COOH, –NH2 , or –SH) on a water-soluble QD surface can be further converted to other functional groups to allow more versatile conjugation of QDs to biomolecules (site-specific conjugation, molecules that are sensitive to EDC or SMCC modification). For instance, carboxylic acids on QDs have been converted to hydrazides, allowing attachments of biomolecules containing sugar groups [49]. Recently, Zhang et al. [53] presented another strategy for site-specific conjugation of biomolecules to the QD surface. This study utilized the specific and stable binding between HaloTag protein (HTP) and its ligand. QDs were first functionalized with HaloTag ligands, and the protein of interest (e.g., Renilla luciferase) was genetically fused to a HTP. When mixed together, QDs and Rluc8 can be immobilized on QDs through the HTP–HaloTag ligands linkage. In most cases, the biological functions of these molecules have been preserved during the conjugation process.

7.7 ANIMAL IMAGING APPLICATIONS OF QUANTUM DOTS Fluorescent proteins and small organic dyes have been used as a fluorescent contrast agent for living animal imaging. However, compared to other imaging modalities, such as PET and MRI, fluorescent imaging is still limited by the poor transmission of visible light through biological tissue. One way to attenuate this penetration problem is to use near-infrared (NIR, 650–900 nm) light, since hemoglobin and water, the major absorbers of visible and infrared light, respectively, have their lowest absorption coefficient in this region [53]. Few organic dyes, however, are available that emit brightly in this spectral region, and they suffer from the photobleaching problem. On the contrary, the novel optical properties of QDs allow the synthesis of bright and stable fluorescent labels that emit in the near-infrared spectrum by adjusting their size and composition. Because visible QDs are more synthetically advanced, most animal imaging studies implementing quantum dots have used CdSe/ZnS QDs that emit visible light and a few recent studies have started using near-infrared dots [54, 55]. Although still far from its mature stage, these studies have demonstrated the great performance and promise of QDs as fluorescent imaging agents in living animals, due to their superior ability to remain photostable and bright. In the following section, we first review the major successful applications of quantum dots in animal imaging and then discuss the current issues and limitations followed by future perspectives.

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7.7.1 Nontargeted Animal Imaging QDs have been used for nontargeted imaging in various animal models [55–57]. Ballou and co-workers [57] injected PEG-coated QDs into the mouse bloodstream and investigated how the surface coating would affect their circulation lifetime. In contrast to small organic dyes, which are eliminated from the circulation within minutes after injection, PEG-coated QDs were found to stay in the blood circulation for an extended period of time (half-life more than 3 h). This long-circulating feature is due to the relatively large size of PEG-coated QDs, which falls within an intermediate size range: they are small enough and sufficiently hydrophilic to slow down opsonization and reticuloendothelial uptake, but are large enough to avoid renal clearance. Amazingly, these QDs maintained their fluorescence even after 4 months in vivo. In 2003, Larson et al. [5] intravenously injected green QDs (550 nm) in a living mouse and visualized them dynamically through the skin (in capillaries hundreds of micrometers deep) by using two-photon microscopy (Fig. 7.3a). In addition to the superior brightness and photostability, this study also found that QDs have two-photon excitation cross sections as high as 47,000 Goeppert-Mayer units, by far the largest of any label used in multiphoton microscopy. Two-photon excitation allows greater tissue penetration due to excitation within the NIR spectral range (e.g., 900 nm), but few fluorophores are bright

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FIGURE 7.3 Animal imaging applications using quantum dots. (a) Two-photon fluorescence imaging of capillaries at the base of the dermis: (a) QD (1 ␮M) intravenously delivered to the animal and (b) fluorescin (40 ␮M) (adapted from Larson et al. [5]). (b) NIR QD fluorescence guided dissection of a lymph node in a pig; 400 pmol of NIR QDs were injected intradermally in the right groin [56]. (c) Targeted tumor imaging using quantum dots antibody conjugates [18]. (d) Tumor imaging using QD705 RGD bioconjugates; arrows indicate tumor sites [55].

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enough for these purposes; QDs, with their large multiphoton excitation cross section, appear to be ideal probes for multiphoton microscopy imaging. Improved tissue penetration can also be achieved by tuning the QD emission to the NIR window [55, 56]. Kim et al. [56] prepared a core–shell nanostructure called type II QDs with fairly broad emission at 850 nm and a moderate quantum yield of ∼13%. In contrast to conventional QDs (type I), the shell materials in type II QDs have valence and conduction band energies both lower than those of the core materials. As a result, the electrons and holes are physically separated and the nanoparticles emit light at reduced energies (longer wavelengths). NIR QDs were injected intradermally in the left paw of a living mouse. Their results showed rapid uptake (5 min) of QDs into nearby lymph nodes and that they could be imaged virtually background-free (Fig. 7.4b). The same study also demonstrated the feasibility of image-guided surgery in a big animal model. Thus 400 pmol of NIR QDs injected intradermally permits sentinel lymph nodes 1 cm deep to be imaged easily in real time (and removed surgically) using an excitation fluorescence rate of only 5 mW/cm2 . QDs have also been used for cell tracking studies [17, 36, 57, 58]. QDs were delivered into live mammalian cells via three different mechanisms: nonspecific pinocytosis, microinjection, and peptide-induced transport (e.g., using the protein transduction domain of HIV-1 Tat peptide, Tat-PTD) [59]. A surprising finding was that two billion QDs could be delivered into the nucleus of a single cell, without compromising its viability, proliferation, or migration [35, 58, 59]. The ability to image single-cell migration and differentiation in real time is expected to be important to several research areas such as embryogenesis, cancer metastasis, stem-cell therapeutics, and lymphocyte immunology. These studies demonstrated the potential of using QDs to enable biologists to track cell (one polymer-coated QD is about one-thousandth of a mammalian cell), tissue, and organ developments over extended periods of time, a task not possible with small molecule organic dyes with fast photobleaching [60, 61]. 7.7.2 Targeted Animal Imaging The above studies demonstrated the capability of QDs for living animal imaging, but had only demonstrated image contrast at the tissue/organ level. The goal of molecular imaging is to generate image contrast due to the molecular difference in different tissues and organs. This requires a probe that has a targeting moiety to generate contrast only in locations specified by the targeting probe. Akerman et al. [49] were the first to explore the possibility of using QD–peptide conjugates to target tumor vasculatures in vivo. QDs coated with peptides targeting the lung vasculature, blood vessel and tumor cell, or lymphatic vessels were injected systematically into mouse. Although they were not able to image the QDs in a living animal, histological sections of different organs after 5 or 20 min of circulation revealed that QDs homed to tumor vessels guided by the peptides, but not to surrounding tissues, probably due to their larger size relative to organic dyes (which would stain surrounding tissues). Whole animal imaging of molecular-level detection was realized by Gao et al. [18] in 2004 using red fluorescent QDs conjugated to antibodies specific to prostate-specific membrane antigen (PSMA) on a human prostate cancer induced in a mouse. QD conjugates used in this study contained an amphiphilic triblock copolymer for in vivo protection, targeting ligands (anti-PSMA) for tumor antigen recognition, and multiple PEG molecules for improved biocompatibility and circulation. The QD-tagged PSMA antibodies recognized and bound at the tumor site and were clearly imaged in vivo (Fig. 7.3c). Control experiments with QD-PEG (no targeting ligand) or QD only confirmed

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the binding was specific and the accumulation of nontargeted QDs at tumor sites was marginal compared with the targeted ones. There are two possible mechanisms for the preferential accumulation of QDs at tumor sites: passive targeting due to the enhanced permeability and retention (EPR) effect and active targeting because of the antibody, PSMA, which is a cell surface marker for both prostate epithelial cells and neovascular endothelial cells. Because the QDs emit in the visible range, a spectral unmixing algorithm was used to separate QD signal from background autofluorescence. More recently, Cai and co-workers used near-infrared QDs for tumor imaging by targeting angiogenesis, the formation of new blood vessels from preexisting vasculature [55, 62, 63]. Amine-modified QD705 (emission maximum at 705 nm) was conjugated to ␣v ␤3 integrin-targeting cyclic RGD peptide and injected intravenously into living mice. Tumor fluorescence reached a maximum at 6 h postinjection with good contrast (Fig. 7.3d) Since angiogenesis is common to all tumors, this technique may aid cancer detection and management in general. It is worth noting that, in both studies, a significant portion of the injected QDs went to the RES system, including the liver, spleen, and lung [64].

7.8 RESONANCE ENERGY TRANSFER BASED DETECTION AND IMAGING 7.8.1 QD Fluorescence Resonance Energy Transfer Fluorescence resonance energy transfer (FRET) is a process in which energy is transferred from an excited donor to an acceptor via a resonant, near-field dipole–dipole interaction [65]. FRET is very sensitive to the distance between donor and acceptor and has been used to study biomolecule conformation, dynamics, and interactions. Several problems associated with FRET with organic fluorophores include fast photobleaching, and spectral overlap between the acceptor and donor. There are several examples of QDs used for FRET in biological systems [39, 44, 66, 67]. Medintz and co-workers engineered a histidine tag on maltose binding protein (MBP) that finally bound electrostatically to QDs, which served as the FRET donor [68] (Fig. 7.4A). With a quencher bound to the maltose-binding site, QD fluorescence was inhibited but could be recovered by adding maltose to displace the quencher. In 2005, the same group did an in-depth study using the MBP system by varying the QD size and the quantity of acceptor dye [66]. Increasing the amount of acceptor dye, or the degree of spectral overlap (by changing QD size), caused a substantial enhancement in energy transfer efficiency. This study demonstrated that QDs can be used as efficient energy donors in the FRET system and showed that by tuning their size, QDs can transfer energy to a number of organic dye molecules. From these studies, the unique advantages of QDs as FRET donors are apparent: size-tunable emission spectra can be used to improve spectral overlap with a particular acceptor dye, and the multivalency ability allows multiple acceptor dye molecules on a single QD donor and this will substantially improved FRET efficiency. QDs could even drive biosensors through a two-step FRET mechanism overcoming inherent donor–acceptor distance limitations, as demonstrated by another work by the same researchers. In this assembly, each 530-nm QD is surrounded by ∼10 MBPs (each monolabeled with Cy3). A second dye, ␤-cyclodextrin-Cy3.5 (␤-CD-Cy3.5), specifically binds in the MBP central binding pocket. Excitation of the QD leads to FRET excitation of the MBPCy3, which, in turn FRET-excites the ␤-CD-Cy3.5, overcoming the low direct QD-Cy3.5 FRET efficiency. Recently, QD-FRET has been successfully applied on proteolytic activity detection by the same group [40]. In this scenario, a rationally designed multifunctional

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FIGURE 7.4 Resonance energy transfer based detection and imaging using QDs. (a) QD FRET. Formation of QD (560)-MBP-␤-CD-QSY9 (maximum absorption ∼565 nm) results in quenching of QD emission. Added maltose displaces ␤-CD-QSY9 from the sensor assembly, resulting in an increase in direct QD emission [68]. (b) QD BRET quantum dot (655) is covalently linked to a BRET donor, Rluc8. The bioluminescence energy of Luc8-catalyzed oxidation of coelenterazine is transferred to quantum dots in close proximity, resulting in QD emission (655 nm) [72].

peptide sequence was engineered between the donor (QD) and acceptor (organic dye). The peptide sequence consists of four functional domains including (1) polyhistidine sequence for self-assembly onto the DHLA capped QDs; (2) a helix-linker spacer to provide rigidity; (3) an exposed protease recognition/cleavage sequence; and (4) a C-terminal site-specific location (cysteine thiol) for dye attachment. When the donor–acceptor hybrid is integrated, FRET takes place and the emission spectrum is a composite of QD emission and the acceptor emission. Upon the addition of the active protease (caspase-1, collagenase, chymotrypsin, or thrombin), the acceptor dye is released from the donor (QD), and the observed fluorescent emission is purely QD emission. Since all the results so far are from experiments performed in buffers, and not in biological samples such as blood or serum, it remains to be determined whether these QD–FRET nanoassemblies can be applied for in vivo imaging of proteolytic activity. Paradoxically, the excellent QD donor properties (long fluorescence lifetime, broad absorption, and high extinction coefficient) may almost preclude their role as FRET acceptors for organic dyes [66]. However, this does not limit their use as acceptors in other applications such as bioluminescence resonance energy transfer as discussed below. 7.8.2 QD BRET Based Animal Imaging Bioluminescence resonance energy transfer (BRET) is a naturally occurring phenomenon whereby a light-emitting protein (the donor, e.g., R. reniformis luciferase) nonradiatively

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transfers energy to a fluorescent protein (the acceptor, e.g., GFP) in close proximity. BRET is analogous to FRET except that the energy comes from a chemical reaction catalyzed by the donor enzyme (e.g., R. reniformis luciferase-mediated oxidation of its substrate coelenterazine) rather than absorption of excitation photons. Compared to fluorescence imaging, bioluminescence has extremely high sensitivity for in vivo imaging purposes [69]. BRET has been used to study protein–protein interaction in living cells and animals [70, 71]. In a study by De and Gambhir, a hRluc8 (donor)–GFP2 (acceptor) BRET pair was utilized to study FKBP12 (fused with hRluc8) and FRB (fused with GFP2 ) interactions inside living mice [70]. BRET signal could be detected by using a cooled charge-coupled device (CCD) camera. Recently, work in the Rao laboratory has demonstrated the feasibility of using QDs as the acceptor in a bioluminescence resonance energy transfer system [72]. In this study, QDs were covalently conjugated by EDC coupling to the donor, Rluc8 protein, an eightmutation variant of the bioluminescence R. reniformis luciferase developed in the Gambhir lab [73]. The protein emits blue light with a peak at 480 nm upon the addition of the substrate, coelenterazine. If the QDs are in close proximity of the protein, they can be excited and emit at its emission maximum (Fig. 7.4B). The advantage of using bioluminescence versus fluorescence lies in the fact that no external excitation is needed. This “self-illuminating” feature allows cancer imaging in deeper tissue where light resources are limited. Since no excitation light is needed, the autofluorescence problem is automatically solved. In addition, since this coupling method is generic, any QDs with carboxyl groups (including those emitting in the NIR region, e.g., 705-, 800-nm QDs have been tested) can be used as a BRET accepter and therefore allows for multiplexed imaging. Compared with existing QDs, self-illuminating QD conjugates have greatly enhanced sensitivity in small animal imaging, with an in vivo signal-to-background ratio of >1000 for 5 pmol of conjugates subcutaneously injected. One critical issue in making these self-emitting QDs is the size of the nanoparticle, since like FRET, BRET is also a distance-sensitive process. Increase in the QD conjugate size results in greater distance between the protein and the fluorescent semiconductor core, and thereby decreases the energy transfer efficiency significantly. For instance, the authors have observed that increasing the protein–nanoparticle distance by only 2–3 nm causes the BRET ratio to drop from 1.29 to 0.37 [72]. By attaching targeting moieties such as tumor homing antibodies or peptides to the BRET assembly, it is possible to use BRET for targeted tumor imaging in living animals. More recently, QD-BRET was successfully applied to proteolytic activity detection in buffer with a slightly different coupling scheme [74]. In this approach, the luciferase–protease substrate recombinant protein was genetically modified with an additional intein segment. Carboxylated QDs were functionalized first with adipic dihydrazide because hydrazides are excellent nucleophiles to attack the thioester intermediate of inteins. The reaction proceeds rapidly upon mixing the two together and results in the cleavage of the intein and ligation of the C terminus of the recombinant protein to the QDs. Using this approach, the authors have successfully applied this method to synthesize a series of nanosensors for sensitive detection of MMP-2, MMP-7, and urokinase-type plasminogen activator (uPA). These prepared nanosensors can not only detect these proteases in complex biological media such as mouse serum and tumor lysates with a sensitivity of as low as 1 ng/mL, but they can also detect multiple proteases present in one sample. Considering that irregulation of proteolytic activity is an important hallmark of various diseases such as cancer progression and the development of athersclerotic plaques [75, 76], these types of QD-BRET probes have broad potential applications as in vivo molecular imaging agents.

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7.9 RECENT DEVELOPMENTS 7.9.1 Reducing the Size One important issue with targeted in vivo imaging with nanoparticles is the opsonization and strong uptake by the RES system, which entraps the majority of the particles shortly after they enter the circulation system, thus preventing them from reaching the targted sites efficiently. RES uptake is influenced by many factors including particle size, shape, and surface chemistry. According to a study by the Frangioni laboratory [77], the hydrodynamic size (HD) of a nanoparticle has to be equal to or less than 5.5 nm in order to completely evade the RES organs (no accumulation in the liver, spleen, or lung) and be cleared by the renal system. However, the QDs are at least over 10 nm in HD. One approach to downsize QDs is through engineering the coating. Recently, Smith and Nie reported a new class of multifunctional multidentate polymer ligands, which not only minimized the HD of QDs but also preserved the colloidal stability and photobleaching/signal brightness [78]. A major finding is that a mixed composition of thiol (–SH) and amine (–NH2 ) groups grafted to a linear polymer chain can lead to a highly compact QD with long-term colloidal stability, strong resistance to photobleaching, and high fluorescence quantum yield. In contrast to the standing brush-like conformations of PEGylated dihydrolipoic acid ligands and monovalent thiols, these multidentate polymer ligands can wrap around the QD in a closed “loops-and-trains” conformation. Using this method, a new generation of bright and stable CdTe QDs with small hydrodynamic diameters between 5.6 and 9.7 nm, with fluorescence emission tunable from the visible (515 nm) to the near-infrared (720 nm), were prepared. In addition to CdTe nanocrystals, the same coating method is applicable to a broad range of core nanocrystals as well as core/shell nanostructures including CdS, ZnSe, CdSe/ZnS, and CdTe/CdS. The in vivo behavior and utility for imaging these nanoparticles remain to be evaluated.

7.9.2 Reducing/Eliminating Toxicity Cell culture studies [79] indicate that CdSe QDs are highly toxic to cultured cells under UV illumination for extended periods of time. This is not surprising because the energy of UV irradiation is close to that of a covalent chemical bond and dissolves the semiconductor particles in a process known as photolysis, releasing toxic cadmium ions into the culture medium. In the absence of UV irradiation, QDs with a stable polymer coating are likely to be much less toxic to cells and animals and this has been confirmed by a number of in vivo animal studies [17, 56]. Still, the perceived toxicity of cadmium has cast a doubtful future for cadmium-based QDs and intrigued the development of noncadmium QD alternatives. Among these alternatives are InAs QDs (will be discussed in Section 7.9.3), doped QDs, and carbon QDs. Doped QDs (d-dots) that do not contain toxic heavy metal ions have been actively studied by the Peng laboratory [80–82]. A d-dot often consists of a semiconductor nanocrystal core, such as ZnSe, doped with a transition metal ion such as Cu or Mn. In 2005, Peng and co-workers reported a method that yields doped ZnSe with high purity and tunable emission. The key feature of the related synthetic chemistry is decoupling the doping from nucleation and/or growth through nucleation-doping and growth-doping strategies [80]. Further optimization of the doping chemistry improved the PL quantum yield of the Mn-doped ZnSe d-dots (Mn:ZnSe d-dots) to the level of typical CdSe q-dots, as high as 60–70% PL QY measured against organic dyes [81]. Mn:ZnSe d-dots with a tunable

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photoluminescence peak position were made water soluble by coating with a monolayer of mercaptopropionic acid. The overall size of such d-dots/ligand complexes is only about 7–8 nm, implying an excellent permeability in biological issues. Carbon quantum dots are produced via laser ablation of a carbon target in the presence of water vapor with argon as carrier gas. The as-produced samples were dominated by nanoscale carbon particles in aggregates of various sizes and showed no detectable photoluminescence. Further acidic treatment and subsequent surface passivation resulted in bright luminescence emissions [83]. These dots are about 5 nm in diameter under TEM observation and have a continuous absorption spectra and can emit light from 450 to 700 nm depending on the excitation light. A follow-up study discovered the two-photon emissions and showed that they’re bright enough for cellular imaging [84]. Both the d-dots and carbon dots are promising nontoxic alternatives to the Cd-based QDs; however, more work is needed to make them comparable (optical properties, stability, and surface functionlization) to the CdSe QDs as in vivo molecular imaging probes. 7.9.3 Shifting to the Red: NIR QDs As mentioned earlier, the ideal QD emission for in vivo molecular imaging applications should be in the near-infrared (700–900 nm) [53]. This issue was noticed when QDs were first applied for in vivo molecular imaging and high-quality NIR QDs have been under active development since then. By far, there are two major caetgories of NIR QDs tested for in vivo imaging purposes: Cd-based (e.g., type II QDs have CdTe core [23, 55]) and non-Cd-based (InAs, InP, etc.) [85, 86]. The Bawendi laboratory at MIT and the Frangioni group at Harvard are the pioneers in developing NIR QDs and applying them for in vivo imaging [55, 85, 86]. Different formulations have been developed and tested in vivo. Type II QDs consisting of CdTe had been applied for sentinel lymph node mapping in 2004 [56]. Most recently, they reported the synthesis of a size series of (InAs)ZnSe (core)shell QDs that emit in the near-infrared and exhibit hydrodynamic size <10 nm [86] and demonstrated their utility in vivo by imaging multiple, sequential lymph nodes and showing extravasation from the vasculature in rat models. These QDs consist of an InAs core of <2 nm and a ZnSe shell and quantum yields were 6–9% in water. Coating DHLA-PEG renders water solubility and the total hydrodynamic size became 8.8 nm. When injected subcutaneously into the paw of a mouse or rat, the QDs rapidly migrated to the SLN, as previously reported for two other QD formulations. However, the new QDs did not get trapped completely in the SLN, but instead migrated further into the lymphatic system, imaging up to 5 lymph nodes sequentially, as well as the channels between nodes. A second new observation was the extravasations of the QDs from the vasculature when injected intravenously. And these were attributed to the unusual small size and the resistance to biofouling of the this new type of QDs. A more recent study showed that InAs-based QDs had longer circulating half-life with significant renal clearance. Xie et al. [87] reported the synthesis of small InAs /InP/ZnSe QDs (below 10 nm) that emit NIR light. After being administered to nude mice intravenously, the particle remained in the bloodstream for a prolonged time (110 min, circulation half-time vs. 6 min for its commercial peer) and showed evidence of improved renal clearance as compared to commercial NIR QDs both from wholebody imaging and ex vivo imaging of the different organs. Considering that InAs and InP are also more “tolerant” (compared to Cd-, Pd-, Hg-based QDs) in the body, these studies collectively have shown that InAs-based QDs are most likely the future for QD-based in vivo molecular imaging. As for Cd-based NIR dots, a recent study by Smith and co-workers [88]

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demonstrated that the epitaxial deposition of a compressive shell (ZnS, ZnSe, ZnTe, CdS, or CdSe) onto a soft nanocrystalline core (CdTe) to form a lattice-mismatched quantum dot can dramatically change the conduction and valence band energies of both the core and the shell. In particular, standard type-I quantum-dot behavior is replaced by type-II behavior, which is characterized by spatial separation of electrons and holes, extended excited-state lifetimes, and giant spectral shifts. Moreover, the strain induced by the lattice mismatch can be used to tune the light emission—which displays narrow linewidths and high quantum yields—across the visible and near-infrared part of the spectrum (500–1050 nm). Owing to their near-infrared emission spectra, small sizes, and low cadmium content (only the CdTe core contains cadmium), this new class of QDs is also attractive for multicolor molecular, cellular, and in vivo imaging. 7.9.4 Getting More from One Particle: QDs in Multimodality Molecular Imaging The recognition that many biomedical imaging modalities provide complementary information has stimulated intense interest in multimodality imaging, using more than one modality to probe a sample of interest. For example, a marriage of MRI and optical techniques will provide anatomical/biodistribution info (MRI) and detailed information at the subcellular level (optical). QDs have been explored for multimodality molecular imaging by several groups around the world. Moulder et al. [89, 90] from The Netherlands has developed an MRI/optical dual-modality imaging system by encapsulating QDs with paramagnetic lipids such as high-density lipids (HDLs) and applied for the imaging of atherosclerotic plaques. In another example, a positron emission tomography/near-infrared fluorescence dual-modality imaging probe was made by conjugating NIR QDs with DOTA, which chelates 64 Cu, which is a radiotracer for PET imaging [91]. While multimodal probes can be made by simply conjugating probes of different functionality, a more elegant solution is to incorporate multiple functionalities in a single probe. Wang and co-workers reported the synthesis of core–shell quantum dots with high relaxivity and photoluminescence [92]. The dual-modality QDs have a composition of CdSe/Zn1-x Mnx S and can be made into watersoluble nanoparticles by capping the core/shell particles with amphiphilic polymer, and the QY values in water reached 21%. These materials also demonstrated high relaxivity with r1 values in the range of 11–18 mM-1 s-1 (at room temperature, 7 T). In vitro cell culture results showed that the QY and manganese concentration in the particles were sufficient to produce contrast for both modalities at relatively low concentrations of nanoparticles. The in vivo performance of these particles remains to be tested.

7.10 ISSUES AND PERSPECTIVES 7.10.1 Less RES Uptake and Longer Circulation Time The first issue is their relatively short circulation half-life, preventing long-term imaging or cell tracking studies; and this applies to all nanostructures injected systematically into a living body. Literature on the in vivo studies using QDs for imaging have revealed that their circulation half-time is influenced strongly by the surface chemistry and that they are cleared from the circulation primarily by phagocytosis of the nanoparticle by RES in the liver, spleen, and lymph nodes [17, 56]. Coating with PEG increases the circulation half-life, and

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attachment to targeting moieties (such as antibodies) reduces the dose needed to generate contrast between normal and tumor tissue [17]. Even with these strategies incorporated, the majority of nanoparticles still end up in the RES. A recent study using commercial nontargeted amino (PEG) QD705 for imaging glioma in rat suggested that maximal RES phagocytosis of QDs was reached between 3.4- and 8.5-nmol doses [93]. Significant tumor uptake of QDs was observed for doses higher than 8.5 nmole. Such a high dose might pose health hazards to living subjects. Thus there is an urgent need for engineering strategies to improve the circulation half-time of QDs, since the longer they can stay in the circulation system the better are the chances that they may be able to get to the tumor site. This might be achieved by minimizing the opsonization or other components of the RES. Intending to attack this problem, a few groups have started studying the interaction between blood components and nanoparticles [94, 95]. The rationale is that nanoparticles, once entered into the bloodstream, are immediately covered by plasma proteins; what the RES system “sees” and what defines the identity of the nanoparticle is largely the protein corona around the particle, not the core material. Elucidation of the profiles of adsorbed proteins on nanoparticles has the potential to facilitate engineering a surface chemistry that is less prone to opsonization and RES uptake. Another issue with nanoparticle-based molecular imaging is in vivo delivery and efficient extravasation to tumor sites. Using intravital microscopy, Smith et al. [96] discovered that QDs with tumor targeting arginine–glycine–aspartic acid (RGD) peptide were not able to extravasate in an SKOV-3 mouse ear tumor model; specific binding only occurred in the tumor neovasculature with aggregated QD conjugates. 7.10.2 Deeper Tissue Penetration Although the superior brightness and photostability of QDs made them attractive candidates for in vivo animal imaging, most of the current QDs still emit within the visible range. The ideal QDs for deep tissue imaging, that is, high-quality QDs with near-infrared-emitting properties, are not yet available. Recent developments include a promising water-based synthesis method that yields particles that emit from the visible to the NIR spectrum and are intrinsically water soluble, but the particles have yet to be tested in biological environments. Most materials (e.g., PdS, PdSe, CdHgTe, and CdSeTe) are either not bright enough or not stable enough for biomedical imaging applications. As such, there is an urgent need to develop bright and stable near-infrared-emitting QDs that are broadly tunable in the farred and infrared spectral regions [17]. In addition to high-quality NIR QDs, multiphoton fluorescence microscopy and novel illuminating mechanisms such as bioluminescence energy transfer can all be used to achieve deeper tissue penetration. 7.10.3 Perspectives Quantum dots as novel fluorescent probes have proved to be tremendously useful in many areas of biological and medical research, especially multiplexed tissue/cell labeling and live cell imaging as well as in vivo imaging. However, as an in vivo imaging agent, QDs are still at a very premature stage of development. Several issues (including toxicity, size, and trapping by the RES) remain to be solved before their potential can fully be exploited in this arena and applied to human subjects. Although the performances of visible QDs are greatly improved over conventional fluorophores, an ideal QD fluorophore should emit in the near-infrared/far-red region with high quantum yield and excellent stability. Apart from being an optical imaging agent, QD-based or multimodality molecular imaging have also

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merged and will be the future of molecular imaging. In summary, quantum dot technology for in vivo cancer imaging is still an area of active research and will require the ongoing collaboration of chemists, biologists, and material scientists for its full potential to be released.

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76. Malemud, C. J. Matrix metalloproteinases (MMPs) in health and disease: an overview. Front. Biosci. 2006, 11 (May 1), 1696–1701. 77. Choi, H. S.; Liu, W.; Misra, P.; Tanaka, E.; Zimmer, J. P.; Itty, I. B.; Bawendi, M. G.; Frangioni, J. V. Renal clearance of quantum dots, Nat. Biotechnol. 2007, 25, 1165–1170. 78. Smith, A. M.; Nie, S. Minimizing the hydrodynamic size of quantum dots with multifunctional, multidentate polymer ligands. J. Am. Chem. Soc. 2008, 130(34), 11278–11279. 79. Jackson, H.; Muhammad, O.; Daneshvar, H.; Nelms, J.; Popescu, A.; Vogelbaum, M. A.; Bruchez, M.; Toms, S. A. Quantum dots are phagocytized by macrophages and colocalize with experimental gliomas. Neurosurgery 2007, 50, 524–529. 80. Pradhan, N.; Goorskey, D.; Thessing, J.; Peng, X. An alternative of CdSe nanocrystal emitters: pure and tunable impurity emissions in ZnSe nanocrystals J. Am. Chem. Soc. 2005, 127(50), 17586–17587. 81. Pradhan, N.; Battaglia, D. M.; Liu, Y.; Peng, X. Efficient, stable, small, and water-soluble doped ZnSe nanocrystal emitters as non-cadmium biomedical labels. Nano Lett. 2007, 7, 312–317. 82. Pradhan, N.; Peng, X. Efficient and color-tunable Mn-doped ZnSe nanocrystal emitters: control of optical performance via greener synthetic chemistry. J. Am. Chem. Soc. 2007, 129(11), 3339–3347. 83. Sun, Y. P.; Zhou, B.; Lin, Y.; Wang, W.; Fernando, K. A.; Pathak, P.; Meziani, M. J.; Harruff, B. A.; Wang, X.; Wang, H.; Luo, P. G.; Yang, H.; Kose, M. E.; Chen, B.; Veca, L. M.; Xie, S. Y. Quantum-sized carbon dots for bright and colorful photoluminescence. J. Am. Chem. Soc. 2006, 128(24), 7756–7757. 84. Cao, L.; Wang, X.; Meziani, M. J.; Lu, F.; Wang, H.; Luo, P. G.; Lin, Y.; Harruff, B. A.; Veca, L. M.; Murray, D.; Xie, S. Y.; Sun, Y. P. Carbon dots for multi-photon bioimaging. J. Am. Chem. Soc. 2007, 129(37), 11318–11319. 85. Kim, S. W.; Zimmer, J. P.; Ohnishi, S.; Tracey, J. B. Frangioni, J. V.; Bawendi, M. G. Engineering InAsxP1-x/InP/ZnSe III–V alloyed core/shell quantum dots for the near-infrared. J. Am. Chem. Soc. 2005, 127(30), 10526–10532. 86. Zimmer, J. P.; Kim, S. W.; Ohnishi, S.; Tanaka, E.; Frangioni, J. V.; Bawendi, M. G. Size series of small indium arsenide–zinc selenide core–shell nanocrystals and their application to in vivo imaging. J. Am. Chem. Soc. 2006, 128, 2526–2527. 87. Xie, R.; Chen, K.; Chen, X.; Peng, X. InAs/InP/ZnSe core/shell/shell quantum dots as near-infrared emitters: bright, narrow-band, non-cadmium containing, and biocompatible InAs/InP/ZnSe QDs. Nano Res. 2008, 1(6), 457–464. 88. Smith, A. M.; Mohs, A. M.; Nie, S. Tuning the optical and electrical properties of colloidal nanocrystals by lattice strain. Nat. Nanotechnol. 2009, 4(1), 56–63. 89. Mulder, W. J.; Castermans, K.; van Beijnum, J. R.; Oude Egbrink, M. G.; Chin, P. T.; Fayad, Z. A.; L¨owik, C. W.; Kaijzel, E. L.; Que, I.; Storm, G.; Strijkers, G. J.; Griffioen, A. W.; Nicolay, K. Molecular imaging of tumor angiogenesis using alphav beta3 -integrin targeted multimodal quantum dots. Angiogenesis 2009, 12(1), 17–24. 90. Cormode, D. P.; Skajaa, T.; van Schooneveld, M. M.; Koole, R.; Jarzyna, P.; Lobatto, M. E.; Calcagno, C.; Barazza, A.; Gordon, R. E.; Zanzonico, P.; Fisher, E. A.; Fayad, Z. A.; Mulder, W. J. Nanocrystal core high-density lipoproteins: a multimodality contrast agent platform. Nano Lett. 2008, 8(11), 3715–3723. 91. Chen, K.; Li, Z. B.; Wang, H.; Cai, W. B.; Chen, X. Dual-modality optical and positron emission tomography imaging of vascular endothelial growth factor receptor on tumor vasculature using quantum dots. Eur. J. Nucl. Med. Mol. Imaging 2008, 35, 2235–2244. 92. Wang, S.; Jarrett, B. R.; Kauzlarich, S. M.; Louie, A. Y. Core/shell quantum dots with high relaxivity and photoluminescence for multimodality imaging. J. Am. Chem. Soc. 2007, 4; 129(13), 3848–3856.

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93. Jackson, H.,Muhammad, O.; Daneshvar, H.; Nelms, J.; Popescu, A.; Vogelbaum, M. A.; Bruchez, M.; Toms, A. Quantum dots are phagocytized by macrophages and colocalize with experimental gliomas. Neurosurgery 2007, 50, 524–529. 94. Cedervall, T.; Lynch, I.; Lindman, S.; Bergg˚ard, T.; Thulin, E.; Nilsson, H.; Dawson, K. A.; Linse, S. Understanding the nanoparticle–protein corona using methods to quantify exchange rates and affinities of proteins for nanoparticles. Proc. Natl. Acad. Sci. U.S.A. 2007, 104, 2050–2055. 95. Kim, H. R.; Andrieux, K.; Delomenie, C.; Chacun, H.; Appel, M.; Desma¨ele, D.; Taran, F.; Georgin, D.; Couvreur, P.; Taverna, M. Analysis of plasma protein adsorption onto PEGylated nanoparticles by complementary methods: 2-DE, CE and Protein Lab-on-chip system. Electrophoresis 2007, 28, 2252–2261. 96. Smith, B. R.; Cheng, Z.; De, A.; Koh, A. L.; Sinclair, R.; Gambhir, S. S. Real-time intravital imaging of RGD–quantum dot binding to luminal endothelium in mouse tumor neovasculature. Nano Lett., 2008, 8(9), 2599–2606.

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CHAPTER 8

Biopolymer, Dendrimer, and Liposome Nanoplatforms for Optical Molecular Imaging DAVID PHAM, LING ZHANG, BO CHEN, and ELLA FUNG JONES Center for Molecular and Functional Imaging, Department of Radiology and Biomedical Imaging, University of California, San Francisco, California, USA

8.1 INTRODUCTION Recent advances in human genome mapping, molecular biology, and high-throughput screening have led to new discoveries of biomarkers and their corresponding binding ligands (peptides or small molecules) [1] for the development of targeted pharmaceuticals [2, 3]. However, the in vivo utility of small molecules as therapeutic or imaging agents often faces issues concerning their effective dose, stability, and rapid clearance. The appeal of macromolecular nanoconstructs as carriers is their multivalent and multifunctional attributes that allow high payload of drug molecules and active targeting groups to be incorporated to enhance therapeutic or imaging efficacy [4–6]. In addition, the ability to engineer the carrier surface with benign pharmacokinetic and/or biodistribution modifiers may provide a “stealth” environment to prevent uptake by macrophages and prolong in vivo blood half-life [7–9]. Biopolymers (60–250 nm), dendrimers (5–200 nm), and liposomes (20–250 nm) are common drug delivery systems. These nanoconstructs have long plasma half-life, and they can passively target aberrant tissues through the enhanced permeability and retention (EPR) effect [10]. These platforms set the precedent of effective drug delivery to diseased tissues while maintaining stability and therapeutic efficacy in circulation [11]. In a tumor environment, the defective leaky vessels and the lack of lymphatics favor the permeability and retention of nanoscale therapeutic and imaging agents [1, 12]. Normal vasculature with a uniform network has diameters ranging from 8 to 10 ␮m and in between the vascular endothelial cells, there are gaps ranging from 5 to 10 nm. Depending on the tumor types, these gaps can be as large as 100–780 nm in a highly irregular network with vessel diameters ranging from 20 to 100 ␮m [13–15]. Nanoscale pharmaceuticals take advantage of these physiological differences to differentiate pathological and normal tissues. However, with the rapid cell proliferation in tumorous regions, the blood vessels are

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not always consistently leaky. In some regions, the vessels may be compressed or collapsed, and others may be leaky or even completely normal [16]. Nanoconstructs with differing chemical constitutions, size, shape, surface charge, and solubility may result in a wide range of pharmacokinetic and biodistribution profiles [11, 15, 17, 18]. It is therefore prudent to fully understand not only the in vivo properties of the nanoplatform itself, but also the intended disease target and clearance for the development of an effective therapeutic or imaging agent. In this chapter, we focus on the use of biopolymers, dendrimers, and liposomes as nanoplatforms for imaging. An overview of each type of platform and examples of their use in optical imaging are presented.

8.2 BIOPOLYMERS Biopolymers have been developed for medicinal use since the 1950s [19]. This class of nanoconstruct comprises three subtypes including (1) pseudosynthetic polymers, (2) synthetic polymers, and (3) natural polymers (see examples listed in Table 8.1) [9]. The application of biopolymers in therapeutics and imaging is limited by their consistent synthetic modification due to their polydispersity and entangling structure [37]. Nonetheless, biopolymers are well suited for use in controlled drug release in areas of cardiology, ophthalmology, endocrinology, orthopedics, and oncology [1]. Applications in optical imaging were not realized until the recent development of near-infrared (NIR) molecular beacons by Weissleder and co-workers [38]. 8.2.1 Optical Probe with a Pseudosynthetic Biopolymer Backbone Perhaps the most iconic application of biopolymers in optical imaging is the proteasesensing molecular beacons [38]. This class of optical nanoprobes consists of three components: the poly-(l-lysine) polymeric backbone, the self-quenching NIR cyanine dye, and the peptide substrate specific to protease known to overexpress in diseased tissues. These probes are assembled in a modular fashion. The PEGylated poly-l-lysine backbone is grafted with an optimal number of peptide substrates with an NIR fluorescent dye molecule conjugated at the N terminus. The resulting probe is fine-tuned with proteolytic cleavage sites for TABLE 8.1 Examples of Each Subtype of Biopolymer Biopolymers

Examples

Applications

Reference

Pseudosynthetic

Poly(amino acids), poly(glutamic acid) (PGA), poly(malic acid), and poly(aspartamides)

Anticancer Antiviral Tissue engineering

20, 21 22, 23 24, 25

Synthetic

N-(2-hydroxypropyl) methacrylamide (HPMA) copolymers, poly(vinylpyrrolidone) (PVP) and poly(ethyleneimine) (PEI)

Anticancer Multiple sclerosis Cholesterol control Drugs

26–29 30 31, 32 33

Natural

Dextran (␣-1,6 polyglucose), dextrin (␣-1,4 polyglucose), hyaluronic acid, and chitosans

Renal control HIV drug Antiviral

34 35 36

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efficient probe activation, while incorporating an optimal number of NIR dye molecules for effective self-quenching. In the absence of targeted protease, the self-quenched probe is optically silent, but upon exposure to targeted protease, the peptide substrate is cleaved and fragments containing dye molecules are released from the nanoprobe to give detectable fluorescent signal. Through phage display [39] and high-throughput screening, discoveries of potent peptide sequences have been identified for different diseases and cellular processes. The adaptation of various potent peptide sequences in the modular probe construct makes it an attractive approach to interrogate a wide variety of diseases such as cancers, rheumatoid arthritis, and arthrosclerosis [18]. 8.2.2 Nanoparticles Based on Synthetic Polymer Synthetic polymers have been used as protective coatings to improve the safety and stability of drugs. Gao and co-workers adapted the same strategy and created a triblock copolymer encapsulated quantum dot (QD) targeted optical imaging reporter [40]. The CdSe-ZnS QD is coated with a poly(butylacrylate)-b-poly(ethylacrylate)-b-poly(methylacrylic acid) block copolymer. The polymer-coated QD was then pegylated and functionalized with monoclonal antibody ligands that target prostate-specific membrane antigen (PSMA). When compared to the nontargeted PEGylated QD construct in mouse xenografts, the targeted probe showed a much stronger signal in the tumor. To fully take advantage of polyvalency [4], Cai and co-workers opted to use a PEG-coated QD conjugated with a smaller RGD peptide [41] and demonstrated a four-fold stronger signal-to-background ratio in in vivo imaging of integrin ␣v ␤3 -positive models than the corresponding nontargeted QD probe [42]. A novel reversibly photoswitchable polymeric NIR probe based on Nisopropylacrylamide (NIPAM) interspersed with styrene and divinylbenzene units was reported by Zhu and co-workers [43]. The polymer was conjugated with perylene diimide (PDI; ␭ex ≈ 480 nm, ␭em ≈ 535 nm) and spiropyran (SP). Spiropyran has two interchangeable conformations: the emissive opened merocyanine (mero form) (␭ex ≈ 588 nm, ␭em ≈ 600–750 nm) and the nonemissive spiropyran (spiro form). The ring opening to merocyanine is induced by UV light, and the nonemissive spiro ring closure forms upon mild heating under visible light (Fig. 8.1). Although in water SP is not appreciably fluorescent in either form, the mero form is highly fluorescent in a hydrophobic environment. Upon UV irradiation, these particles produce a red emission by virtue of the fluorescence resonance energy transfer (FRET) effect from perylene dimide (PDI). The reverse effect is achieved through mild heating and exposure to visible light to give a green emission from the PDI molecules. A potential application of such a probe is to remove ambiguities in fluorescence microscopy stemming from tissue autofluorescence. 8.2.3 An NIR Theragnostic Agent Based on Natural Polymers Kwon and co-workers reported the derivatization of several glycol chitosan (GC) polymers (MW = 20 kDa, 100 kDa, and 250 kDa) with hydrophobic cholanic acid [44]. These nanoparticles were subsequently tagged with NIR dye Cy5.5 for in vivo pharmacokinetic and biodistribution studies, which showed greater tumor accumulation and longer blood circulation correlating with the higher molecular weight nanoparticles. This result led to the recent development of an optical theragnostic agent, an agent with both therapeutic and diagnostic attributes, based on the 250 kDa GC particle encapsulated with paclitaxel [45]. In mouse xenografts implanted with SCC7 tumor cells, preferential accumulation of the

blue excitation

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FIGURE 8.1 An illustration of the reversible dual-color fluorescence mechanism of the spiropyran polymeric nanoparticles. (Reproduced with permission from Zhu et al. [43]. Copyright © 2007 American Chemical Society.)

X

green fluorescence

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theragnostic agent in the tumor was observed by in vivo fluorescence imaging after 12 hours with signal persistence beyond 72 hours. Furthermore, the paclitaxel-loaded nanoparticle effectively inhibited tumor growth after the fifth administration on the 12th day.

8.3 DENDRIMERS Dendrimers are perfectly branched polymeric macromolecules that are synthetically constructed stepwise, layer-by-layer, emanating from a multifunctioned core or vice versa, known as divergent and convergent synthesis, respectively. Each additional layer, denoted as a generation, can drastically increase the number of the surface reactive functional groups. Unlike the traditional polymerization processes that result in polydispersed molecular weight, strict control over the synthesis of dendrimers results in monodispersed molecules, an important factor for reproducibility in pharmacological applications. Along with low polydispersity, dendrimers have well-defined surface functionalities for multivalency, and tunable parameters for optimization of pharmacokinetics. Dendrimers are typically 2–15 nm in diameter with a molecular weight ranging from 15 to 4000 kDa [46]. Pharmaceutically relevant dendrimers are based on polyamidoamines, polyamines, polyamides, poly(aryl ethers), polyesters, carbohydrates, and DNA [47]. Lower generation dendrimers are floppy disk-like structures, but beyond the fourth generation, they conform into the globular geometry similar to that of proteins [48]. The hydrodynamic size of a dendrimer depends on the compressibility of the chemical constituents, but the general trait for all dendrimers is that they become increasingly less compressible with increasing generations and steric hindrance. The shape and compressibility, along with biodegradability, are determining factors of the pharmacokinetics and feasibility of dendrimers in biological applications. Dendrimers of similar molecular weights with different geometrical shapes, like “bow-tie” [49] or “rigid rod” [50], and different sizes and cores [51] significantly alter the plasma half-life and biodistribution profiles [52]. PEGylated dendrimers prepared by conjugating poly(ethylene glycol) (PEG) to the dendrimer core have played an increasingly important role in biomedical research owing to their low toxicity, hemolytic properties, long circulation time, and high accumulation in tumor tissues [49]. 8.3.1 Optical Probes with a Dendrimer Backbone The earliest application of dendrimers in optical imaging was reported by McIntyre and co-workers on the development of fluorogenic peptidic dendrimer for the detection of matrix metalloproteinase-7 (MMP-7) activity in colon cancer [53]. Much like the selfquenched probe discussed earlier, this probe uses a fourth generation polyamidoamine (PAMAM) dendrimer as a backbone with fluorescein (␭ex = 493 nm; ␭em = 517 nm) and tetramethylrhodamine (␭ex = 554 nm; ␭em = 572 nm) fluorophores to promote energy transfer/self-quenching mechanisms. Although the use of this particular dye pair has limited tissue penetration for imaging, the synthetic versatility of dendrimer over a wide range of molecular weight for optimization of clearance, quenching, or proteolytic cleavage makes it an excellent design strategy for imaging. Besides probing for proteolytic activities, dendrimers have also been used as a carrier to characterize the tumor microenvironment. Almutairi and co-workers reported a pH-sensing NIR nanoprobe capable of reporting fluorescence lifetime under acidic conditions in vivo. The probe was synthesized by reacting the hydrazine on the NIR dye (cypate) with the

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ketone on the pegylated dendrimer to form acid-sensitive hydrazone linkers. At neutral pH, the NIR dye is “silent” where the fluorescence intensity is insensitive to the local environment. Under an acidic environment, hydrolysis of the hydrazone linkers releases the NIR dyes from the probe, thus increasing the intensity of the fluorescence signal [54]. 8.3.2 Multimodal Molecular Probes Radioscintigraphy and magnetic resonance (MR) are two imaging modalities with highest detection sensitivity and spatial resolutions respectively [55], but dendrimers provide a good platform for combining them into multimodal imaging probes. Dendrimerbased dual-modal probes combining optical imaging reporters with a MR contrast agent [56, 57] and a radionuclide [58] have been developed for in vivo lymphography. The MR/NIR optical probe was built on a sixth generation PAMAM (G6-PAMAM) dendrimer functionalized with a cyanine NIR dye (Cy5.5) and saturated with gadolinium 2-(4-isothiocyanatobenzyl)-6-methyl-diethylenetriaminepentaacetic acid (Bz-DTPA). The resulting dual-modal nanoprobe easily identified all sentinel lymph nodes, and the nodes can be resected under NIR-guided surgery. Similarly, the radionuclide/NIR probes were synthesized based on the same G6-PAMAM dendrimer. The dual-modal radionuclide/NIR probe was built with a modified Bz-DTPA chelater. With radioscintigraphy being much more sensitive than MR, only a single Bz-DTPA is chelated to 111 In and the rest of the ligands were left unlabeled. In addition to the 111 In/Cy5.5 probe, four other 111 In/NIR optical probes with different emission wavelengths were prepared. in vivo visualization was carried out on a mouse model, where the five probes were injected intracutaneously at five different sites and were visualized with a multiexcitation spectral fluorescence imaging station. Impressively, the report showed five lymphatic drainages simultaneously in different fluorescent hues that correlated with both in vivo and ex vivo radionuclide scintigraphic images. 8.3.3 Targeted Nanoprobes and Therapeutics The ease of conjugating different imaging reporters on the surface of dendrimers also pave the way to combine imaging with therapeutic properties in a single platform. Backer and co-workers synthesized a fifth generation PAMAM decorated with ∼110 decaborate (G5DB) moieties to be used as a boron neutron capture therapy agent for cancer treatment [59]. To this G5-DB nanotherapeutic construct was conjugated an NIR reporter (G5-DB-Cy5). Multiple G5-BD-Cy5 constructs were then linked to a vascular endothelial growth factor (VEGF). in vivo imaging of 4T1 mice bearing breast carcinoma showed selective uptake of VEGF-(G5-BD-Cy5)4 at the periphery of the tumor, whereas the nontargeted G5-BD-Cy5 construct show no accumulation in the tumor. Baker and co-workers [60, 61] and Brechbiel and co-workers [62] demonstrated that nanoprobes with targeting moieties, such as folic acid (FA), monoclonal antibodies for human epidermal growth factor receptor-2 (anti-HER2 mAb), and cyclic-(arginine-glycineaspartic acid) (RGD), can be incorporated with therapeutic drug molecules. Among these nanoprobes is the first dendritic therapeutic agent that not only demonstrated in vivo targeting of cancerous tissues, but also has increased therapeutic efficacy compared to the anticancer drug methotrexate (MTX) alone [60]. The Baker group’s dendritic prodrug was built on G5-PAMAM functionalized with five folic acid targeting moieties, five fluorescein isothiocyanate (FITC) reporters, and five MTX groups. In mice bearing human epithelial

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cancer cells, there is a 50% survival rate in mice treated with the dendritic prodrug after 67 days as opposed to no survival in the population treated with MTX. Furthermore, at the termination of the trial, the surviving mice showed no sign of toxicity in the heart or development of myopathy. The mice treated with the prodrug also show significant necrosis in the tumor.

8.4 LIPOSOMES Liposomes are spherical vesicles formed by an assembly of phospholipids containing hydrophilic head groups and hydrophobic aliphatic chains. They are usually prepared from biocompatible lipids that can be metabolized and cleared. In the past two decades, various forms of these lipid vesicles have been used to encapsulate drug molecules to improve in vivo delivery to disease targets [63]. Since the approval by the U.S. Food and Drug Administration (FDA), liposomes have become one of the most widely used drug delivery systems. However, an effective drug delivery using liposomes is not without challenges. The liposome’s size and surface charge are important elements that influence in vivo phagocytic events and biodistribution. Particles of appropriate size are phagocytized by cells in the reticuloendothelial system (RES): for example, Kupffer cells in the liver and macrophages in the spleen [64]. Additionally, the electrostatic charge on the liposome surface, whether it is neutral, positive, or negative, may also alter the overall pharmacokinetics, clearance, and uptake by disease targets [63]. Surface modification with hydrophilic polymers, such as PEG, is by far the most common strategy to improve the in vivo properties of liposomes. The introduction of PEG alters the surface hydration and steric factors, creating a “stealth” environment to bypass phagocytosis by the RES, and such strategy has been adapted for the development of a number of liposome-based imaging agents [65–68]. 8.4.1 Incorporation of Molecules in Liposomes Owing to the amphipathic nature of liposomes, a wide range of molecules may be incorporated into liposomes based on their interactions with the hydrophilic core, the hydrophobic shell, or the charge on the liposome surface. Additionally, desired ligands may be tethered onto the liposome surface through direct conjugation chemistry (Fig. 2). All these modes of molecular engineering have been explored for optimization of drug delivery [69], development of new imaging agents [70–72], and enhancement of imaging efficacy [73–79]. 8.4.2 Optical Imaging with Liposome-Based Nanoprobes Conjugation of organic dyes onto the liposomes has yielded liposome-based optical probes for both in vitro and in vivo imaging. Rhodamine [80, 81], fluorescein [81, 82], and Cy5.5 [83, 84] are among the most common fluorescent dye molecules being used. Besides organic dyes, CdSe/CdTe QD have also been incorporated for optical imaging [71]. Owing to their hydrophobic surface and potential in vivo toxicity [85, 86], the clinical application of QDs is often faced with skepticism. The encapsulation of QDs in liposomes therefore presents an attractive approach to improve QDs’ biocompatibility and toxicity [87].

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(b)

(a)

Hy m dro ole ph cu ilic les

c bi ho s op le dr lecu y H o m

ic at st n ro io ct ract e El nte I

co Dir nju ec ga t tio n

(c)

(d)

FIGURE 8.2 Incorporation of a variety of molecules in liposomes via (a) encapsulation into the hydrophilic core vesicle, (b) embedding hydrophobic molecules in the aliphatic shell, (c) electrostatic interactions, and (d) direct conjugation to the phospholipid.

Recently, it was also reported that liposomes consisting of hepatitis B surface antigen and fluorescently labeled polystyrene beads or plasmids were used for tissue and cellular targeting with high specificity [88]. Another method of liposome ligation was reported by Reulen et al. [89], which showed fluorescent proteins covalently coupled to cysteinefunctionalized phospholipids on the liposome surface for optical imaging of collagen.

8.4.3 Multimodal Functionality of the Liposome Platform The ease of incorporating various functional molecules in the core vesicle, the aliphatic shell, or the lipid surface makes the liposome an ideal platform for multimodal targeted imaging and therapy. As a delivery vehicle, the liposome is known for its ability to passively accumulate at tumor regions via the EPR effect. By conjugating active targeting ligands on the lipid surface, researchers have further improved liposomes’ delivery efficacy. Park and co-workers have shown that, by incorporating active targeting ligand such as antiHER2 (human epidermal growth factor receptor 2) or anti-EGFR (epidermal growth factor receptor) fragments, liposomes can not only accumulate at tumor sites but these ligands also facilitate active internalization by tumor cells [90, 91]. This combined passive and active targeting strategy has been further adapted with other disease binding ligands such as the RGD peptide for ␣v␤3-integrin [80], CREKA peptide for clotted plasma proteins [82], and GE11 peptide for EGFR [83]. The multimodal functionality is not only limited to targeting; multimodality imaging with optical and MR is often sought after using the liposome platform. For example, Mulder and co-workers reported the use of paramagnetic lipid to coat quantum dots for MR and optical dual imaging [66, 67, 80]. Alternatively, Kamaly et al. [68] synthesized a fluorescently labeled gadolinium lipid for cell labeling and tumor imaging. Another variation includes a dual-modal probe with fluorescence and magnetic properties constructed by

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linking the near-infrared fluorescent transferrin conjugate (Tf NIR ) on the surface of gadolinium encapsulated cationic liposome to image breast cancer [92].

8.5 PERSPECTIVE As shown in the examples presented here, biopolymers, dendrimers, and liposomes have proved to be versatile platforms for targeted delivery, controlled tissue distribution, and clearance. These nanoplatforms will continue to play a vital role in new agent development. With their characteristic multivalency and multifunctional attributes, highly sensitive and specific nanoagents will soon be applied in the clinic and the ability to simultaneously obtain in vivo disease information from different imaging modalities will soon be realized. Finally, the “stealth” environment provided by these platforms also opens new avenues to evaluate new materials that otherwise would be biologically incompatible (e.g., iron oxides and quantum dots). These platforms will allow a safe entry of many new agents for clinical evaluations in the future.

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85. Dubertret, B.; Skourides, P.; Norris, D. J.; Noireaux, V.; Brivanlou, A. H.; Libchaber, A. In vivo imaging of quantum dots encapsulated in phospholipid micelles. Science 2002, 298(5599), 1759–1762. 86. Chen, C.-S.; Yao, J.; Durst, R. Liposome encapsulation of fluorescent nanoparticles: quantum dots and silica nanoparticles. J. Nanopart. Res. 2006, 8(6), 1033–1038. 87. Dudu, V.; Ramcharan, M.; Gilchrist, M. L.; Holland, E. C.; Vazquez, M. Liposome delivery of quantum dots to the cytosol of live cells. J. Nanosci. Nanotechnol. 2008, 8, 2293–2300. 88. Jung, J.; Matsuzaki, T.; Tatematsu, K.; Okajima, T.; Tanizawa, K.; Kuroda, S. I. Bio-nanocapsule conjugated with liposomes for in vivo pinpoint delivery of various materials. J. Control. Release 2008, 126(3), 255–264. 89. Reulen, S. W. A.; Brusselaars, W. W. T.; Langereis, S.; Mulder, W. J. M.; Breurken, M.; Merkx, M. Protein–liposome conjugates using cysteine-lipids and native chemical ligation. Bioconjug. Chem. 2007, 18(2), 590–596. 90. Kirpotin, D. B.; Park, J. W.; Hong, K.; Shao, Y.; Shalaby, R.; Colbern, G.; Benz, C. C.; Papahadjopoulos, D. Targeting of liposomes to solid tumors: the case of sterically stabilized anti-Her2 immunoliposomes. J. Liposome Res. 1997, 7(4), 391–417. 91. Mamot, C.; Drummond, D. C.; Noble, C. O.; Kallab, V.; Guo, Z.; Hong, K.; Kirpotin, D. B.; Park, J. W. Epidermal growth factor receptor-targeted immunoliposomes significantly enhance the efficacy of multiple anticancer drugs in vivo. Cancer Res. 2005, 65(24), 11631–11638. 92. Shan, L.; Wang, S. S., R.; Bhujwalla, Z. M.; Wang, P. C. Dual probe with fluorescent and magnetic properties for imaging solid tumor xenografts. Mol. Imaging 2007, 6(2), 85–95.

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CHAPTER 9

Nanoplatforms for Raman Molecular Imaging in Biological Systems ZHUANG LIU Institute of Functional Nano & Soft Materials, Soochow University, Suzhou, Jiangsu, China

9.1 INTRODUCTION Traditional optical imaging of biological samples mostly relies on bioluminance and fluorescence imaging. The latter, in particular, has been widely utilized to study biological systems in vitro and in vivo. Despite the powerfulness of this technique, a few factors still limit its applications [1]. Rapid photobleaching is a problem for most organic fluorescent dyes, preventing long-term imaging and monitoring of fluorescently labeled substances [2]. The autofluorescence background of biological tissues also restricts the sensitivity of fluorescence imaging [3]. Moreover, fluorescent molecules normally have wide emission peaks with full-width at half-maximum (FWHM) as large as 100 nm, which leads to spectral overlay between different dyes and limits the imaging multiplicity. This is especially problematic for in vivo imaging, in which the ideal excitation and emission wavelengths are in a narrow tissue transparency window in the near-infrared (NIR) region (700–900 nm) [4, 5]. Therefore new imaging techniques as well as novel contrast agents are desired to circumvent those problems. As a photon scattering phenomenon, Raman spectroscopy has distinctive features from fluorescence emission. The Raman scattering of molecules exhibits narrow spectral fingerprints with FWHM within a few nanometers and thus allows a dramatically increased imaging multiplicity compared with fluorescent imaging [6]. Tissue autofluorescence is no longer a problem because the sharp peaks in Raman spectra can easily be distinguished from the fluorescence background, imparting high imaging and detection sensitivity. In addition, Raman scattering signals are generally more robust than fluorescence with much less photobleaching. Therefore Raman imaging is a promising technique for the next generation of biomedical imaging [7–9]. Despite the above advantages in Raman imaging, the photon efficiency in Raman scattering is very low. Traditional Raman spectroscopy measurement of small molecules has very low detection sensitivity due to poor Raman scattering signals. In the past decade, the development of nanotechnology has provided a number of breakthroughs to Raman Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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imaging by enhancing Raman scattering intensities through various mechanisms including resonance Raman scattering and SERS, boosting applications of this imaging modality in biomedical research. In this chapter, resonance Raman scattering imaging with single walled carbon nanotubes (SWNTs) in biological systems is reviewed. Recent advances of SERS Raman imaging based on nanoplatforms are summarized. Nanotechnology, in combination with modern optical techniques in Raman spectroscopy, is likely to bring novel opportunities to the realm of molecular imaging.

9.2 RAMAN SCATTERING For light scattering by particles much smaller than the wavelength, a photon is incident on a molecule and interacts with the electric dipole of the molecule, inducing an excitation to a virtual state, followed by nearly coincident deexcitation by a photon emission [10]. If the final state of the molecule is at the same vibration level as the initial state, the scattering photon will have exactly the same energy as the incident photon and this scattering is Rayleigh scattering. In Raman scattering, in contrast, the final state has a different vibration level from the initial state, causing a shift of wavelength in the scattered photon. For most molecules, Rayleigh scattering normally overwhelms Raman scattering. In Stokes Raman scattering, molecules in the ground state (the dominant population at room temperature) will be excited to the virtual state and relax back to the excited vibrational state with higher energy. Therefore the scattered light will show a red shift, which is called the Stokes Raman shift (Fig. 9.1). On the other hand, in rare cases if molecules are in the excited vibrational state before photon excitation and relax to the ground state after photon scattering, the scattered photon will have a blue shift with a wavelength lower than the excitation photon. This anti-Stokes Raman scattering has much lower intensity than Stokes Raman scattering but no fluorescence background. When measuring samples with high fluorescence emissions, anti-Stokes Raman scattering may be used because of the higher signal-to-noise ratio despite its weak scattering intensity. Raman scattering of small molecules usually has very low photon efficiency because the majority of incident photons are elastically scattered by Rayleigh scattering. Only a very small fraction of scattered light (approximately one out of a million photons) is Raman scattering [10]. Therefore the spontaneous Raman scattering of molecules is normally too weak to be used in biological detection and imaging, unless enhanced through certain mechanisms including resonance Raman scattering and SERS, as is discussed in the following two sections.

9.3 RESONANCE RAMAN SCATTERING OF SWNTs FOR IN VITRO AND IN VIVO MOLECULAR IMAGING 9.3.1 Resonance Raman Scattering of SWNTs In resonance Raman scattering, the wavelength of the incoming photon is adjusted to match the electronic transition energy of the molecule. Instead of being excited to a virtual energy state, the molecule could undergo a real electronic transition after excitation by an incident photon [11]. When the laser beam has a frequency close to an electronic

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Virtual energy level

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ve

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FIGURE 9.1 Rayleigh and Raman light scattering. In Rayleigh scattering, the scattered photon has the same energy as the excitation light. In Raman scattering, the final state of the molecule has a different vibrational energy level from the initial state, affording a frequency shift of the scattered light, which is red-shifted for Stokes Raman scattering and blue-shifted for anti-Stokes Raman scattering.

transition (in resonance), the Raman scattering of vibration models associated with that particular electronic transition will be greatly enhanced by many orders of magnitude. The largely increased Raman scattering intensity allows much higher detection sensitivity in biological systems. Since different molecules have different electronic transitions, the resonance Raman spectroscopy technique became more popular after the advent of tunable lasers in the 1970s. SWNTs are one-dimensional quantum wires whose density of states (DOS) is characterized by the so-called van Hove singularities (VHSs) (Fig. 9.2a) [13, 14]. The sharp VHS defines narrow energy ranges where the DOS intensity is very high, allowing a single nanotube to behave like a molecule with well-defined electronic energy levels at each VHS. The electronic structures of SWNTs are associated with their chiralities and diameters, with electronic transition energies Eii (E11 and E22 are in the near-infrared to visible regions) varying for nanotubes with different chiral indices (n, m) (Fig. 9.2b). For a single SWNT, if the visible or NIR laser frequency is close enough to the E11 or E22 energy, an electronic transition will be induced, followed by the strongly enhanced resonance Raman scattering. Most of as-prepared SWNT samples have various populations of nanotubes with different (n, m). Therefore only a certain fraction of SWNTs in a bulk or solution sample will be in resonance with the laser. Nonresonance Raman signals from other populations of nanotubes are relatively weak. SWNTs prepared by different methods have different chirality and diameter distributions. To obtain the highest resonance effect in Raman spectroscopy, the laser wavelength could be chosen to match the electronic transition energy of the most populated nanotubes at certain chiral indices (n, m). For example, SWNTs prepared by the high-pressure carbon monoxide (Hipco) method exhibit strong Raman scattering signals under a 785-nm laser excitation because many nanotubes including (10, 5), (9, 7), and (11, 3) ones in a Hipco SWNT sample are in or close to the resonance of the 785-nm laser [15]. SWNTs have multiple Raman peaks. The radial breathing model (RBM) and tangential G-band are two important and usually the strongest features in the SWNT Raman spectrum

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FIGURE 9.2 Electronic structure of SWNTs. (a–c) Theoretically calculated density of electronic state for an armchair (10, 10) SWNT (a), chiral (11, 9) SWNT (b), and zigzag (22, 0) SWNT (c). (d) Electronic transition energies Eii for all the (n, m) SWNTs with diameters from 0.4 and 3.0 nm using a simple first-neighbor tight binding model [12]. (Reproduced with permission from Jorio et al. [13]. Copyright © 2003 New Journal of Physics.)

[13]. The RBM peak of SWNTs occurring at wave numbers between 100 cm−1 and 300 cm−1 corresponds to nanotube “breathing,” the vibration of carbon atoms in the radial direction (Fig. 9.3a). The vibration energy and thus the RBM peak position is determined by the diameter of nanotubes. The G-band peak at around 1580 cm−1 is due to tangential vibration of carbon atoms in the graphitic plane (Fig. 9.3b). Depending on the vibrational direction, the SWNT G-band has two subbands, G+ and G− peaks, which correspond to

FIGURE 9.3 Schematic drawings of atom vibrations in RMB model and G-band model of SWNTs. Depending on the vibration direction, the SWNT G-band has two subbands—G+ band vibration in parallel with the tube axis and G− band vibration perpendicular to the tube axis.

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vibrations along the tube axis direction and the circumferential direction, respectively. For semiconducting nanotubes, the G+ peak at ∼1590 cm−1 is normally stronger than the G− peak at ∼1570 cm−1 . Unlike RBM peaks, the SWNT G-band peak is less affected by the nanotube diameter. Both RBM and G-band SWNT Raman peaks have been used in Raman imaging studies in biological systems.

9.3.2 Raman Imaging with SWNTs in Biological Systems In recent years, carbon nanotubes have been widely explored for applications in biomedicine, including drug delivery, biological sensing, and molecular imaging (see Chapter 20 of this book). Studies have shown that functionalized SWNTs taken up by living systems escape from cells by exocytosis [16], and from mice mainly via biliary and possibly via renal excretions [17], without exhibiting obvious toxicity [18] in both cases. Carbon nanotubes are able to enter cells via the endocytosis mechanism, shuttling various biological molecules including proteins, DNA, RNA, and drugs into cells. In vivo cancer treatment in an animal model has been realized with a nanotube drug carrier [19]. Ultrasensitive ex vivo protein microarrays have been developed utilizing the resonance Raman scattering of SWNTs in combination with SERS, achieving a protein detection limit down to 1 fM [20]. Moreover, the intrinsic optical properties of SWNTs including resonance Raman scattering and photoluminance in the NIR region have also been used in biological imaging in vitro and in vivo. Raman imaging of SWNTs in live cells was first reported by Heller et al. [21] in 2005 [21]. Hipco SWNTs were functionalized by DNA and used to incubate with 3T3 fibroblast and myoblast stem cells. A 785-nm laser was used as the excitation source for Raman spectroscopic mapping. The RMB peak intensity at each mapping pixel was integrated and color-scaled to create a Raman image. Raman imaging of those cells showed high SWNT Raman signals inside cells, a direct proof of nanotube cellular entry without using additional fluorescent labels. Interestingly, they observed ultrahigh photostability of SWNTs, many orders of magnitudes better than organic fluorescent dyes and NIR quantum dots. The high photostability is due to the inert chemical structure of SWNTs and allows long-term tracking and imaging of SWNTs in biological systems. This was further shown by Liu et al. [17] in a later work, in which ex vivo Raman spectroscopy measurement was used to track SWNT biodistribution in mice over 3 months. Besides in vitro imaging of cells, in vivo Raman imaging of SWNTs in live animals has been achieved by the Gambhir group at Stanford [22, 23]. Hipco SWNTs functionalized by phospholipid–polyethylene glycol (PL-PEG) were conjugated with a targeting Arg–Gly–Asp (RGD) peptide to recognize integrin ␣v ␤3 receptors overexpressed on tumor vasculature and U87MG human glioblastoma cancer cells [24]. U87MG tumor bearing mice were intravenously injected with SWNT-RGD conjugate before Raman imaging. Whole mouse Raman imaging was conducted on a commercial Renishaw Raman microscope with small modifications. Raman images of tumors on live mice revealed time-course information of SWNT levels in tumors. As expected and consistent to a earlier study using position emission tomography (PET) to image radiolabeled SWNTs in mice [24], high SWNT Raman signals were observed on tumors of mice injected with SWNT-RGD but not from those injected with plain SWNTs, owing to the specific binding between the RGD peptide conjugated on SWNTs and the integrin ␣v ␤3 protein expressed on cancer cells

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(a)

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FIGURE 9.4 In vivo Raman imaging with SWNTs. (a) A digital photo of tumor bearing mouse used for Raman imaging. The tumor area in the black box was scanned by a Raman spectroscopic microscope. (b) Panel of tumor images from mouse receiving SWNT-RGD at various time points postinjection starting from left to right with 2 h, 8 h, 24 h, 48 h, and 72 h. (c) Panel of tumor images from mouse receiving plain SWNTs without RGD conjugation at various time points postinjection starting from left to right with 2 h, 8 h, 24 h, 48 h, and 72 h. Strong SWNT G-band signals were observed on tumors from mice injected with SWNT-RGD but not from those injected with plain nanotubes. (Reproduced with permission from Zavaleta et al. [22]. Copyright © 2008 American Chemical Society.)

and the vasculature (Fig. 9.4). This is the first proof-of-principle study achieving in vivo molecular imaging using Raman spectroscopy.

9.3.3 Multiplexed Raman Imaging with SWNTs The frequency of carbon–carbon bond vibration in the SWNT G-band model is dependent on the atomic mass. Changing of carbon isotope from normal C12 to C13 changes the position of the SWNT Raman G-band peak, lowering the Raman shift by ∼62 cm−1 . It has also been found that the SWNT G-band frequency shifts as the variation of C13 isotope composition ratios in nanotubes. Utilizing the isotopically modified SWNTs, multiplexed multicolor Raman imaging has been realized by Liu et al. [25] in Dai group at Stanford. Pure C12, pure C13, and mixed C12/C13 SWNTs synthesized by the chemical vapor deposition (CVD) method exhibiting Raman G-band peaks at positions of 1590 cm−1 , 1528 cm−1 , and 1544 cm−1 , respectively, are used as three different Raman “colors” for multiplexed imaging (Fig. 9.5a,b). Those nanotubes were functionalized and conjugated with three different targeting ligands including RGD peptide, anti-EGFR/Her1, and

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(a)

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FIGURE 9.5 Multicolor Raman imaging with isotopically modified SWNTs. (a) Schematic SWNTs with three different isotope compositions (C13-SWNT, C12/C13-SWNT, C12-SWNT) conjugated with different targeting ligands. (b) Solution phase Raman spectra of the three SWNT conjugates under a 785-nm laser excitation. Different G-band peak positions were observed. (c) A deconvoluted confocal Raman spectroscopic image of a mixture of three cell lines with different receptor expressions incubated with the three-color SWNT mixture. Scale bar = 100 ␮m. (Reproduced with permission from Liu et al. [25]. Copyright © 2008 American Chemical Society.)

anti-Her2 antibodies to selectively recognize specific receptors on cancer cells. After staining by such a three-color SWNT mixture, cells were imaged by a confocal Raman microscope to obtain spectroscopic maps. The recorded spectrum was deconvoluted into three individual spectra of different SWNT “colors” by curve fitting, obtaining relative intensities of three “colors” at each imaging pixel. Three cancer cell lines with different receptor expressions were well differentiated by their labeled SWNT Raman “colors” in multicolor Raman images (Fig. 9.5c), demonstrating the ability of multiplexed Raman imaging by SWNTs with different isotope compositions, for probing and imaging several biological species simultaneously. By varying C12/C13 ratios in the nanotube synthesis, more SWNT Raman “colors” can be obtained. It is estimated that at least five SWNT Raman “colors” can be imaged and distinguished by this method. Besides the isotopically controlled G-band model shift, the RMB model of SWNTs also shows varied peak positions for nanotubes with different diameters. More Raman colors may be obtained if diameter-separated SWNTs are used [26, 27]. Compared with SERS nanoparticles to be discussed in the next section, Raman signals of SWNTs are weaker than some SERS nanoparticles on a per particle basis [23].

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However, a SERS nanoparticle contains many Raman active molecules absorbed on the surface of a heavy noble metal particle, while a SWNT is a single molecule with much less weight. Therefore if these two are compared at the same weight concentration, the Raman intensity of SWNTs could be close to that of SERS nanoparticles. Taken together, SWNTs are promising Raman tags for multiplexed Raman molecular imaging in vitro and potentially in vivo.

9.4 NANOPLATFORMS FOR SERS RAMAN IMAGING 9.4.1 Surface Enhanced Raman Scattering The development of SERS has generated large impacts in the areas of nanotechnology, single-molecule spectroscopy, as well as detection and analysis in biomedical systems. It has been proved that Raman intensities can be enormously enhanced by as much as 1014 –1015 -fold when Raman active molecules are adsorbed on a rough noble metal surface or nanoparticles [6, 28]. The huge enhancement factor allows ultrasensitive Raman spectroscopic measurement of molecules down to a single-molecule level, useful in biological detection and imaging. The exact mechanisms of the enhancement effect in SERS still require further explorations. Currently, two theories—the long-range electromagnetic effect [29] and the shortrange chemical effect—have been proposed to explain this unique phenomenon. These two mechanisms, although substantially different from each other, may simultaneously contribute to the SERS effect. In the electromagnetic mechanism, the surface plasmon of the metal is excited by incident photons, enhancing the electromagnetic field at the surface. If the laser frequency is in resonance with the surface plasmon, the field enhancement effect will be the highest. Silver and gold are thus the most popularly used metals for SERS because their plasmon resonance frequencies fall in the visible to NIR range. In addition, the plasmon oscillation must be perpendicular to the surface in order for scattering to occur. Therefore a rough metal surface or metal nanoparticles are required for the SERS effect [29, 30]. Although the electromagnetic effect applies for many molecules located close to the metal nanoparticle surface, it cannot fully explain the large magnitude of enhancement observed in many experiments. As reported by Xu et al. [31], the maximal enhancement of the electromagnetic field is about 1011 , which only occurs for a strongly coupled structure such as dimer configurations or sharp protrusions. Another effect, the chemical effect, also contributes to the SERS effect and bridges the gap to the observed maximal SERS enhancement of the order 1014 . In the chemical effect [32–34], the Raman active molecule is bound to a certain active site on the metal surface, allowing a charge transfer between each other [35]. The electronic transition energy between the highest occupied molecular orbital (HOMO) and lowest unoccupied molecular orbital (LUMO) in most molecules is usually higher than the energy of visible and NIR light. The charge transfer could lower the transition energy to be close to the incident laser frequency, inducing an enhancement effect similar to that of resonance Raman scattering. In contrast to the electromagnetic effect, which is effective in a long range (10–100 nm), the chemical effect requires direct chemical bonds between molecules and surface active sites (e.g., terraces and steps) [34], and thus only occurs in a very short range on the atomic scale.

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9.4.2 Fabrication of SERS Nanoprobes Numerous SERS nanoprobes have been developed in the past decade for biological detection and imaging. A typical SERS nanoprobe is composed of four major components (Fig. 9.6a). A metal nanoparticle core is needed to provide the SERS effect. Raman active dye molecules are absorbed on the metal particle to contribute Raman signatures. A surface coating is necessary to afford nanoprobes stability and biocompatibility. Targeting ligands (e.g., antibodies) are conjugated on those nanoprobes to allow selective binding of the nanoparticle to a specific target. Silver and gold are the most commonly used metals for SERS nanoparticles because their plasmon peaks are located in the visible to NIR range. Although silver often offers a higher SERS enhancement factor than gold, gold nanomaterials are preferred for SERS nanoplatforms used in biological imaging especially for in vivo imaging, owing to their plasmon peaks at higher wavelengths, close to or in the NIR window that is ideal for in vivo optical imaging. Besides nanoparticles with single metal composition, Au/Ag alloy nanoparticles [36] and core–shell nanoparticles [37] have also been developed for SERS Raman enhancement. The shape of nanomaterials is not limited to spherical particles. Many nanomaterials with interesting shapes such as gold nanorods [38] and nanocages [39] have also been utilized as SERS platforms. Those nanomaterials have the resonance plasmon in the NIR region and may be particularly useful for in vivo detection and imaging. In Raman active vibration models, the dipole moment of a molecule has to be kept constant during vibration. A number of small molecules, usually aromatic ones, are used as Raman active dyes in SERS nanoplatforms. Pyridine is the first Raman active molecule involved in SERS observation [40]. Many thiol containing aromatic molecules, such as thiophenol, naphthalenethiol, and 4-mercaptobenzoic acid, are widely employed as the Raman dyes on SERS nanoparticles [41–43]. Various fluorescent dyes such as rhodamine derivatives and Cy dyes are also useful Raman reporter molecules in SERS systems [6, 9, 36].

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FIGURE 9.6 Schematic drawings of a few commonly used SERS nanoprobes for biological detection and imaging. (a) A SERS nanoprobe based on a single metal nanoparticle. A coating layer envelops the metal nanoparticle labeled with Raman dyes. Targeting ligands such as antibodies are then linked on the SERS nanoparticle. (b) A SERS bead with many small metal nanoparticles set on a big silica bead. (c) A SERS nano- or microprobe with a nucleus formed by aggregated small metal nanoparticles. Raman dyes (green) are deposited on metal nanoparticles while targeting ligands are conjugated on the surface of the outer layer in (b) and (c).

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Because every molecule has its own unique Raman fingerprint, multiple targets can be identified in a signal measurement using SERS nanoparticles labeled by different Raman active dyes, allowing a high degree of multiplicity in imaging and detection [44]. A coating layer is required on SERS nanoparticles to protect Raman active dyes from falling off, prevent SERS enhancement of Raman signals of nontarget molecules nonspecifically adsorbed on the metal core, and impart nanoparticles biocompatibility. The first generation SERS Raman nanoprobes were synthesized by coattaching of antibodies and Raman dyes on Au nanoparticles without an additional protection layer [41]. Although functioning, those SERS Raman probes are less stable and have substantial nonspecific absorption signals because they are not well shielded from the external environment. A silica or glass layer coating is the most commonly used strategy in the fabrication of SERS nanoprobes [45–48]. Raman dyes are absorbed on metal nanoparticles first. A layer of silica is then introduced to encapsulate the nanoparticle. Undesired binding of interfering Raman active molecules on the metal surface are avoided by the silica coating, which also improves the stability of SERS Raman nanoprobes. However, bare silica-coated nanoparticles still tend to nonspecifically bind to proteins and cell surface due to limited hydrophilicity of the silica surface [49, 50]. An Additional hydrophilic coating such as PEG polymer may be helpful to further improve the performance of this type of SERS nanoprobes. Recently, the Nie group has developed a simple strategy to acquire stable and biocompatible SERS nanoprobes for in vivo Raman imaging [9]. In their work, green isothiocyanate was absorbed on the surface of a gold nanoparticle, which was further protected by thiolated PEG. Addition of PEG-SH did not replace Raman dye molecules absorbed on the gold surface and improved the stability and biocompatibility of nanoparticles. Nonspecific binding of nanoparticles on cells was also minimized. Interestingly, this method applies for other nonsulfur-containing positively charged Raman active molecules, which are absorbed on the gold surface due to an electrostatic force and delocalized ␲-electron interaction. To create “smart” SERS nanoprobes with molecular recognition ability, targeting ligands are conjugated to SERS nanoparticles. Various antibodies have been linked to SERS nanoprobes to detect or image different molecular species in biological systems [9, 45–48, 51, 52]. Oligonucleotides can also be used to functionalize SERS nanoparticles to make gene probes with high detection sensitivities [36, 44]. Besides SERS nanoplatforms with single nanoparticle cores, SERS beads with multiple metal nanoparticles in one bead have also been developed for Raman detection and imaging. In one type of SERS beads, small gold or silver nanoparticles labeled with Raman dyes are anchored on the surface of a big silica bead from 100 nm to a few micrometers (Fig. 9.6b) [45, 53]. The bead is further coated with a layer of silica to acquire better stability. Another type of SERS beads have Raman dye labeled small Ag or Au nanoparticles aggregated in the nucleus, and an outer lay of polymer or silica coating to envelop the beads (Fig. 9.6c) [51, 52]. Composite organic–inorganic nanoparticles (COINs) developed by Intel Corp. fall into this category [54, 55]. The SERS enhancement effect of such a coupled nanostructure in the “hot” bead core with metal nanoparticle aggregates could be higher than that of a single individual metal nanoparticle. However, the near micrometer sizes of some SERS beads are not ideal for certain applications such as in vivo molecular imaging. 9.4.3 Biological Detection and Diagnosis by SERS Detection of biological species by SERS spectroscopy has been widely explored in recent years. The Vo-Dinh group has carried out many pioneer studies in this area [56–59]. In an

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earlier work, they developed a SERS gene probe to selectively detect the DNA of human immunodeficiency virus (HIV) [56]. A primer sequence labeled with a Raman dye was used to amplify the HIV gene by polymerase chain reaction (PCR), affording amplified labeled strands, which were then captured by complementary oligonucleotides anchored on a substrate. Silver was evaporated on the substrate to enhance Raman signals of the dye labeled gene probe through SERS, achieving high sensitivity in the detection of the HIV gene. However, the silver evaporation step used in this protocol is not compatible for a conventional diagnostic lab. Silver colloid deposition was thus employed in a later study to replace the silver evaporation [59]. Similar strategies have been used for SERS detection of various cancer genes on substrates (Fig. 9.7a) [60–63]. Another type of nanoparticle based SERS detection method has been reported by the Mirkin group [36, 44, 64]. In those strategies, Raman dye labeled targeting gold nanoparticles were used to recognize DNA and RNA through hybridization of complementary sequences, or proteins by antibody–antigen binding (Fig. 9.7b). Silver deposition was then

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FIGURE 9.7 Schematic drawings of different SERS detection strategies in biological assays. (a) Raman labeled gene probes for DNA detection. Ag deposition is applied to enhance Raman signals by SERS [56, 60–63]. (b) Gold nanoparticle based SERS detection of DNA or proteins [36, 44, 64]. The scattering of Raman dyes on gold nanoparticles is further enhanced by deposition of silver nanoparticles. (c) A conventional strategy of protein sandwich assay using SERS nanoprobes [41, 52, 54, 65, 66]. (d) Protein microarray assay using SWNTs as the Raman tag [20]. Ultrasensitive detection is achieved by combining the resonance enhancement effect and the SERS enhancement effect. (Reproduced with permission from Cao et al. [44]. Copyright © 2002 Science Magazine, 2008 Nature Publish Group.)

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introduced on those gold nanoparticles captured either on a substrate or a glass bead, to obtain dramatically enhanced SERS effects for Raman detection. Multiplexed detection of six DNA targets with high sensitivity (20 femtomolar) was realized by using six different Raman dyes with easily distinguishable sharp spectroscopic fingerprints [44]. Many other groups have also reported protein or gene detection with SERS nanoprobes (Fig. 9.7c) [41, 52, 54, 65, 66]. Ni et al. [41] used antibody conjugated SERS nanoparticles to replace fluorescence dyes in a sandwich immunoassay and achieved a detection limit at the nanomolar level. In another work by Sun et al. [65], thiolated oligonucleotide functionalized gold nanoparticles were labeled with various nonfluorescent Raman dyes for multiplexed DNA detection. Up to eight different DNA targets were detected simultaneously through SERS Raman spectroscopy with a detection limit at about 10−7 M. COIN nanoprobes have also been used in the protein sandwich assay, showing a high sensitivity of 1 pg/mL in the detection of interleukin-2 protein [54]. In a recent study by Dai and co-workers, the intense resonance Raman scattering of SWNTs was combined with SERS, extending the detection limit of traditional fluorescence assays from approximately 1 pM [67] to the femtomolar level [20]. In this work, the sandwich protein assay was performed on a gold coated substrate with antibody conjugated SWNTs as Raman probes (Fig. 9.7d). After annealing at 400 ◦ C in H2 , the gold film aggregated and formed small particle islands, which afforded a nearly 100-fold increase in SWNT Raman scattering intensity, boosting the detection sensitivity by two orders of magnitude. Multiplexed protein microarray was also realized using isotopically modified SWNTs with shifted G-band peaks. 9.4.4 Molecular Imaging with SERS Nanoprobes SERS nanoprobes conjugated with targeting ligands can be used to selectively recognize specific receptor molecules. In vitro Raman imaging is thus able to reveal the molecular information of targeted cells. Lee and co-workers used antibody conjugated glass coated SERS beads to recognize HER2 and CD10 proteins expressed on cancer cells and showed high targeting specificity [45]. In a later work by the same group, simultaneous multiplexed Raman imaging of two proteins on cells and tissue slices was performed, also using antibody conjugated SERS beads [53]. Another Korean group used Au/Ag core/shell nanoparticles as the SERS platform with rhodamine 6G Raman dye placed between the core/shell structure [37]. Those nanoparticles were further coated with a PEG layer and conjugated with antiPLC␥ 1 antibody for cell labeling and Raman imaging. Instead of using additional Raman dyes, Kneipp et al. [68, 69] were able to image unlabeled gold nanoparticles inside cells by Raman spectroscopy and detect chemical changes in the intracellular environment of these nanostructures. Biological molecules inside cells were adsorbed on the gold nanoparticle surface after endocytosis of particles, affording strong SERS enhanced Raman signatures of those molecules to be detected by Raman spectroscopy. In vivo Raman detection and imaging with SERS nanoprobes have been relatively less explored. To date, there have been two groups reporting on Raman imaging in small animal models using SERS nanoparticles as contrast agents. Nie and co-workers [9] coated Raman dye labeled gold nanoparticles with PEG. and an infrared dye diethylthiatricarbocyanine (DTTC) was used as the Raman reporter. Those nanoparticles were further conjugated with a single chain variable fragment (ScFv) antibody to recognize the epidermal growth factor receptor (EGFR), an important cancer biomarker overexpressed on various types of cancer cells. The obtained SERS nanoprobes exhibited strong Raman signals and high

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FIGURE 9.8 In vivo tumor targeting and Raman spectroscopic detection by using SERS nanoprobes [9] (a, b) Raman spectra obtained from the tumor and the liver locations 5 hours after injection of targeted (a) and nontargeted (b) SERS nanoparticles. High tumor uptake of targeted nanoparticles was evidenced by the strong Raman signals from the tumor. (c) Photographs showing a laser beam focusing on the tumor site or on the anatomical location of liver. (Reproduced with permission from Qian et al. [9]. Copyright © 2008 Nature Publishing Group.)

targeting specificity in vitro on EGFR positive neck squamous carcinoma Tu686 cells. After intravenous injection of SERS nanoprobes into mice bearing Tu686 xenograft tumors, Raman spectra were recorded at the tumor site and the anatomical location of the liver (Fig. 9.8). Mice injected with targeted SERS nanoparticles showed strong Raman signatures of DTTC reporter molecules at the tumor site, indicating high tumor uptake of nanoprobes owing to the specific ScFv antibody–EGFR binding (Fig. 9.8a). Nontargeted nanoparticle injected mice, in contrast, showed low Raman signals in the tumor (Fig. 9.8b). Although nanoparticles exhibited high uptake in the reticuloendothelial system (RES) including liver and spleen, the observed Raman signals from the liver were low due to the limit of tissue penetration depth of light in optical imaging (Fig. 9.8a,b). This work is the first success of using SERS nanoparticles for in vivo detection of cancer biomarkers with potential high sensitivity. In another work, Gambhir and co-workers have used glass coated gold nanoparticles as the SERS platform for in vivo Raman imaging in mice [23]. Several types of SERS nanoparticles labeled by different Raman reporters with distinctive spectral fingerprints were subcutaneously injected into mice for imaging (Fig. 9.9). Raman spectroscopic scanning was carried out on a live mouse, revealing different Raman “colors” after spectral

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FIGURE 9.9 In vivo Raman imaging with SERS nanoparticles [23]. (a, b) Raman spectra of two types of SERS nanoparticles (a, S421; b, S440) with different Raman labels. (c) Raman spectrum of a mixture of S421 and S440 nanoparticles. (d) A multicolor Raman image of a mouse subcutaneously injected with S421 nanoparticles, S440 nanoparticles, and a mixture of the two nanoparticles. Artificial red and green colors were assigned to Raman signals from S421 and S440 nanoparticles, respectively. (e) A whole-body Raman image of a mouse intravenously injected with SERS nanoparticles. (f) A Raman image of liver area showing high uptake of nanoparticles in the liver. (Reproduced with permission from Keren et al. [23]. Copyright © 2008 National Academy Society.)

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unmixing. In addition to the proof-of-principle imaging of SERS probes under the skin, they intravenously injected a mouse with SERS nanoparticles and performed whole-mouse imaging, observing strong Raman signals from the liver owing to the RES uptake of nanoparticles. This is the first study showing the feasibility of the whole-animal Raman imaging in a multiplexed manner. However, targeted multiplexed molecular imaging in vivo with SERS nanoprobes requires further efforts.

9.5 CONCLUSION AND PERSPECTIVES Raman imaging has various advantages over fluorescence imaging. The sharp spectral signatures impart to Raman imaging a higher imaging multiplicity compared with fluorescence imaging. The benefit is the most obvious in the case of in vivo imaging, when excitations and emissions in the narrow NIR window are necessary to allow the maximal tissue penetration of light. The narrow peak features in a Raman spectrum also enable us to easily distinguish Raman peaks and subtract the autofluorescence background from a recorded spectrum. Removal of interfering background allows higher detection and imaging sensitivity compared with the fluorescent technique, whose detection limit is largely affected by the autofluorescence from biological species. Moreover, Raman tags generally have higher photostability than fluorescent molecules, good for long-term imaging and tracking in biological systems. In this chapter, the basic physics of Raman light scattering, resonance Raman scattering, and surface enhanced Raman scattering are briefly introduced. Applications of resonance Raman scattering of carbon nanotubes in biological imaging are summarized. Pioneer works using SWNTs for in vivo Raman molecular imaging and in vitro multiplexed imaging are described. In a related vein, various SERS nanoplatforms and their applications are reviewed. Commonly used strategies of fabrication of SERS nanoprobes, gene and protein assays with SERS, and biological imaging with SERS nanoparticles are discussed. Despite the rapid progress and great potential in Raman imaging, a few challenges still hamper future applications of Raman nanoplatforms. Current Raman imaging suffers from slow scanning speed. Point-by-point scanning is the most commonly used imaging methodology with a spectrum at each mapping point recorded by a spectrometer. This is a relatively slow process limited by both instruments and the nanoplatforms. Regarding instruments, new Raman microscopes with better designs and the capability of fast mapping are on the market now. However, Raman mapping of a large area with high spatial resolution is still a time-consuming process that often requires tens of minutes or up to a few hours. On the other hand, the absolute intensities of Raman scattering of Raman nanoplatforms including SWNTs and SERS nanoparticles are still much weaker than the fluorescence of high quantum yield fluorescent dyes. Brighter Raman nanoplatforms are desired for future imaging applications. Compared with SERS nanoparticles, SWNTs have a weaker Raman signal per particle. However, currently used SWNTs in Raman detection and imaging are mixtures of nanotubes with different chiralities. Therefore only a small percentage of SWNTs are in resonance with the laser source. Separated SWNTs with a single chirality composition are expected to exhibit a much stronger resonance effect when a proper laser wavelength is applied to match the E11 or E22 energy of this particular type of SWNTs. In addition, significant SERS enhancement of SWNT Raman scattering has been reported earlier on the substrate for

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biological detection. Efforts may be put to develop SERS enhanced SWNT nanoplatforms for Raman imaging to achieve stronger signals and faster mapping speeds. In the aggregated noble metal nanoparticle system, the interparticle distance is very small, affording a much more localized region with ultrahigh enhancement of the electromagnetic field [31]. SERS nanoplatforms such as COINs [54] with a core of aggregated metal nanoparticles (Fig. 9.6c) are expected to show stronger SERS enhancement effect than single metal nanoparticles. However, the aggregation of metal nanoparticles is not perfectly controlled, raising potential batch-to-batch variation problems of those SERS platforms. For in vivo Raman imaging, the safety issue of nanoplatforms is a concern. Like other nanomaterials, carbon nanotubes and SERS nanoparticles all tend to accumulate in RES organs after intravenous administration into animals. Better surface coating is able to reduce RES uptake. Although PEGylated SWNTs appear to be safe in mice at tested doses, much more work is required to fully understand the potential side effects of those inorganic nanomaterials to a living subject. Besides conventional Raman spectroscopy, coherent anti-Stokes Raman scattering spectroscopy (CARS), a nonlinear optical process with three laser beams involved, has also been widely used in biological imaging [8, 70–72]. CARS spectroscopy is free of fluorescence background and yields higher total signals than Raman spectroscopy owing to the coherent addition of emissions from molecules in the laser beam focused region. Therefore CARS imaging has a much faster scanning speed and enables real-time imaging of biological systems with chemical selectivity in the absence of additional labels. However, though without interference from the fluorescence, CARS spectroscopy contains an inherent nonresonant background, which limits the detection of CARS signals from molecules at low concentrations. Studies have shown that gold nanoparticles are able to enhance CARS signals of their surrounding molecules, allowing indirect detection of gold nanoparticles in biological samples [73, 74]. Combination of CARS spectroscopy and nanotechnology may offer new opportunities in biological imaging. In summary, Raman imaging with nanoplatforms has been intensively investigated in the past decade. Various problems including relatively weak Raman signals, slow imaging process, and concerns of in vivo toxicity of inorganic nanoparticles remain to be resolved. Nevertheless, Raman imaging is a promising technique for future highly sensitive multiplexed molecular imaging in vitro and in vivo.

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63. Volkan, M.; Stokes, D. L.; Vo-Dinh, T. A sol-gel derived AgCl photochromic coating on glass for SERS chemical sensor application. Sensors and Actuators B—Chemical 2005, 106(2), 660–667. 64. Cao, Y. C.; Jin, R. C.; Nam, J. M.; Thaxton, C. S.; Mirkin, C. A. Raman dye-labeled nanoparticle probes for proteins. J. Am. Chem. Soc. 2003, 125(48), 14676–14677. 65. Sun, L.; Yu, C. X.; Irudayaraj, J. Surface-enhanced Raman scattering based nonfluorescent probe for multiplex DNA detection. Anal. Chem. 2007, 79(11), 3981–3988. 66. Fritzsche, W.; Taton, T. A. Metal nanoparticles as labels for heterogeneous, chip-based DNA detection. Nanotechnology 2003, 14(12), R63–R73. 67. Espina, V.; Woodhouse, E. C.; Wulfkuhle, J.; Asmussen, H. D.; Petricoin, E. F., 3rd; Liotta, L. A. Protein microarray detection strategies: focus on direct detection technologies. J. Immunol. Methods 2004, 290(1-2), 121–33. 68. Kneipp, K.; Haka, A. S.; Kneipp, H.; Badizadegan, K.; Yoshizawa, N.; Boone, C.; Shafer-Peltier, K. E.; Motz, J. T.; Dasari, R. R.; Feld, M. S. Surface-enhanced Raman spectroscopy in single living cells using gold nanoparticles. Appl. Spectrosc. 2002, 56(2), 150–154. 69. Kneipp, J.; Kneipp, H.; McLaughlin, M.; Brown, D.; Kneipp, K. In vivo molecular probing of cellular compartments with gold nanoparticles and nanoaggregates. Nano Lett. 2006, 6(10), 2225–2231. 70. Tolles, W. M.; Nibler, J. W.; Mcdonald, J. R.; Harvey, A. B. Review of theory and application of coherent anti-stokes Raman-spectroscopy (CARS). Appl. Spectrosc. 1977, 31(4), 253–271. 71. Zumbusch, A.; Holtom, G. R.; Xie, X. S. Three-dimensional vibrational imaging by coherent anti-Stokes Raman scattering. Phys. Rev. Lett. 1999, 82(20), 4142–4145. 72. Xie, X. S.; Yu, J.; Yang, W. Y. Perspective—living cells as test tubes. Science 2006, 312(5771), 228–230. 73. Ichimura, T.; Hayazawa, N.; Hashimoto, M.; Inouye, Y.; Kawata, S. Local enhancement of coherent anti-Stokes Raman scattering by isolated gold nanoparticles. J. Raman Spectrosc. 2003, 34(9), 651–654. 74. Tong, L.; Lu, Y.; Lee, R. J.; Cheng, J. X. Imaging receptor-mediated endocytosis with a polymeric nanoparticle-based coherent anti-stokes raman scattering probe. J. Phys. Chem. B 2007, 111(33), 9980–9985.

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CHAPTER 10

Single-Walled Carbon Nanotube Near-Infrared Fluorescent Sensors for Biological Systems JINGQING ZHANG and MICHAEL S. STRANO Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, Massachusetts, USA

10.1 INTRODUCTION Semiconducting single-walled carbon nanotubes (SWNTs) are a unique class of materials that fluoresce in the near-infrared (NIR) range, typically from 820 to above 1600 nm in wavelength, and are unique in that they are one of the very few photostable species that do so. The optical properties of SWNTs have been thoroughly investigated both theoretically and experimentally, and a complete review of their photophysics is beyond the scope of this chapter (readers are referred to Refs. 1–3). Rather, we focus on summarizing the work from our laboratory and others in using near-infrared fluorescence from SWNTs to detect a broad range of biomolecules. This chapter first summarizes the inherent properties of carbon that motivate this work, then describes the development of sensor assays of various types, and examines future efforts and applications of these platforms to solve biological and biomedical problems.

10.2 FAVORABLE PROPERTIES FOR SENSOR APPLICATIONS Semiconducting single-walled carbon nanotubes have several advantages for sensor applications in the life sciences: (1) completely photostable emission, with no blinking or bleaching, even at high fluence [4]; (2) near-infrared emission where many biological tissues are transparent and many others do not autofluoresce [4, 5–8]; (3) sensitivities that enable even a single-molecule docking to the SWNT sidewall to be detected [3, 9–11]; (4) if encapsulated or functionalized appropriately, they are noncytotoxic and demonstrate favorable pharmacological clearance and biodistribution in vivo [12]; (5) they lack surface states and the requirement of surface passivation to emit, meaning that target molecules can Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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access the SWNT surface reversibly in many cases where it can be detected [1, 13]; (6) there are large classes of proteins and enzymes [11, 13], DNA oligonucleotides [14–16], peptides [17, 18], and polymers [19] that can encapsulate or wrap SWNTs to enable selective interaction [19] and detection of target molecules, a subfield that is just beginning.

10.3 TISSUE IMPLANTABLE SENSORS Any fluorescent sensor designed for long-term and real-time monitoring in vivo has two principal figures of merit (FOM) that need to be considered, quantum yield and photostability. The quantum yield determines the maximum emitted signal from a sensor for a given excitation intensity [20–23]. However, photobleaching is a property of all organic [7] and most nanoparticle fluorophores [6] investigated to date except for SWNTs [24]. Fluorophores that undergo photochemistry necessarily have limited viability with respect to sensor lifetime (Table 10.1) [25]. The photostability of the fluorescent probe dictates the lifetime of the optical device. For example, for electrochemical glucose sensors, a maximum baseline drift of 20% is tolerated before recalibration is needed. If this is used as the upper limit on the working lifetime (␶ ), one can express it in terms of the photobleaching rate constant k [25]: ␶ =−

ln (0.8) k

(10.1)

Table 10.1 uses photobleaching rate constants obtained from the literature to estimate lifetimes of some common organic fluorophores and quantum dots [25]. These rate constants are highly dependent on the environment and the experimental setup. However, even highly optimized quantum dots still suffer from prominent photobleaching. SWNTs are the only optical probe investigated to date that possesses no photobleaching threshold [24, 26]. In order to have continuous, real-time monitoring, a nonphotobleaching sensor is necessary. In addition, for any in vivo sensor, tissue transparency is another necessary prerequisite. The absorption and scattering of the emitted light within the tissue is strongly wavelength dependent (Fig. 10.1) [4, 25] and limits the fluorescent probe that can be used for any in vivo detection purpose. SWNT sensors have advantages for sensing in tissues as the fluorescence is of longer wavelength where these materials have greater transparency. In many fluorometric assays, autofluorescent background is a key limitation. Near-infrared TABLE 10.1 Comparison of Photobleaching Tendency of Common Organic and Nanoparticle Fluorophores Fluorophore IR-Dye 78-CA [6] Cy5 [6] Indocyanine Green [7] Type II NIR QD [6] SWNT [24] a

Photobleaching Rate Constant (h−1 )

Fluence (mW/cm2 )

Nominal Sensor Lifetimea

250.92 20.52 0.0412 0.0827 0

600 600 28 600 1.0 × 106

3.2 s 39.1 s 5.4 h 2.7 h ∞

Estimated optical sensor device lifetime calculated from Eq. (10.1).

Source: Barone et al. [25].

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µ (oxy)

100 Cy5 Absorption coefficient

219

µ (deoxy) N-IR Qdots

Average

10

1

Indocyanine Green SWNT

0.1 600

700

800

900

1000

1100

Wavelength (nm)

FIGURE 10.1 Absorption coefficient of human whole blood (the dominant absorber for in vivo applications) for oxygenated, deoxygenated, and the average value. Potential fluorophores are compared. (Reproduced with permission from Barone and co-workers [14, 24, 25].)

fluorophores avoid this background that is typically strong at visible wavelengths but not in the near-infrared where SWNTs fluoresce. Consider an optical sensor implanted a distance d into a tissue sample with absorption coefficient ␮ and quantum yield ␸. A simple one-dimensional, absorption and fluorescence model can be used to estimate the signal intensity of different fluorophores and the lifetime of the device [25]: Is = I0 ␸e−2␮d−k␶

(10.2)

where I 0 = incident excitation intensity I s = fluorescent optical signal outside the body k = pseudo-first-order photobleaching rate constant ␶ = aggregate excitation exposure lifetime For the maximum emitted signal from a sensor implanted a distance d into tissue, the fluorophore properties should be such that the figure of merit (FOM) = ␸e−2␮d is maximized. Table 10.2 reports this value for some common organic and nanoparticle fluorophores. This calculation clearly indicates that visible fluorophores are strongly attenuated in blood samples. Figure 10.2 shows a wavelength window (near-infrared wavelength) at which SWNTs naturally fluoresce, whereas human tissues strongly absorb and autofluoresce at visible wavelengths. 10.3.1 Single-Walled Carbon Nanotubes as Fluorometric Glucose Sensors The first example of a SWNT-based fluorescence sensor was by Barone et al. [13] and introduced the concept of modulating the fluorescence of SWNTs in response to a target analyte such as glucose.

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TABLE 10.2 Comparison of Commonly Utilized Visible and Near-Infrared Organic and Nanoparticle Fluorescent Probes Based on Quantum Yield (Q.Y.) and Absorbance in Human Whole Blood (␮) [4, 25] Standards Cy5 Fluorescein Rhodamine 6G Rhodamine B Indocyanine Green Indocyanine Green Type II NIR QD SWNT a Based

Q. Y. (%)

Conditions for Q. Y. Excitation ␮ (cm−1 ) ␮ (cm−1 ) Measurement (nm) Oxy Deoxy

27 95 95 31 0.266

PBS 0.1 M NaOH, 22 ◦ C Water Water Water (0.15 g/L)

620 496 488 514 820

2 150 200 110 0.96

60 120 105 190 0.77

3.20 × 10−28 5.23 × 10−118 3.30 × 10−133 1.60 × 10−131 4.72 × 10−4

1.14

Blood (0.08 g/L)

830

1.01

0.7788

1.91 × 10−3

13 0.1

PBS PBS

840 1042

1.05 0.889

0.778 0.12

2.09 × 10−2 3.65 × 10−4

FOMa

on a 1-cm implanted sensor device, an average ␮ is utilized.

Source: Adapted from Saito et al. [3].

Because of the large need, with over 194 million diabetics worldwide [27], and the tangible benefits of frequent blood glucose measurements [28], a significant effort has been made toward the development and production of a real-time, continuous glucose sensor capable of in vivo operation. Current glucose sensors operate electrochemically and are the standard, partly because they are already commercially available [29] and have FDA approval for use in humans. The majority of these devices utilize transdermal electrodes that require frequent replacement every few days. Devices in development utilize telemetry, and require some form of surgical implantation [30]. If instead, sensors are based on optical signal transmission, they can either be minimally invasive, subcutaneously implanted fluorescence-based devices or involve completely noninvasive spectroscopic detection

1

8

0.8

6

0.6

4

0.4

2

0.2

Blood Abs.

0 400

Water Abs. 650

900

1150

Normalized fluorescence

Absorbance (cm-1)

10

0

1400

Wavelength (nm)

FIGURE 10.2 Nanotubes fluoresce at near-infrared wavelengths where tissue does not absorb. (Adapted from Wray and et al. [4].)

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techniques. The former are typically based on either FRET between fluorophores [5, 31–33], or some type of competitive binding [34, 35], while the latter currently lack analyte specificity and dynamic range [36], as detailed in the reviews of Moschou, Wickramasinghe, and Wilson [29, 37, 38]. In addition to fluorescence-based techniques, several photonics applications have also been proposed for glucose sensing [39, 40]. However, the high-resolution gratings and notch filters necessary for Raman-based schemes are expensive [39, 40]. A fluorescencebased device would require lower resolution and would be relatively inexpensive due to the affordability of current instrumentation and its amenability to microfabrication techniques [41]. Additionally, a small fluorescence based device could readily be implanted beneath the skin without requiring surgical procedures. In an effort to produce a glucose sensor based on SWNT fluorescence, the Strano lab has explored noncovalent methods of modulating the nanotube emission. In this work, new types of noncovalent functionalization using electron-withdrawing molecules were shown to provide sites for transferring electrons in and out of the nanotube. More specifically, they have shown [13] that it is possible to link enzyme activity or protein binding to electron withdrawal from or donation to a nanotube through a variety of surface chemistries [42, 43]. In this first glucose-sensing work, an electroactive species such as potassium ferricyanide, K3 Fe(CN)6 , can irreversibly adsorb at the surface of the nanotube and shift the Fermi level into the valence bands, or quench the emission after photoexcitation. This quenched state can be partially reversed by hydrogen peroxide (H2 O2 ), resulting in a fluorescence recovery. As glucose oxidase catalyzes the reaction of ␤-d-glucose to the d-glucono-1,5lactone with a H2 O2 coproduct, it is possible to couple the carbon nanotube fluorescence to the glucose concentration [13]. Because the ferricyanide adsorption is irreversible, the sensing solution can be loaded into a sealed dialysis capillary where glucose is free to diffuse across the capillary boundary while the nanotube sensing complex is retained (Fig. 10.3). The nanotube fluorescence from such a capillary is easily imaged in the NIR through a human epidermal tissue sample. The fluorescence of one nanotube is shown to respond to glucose after an 80-s transient, even in a whole blood specimen maintained at 37 ◦ C and pH 7.4. The calculated detection limit is approximately 2.2 molecules/nm of nanotube length. The results demonstrate new opportunities for nanoparticle optical sensors that operate in strongly absorbing media of relevance to medicine or biology. In addition, Barone et al. [25] have compared a direct concentration measuring sensor with flux-based sensors and concluded that the former is more stable to membrane biofouling.

10.4 DNA SENSORS Fluorescence detection of specific DNA sequences through hybridization with a complementary DNA probe has many applications in biological and medical sciences. Jeng et al. [14] have demonstrated the optical detection of DNA hybridization on the surface of single-walled carbon nanotubes (SWNTs) in solution through a SWNT bandgap fluorescence modulation. In this work, they first developed a novel method to suspend SWNTs using DNA and remove free DNA through a two-step dialysis method. The addition of complementary DNA (cDNA) to this DNA–SWNT suspension caused a 2-meV hypsochromic shift in the (6,5) SWNT fluorescence. In contrast, the addition of a noncomplementary DNA (nDNA) does not introduce any significant shift. The normalized energy shift can be

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(a)

HO HO

O OH OH OH

Enzyme/ SWNT solution

Porous dialysis capillary 13 kDa MW cutoff 105 5

nIR image of filled capillary beneath tissue sample

4.5

200 µm

3.5 3 2.5 2 1.5 1

Epidermal Tissue Sample

0.5

(b)

0

1 cm

1.10

(i)

Relative Intensity

(c) 4.2 mM

0.80

2.4 mM

0.50

1.4 mM 0.20 0

10

Time (min)

Relative Intensity

4

20

0.80 0.60 0.40 0.20 0.00 0

5

10

Glucose concentration (mM)

FIGURE 10.3 (a) A 200 ␮m × 1 cm, 13-kDa microdialysis capillary, shown to scale on a human finger, is loaded with the nanotube solution allowing glucose to diffuse through the membrane with the containment of the sensing medium. Placing the capillary beneath a human epidermal tissue sample (above) and the nanotube fluorescence can be clearly mapped from the capillary, seen in the two-dimensional (2D) profile. (b) After ferricyanide surface reaction at 37 ◦ C and pH buffered at 7.4 (i) the fluorescence response to the addition of 1.4 mM, 2.4 mM, and 4.2 mM glucose is scaled by the difference between minimum and maximum intensities. (c) The response function relates the normalized intensity to the local glucose concentration in the range of blood glucose detection with a type I absorption isotherm. (Adapted from Barone’s work [1, 3, 13].)

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fitted to a normalized exciton binding energy, which is a function of the surface coverage of DNA on the SWNT, yielding an estimated initial fractional coverage of 0.25, and a final coverage of 0.5 due to DNA hybridization. As the exposed area on the SWNT increases, the local dielectric environment of the SWNT changes, resulting in a decrease in the SWNT emission energy [43]. The hybridization process on the nanotube is found to be significantly slower than hybridization of the same sequence free in solution. Therefore the kinetics and thermodynamics of this hybridization mechanism have been studied to elucidate the reasons for this discrepancy in hybridization rate. The kinetic rates for hybridization were fitted using a two-step mechanism in which the DNA first adsorbed on the nanotube in a fast step, followed by a much slower hybridization step. The rates were also measured at different temperatures so the energetics could be determined. The conclusion of this study was that the cause for the slow hybridization was energetic. Hybridization on the SWNT was compared to transition state thermodynamics where the transition state was represented by the stage when DNA was adsorbed to the SWNT, but had not yet hybridized. Both the original DNA and cDNA strands are adsorbed to the SWNT before hybridization, resulting in a reduced initial free energy. Therefore the effective activation energy is increased between the unhybridized and transition states for DNA–SWNT (Fig. 10.4). As a follow-up work, Jeng et al. [45] have shown for the first time that the detection of a single nucleotide polymorphism (SNP) is indeed possible using SWNT fluorescence. It is known that genomic mutations could lead to genetic disorders [46] and predisposition to diseases [47, 48], such as cancer. The most common mutation investigated so far is the single nucleotide polymorphism (SNP) [47–49], a type of variation where one base can be occupied by either of two different bases [46, 47, 50]. In this work, a SNP sensor was synthesized by adsorbing a probe DNA on the SWNT surface via the two-step dialysis described earlier. The presence of the SNP DNA strand can be detected through an emission energy increase in the peak maximum of the (6,5) nanotube fluorescence, and distinguished from the response caused by completely complementary DNA, as shown in Figure 10.5. Kinetics experiments suggested a two-step model including a fast adsorption step, followed by a slower hybridization step [51]. In addition, the slower

2.5 Normalized Intensity

ΔE (meV)

2 1.5 1 0.5 0 -0.5

0

500 1000 cDNA or nDNA (nM) (a)

1500

1.05 1 0.95 0.9 0.85 987

992 997 Wavelength (nm) (b)

FIGURE 10.4 (a) Addition of complementary DNA (cDNA) causes an increase in energy of the steady state (6,5) fluorescence peak while there is negligible energy change with noncomplementary DNA (nDNA). The solid line is a fit of the dielectric model to the cDNA energy shifts. (b) Sample spectra of the fluorescence peak blue shift with cDNA addition. (Reproduced with permission from Jeng et al. [14].)

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(a)

(b)

FIGURE 10.5 The presence of a single nucleotide polymorphism (SNP) in the complement results in incomplete hybridization on the nanotube surface and a smaller surface coverage. (a) Transient detection of the energy shifts for cDNA and SNP at 37 ◦ C were fitted with a two-step model of adsorption followed by hybridization of the cDNA or SNP. While the equilibrium constants are the same, the partial hybridization of SNP comes to steady state slower than cDNA hybridization. (b) A diagram showing the different nanotube coverage that can result from complementary DNA (cDNA), and cDNA containing a SNP.

path to reach steady state of the SNP strand is consistent with a disruption in hybridization at the mismatched nucleobase pair in the center of the DNA sequence. The diagram in Figure 10.5b illustrates how full hybridization of the complement with the probe DNA that is already on the nanotube would result in denser DNA packing and more expulsion of water than the partial hybridization that would occur between the SNP and probe DNA. Our observations, that the cDNA + probe DNA on the SWNT result in a higher fluorescence energy than the SNP + probe DNA on the SWNT is consistent with work by Ohno and others [52, 53]. Although the (6,5) nanotube can be used to detect SNP for this particular sequence, this method cannot yet be applied to any sequence generically. One reason for this discrepancy is the sequence-dependent secondary structures that can form. More work is needed to develop a generic method to detect any DNA sequence using SWNT fluorescence.

10.5 SENSORS FOR CELLULAR IMAGING As mentioned earlier, a fluorescence-based SWNT senor has the capability for real-time, long-term cellular and tissue-based monitoring. This is made possible because fluorescence

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does not blink or photobleach and lies within the NIR wavelength window where water or media do not absorb, scatter, or autofluoresce. Heller et al. [15] used the fluorescence of SWNTs to study the B–Z transition of DNA. This work identified the transition of DNA secondary structure from an analogous B to Z conformation, which modulates the dielectric environment of the SWNT around which it is adsorbed. The SWNT bandgap fluorescence undergoes a red shift when an encapsulating 30-nucleotide oligomer is exposed to cations that bind to the nucleobases and screen the charged backbone of DNA. The transition is thermodynamically identical for DNA on and off the nanotube, except that the propagation length of the former is shorter by fivesixths. The magnitude of the energy shift is described by using an effective medium model and the DNA geometry on the nanotube sidewall. The detection of the B–Z change was demonstrated in whole blood, tissue, and from within living mammalian cells. Each type of SWNT has two distinct optical modes for sensing, intensity modulation, and emission wavelength shift. Heller et al. [16] showed that a pair of single-walled nanotubes provides at least four modes that can be used to uniquely fingerprint agents by the degree to which they alter both the emission band intensity and wavelength. They validated this identification method in vitro by demonstrating the detection of six genotoxic analytes, including chemotherapeutic drugs and reactive oxygen species, which are spectroscopically differentiated into four distinct classes, and also by demonstrating single molecule sensitivity in detecting hydrogen peroxide. Finally, the SWNT sensors detected and identified these analytes in real time within live 3T3 cells, demonstrating multiplexed optical detection from a nanoscale biosensor and the first label-free tool to optically discriminate between genotoxins (Fig. 10.6). A major challenge in the synthesis of nanotube or nanowire sensors is to impart selective analyte binding through means other than covalent linkages, which compromise electronic and optical properties. Kim et al. [19] synthesized a 3,4-diaminophenylfunctionalized dextran (DAP-dex) wrapping for single-walled carbon nanotubes (SWNTs) that imparts rapid and selective fluorescence detection of nitric oxide (NO), which is a gaseous free radical playing an important role as an intracellular and intercellular messenger for signaling. The near-infrared (NIR) fluorescence of SWNTDAP-dex is immediately and directly bleached by NO, but not by other reactive nitrogen and oxygen species. It is found that the fluorescence bleaching of SWNTDAP-dex is reversible and caused by electron transfer from the top of the valence band of the SWNT to the lowest unoccupied molecular orbital (LUMO) of the NO radical. The resulting optical sensor is able to spatiotemporally detect NO produced by stimulating inducible NO synthase (iNOS) in Raw 264.7 macrophage cells, as shown in Figure 10.7. They have also demonstrated the potential of the optical sensor for in vivo detection of NO in a tissue model. The ability of SWNTDAP-dex optical sensor to provide temporal and spatial distribution of NO at nanomolar ranges in a living organism demonstrates its potential for a variety of biological applications.

10.6 SINGLE-MOLECULE SENSORS As mentioned earlier, SWNTs are very sensitive to perturbation of their electronic structure. This type of sensitivity can go down to the single-molecule level. Cognet et al. [10] observed single-molecule adsorption and desorption events by studying single step quenching and brightening of fluorescence of sodium dodecylbenzenesulfonate wrapped SWNT. Jin et al. [11] have developed a type 1 collagen film, similar to those used as 3D cell scaffolds for tissue engineering, embedded with SWNT sensors capable of reporting

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A

OH NH2

O RO P O O–

O O P –O O

O

O

O O P O O– N

N O

NH

N O NH H2N

O

O RO P O O OR O–

N

N

O O P –O O

O

O O P O O– N

O

O RO P O O–

O OR

N O

Cl

NH

O

N O

NH

N H2 N

O

NH

O

N O

NH

O

1015

•OH

1050

980

Wavelength (nm)

Alkylating agent

1.0

E

O O P O O–

N O

NH

F

0.9

H2O 2

1.00

0.96

(6,5)

0.94

0.7

0.92

(7,5)

O

100

200

300

O O P O O– N N NH

X O

O RO P O O–

O OR N O

NH

G

1.0 0.8

1O

H

2

0.6 0.4

10

20

30

O

O O P –O O

NH

O

O O P O – O O

O N

NH

N H2N

NH

O

O OR

O N NH O

100

200

300

0.998 0.997

1.00

Time (min)

0.999 0.998 0.997

Time (min)

0.99

200

300

400

0.0

0.5

0.0

0.5

1.0

1.5

2.0

2.5

1.0

1.5

2.0

2.5

1.000

0.98

0.996

0.996

0.6

400

Norm. Energy

0.999

I

0.8

1.001

1.000 Norm. Energy

1.000

•OH

1.0

0.4 0

40

1.001

100

O

N O

O

0.2 0

400

1.001

0

NH

O

2

N H2N

O

0.90 0

N

1050

0.98

0.8

O NH

H2N

Wavelength (nm)

0.6

Norm. Energy (E/Emax )

O

Norm. Energy

Norm. Intensity (I/Imax)

1.1

1015

N

O

O

•OH O RO P O O–

D

NH

O

1

C

(7,5) 2

N

Norm. Energy

Relative Int. 980

H2O 2

OR

O

H2 O2

Norm. Energy

Relative Int.

Alk. B agent

1O

O O P O O– N

O

N

Alkylating agent

(6,5)

O O P O O–

O

N

0.999 0.998 0.997 0.996

0

10

Time (min)

20

30

40

Time (min)

0

100

200

300

Time (min)

400

Time (min)

All Data with Multiplexed PC1 vs PC2

•OH

Alkylating agent 0.5

Principal Component 2

1O 2

0.1 –0.6

H2O2

-0.1

0.4

0.9

1.4

J

Alkylating Agent Conc Peroxide Conc Singlet Oxygen Conc Hydroxyl Radical Conc Alkylating Agent Trans Peroxide Trans

–0.3 Singlet Oxygen Trans Hydroxyl Radical Trans Multiplexed Trans

-0.7 Principal Component 1

FIGURE 10.6 Multimodal detection of four reaction pathways. (A) Scheme of interactions on the DNA–SWNT complex: alkylating agent reaction with guanine, hydrogen peroxide (H2 O2 ) adsorption on the nanotube sidewall, singlet oxygen (1 O2 ) reaction with DNA, and hydroxyl radical (•OH) damage to DNA. (B) DNA–SWNT photoluminescence spectra before (blue) and after (green): introducing mechlorethamine, (C) hydrogen peroxide, (D) singlet oxygen, and (E) hydroxyl radicals. (F) Transient responses of photoluminescence intensity (top) and energy (bottom) of the (6,5) nanotube (black) and (7,5) nanotube (red) upon introducing mechlorethamine (blue border), (G) hydrogen peroxide (magenta border), (H) singlet oxygen (orange border), and (I) hydroxyl radicals (green border). (J) Plot of first two principal component scores of transient (closed points) and concentration-dependent (open points) detection data (from Supplementary Figure S1). Black crosses represent simultaneous singlet oxygen and hydroxyl radical generation (from Supplementary Figure S2b). Area-minimized ovals encompass all data sets taken for each analyte, including those not shown. Arrows denote direction of increasing concentration or time. (Reproduced with permission from Heller et al. [15].)

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FIGURE 10.7 In vitro NO detection using SWNTDAP-dex sensor. (a) NIR fluorescence images of Raw 264.7 cells incorporating SWNTDAP-dex before and after NO addition. After the cells were first incubated with SWNTDAP-dex (1 ␮g/mL) for 12 h, and washed with PBS to remove SWNTs in the growth medium, an extra solution of NO (5 ␮M, PBS) was added into the cell media at 330 s. A control experiment was also conducted without NO addition. Fluorescence was monitored in real time with 658-nm laser excitation (35 mW) and 63× objective. (b) Real-time tracking of NIR fluorescence response inside Raw 264.7 cells with an addition of NO solution (5 and 0.5 ␮M) at around 330 s. Fluorescence bleaching of SWNTDAP-dex was observed. (c) Averaged NIR fluorescence response to NO produced by iNOS in Raw 264.7 cells stimulated with LPS (20 ng/mL) and IFN-␥ (20 U/mL). The NIR fluorescence images were taken 0 and 12 h after stimulation. As a control experiment, the NIR fluorescence image was taken 12 h after incubation of Raw 264.7 cells without stimulation. (d) Averaged fluorescence intensity from each cell region responding to NO produced by iNOS in Raw 264.7 cells stimulated with LPS and IFN-␥ . After the cells (blue control, without any treatment; red, stimulated with LPS and IFN-␥ ; green, pretreated with NMA (2 mM) for 1 h, and then stimulated with LPS and IFN-␥ ) were monitored at a 2-h interval 6 h after stimulation. (Reproduced with permission Kim et al. [19].)

single-molecule adsorption and desorption. Hidden Markov modeling (HMM) is used to link single-molecule adsorption events to rate constants, as they have demonstrated for H2 O2 , H+ , and Fe(CN6 ) [1, 11]. Of the three kinds of reactant molecules studied, H2 O2 showed the highest quenching equilibrium constant of 1.59 at 20 ␮M while H+ was so insensitive that a similar equilibrium constant was achieved at a concentration as high as 0.1 M (pH 1). These results were self-consistent as reverse (dequenching) rate constants (600 ␮s−1 for H2 O2 , 1130 ␮s−1 for H+ , and 400 ␮s−1 for Fe(CN6 )3− ) were concentration independent while the forward (quenching) rate constants varied with concentration. These developments provide the groundwork for SWNTs to function as single-molecule stochastic biosensors.

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In another work by Jin et al. [54], human carcinoma epidermal A431 cells were grown on the porous collagen–SWNT film (Fig. 10.8a), and relatively stable Reactive Oxygen Species (ROS) generated from the cells like H2 O2 can diffuse into the film. The definition of ROS is broad. Species including singlet oxygen, superoxide anion radical, hydroxyl radical, hydrogen peroxide, and nitric oxide are all considered to be ROS. However, most of these species are highly unstable with a lifetime between a nanosecond and a microsecond in solution [55, 56]. For instance, the lifetime of singlet oxygen is calculated to be 4 ␮s in vivo [55]; the lifetimes of superoxide radical anion and hydroxyl radical are 1 ␮s and 1 ns, respectively [56]. As a result, only the relatively stable species can diffuse into the film and act as quenchers to the nanotubes over hours of experiment. The SWNTs embedded in the film serve as single-molecule sensors, reporting the fluorescence quenching reactions [57] with the H2 O2 (Fig. 10.8a,c) from EGF–EGFR binding [58]. The existence of the cells causes stepwise quenching and dequenching reactions to occur (top green trace,

FIGURE 10.8 (a) A431 cell was cultured on the collagen–SWNT film where the SWNT sensors were excited by a 658-nm excitation laser at 1 mW of the sample through an Alpha Plan-Apo 100x/1.46 oil emersion objective. The fluorescence emission from the SWNT was collected by the InGaAs nIR camera at 1 frame per second. The quenching activity of each sensor from hydrogen peroxide was binned into 16 categories represented by 16 different color bars with red the highest and black the lowest number of quenching transitions. Scale bars represent 10 microns. (b) Representitive confocal image for A431 cells with EGFR (red) labeled with rabbit polyclonal to EGFR and Alexa Fluor 568 donkey anti-rabbit IgG. The nuclei (blue) is labeled with 4 ,6-diamidino-2-phenylindole (DAPI). (c) Example trace of one SWNT sensor that shows stepwise fluorescence quenching from hydrogen peroxide adsorption, compared to the stable fluorescence trace of SWNT sensor in the absence of a cell. The quenching is reversible. (d) Hydrogen peroxide molecules are detected from the real-time quenching of SWNT sensors for both live and fixed A431 cells with the quenching rate (#/sensor/s) documented in real time. (Adapted from Jin et al. [54].)

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Fig. 10.8c), compared to the control experiment where there is no cell (bottom green trace, Fig. 10.8c). As has been detailed in our previous study [57], HMM can be used to recover the actual steps (red traces, Fig. 10.8c) from the data (green traces, Fig. 10.8c). The reversible quenching of collagen-wrapped SWNT by H2 O2 is the most predominant among other relative stable ROS as has been discussed previously. In order to image EGFR, A431 cells were immunostained using rabbit polyclonal to EGFR as the primary antibody and Alexa Fluor 568 donkey anti-rabbit IgG as the secondary antibody. As can be seen in the representative confocal image (Fig. 10.8b), A431 cells express a large amount of EGFR, which is considered to be generating H2 O2 upon biding to EGF [58]. The quenching rate was calculated in real time for EGF stimulation (500 ng/mL EGF was added at t = 0) on live/fixed A431 cells (Fig. 10.8d). Compared to the control where there is no EGF stimulation, the quenching rate of both live and fixed A431 cells has been increased from EGF stimulation. As can be seen from Figure 10.8d, the behaviors of single A431 cells after EGF stimulation are similar: the quenching rate increased rapidly right after stimulation. However, the time point of maximal response ranges from 600 to 1800 s after stimulation. There is no significant difference between live and fixed A431 cells. Compared to the ensemble measurement on thousands of cells, this platform allows the real-time measurement on a single-cell level for the first time.

ACKNOWLEDGMENTS We thank Paul W. Barone, Daniel A. Heller, Esther S. Jeng, Hong Jin, and Jong-Ho Kim for providing valuable input into this chapter.

REFERENCES 1. O’Connell, M. J.; Bachilo, S. M.; Huffman, C. B.; Moore, V. C.; Strano, M. S.; Haroz, E. H.; Rialon, K. L.; Boul, P. J.; Noon, W. H.; Kittrell, C.; Ma, J. P.; Hauge, R. H.; Weisman, R. B.; Smalley, R. E. Band gap fluorescence from individual single-walled carbon nanotubes. Science 2002, 297(5581), 593–596. 2. Bachilo, S. M.; Strano, M. S.; Kittrell, C.; Hauge, R. H.; Smalley, R. E.; Weisman, R. B. Structure-assigned optical spectra of single-walled carbon nanotubes. Science 2002, 298(5602), 2361–2366. 3. Saito, R.; Dresselhaus, M. S.; Dresselhaus, G. Physical Properties of Carbon Nanotubes. Imperial College Press: London, 1998. 4. Wray, S.; Cope, M.; Delpy, D.; Wyatt, J.; Reynolds, E. Characterization of the near infrared absorption spectra of cytochrome aa3 and haemoglobin for the non-invasive monitoring of cerebral oxygenation. Biochim. Biophys. Acta 1988, 933, 184–192. 5. McCartney, L. J.; Pickup, J. C.; Rolinski, O. J.; Birch, D. J. S. Near-infrared fluorescence lifetime assay for serum glucose based on allophycocyanin-labeled concanavalin A. Anal. Biochem. 2001, 292(2), 216–221. 6. Kim, S.; Lim, Y.; et al. Near-infrared fluorescent type II quantum dots for sentinel lymph node mapping. Nat. Biotechnol. 2004, 22(1), 93–97. 7. Saxena, V.; Sadoqi, M.; Shao, J. Degradation kinetics of indocyanine green in aqueous solution. J. Pharm. Sci. 2003, 92(10), 2090–2097.

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8. Klonis, N.; Quazi, N. H.; Deady, L. W.; Hughes, A. B.; Tilley, L. Characterization of a series of far-red-absorbing thiobarbituric acid oxonol derivatives as fluorescent probes for biological applications. Anal. Biochem. 2003, 317(1), 47–58. 9. Durkop, T.; Getty, S. A.; Cobas, E.; Fuhrer, M. S. Extraordinary mobility in semiconducting carbon nanotubes. Nano Lett. 2004, 4(1), 35–39. 10. Cognet, L.; Tsyboulski, D. A.; Rocha, J. D. R.; Doyle, C. D.; Tour, J. M.; Weisman, R. B. Stepwise quenching of exciton fluorescence in carbon nanotubes by single-molecule reactions. Science 2007, 316(5830), 1465. 11. Jin, H.; Heller, D. A.; Kim, J.-H.; Strano, M. M. Stochastic analysis of stepwise fluorescence quenching reactions on single-walled carbon nanotubes: single molecule sensors. Nano Lett. 2009, 8(12), 4299–4304. 12. Liu, Z.; Chen, K.; Davis, C.; Sherlock, S.; Cao, Q.; Chen, X.; Dai, H. Drug delivery with carbon nanotubes for in vivo cancer treatment. Cancer Res. 2008, 68(16), 6652. 13. Barone, P. W.; Baik, S.; Heller, D. A.; Strano, M. S. Near-infrared optical sensors based on single-walled carbon nanotubes. Nat. Mater. 2005, 4(1), 86–U16. 14. Jeng, E. S.; Moll, A. E.; Roy, A. C.; Gastala, J. B.; Strano, M. S. Detection of DNA hybridization using the near-infrared band-gap fluorescence of single-walled carbon nanotubes. Nano Lett. 2006, 6(3), 371–375. 15. Heller, D. A.; Jeng, E. S.; Yeung, T. K.; Martinez, B. M.; Moll, A. E.; Gastala, J. B.; Strano, M. S. Optical detection of DNA conformational polymorphism on single-walled carbon nanotubes. Science 2006, 311, 508–511. 16. Heller, D. A.; Jin, H.; Martinez, B. M.; Patel, D.; Miller, B. M.; Yeung, T. K.; Jena, P. V.; Hartner, C.; Ha, T.; Silverman, S. K. Multimodal optical sensing and analyte specificity using single-walled carbon nanotubes. Nat. Nanotechnol. 2008, 4(2), 114–120. 17. Pender, M. J.; Sowards, L. A.; Hartgerink, J. D.; Stone, M. O.; Naik, R. R. Peptide-mediated formation of single-wall carbon nanotube composites. Nano Lett. 2006, 6(1), 40–44. 18. Witus, L. S.; Rocha, J. D. R.; Yuwono, V. M.; Paramonov, S. E.; Weisman, R. B.; Hartgerink, J. D. Peptides that non-covalently functionalize single-walled carbon nanotubes to give controlled solubility characteristics. J. Mater. Chem. 2007, 17(19), 1909–1915. 19. Kim, J. H.; Heller, D. A.; Jin, H.; Barone, P. W.; Song, C.; Zhang, J.; Trudel, L. J.; Wogan, G. N.; Tannenbaum, S. R.; Strano, M. S. The rational design of nitric oxide selectivity in single-walled carbon nanotube near-infrared fluorescence sensors for biological detection. Nat. Chem. 2009, 1(6), 473–481. 20. Mujumdar, R.; Ernst, L.; Mujumdar, S.; Lewis, C.; Waggoner, A. Cyanine dye labeling reagents: sulfoindocyanine succinimidyl esters. Bioconjug. Chem. 1993, 4, 105–111. 21. Benson, R.; Kues, H. Fluorescence properties of indocyanine green as related to angiography. Phys. Med. Biol. 1978, 23(1), 159–163. 22. Lakowicz, J. Principles of Fluorescence Spectroscopy, 2nd ed. Academic/Plenum Publishers: New York, 1999. 23. Magde, D.; Rojas, G.; Seybold, P. Solvent dependence of the fluorescence lifetimes of xanthene dyes. Photochem. Photobiol. 1999, 70, 737. 24. Heller, D. A.; Baik, S.; Eurell, T. E.; Strano, M. S. Single-walled carbon nanotube spectroscopy in live cells: towards long-term labels and optical sensors. Adv. Mater. 2005, 17, 2793–2798. 25. Barone, P.; Parker, R.; Strano, M. S. In-vivo fluorescence detection of glucose using a single walled carbon nanotube optical sensor: design, fluorophore properties, advantages and disadvantages. Anal. Chem. 2005, 77, 7556–7562. 26. Cherukuri, P.; Bachilo, S. M.; Litovsky, S. H.; Weisman, R. B. Near-infrared fluorescence microscopy of single-walled carbon nanotubes in phagocytic cells. J. Am. Chem. Soc. 2004, 126, 15638–15639.

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CHAPTER 11

Microparticle- and Nanoparticle-Based Contrast-Enhanced Ultrasound Imaging ¨ NIRUPAMA DESHPANDE and JURGEN K. WILLMANN Department of Radiology and Molecular Imaging Program at Stanford, Stanford University School of Medicine, Stanford, California, USA

11.1 INTRODUCTION TO ULTRASOUND IMAGING IN MEDICINE Ultrasound refers to sound waves of 20,000 or more vibrations per second, which are far above the frequency heard by the human ear. Ultrasound imaging in medicine is based on the ability of these sound waves to reflect back from tissues in the human body, allowing an anatomical picture to be drawn on a screen of what is inside the body. Thus in technical terms ultrasound imaging is based on the reception, analysis, and display of acoustic signals produced by reflection or backscatter of sound (echo) above the audible frequency range for humans (20 Hz to 20 kHz). Ultrasound is primarily a tomographic modality, meaning that it generates an image that is typically a cross section of the tissue volume under investigation. It is also a soft-tissue modality, given that current ultrasound methodology does not provide useful images of or through bone or bodies of gas, such as found in the lung and bowel. In the past decade ultrasound imaging has gained immense popularity in medical diagnostic and therapy due to at least four characteristics: (1) it is a noninvasive, real-time imaging modality, (2) it does not utilize ionizing irradiation for image acquisition, (3) it is a quantitative imaging modality, and (4) it has a relatively low cost compared with other imaging modalities. 11.1.1 Historical Review Karl Theodore Dussik of Austria published the first paper on medical ultrasound in 1942, attempting to locate the cerebral ventricles in the brain by measuring the transmission of ultrasound beams through the skull. The first experimental investigations were performed with A-mode (amplitude modulation) on the cathode-ray tube display made up of a vacuum tube containing a source of electrons and a fluorescent screen used to create an image. On the cathode-ray tube display one axis represents the time required for the return of the echo and the other corresponds to the strength of the echo: thus providing a very simple display Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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plotted as a series of peaks, the height of which represents the depth of the echoing structure from the transducer. In the early 1950s the first B-mode (brightness modulation) images were obtained. In B-mode, the position of a spot on the cathode-ray tube display corresponds to the time elapsed as well as the position of the echogenic surface and the brightness of the spot corresponds to the strength of the echo. Movement of the ultrasound transducer over tissue produces a sweep of the ultrasound beam and a tomographic scan of a cross section of the body. After 1960, two-dimensional (2D) compound procedures were developed with ultrasound information generated from multiple angles being combined in one 2D image. This technique led to a boom in obstetric and gynecologic ultrasound worldwide. In the 1970s B-mode gray scale imaging became available, wherein the returning sound pulses have different shades of gray depending on their intensities, with solid areas appearing white and liquid areas appearing black. With progress of computer technology, ultrasonic imaging became better and faster and a wider range of clinical applications evolved over the following years. Real-time B-mode imaging, high-resolution scanners with digital beam forming, high transducer frequencies, multichannel focus, and broadband transducer technology have developed. Color Doppler and Duplex became available and sensitivity for low flow has continuously improved [1]. In Doppler ultrasonography, shifts in frequency between emitted ultrasonic waves and their echoes are used to measure the velocities of moving objects. Based on the principle of the Doppler effect, this technique is frequently used to examine cardiovascular blood flow but cannot be used to assess tissue microvascular blood flow. Duplex ultrasonography combines the standard real-time B-mode display with pulsed Doppler signals, allowing analysis of frequency shifts in an ultrasonographic signal, reflecting motion within a tissue (e.g., blood flow). This led to the invention of the power Doppler technique that is sensitive to low blood flow, enabling detailed visualization of vascular blood structure. The flow information is based on the amplitude of echoes received from moving cells and not on frequency shifts. Today, machines with advanced ultrasound system performance are equipped with real-time compound imaging. Furthermore, with the advent of contrast agents that enhance ultrasound signal, many techniques such as tissue harmonic imaging, contrast harmonic imaging, pulse inversion imaging, and 3D and 4D ultrasound with panoramic view became available. Harmonic imaging and pulse inversion imaging utilize the oscillation of gas-filled contrast microbubbles that are circulating in the blood flow, leading to scattering of the ultrasound wave. When an acoustic wave encounters a microbubble, it alternately compresses the microbubble on the positive pressure and expands it on the negative pressure. On the positive portion of the wave, the microbubbles are compressed in a different fashion from the way they expand in the negative portion. This results in an asymmetric, nonlinear bubble oscillation. Instead of producing a sinusoidal echo with a clean frequency spectrum, it produces an echo with asymmetric top and bottom. It is this asymmetry that produces harmonics and can be utilized to enhance the signals from the microbubble contrast agents. Ultrasonic waves of very low sound intensity lead to a linear (resonant) oscillation of the microbubbles. Greater sound amplitudes cause nonlinear (asymmetrical) oscillation of the microbubbles, during which other frequencies besides the output frequency are also emitted during the signal response. These other frequencies are the so-called harmonic frequencies, or overtones and undertones. The frequency components that are harmonic overtones are multiples of (i.e., double, triple, quadruple, etc.) the fundamental frequency used and/or its half, third, quarter, and so on in the case of undertone frequency component. The latter are also called subharmonic frequencies. Among the harmonic overtone frequencies, the second harmonic overtone frequency is of special importance in contrast-agent-specific

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sonography, due to its high amplitude. The basic frequency received from tissue and the microbubble contrast agent is suppressed by special filtering techniques during signal processing and only the contrast-agent-specific parts of the second harmonic frequency are used to generate the image. This procedure results in a clear reduction of the tissue signal portion, resulting in an improvement of the signal-to-noise ratio. In pulse inversion imaging technology, two successive ultrasonic pulses are emitted, whereby the phase of the second pulse is inverted with the phase of the first. From the received (backscattered) signals the two sequential waves (positive and negative waves) are subtracted from each other, whereby mathematically, the results cancel each other with linear signals. On the other hand, the nonlinear signals of the microbubbles contrast agent will not be cancelled out and can be visualized [2]. Finally three-dimensional (3D) ultrasound utilizes several 2D images acquired by moving the ultrasound transducer across the body surface, thereby resulting in a volumetric image. Four-dimensional (4D) ultrasound takes 3D ultrasound images and adds the element of time to the progress so that a moving 3D image is generated [3].

11.1.2 Ultrasound Equipment A basic ultrasound imaging system has the following components: (1) the ultrasound transducer is the main part of the ultrasound imaging system from which the ultrasonic waves are transmitted and received. The waves are generated by electrically stimulating a piezoelectric (pressure electricity) crystal, which produces elastic vibrations that are transmitted to the material being investigated. (2) The central processing unit (CPU) is the computational unit of the ultrasound machine that contains a microprocessor, memory, and amplifiers. It also contains the electrical power supplies for itself and the ultrasound transducer. The CPU does the calculations and data processing and forms the image on the monitor. The CPU also stores the processed data and/or images. (3) The transducer pulse control controls the changes in amplitude, frequency, and duration of the ultrasound pulses emitted from the ultrasound transducer. (4) Keyboard cursors are used for data input and measurement of images. (5) Finally, the monitor displays the images from the ultrasound data processed by the CPU [3].

11.1.3 Safety and Side Effects of Ultrasound Imaging There have been no substantiated side or ill effects of ultrasound imaging documented in studies in both humans and animals. Nevertheless, since ultrasound produces energy, there have been some concerns about the safety of ultrasound due to the development of heat or thermal effect. The thermal effect of ultrasound is caused by absorption of the ultrasound beam energy. As the ultrasound waves are absorbed, their energies are converted into heat. Energy of an ultrasound wave is described by pressure, temperature, density, and particle displacement. An ultrasound wave propagating through a medium loses its energy as a function of distance via reflection, absorption, and scattering. This lost energy is converted into heat. The thermal effect is highest in tissue with a high absorption coefficient, particularly in bone, and is low where there is little absorption such as fluid. The temperature rise is also dependent on the thermal characteristics of the tissue (conduction of heat and perfusion), the ultrasound intensity, and the length of examination time [4].

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11.1.4 General Diagnostic Indications of Ultrasound Imaging Ultrasound imaging in medicine is the most widespread imaging technique worldwide; it is easy to use, and the equipment is small and portable and therefore suitable for a doctor’s private practice use and ambulance, field, or emergency room applications. One of the main advantages of ultrasound imaging is that it is relatively inexpensive compared to other imaging modality such as magnetic resonance imaging (MRI), computed tomography (CT), or positron emission tomography (PET) imaging (Fig. 11.1) [5]. Diagnostic ultrasonography today is used in many medical fields as the main advantage is that certain structures can be observed without using radiation and it is relatively inexpensive with easy portability. Ultrasound produces real-time images of soft tissue and can capture structure of internal organs in the abdomen, pelvis, and thoracic cavity. For example, in obstetrics and gynecology, ultrasound is the first-line imaging modality for assessment of the female pelvis, for characterization of adnexal and uterine lesions, for diagnosing pregnancy, and monitoring fetal development. In cardiovascular imaging, ultrasound is used, for example, to evaluate the presence of blood flow by Doppler ultrasound through the major arteries and veins, to reveal stenosed or occluded arteries or veins, and to identify aneurysms and thrombus formations. Other imaging indications for ultrasound include assessment of size, shape, location, and diseases of solid organs in the abdomen including diseases of the liver, gallbladder, spleen, pancreas, kidneys, and pancreas. In addition, ultrasound is the first-line imaging modality for evaluation of the thyroid. In musculoskeletal radiology, ultrasound is used for assessment of muscles and ligaments and allows for accurate diagnosis of ligamentous tears in the shoulder. There is also an established use of ultrasound as a rapid imaging tool for diagnosis in the emergency room including the fast diagnosis of free intraperitoneal fluid from solid organ injuries [6].

FIGURE 11.1 Summary of advantages and disadvantages of some of the most common imaging modalities that are used for preclinical molecular imaging research. All of the imaging modalities have the potential for clinical translation. Compared with other imaging modalities, ultrasound is widely available in the imaging community throughout the world, is portable, and is relatively cheap compared with other imaging modalities. Due to its advantages, there is an increasing interest in evaluating ultrasound as a molecular imaging tool. (Reproduced with permission from Willmann and van Bruggen [5].)

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11.2 CONTRAST-ENHANCED ULTRASOUND IMAGING Contrast-enhanced ultrasound imaging can be defined as the application of ultrasound contrast agents to traditional medical ultrasonic imaging. Ultrasound waves travel through tissue or media under the influence of sound pressure and because molecules of tissue are bound elastically to one another, the excess pressure results in a wave propagating through the tissue. Thus acoustic impedance is important in the determination of acoustic transmission and reflection at the boundary of two materials having different acoustic impedances. The acoustic impedance (Z) of a material is defined as the product of its density (p) and acoustic velocity (V), given by the equation, Z = p × V. Ultrasound contrast agents are designed on the principle of acoustic impedance. Contrast agents usually consist of a gaseous core enclosed by an outer shell for stability. These agents act as echo enhancers when circulating in blood flow due to the high difference in acoustic impedance at the interface between gas and blood. The enhanced echo intensity is proportional to the change in acoustic impedance as the sound beam crosses from the blood to the gas. The contrast agent being gas filled is much more compressible than surrounding aqueous biological fluid; therefore ultrasound pressure waves in the megahertz range cause compression and expansion of contrast agents resulting in effective ultrasound backscattering [7]. These scattered ultrasound waves can then be detected by the ultrasound transducer. The contrast agents are administered intravenously into the systemic circulation and remain within the vasculature due to their size of several microns (1–4 ␮m) with no leakage and hence have access to the microvascular level. Several types of contrast-enhanced agents have been developed; the properties and application of each of these contrasts are discussed in the Section 11.3. Contrast-enhanced ultrasound can be classified into two general approaches: (1) contrastenhanced ultrasound imaging using nontargeted contrast agents, which are already being used in clinical practice, and (2) contrast-enhanced ultrasound imaging using molecularly targeted ultrasound contrast agents that are currently still restricted to preclinical research.

11.2.1 Nontargeted Contrast-Enhanced Ultrasound Imaging Nontargeted contrast-enhanced ultrasound imaging makes use of contrast agents, such as microbubbles, consisting of a gaseous core stabilized by a biocompatible outer shell and injected intravenously into the systemic circulation in a small bolus. Microbubbles once injected remain in circulation for a period of time, during which ultrasound waves are directed at the area of interest. When microbubbles in the blood flow past the area, which sees an ultrasound wave, they reflect a unique echo that stands in stark contrast to the surrounding tissue due to the orders of magnitude mismatch between microbubble and tissue echogenicity. The ultrasound system converts the strong echogenicity into a contrastenhanced image of the area of interest. In this way, the bloodstream’s echo is enhanced, thus allowing distinguishing blood from surrounding tissues. There are several applications of nontargeted contrast agents, which are listed below. 1. Edge Delineation in the Heart. Accurate detection of differences in tissue structures is crucial in ultrasound imaging of the heart, where a thinning, thickening, or irregularity in the heart wall indicates a serious heart condition that requires either monitoring or treatment. Microbubbles enhance the contrast at the interface between the heart

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tissue and the circulating blood and allow better delineation of morphological cardiac pathologies. 2. Assessment of Blood Volume and Perfusion. Contrast-enhanced ultrasound holds the promise for evaluating the degree of blood perfusion and blood volume in an organ or area of interest. The body consists of 90% water and is therefore acoustically relatively homogeneous. Blood and surrounding tissues have similar echogenicities; thus it is difficult to clearly discern the degree of blood flow through an organ (perfusion), or the interface between the tissue and blood using traditional ultrasound [8]. Color Doppler ultrasound that measures velocity of every voxel (3D equivalent of pixel) in a given region is conventionally used for detection of blood flow. The use of ultrasound contrast agents enhances Doppler flow signals by providing more and better acoustic scatterering. This results in improved detection of blood flow from vessels by providing high-resolution real-time assessment of flow without color Doppler artifacts. 3. Diagnosing Liver Diseases. Ultrasound is a widely used modality for imaging liver pathologies. By using nondestructive low-acoustic-power ultrasound with perfluorocarbon or sulfur hexafluoride-filled microbubbles, contrast-enhanced ultrasound has allowed improved detection and characterization of focal liver lesions. By analyzing the enhancement patterns of different focal liver lesions following intravenous administration of contrast agents, focal liver lesions can often be characterized by ultrasound without the need for further follow-up procedures such as biopsy. For example, malignant liver lesions such as hepatocellular carcinoma (HCC) are usually hypoechoic compared to normal liver parenchyma on delayed phase images (acquired several minutes after contrast administration), whereas benign lesions such as hemangiomas remain iso- or hyperechoic without washout of contrast on delayed phase images [9]. 11.2.2 Targeted Contrast-Enhanced Ultrasound Imaging (Molecular Ultrasound Imaging) The latest ultrasound approach combines ultrasound imaging technology with actively targeted ultrasound contrast agents; that is, agents to which one or more biologically active molecules have been attached such that the agents localize in a specific area of interest. Detection of bound targeted contrast agents can then show the area of interest expressing that particular molecule, which can be indicative of a certain disease state, or identification of a particular cell type. Using targeted agents, clinical and experimental roles of ultrasound can even further be expanded into applications including noninvasive detection of pathology using disease-associated molecular signatures, detection of gene expression, investigation of drug localization, and delineation of molecular mechanisms of disease. Targeted contrastenhanced ultrasound imaging is being developed for a variety of medical applications and is detailed in subsequent parts of this chapter. 11.2.3 Disadvantages of Contrast-Enhanced Ultrasound Imaging The introduction of contrast agents can provide nuclei for the generation of acoustic cavitation [10]. Cavitation is any activity of highly compressible transient or stable bubbles of gas or vapor, generated by ultrasonic power in the propagation medium. One of the side

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effects of ultrasound energy is the production of heat and increasing heat is associated with the risk of cavitation. However, the threshold for cavitation is high and does not occur at current levels of diagnostic ultrasound. Cavitation can be described as inertial or noninertial. Inertial cavitation, formerly called “transient” cavitation, occurs when the acoustic pressure amplitude is sufficiently high and above a threshold level. Under this condition the microbubble contrast agents will first grow in volume and then implode violently. If the core of a microbubble is gas with high specific heat ratio (constant pressure/constant volume), high temperature may result during implosions and highly reactive free radicals may be generated. Thus inertial cavitation has the most potential to damage tissue and cause local microvasculature ruptures and hemolysis [11]. Ultrasound produces more heat as the frequency increases. Therefore the ultrasonic frequency must be carefully monitored. Microbubbles burst at low ultrasound frequencies and at high mechanical indices (MIs—defined as the measure of the acoustic power output of the ultrasound imaging system). Increasing MI increases image quality, but there are trade-offs with microbubble destruction and thus inertial cavitation. Noninertial cavitation is the process in which microbubbles in a liquid are forced to oscillate asymmetrically in the presence of an acoustic field, when the intensity of the acoustic field is insufficient to cause total bubble collapse. Asymmetric contraction of an inertial cavitation bubble generates a microjet-like small stream that has a localized but unfavorable mechanical effect, such as membrane damage of a cell [12]. Targeted contrast agents to date have been restricted to preclinical trials due to the binding chemistry used so far for targeting microbubbles. Introduction of currently used targeting ligands can be immunogenic in human subjects as the binding ligands such as monoclonal antibodies are often derived from animal culture and the often used coupling strategy with biotin–streptavidin can cause allergic and immunogenic reactions [13]. However, humanized and chimeric antibodies or other small peptides can be used as an alternative source of ligands to functionalize microbubbles for clinical translation, such as synthetically manufactured targeting peptides that perform the same function, but without being immunogenic [14]. In addition, couple strategies other than biotin–streptavidin interactions are currently explored to further help moving targeted ultrasound contrast agents into the clinic [15, 16].

11.3 ULTRASOUND CONTRAST AGENTS 11.3.1 Background Blood is two to three orders of magnitude less echogenic than tissue due to the relatively small acoustic impedance difference between red blood cells and plasma. This low bloodto-issue signal ratio hinders the detection of small blood vessels. Ultrasound contrast agents, typically between 0.1 and 5 ␮m in diameter, increase the echo amplitude from the blood pool and allow the detection of vessels even as small as capillaries. The earliest introduction of contrast agents was by Gramiak and Shah [17], when they injected agitated saline into the ascending aorta and cardiac chambers during echocardiography to opacify the left heart chamber. Strong echoes were produced due to the acoustic mismatch between the air bubbles and surrounding blood. Today the contrast agents are a long way away from being produced by simple agitation, which causes the air bubbles to quickly dissipate and dissolve in the blood. The basic principle of contrast agents is using gas bubbles,

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which are encapsulated with a shell for stabilization. The ideal qualities in a contrast agent are high echogenicity, low attenuation, low blood solubility, low diffusivity, and lack of biological side effect. Basic principles for fabrication of various types of contrast agents are summarized in later sections. The commercial development of contrast agents began in the 1980 with a great effort on stabilization of the small microbubbles. Now, ultrasound contrast media sales have been growing at a compound rate of over 30% per year since 2004, with sales of $77.1 million in 2006 [18]. However, FDA approved indications are still limited. In addition, currently there is no FDA approved targeted contrast agent available. 11.3.2 Types of Ultrasound Contrast Agents Ultrasound contrast agents are biocolloids (colloidal particles made from biocompatible materials) and can be divided into two main classes: (1) microbubble based contrast agents, and (2) non-microbubble based contrast agents. Microbubbles are gas–liquid emulsions whereas non microbubble based contrast agents include liquid–liquid emulsions (nanodrops), solid nanoparticles, and liposomes [19]. Contrast agents measured for performance are based on factors such as minimal toxicity in vivo, ease of administration, and contrast to tissue ratio (CTR), which is the ratio between the acoustic backscatter coming from the contrast agent and tissue. The degree of acoustic backscatter depends on the intrinsic properties of the contrast agent, such as the compressibility of the contrast agent, the density difference between the contrast agent and surrounding tissue, and nonlinear effects (see Section 11.1.1) and resonance. The latest ultrasound technique is the development of targeted ultrasound contrast agents. Targeted ultrasound contrast agents are specially modified such that they recognize specific molecular markers and localize in a specific area of interest. Targeted contrast agents provide a powerful tool to evaluate and treat tissue in disease states in whole animal models and patients. The idea that contrast agents can be coupled to ligands for specific proteins and imaged to report molecular targets was developed in the late 1990s, therefore giving birth to the principle of molecular ultrasound. Targeted contrast agents can be generated by either passive or active targeting. Passive targeting is a nonspecific accumulation of contrast agents at the target site. The desired mechanism of passive targeting can be specified by manipulating the properties of the contrast agent, such as size, chemical properties such as type of encapsulated gas, electrical charge of the shell, the route of injection, and physiological processes of the immune system. Lindner et al. [20] were able to show during inflammatory processes that activated leukocytes engulf and absorb microbubbles, which are surrounded with an albumin or a lipid shell. The adhesion of lipid or albuminencapsulated microbubbles to activated leukocytes is due to binding to ␤2-integrin, Mac 1, and C3, all of which are molecular markers of inflammatory process. Fischer Scientific (Hanover Park, IL) [21] manufactured lipid encapsulated microbubbles with negative and neutral surface charges. In a subsequent ultrasonic investigation in the myocardium of dogs, a myocardial contrast could only be achieved with the negatively charged microbubbles 10 min after intravenous injection. On the other hand, microbubbles that have no charge did not show any contrast enhancement in this case. Oussoren et al. [22] examined the influence of the size of liposomes with respect to their transport from the injection site into the lymphatic vessels and their enhancement in regional lymph nodes (popliteal lymph nodes and iliac lymph nodes) in rats during the period of 52 h after subcutaneous injection. Accordingly, they observed that after 52 h only 26% of the liposomes with a size of 40 nm

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were present; however, there were still 95% of the liposomes with a size of 400 nm in the area of the injection site. In contrast to this, the difference in enhancement was not as clear in the regional lymph nodes, which indicates that larger liposomes become more enhanced in the lymph node than do smaller ones. Thus these experiments suggest that liposomes can be passively targeted to an area of interest based on size and route of injection. Active targeting requires modification of the contrast agent shell to allow selective binding to cellular epitopes or other receptors of interest. The attachment of targeting ligands can either be integrated into the shell during the production process or after manufacturing of the contrast agent. The ligands can be attached through covalent or noncovalent procedures [11, 13, 23]. Furthermore, the attachment of ligands can either be through direct coupling to shell or indirect coupling wherein a linker is used to connect the contrast agent shell and the ligand. This is necessary in cases where there is no direct coupling between contrast agent shell and ligand possible or when the ligand does not extend far enough from the encasement into the outside environment after direct coupling and thus is not in a position to connect to the molecular target structures. In this case, the linker works as a spacer to the contrast agent surface. These ligands can be attached to the contrast agent shell using several binding approaches such as carbodiimide, maleimide, or biotin–streptavidin coupling [11]. Biotin–streptavidin is the most popular coupling strategy because biotin’s affinity for streptavidin is very strong and it is easy to label the ligands with biotin. Currently, in preclinical studies these ligands are often monoclonal antibodies produced from animal cell cultures that bind specifically to receptors and molecules expressed by the target cell type. Peptides, which have the ability to recognize cell epitopes, can also be directly incorporated in the microbubble shell or be conjugated onto the microbubble shell [14, 15]. Visualization and quantification of various molecular molecules by ultrasound contrast agents has been shown for several pathological processes such as angiogenesis, inflammation, or thrombus formation, which will be explained in more detailed in Section 11.3.3.

Microbubbles Microbubbles are a gas–liquid emulsion with sizes ranging from 0.1 to 5 ␮m in diameter. The microbubble signal is dependent on the compressibility and density of the gas core, the viscosity and density of the surrounding medium, the frequency and power of ultrasound applied, and the bubble size. A microbubble is an ideal ultrasound contrast agent due to its excellent ultrasound response. A microbubble undergoes oscillation due to its gas core, which allows it to shrink and expand with the passage of an acoustic wave, resulting in the generation of strong acoustic signals that greatly exceed conventional ultrasound backscatter produced by reflection in acoustic impedance. The natural reaction time of a microbubble to a rapid pressure variation is on the order of microseconds; thus they resonate at frequencies typically used in ultrasound imaging, and resonance can be exploited to generate a very strong echo. Due to this high degree of backscatter in comparison to plasma and blood cells, an ultrasound system is capable of detecting the signature from a single microbubble in an in vitro setting. Klibanov et al. [24] imaged very dilute microbubble dispersion in plastic bags as well as attached microbubbles on a Petri dish. On imaging the plastic bags and Petri dishes, distinct white foci were observed. Concentration at the site of imaging was consistent with signals from individual microbubbles as determined by a Coulter counter. These results imply that the ultrasound system can resolve backscatter signals from individual microbubbles both in solution as well as attached to a surface if the surrounding background signal is kept to near zero.

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Types of Microbubbles Microbubble particles have a protective shell that prevents gas loss and particle fusion. There are a variety of microbubble contrast agents depending on their shell makeup, gas core makeup, and whether or not they are targeted. Microbubble shell material determines how easily the microbubble is taken up by the immune system. A more hydrophilic material tends to be taken up more easily, which reduces the microbubble residence time in the circulation. This reduces the time available for contrast imaging. The shell material also affects microbubble mechanical elasticity. The more elastic the material, the more acoustic energy it can withstand before bursting [25]. Currently, microbubble shells are composed of albumin, galactose, lipid, or polymers [8]. Ultrasound contrast agents can also be classified according to the rigidity of their shell [26]. Soft-shell microbubbles either have no coating or are covered by a thin monolayer of surfactant molecules such as palmitic acid or phospholipids and are very sensitive to pressure changes. Hard-shell microbubbles have a more rigid shell made of polymers or denatured albumin, which can increase their stability to a certain degree. Introduction of a grafted arm of polyethylene glycol (PEG) polymer onto the microbubble shell provides additional stearic protection, prevents microbubble aggregation, and makes the microbubble more nonreactive. It temporarily “hides” the microbubble from immune system uptake, increasing the amount of circulation time and hence imaging time (Fig. 11.2) [11]. The microbubble gas core is the most important part of the ultrasound contrast microbubble because it determines the echogenicity. Gas cores can be composed of air or high molecular weight or heavy gases like perfluorocarbon, sulfur hexafluoride, or nitrogen [8]. Heavy gases are less water soluble so they are less likely to leak out from the microbubble to impair echogenicity and thus prolonging their effective life in circulation [25]. Various methods have been employed to fabricate microbubbles for ultrasound imaging. The most popular method is dispersion of a gas through an aqueous phase by mechanical agitation of the gas–liquid interface. Dispersion is performed with the aid of a sonicator (sonication) or by high-shear mixing (amalgamation). These techniques rely on stochastic events that produce a polydispersed size distribution, generally ranging between submicrometer to tens of micrometers in diameter. Size fractionation techniques can be employed, which are based on buoyancy. Several methods have been described to encapsulate gas in a polymer shell. Dispersion and ionic gelation have been used to create alignate-shelled microbubbles. Organic solvents have been used to dissolve and disperse the polymer, which is then resuspended to form hollow polymer capsules. Polymerization at the air–liquid interface during agitation of an acidic medium was also used. Each of these methods has produced microbubbles with enhanced stability [19]. Lastly, there is a category of targeted microbubbles which retain the same general features as untargeted microbubbles, but they are outfitted with ligands that bind specific receptors expressed by cell types of interest, such as inflamed cells or angiogenic endothelial cells. This is explained in more detail in Section 11.3.3. Once in the blood flow, bound targeted contrast agents need to be differentiated from freefloating agents. This can be achieved if the targeted agent has a long lifetime at the target site, while nonadherent agents have a short circulating half-life. Another approach to differentiate bound from nonbound microbubbles is through the destruction and replenishment method described by Chomas [27] and Pollard [28]. After intravenous injection of the targeted contrast agent, ultrasound imaging is paused for several minutes to allow retention of contrast agents, following which a series of image frames are captured to obtain a signal from the tissue from adherent as well as freely circulating contrast agents. A single or a series of high-power destructive pulses are then given to destroy all contrast agents within

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FIGURE 11.2 Schematic representation of a microbubble in a blood vessel. A microbubble consists of a gas core enclosed by a shell. Various types of gas and shell compositions can be used. The gas core can consist of perfluorocarbon, sulfur hexafluoride, nitrogen, or air. The shell can be made of various substances such as phospholipid molecules held together with hydrophobic interactions or a protein like albumin with disulfide bonds to hold the molecules together. The shell can be modified to create a targeted microbubble by incorporating peptides that recognize specific ligands of interest or antibodies linked either directly onto the shell or indirectly through streptavidin–biotin moieties. Adding a polyethylene glycol (PEG) arm onto the shell provides stability, prevents microbubbles from aggregation, and helps prevent uptake by the immune system.

the beam elevation. After a brief period to allow freely circulating microbubbles to replenish the tissue of interest, a series of image frames are again captured. The signals before and after destruction are averaged and digitally subtracted from the initial predestruction frames to derive the signal attributable on adherent contrast agent only (Fig. 11.3).

Biodistribution of Microbubbles Biodistribution studies revealed that microbubbles have low circulation residence times as they are rapidly removed by the reticuloendothelial system (RES), which is part of the immune system, consisting of the phagocytic cells located in reticular connective tissue, primarily monocytes and macrophages in the liver (called Kupffer cells), spleen, and bone marrow. Walday et al. [30] studied the organ distribution of nontargeted air-filled iodine-125-labeled albumin microspheres (Albunex) in rats and pigs up to 90 min after intravenous administration. Three minutes after administration in rats, about 60% of the radiolabeled albumin microspheres were recovered in the liver, 9% of them were recovered in the spleen, and 5% of them were recovered in the lungs, with only small amounts in other organs such as the brain, heart, and kidneys. In contrast, in pigs, more than 90% of the albumin microspheres were recovered in the lungs, most likely because of the presence of intravascular macrophages in the porcine lungs. Biodistribution

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on air-filled iodine-123-labeled human serum albumin microspheres (Quantison) was studied in humans for 58 h by using radionuclide imaging by Perkins et al. [31]. One hour after intravenous administration, the greatest amount of radiolabeled microsphere uptake was in the liver (41.8%), followed by the spleen (11%) and the lungs (3.4%). Small amounts of radiotracer were also visible in bone and heart tissue in that study. Willmann et al. [32] performed in vivo analysis in living mice for lipid shell perfluorocabon-filled microbubbles targeted to VEGFR2 via anti-VEGFR2 antibodies by using dynamic micropositron emission tomography (micro-PET). Anti-VEGFR2 antibodies were radiolabeled by conjugating radiofluorination agent N-succinimidyl-4-[18 F]fluorobenzoate to the antibodies as a tracer for in vivo assessment of targeted microbubble biodistribution using micro-PET imaging [32]. In that study it was shown that 50% of targeted microbubbles cleared from the blood pool after ∼3.5 min and ∼95% were cleared from the blood circulation after 30 min. These findings also agree with results from repeated in vivo targeted contrast-enhanced ultrasound imaging, which have shown that the US imaging signal in blood vessels returns to almost background levels several minutes after administration of targeted microbubbles [33, 34]. Palmowski et al. [35] studied the biodistribution of hard-shelled microbubbles. The microbubbles consisted of cyanoacrylate and were targeted to VEGFR2 using monoclonal antibodies and streptavidin–biotin interactions. In that study it was observed that 1 min after injection of these microbubbles, 90% of the microbubbles were cleared from the blood and pooled in the lungs, liver, and spleen. VEGFR2 targeted microbubbles also accumulated significantly within the tumor vasculature. A biphased distribution of the microbubbles was observed in the blood. After injection of microbubbles, an initial increase was first seen in the blood, followed by a high concentration in the lungs, which then decreased during the ← FIGURE 11.3 (a) Schematic representation of targeted microbubbles (blue) attached to ligands in tumor vasculature of a solid tumor after intravenous administration. Tumor vasculature consists of vessels with neoplastic endothelial cells (orange). Here, both the tumor cells (gray) and the endothelial cells express an angiogenesis marker such as vascular endothelial growth factor receptor type 2 (VEGFR2). The microbubbles remain predominantly in the vasculature due to their size (several micrometers) and thus adhere only to the endothelial cell and not tumor cells. Not all microbubbles adhere; some are floating freely in the circulation system. After a high-power destructive pulse, adherent microbubbles are destroyed and freely circulating microbubbles replenish from outside the imaging plane after several seconds. (Reproduced with permission from Willmann et al. [29].) (b) This graph represents the method to calculate the signal from targeted microbubbles attached to vascular markers on angiogenic tumor vessels using the destruction–replenishment method. After intravenous administration, microbubbles float into the tumor and start binding to molecular markers to which they are targeted such as VEGFR2. In this example, a 4-min interval is allowed to pass for microbubbles to attach to the vascular target. The measured ultrasound signal consists of the background signal from the tissue, and signal from still freely floating microbubbles and microbubbles attached to the vascular targets. All microbubbles within the beam elevation are then destroyed by an external destructive pulse. The difference in imaging signal before and several seconds after destruction (to allow freely floating microbubbles to recirculate into the imaging plane) corresponds to the imaging signal from microbubbles attached to the vascular markers. (Reproduced with permission from Willmann et al. [29].) (c) Quantitative representation of molecular ultrasound imaging experiment using the destruction–replenishment method. The measured video intensity from the ultrasound image is plotted against the number of imaging frames. The difference between the predestruction and postdestruction imaging signal is automatically calculated and corresponds to the imaging signal from attached microbubbles.

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subsequent 2 h. The decreasing microbubble concentration in the lungs was accompanied by a decent recirculation of microbubbles in the blood. These studies show that, after the first pass, most nontargeted and targeted microbubbles are cleared from the blood circulation and accumulate in the RES. However, due to variable characteristics such as size or shell thickness within the population of injected microbubbles, the microbubbles may contribute to varying levels of stability.

Non-Microbubble Based Contrast Agents Non-microbubble based contrast agents are submicron and nanosized contrast agents that are attracting considerable interest as contrast agents for molecular ultrasound imaging. Their small size makes them also suited to targeting cells outside the capillary vasculature, such as cancer cells [36, 37]. Nonmicrobubble based contrast agents can be defined as particles ranging in size between 10 and 1000 nm. Their unique size range promotes long circulation half-lives and high tissue extravasation rates (the process of exuding or passing out of a vessel into surrounding tissues). Larger particulates greater than 1 ␮m in diameter are quickly recognized and cleared by the body’s reticuloendothelial system (RES) after intravenous injection [38]. Individual nanoparticles when circulating in the bloodstream cannot be detected with ultrasound due to their poor echogenicity. This is due to the inability of nanoparticles to oscillate significantly in an ultrasound field, as their interior core, being mostly liquid, is incompressible. But as the nanoparticles accumulate and aggregate on a surface, the coating on the surface changes the acoustic impedance difference between blood and the surface, making the surface acoustically more reflective and detectable with ultrasound. At diagnostic frequency, nanoparticles are less echogenic than microbubbles but are more suitable for high-frequency imaging approaches used for intravascular imaging, providing near microscopic resolution [38]. Nanoparticles of particular relevance according to their submicron size are perfluorocarbon emulsion (PFC) nanodroplets, polylactide acid (PLA) nanobubbles, solid nanoparticles, and echogenic liposomes (Fig. 11.4) that are discussed in the following sections. Perfluorocarbon Emulsion Nanodroplets Perfluorocarbon emulsion (PFC) nanodroplets are liquid–liquid emulsions consisting of a PFC liquid core encapsulated by a phospholipid monolayer with typical diameters of 200–400 nm. The liquid core, which is 98% by volume of the nanodroplets, can have any one of the numbers of PFCs—namely, perflurooctylbromide (PFOB), perfluoro-15-crown-5 ether (CE), and perfluorodichlorooctane, with PFOB the most commonly used. PFC nanodroplets are stabilized for use in vivo with surfactants, the most common being phospholipids, which restrict the particles to coalesce with one another. PFC nanodroplets are formed by mechanical diminution of the fluorocarbon liquid into an aqueous phase. The dispersed phase consisting of liquid fluorocarbon is immiscible with the continuous aqueous phase and is highly hydrophobic; thus PFC nanodroplets exhibit robust stability against handling, pressure, heat, and shear. The size of these particles can be controlled by extrusion. Individual PFC nanodroplets are less sensitive contrast agents due to poor echogencity of its perfluorocarbon core. However, when collective deposition occurs on the surfaces of tissues or a cell in a layering effect, these particles create a local acoustic impedance mismatch that produces a strong ultrasound signal without a concomitant increase in the background level [39]. Thus PFC nanodroplets are less liable to nonspecific signal enhancement as larger binding events take place for a detectable signal. The nongaseous nature of PFC nanodroplets has advantages for ultrasound imaging as their liquid composition

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FIGURE 11.4 Schematic representation of various kinds of non-microbubble based contrast agents. Perfluorocarbon emulsion nanodroplets (PFC nanodroplets) are liquid–liquid emulsions, polylactide acid (PLA) nanobubbles are gas–liquid emulsions, liposomes consists of a phospholipid bilayer enclosing an air pocket, and solid nanoparticles are solid amorphous substances that have gas or air entrapped in their pores or fissures.

makes the nanodroplets resistant to pressure or mechanical stress, they are not easily deformed, and they do not cavitate during the imaging process [40]. Another advantage of PFC nanodroplets includes the fact that, based on the properties of perfluorocarbon in their liquid cores, they can be vaporized into gas bubbles visible with standard ultrasound imaging [41]. When low boiling point perfluorocarbons such as dodecaperfluoropentane are used in nanodroplets, they spontaneously vaporize into gas pockets caused by the temperature increase due to a body temperature of ∼37 ◦ C in a living subject. The challenge for these low boiling point nanodroplets is for enough contrast agents to have accumulated at the desired target site before spontaneous vaporization occurs. In order to decrease spontaneous vaporization, high-boiling point nanodroplets using liquid perfluorocarbons such as perflurohexyl bromide have been synthesized to minimize spontaneous vaporization at physiological body temperature. However, with these types of high boiling point nanodroplets, higher energies have to be applied to transfer them into visible gas pockets. To overcome this trade-off, Amirriazi et al. showed in vitro that nanodroplets filled with high boiling point perfluorcarbons can be modified by adding iron oxide nanoparticles as nucleation sites inside the nanodroplets, thereby substantially lowering the boiling point of nanodroplets (high enough to minimize spontaneous vaporization and low enough to decrease the required energy for vaporization) [42]. Future studies are warranted to assess in vivo targeting efficiency of these contrast agents. biodistribution of perfluorocarbon emulsion nanodroplets The circulatory halflife period of PFC nanoparticles is in excess of 1 h and the tissue half-life residency is many

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days. Studies done to visualize PFC nanoparticles, which were used as a substitute to create artificial blood in the vascular system of rats using MRI, showed retention of PFC for several days. Images obtained at intervals of approximately 2 h, 2 days, 2 weeks, and 2 months after infusion of PFC nanoparticles showed identification of PFC in the heart, lung, liver, spleen, and large vessels both in vivo and postmortem [43]. Furthermore, it was shown that the retention of PFC nanoparticles in the circulation was prolonged more than 300% when coated with lecithin [44]. This prolonged retention in the circulation resulted in a marked decrease in the amount of perfluorochemicals retained in the liver and spleen.

Polylactic Acid (PLA) Nanobubbles Nanobubbles are gas–liquid emulsions enclosed with a biodegradable polymer, polylactic acid (PLA), ranging from 40 to 200 nm in size. Use of targeted nanobubbles was tested by Liu et al. [45]. The surface of PLA nanobubbles was conjugated to an anti-Her2 antibody (i.e., Herceptin) and the nanobubbles were tested in vitro for specific binding to breast cancer cells that overexpress Her2 receptors. In high-resolution ultrasound B-mode images, the average gray scale of the Her2-positive cells was consistently and significantly higher after nanobubble treatment than negative controls in vitro. Rapoport et al. [46] describe a novel method of combining ultrasound imaging of cancer with ultrasound enhanced drug delivery treatment. Nanobubbles filled with the chemotherapy drug doxorubicin were injected into mice. The bubbles accumulated in the tumors, where they combined to form larger “microbubbles.” When exposed to ultrasound, the bubbles generated echoes, which made it possible to image the tumor. Furthermore, a destructive ultrasound wave destroying the bubbles caused release of the drug that suspended tumor growth [46]. At the time of this writing, biodistribution studies of nanobubbles have not yet been published. Solid Nanoparticles Solid nanosized particles, which include amorphous solid particles like silica or iron oxide particles that contain gas pockets in their pores and fissures, have been discovered to produce detectable backscatter for ultrasound imaging. One of the advantages of these solid nanosized particles is the high tissue extravasation rate due to their unique size range and therefore their potential use to target molecular markers outside the vasculature. In addition, these solid nanoparticles remain stable during insonication; thus they can also generate good tissue enhancement, owing to the acoustic-impedance mismatch between solids and soft tissue compared to that between liquids and soft tissue. Liu et al. [47] tested silica nanospheres, 100 nm in diameter, in a saline suspension and intravenously injected in mice for imaging the liver with a transducer frequency of 30 MHz. It was shown that the mean acoustic intensity increases in the liver following particle administration. Also, these silica nanospheres were tested in vitro on tissue phantoms consisting of particles dispersed in agarose gel. Intensity was seen to increase with increasing particle size and concentration. Targeted solid nanoparticles are yet to be designed and could be the next step since nanoparticles can be utilized effectively as targeted contrast agents because they can be efficiently surface modified with biorecognition molecules given their relatively large surface area. Surface modification could result in nanoparticles escaping the RES, leading to prolonged half-lives in circulation and an increased number of passes through the vasculature to achieve satisfactory targeting. In addition, tumor vasculature exhibits a permeability and retention effect that results in exaggerated extravasation and retention of particles that are smaller than the pore size of tumor endothelia [48], making these solid nanoparticles very

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useful for cancer imaging. It was shown that superparamagnetic iron oxide nanoparticles, another solid nanoparticle agent, provided unexpected enhancement of the delineation of ex vivo rat brain tumors [49] and central nervous system lesions [50] in sonograms. These studies suggest that the potential for using nanoparticles as molecularly targeted ultrasound imaging agents should be explored. Toxicity and biocompatibility issues will also need to be considered if nanoparticles are to be considered for clinical imaging applications.

Liposomes The discovery of liposomes or lipid vesicles emerged from self-forming enclosed lipid bilayers with an aqueous core upon hydration. Liposomes can be unilamellar or multilamellar, and their size can be controlled by sonication and extrusion (forcing the liposomes through microscale pores) to range from ∼20 nm to >10 ␮m in diameter. A semipermeable, hydrophobic membrane separates the interior and exterior compartments of a liposome. Echogenic liposomes contain air pockets within the lipid bilayer to generate acoustic reflectivity. The mechanism of echo contrast appears to be the backscatter from entrapped pockets of air within the liposomes that form during rehydration of freeze-dried liposomes. The size, charge, and surface properties of liposomes can easily be changed by adding ingredients to the lipid mixture before liposome preparation and/or by varying the preparation methods. The ability of liposomes to easily entrap different substances into both the aqueous phase and the liposome membrane compartments makes them suitable for manipulations to generate targeted contrast agents. Echogenic liposomes can be prepared by incorporating gases such as air, argon, or nitrogen into the liposome or by inducing the formation of gas bubbles directly inside the liposome as a result of a chemical reaction [51]. Liposomes composed of phospholipids are not echogenic, but when made with phosphatidylcholine (PC), phosphatidylethanolamine (PE), phosphatidylglycerol (PG), and cholesterol [52] they are highly acoustically reflective. These liposomes improved intravascular contrast by approximately 300% relative to blood and 150% relative to agitated saline when injected in swine [53]. Antibodies can be conjugated on these lipid shells without hindering the echogenecity. biodistribution of liposomes Phospholipid liposomes introduced into the circulation system are very rapidly sequestered by the cells of the reticuloendothelial system (RES) [54]. The clearance half-time of the liposome is usually within 30 min. A large number of liposomes need to accumulate at the target tissue to render it acoustically reflective. This can be achieved through multiple passes of the blood containing the liposomes over the target tissue. A prolonged blood circulation of the liposomes was achieved with the addition of a polyethylene glycol (PEG) coating, which efficiently minimizes their removal by macrophages of the reticuloendothelial system [55]. Marik et al. [56] conducted PET imaging and biodistribution studies of liposomes. To track liposomes, radioactive [18 F]fluorodipalmitin ([18 F]FDP) was incorporated into the lipid molecule of the phospholipid bilayer of the liposome. As a control, free [18 F]FDP was also injected in rats. Maximum intensity projection images obtained from 90-min continuous bed motion scans were used to illustrate the full body distribution of free and liposome-encapsulated [18 F]FDP. Freely injected [18 F]FDP showed an initial concentration of 3% ID/cc and was cleared from the blood within a few minutes. The greatest concentration of free [18 F]FDP was observed in the liver, 5.5% ID/cc and in the spleen at a concentration of 4.2% ID/cc. Liposomal [18 F]FDP remained in the blood circulation at near constant levels for at least 90 min, with a peak concentration near 2.5% ID/cc. In addition, liposomal [18 F]FDP quickly reached a

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steady state in the liver, spleen, kidney, and lungs, which was maintained throughout the scan, reflecting the activity within the vasculature of these organs.

11.3.3 Applications of Molecular Ultrasound Imaging Molecular ultrasound imaging with the advent of targeted contrast agents has opened a variety of diagnostic and therapeutic potentials. Using targeted agents, the clinical and experimental roles of ultrasound are further explored, into applications including noninvasive detection of pathology using disease-associated molecular signatures, detection of gene expression, investigation of drug localization, and delineation of molecular mechanisms of disease. Furthermore, the ability to differentiate between different disease-indicative molecular signatures could allow early assessment of pathology and expedite the design of customized treatments. There are many potential targets available for study by targeted contrast imaging. These molecular signatures can be used to localize ultrasound contrast agents through the use of complementary receptor ligands attached to the contrast agent shell such that the ligand–receptor interaction tethers the agent to the cell of interest. Studies published on targeted ultrasound contrast agents are briefly summarized below and in Tables 11.1 and 11.2.

Molecular Ultrasound Imaging of Tumor Angiogenesis Angiogenesis, the process of new blood vessel formation, plays an important role in tumor growth and metastasis and diverse disease processes such as atherosclerotic plaque growth and adaptation to chronic ischemic disease. Novel molecular imaging strategies that allow direct visualization and quantification of expression levels of key molecular markers of these diseases would be ideal tools for detection as well as monitoring therapeutic treatment in patients. To illustrate this concept, many molecular markers such as integrins or vascular endothelial growth factor receptor type 2 (VEGFR2) have been shown to be upregulated on angiogenic and metastatic endothelial cells in an actively growing tumor vasculature specifically (Fig. 11.5). The introduction of microbubbles targeted to key molecular markers of tumor angiogenesis is now being explored as a novel molecular imaging strategy with ultrasound [72, 73]. A number of studies have addressed assessment of tumor angiogenesis with molecular ultrasound in preclinical studies. Integrins, which are extracellular matrix molecules, have been extensively evaluated for targeting of imaging agents, drugs, and particles to the tumor endothelium. Integrin ␣v ␤3 , in particular, has received a lot of attention as it is highly expressed on activated endothelium and almost absent on normal vessels, making it very useful for detection of tumor formation. Various strategies were used to generate integrin ␣v ␤3 targeted microbubbles, namely, the use of streptavidin–biotin coupling chemistry along with monoclonal antibodies for integrin ␣v ␤3 , and the use of a cyclic RRL (arginine–arginine–leucine) peptide and echistatin as ligands on the microbubble shell that bind integrin ␣v ␤3 . Echistatin is a viper venom disintegrin containing an RGD (arginine–glycine–aspartic acid) peptide that binds integrin receptors. Ellegala et al. [57] imaged malignant gliomas in athymic rats by intracerebral implantation of U87MG human glioma cells. An increase in the acoustic reflectivity of malignant gliomas was demonstrated with the use of echistatin incorporated microbubbles targeted to ␣v ␤3 . Weller et al. [58] used the RRL peptide incorporated microbubbles to show a significant accumulation of targeted microbubbles within subcutaneously implanted, human prostate carcinoma xenografts in mice. Although the RRL peptide is known to bind

Human KDR/ VEGFR2 Endoglin

Angiogenesis

Inflammation Inflammation Inflammation Inflammation Inflammation Inflammation Arteriosclerosis Transplant rejection Transplant rejection Thrombus Thrombus Lymph node

Angiogenesis

Angiogenesis VEGFR2, Integrin alphavbeta3, Endoglin Leukocytes P-selectin P-selectin P-selectin P-selectin MadCAM-1 VCAM-1 Leukocyte ICAM-1 GP IIb/IIIa GP IIb/IIIa L-selectin

Integrin ␣v ␤3 VEGFR2 VEGFR2 VEGFR2 Human KDR/ VEGFR2 VEGFR2

Angiogenesis Angiogenesis Angiogenesis Angiogenesis Angiogenesis

mAb RGD peptide RGD peptide MECA-79 ligand

Microbubble Microbubbles Microbubbles Microbubbles Microbubbles Microbubbles Microbubbles Microbubbles Microbubbles Microbubbles Microbubbles Microbubbles

Microbubbles

mAb

mAb Sulfo-Le-AAA peptide mAb mAb mAb mAb

Microbubble

Microbubbles

Microbubble

Microbubble Microbubble Microbubble Microbubble Microbubbles

Microbubble Microbubble

Contrast Agent

Heterodimeric KDR-targeted peptide mAb

Echistatin peptide Tumor endothelial cell (target was not identified) Knottin peptides mAb mAb mAb Heterodimeric KDR-targeted peptide mAb

Ligand

Animal Model

Tissue necrosis factor-treated cremaster muscle of mice Mouse model for postischemic injury in kidney Rat model for postischemic injury in the myocardium Mouse model for postischemic injury Mouse model for postischemic injury Mouse model for inflammatory bowel disease Mouse model for arteriosclerotic plaque in ApoE deficient mouse Rat cardiac transplantation model Rat cardiac transplantation model Mouse cremaster muscle model for arteriolar and venular clots Thrombi formation in dogs Lymph nodes of mice and dogs

Subcutaneous tumor model in mouse for human pancreatic cancer and orthotopic pancreatic cancer model in mice Subcutaneous tumor model in mouse for human breast, ovarian and pancreatic cancer

Subcutaneous tumor model in mouse for human pancreatic cancer and orthotopic pancreatic cancer model in mice Subcutaneous human colon cancer xenografts in mice

Subcutaneous tumor model in mouse for human ovarian cancer Subcutaneous tumor model in mouse for angiosarcoma and glioma Subcutaneous tumor model in mouse for human melanoma Subcutaneous tumor model in mouse for murine breast cancer Orthotopic rat breast cancer model in rats

Subcutaneous tumor model in rat for human glioma Subcutaneous tumor model in mouse for human prostrate carcinoma

66 67 68 33 76 69 70

20 63 64 65

62

34

15

34, 35

14 59 60 61 16

57 58

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Angiogenesis

Integrin ␣v ␤3 Integrin ␣v ␤3

Molecular Target

Angiogenesis Angiogenesis

Disease

TABLE 11.1 Summary of Published Studies on the Use of Targeted Microbubbles for Ultrasound Imaging of Angiogenesis, Inflammation, Thrombus Formation, and Lymph Nodes

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TABLE 11.2 Summary of Published Studies on the Use of Targeted Nanoparticles for Ultrasound Imaging of Inflammation and Thrombus Formation Molecular Target

Disease

Ligand

Contrast Agent

Atherosclerosis ICAM-1, Fibrinogen

mAb

Liposome

Thrombus

ICAM-1

mAb

Liposome

Thrombus

VCAM-1

mAb

Liposome

Thrombus

Fibrin

mAb

PFC nanoparticles

Thrombus

Stretch induced

Tissue factor mAb

PFC nanoparticles

Animal Model

References

Yucatan miniswine model for atherosclerosis Yucatan miniswine for different stages of atheroma Yucatan miniswine for different stages of atheroma Thrombi formation in the carotid artery of pigs Pig model for balloon stretch induced injury to carotid artery

53

71

71

38

72

tumor endothelium, the molecular target is still not identified. Recent studies conducted by Willmann et al. [14] showed the use of a novel peptide—knottin—that binds to integrin ␣v ␤3 with high affinity. Knottins are small, compact peptides (20–60 amino acids) that consist of a core of at least three disulfide bonds that are interwoven into a “knotted” conformation. The in vivo molecular ultrasound imaging signal seen with knottin peptide when conjugated to the shell of contrast microbubbles was similar or higher when compared with microbubbles

(a)

(b)

FIGURE 11.5 (a) Transverse B-mode ultrasound image of a subcutaneous human colon adenocarcinoma xenograft tumor (arrows) in a nude mouse. (b) Targeted contrast-enhanced ultrasound image of tumor angiogenesis in same xenograft tumor (arrows) 4 min after intravenous injection of microbubbles targeted to VEGFR2; imaging signal from contrast microbubbles bound to VEGFR2 is shown as green overlay on B-mode image. Note: Molecular ultrasound imaging signal was measured using the destruction replenishment approach described in Figure 11.3.

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conjugated to a monoclonal antibody for integrin ␣v ␤3 . Knottins have great potential for in vivo applications due to their resistance to proteolysis and their high thermal stability and are thought to be nonimmunogenic. Finally, knottins may be a promising platform for designing novel contrast agents for molecular ultrasound imaging [14]. Another angiogenic marker, VEGFR2, has also been studied as a target for molecular ultrasound imaging using high-frequency ultrasound. Overexpression of VEGFR2 has been associated with tumor progression and poor prognosis in several tumors including colorectal, gastric, and pancreatic carcinomas; angiosarcoma; breast, prostate, and lung cancers; malignant gliomas; and melanomas. Several studies reported enhanced contrast images of tumor with the use of VEGFR2 targeted microbubbles [29, 32, 34, 35, 59–62]. Willmann et al. [29] used targeted microbubbles by coupling monoclonal antibodies as the binding ligand against VEGFR2 via streptavidin–biotin coupling chemistry; as a control, microbubbles were also conjugated to isotype matched antibody along with unlabeled microbubbles. The ability of these microbubbles to adhere to the target was first confirmed in cell culture experiments. In vivo ultrasound images were then acquired with a 40-MHz linear transducer in a mouse subcutaneous tumor model for angiosarcoma and malignant glioma. Accumulation of VEGFR2 targeted microbubbles was reported within subcutaneously implanted tumors compared to unspecific control microbubbles, which was further confirmed by immunofluorescence analysis of VEGFR2 expression in both tumor types [29]. Rychak et al. [60] reported similar results on subdermal tumors derived from human melanoma cells in mice. Studies performed by Lyshchik et al. [61] compared expression of VEGFR2 by targeted microbubbles on highly invasive metastatic and nonmetastatic murine models of breast cancer cells demonstrating that targeted ultrasound can be used to characterize angiogenic activity corresponding to the degree of malignancy. They observed a higher accumulation of VEGFR2 targeted microbubbles within the more aggressive tumors. Korpanty and colleagues [34] investigated the use of targeted microbubbles to follow vascular response of therapy in addition to detection of tumor angiogenesis by molecular ultrasound. VEGFR2 targeted microbubble accumulation was assessed to quantify vascular effects of two different antitumor therapies—namely, with antivascular endothelial growth factor (VEGF) monoclonal antibodies and/or gemcitabine. The model systems used were subcutaneous as well as orthotopic pancreatic cancer tumors in mice. They detected decreasing marker densities after tumor- suppressive therapy, which correlated with the observed effects of treatment. In addition, they performed a multimarker imaging, assessing VEGFR2 followed by Endoglin targeted microbubbles in the same tumor after a long interval between scans to ensure passive clearance of previously injected microbubbles. Endoglin is a cell membrane glycoprotein that is involved in vascular development and remodeling and is overexpressed on tumor-associated vascular endothelium. Targeted microbubbles showed significant enhancement of tumor vasculature when compared with untargeted or control IgG targeted microbubbles that correlated with ex vivo expression analysis. In the same light Palmowski and colleagues [35] demonstrated an upregulation of VEGFR2 and of integrin ␣v ␤3 during the growth of untreated tumors, and a downregulation of both markers after antiangiogenic therapy using microbubbles conjugated to VEGFR2 antibodies and cyclic RGD peptides. Recently, Pysz et al. [15] evaluated a novel, clinically-translatable microbubble targeted to human kinase insert domain receptor (KDR; the human protein analogous VEGFR-2) for monitoring antiangiogenic therapy in subcutaneous human colon cancer xenografts in mice. In mice receiving antiangiogenic therapy in vivo molecular ultrasound imaging using novel KDR-targeted microbubbles significantly

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decreased as early as 24 hours after initiation of antiangiogenic therapy. In contrast, no difference in molecular ultrasound KDR-targeted imaging signal was observed in nontreated mice. Furthermore, KDR-associated molecular ultrasound signal was observed prior to any changes in tumor size; thus, demonstrating the advantage of early assessment of antiangiogenic therapy using molecular ultrasound imaging prior to overt morphological-anatomical changes become visible in tumors. Another aspect observed by Willmann et al. [59] was that multiple markers conjugated on microbubbles could significantly enhance the imaging signals, which could be a useful tool for detection of tumors early enough for effective therapy. Toward this goal, it was demonstrated that dual-targeted microbubbles carrying antibodies for VEGFR2 and integrin ␣V accumulate to higher intensities compared to single targeted microbubbles in tumor angiogenesis of human ovarian cancer in mice xenografts. These results may be of great significance for early detection of cancer when tumors are too small to cause detectable morphologic changes but large enough to induce tumor angiogenesis (Fig. 11.6). Molecular ultrasound imaging also allows noninvasive mapping of expression levels of angiogenic markers in tumor angiogenesis. In three different tumor types (breast, ovarian, and pancreatic cancer) and using noninvasive molecular ultrasound imaging, Deshpande et al. [62] showed varying expression levels of the three angiogenic markers integrin ␣v ␤3 , endoglin, and VEGFR2 during tumor growth. In this study, Using ex vivo western blotting as reference standard, the study confirmed that molecular ultrasound imaging allows longitudinal noninvasive assessment of the temporal tumor angiogenic molecular marker expression levels in vivo [62]. The results of this study provided further insights into the biology of tumor angiogenesis and may help in defining promising imaging targets for

Endothelial cell

Endothelial cell

Endothelial cell

Endothelial cell

FIGURE 11.6 Single and dual targeted microbubbles and interaction with tumor vessel endothelial cells. Gas-filled microbubbles (MBs) conjugated to ligands bind molecular markers of angiogenesis (e.g., VEGFR2) expressed on tumor vessel endothelial cells. Dual targeted MBs, which carry two types of ligands on their shells, can bind to more molecular markers compared to single targeted microbubbles. This increases the signal intensity from derived bound microbubbles, which, for example, may increase the detection of small tumors.

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both early cancer detection and treatment monitoring of cancer using molecular ultrasound imaging [62].

Molecular Ultrasound Imaging of Inflammation Inflammation is a common physiological process occurring in a vast number of diseases such as ischemia (inadequate blood supply to tissue), arthrosclerosis, or inflammatory bowel disease to mention only a few. One crucial component of inflammation is the activation of leukocytes in the blood pool and their transmigration to the extravascular compartment. The recruitment and transmigration of leukocytes are mediated by the interaction between adhesion molecules on the leukocytes and on the endothelial cell surface [20]. Several molecules including E- and P-selectin mediate the initial capture and consecutive rolling of leukocytes on the inner vessel wall. A firm arrest of the rolling leukocytes, the necessary precondition for transmigration, is promoted by a second group of adhesion molecules, namely, the intercellular adhesion molecule-1 (ICAM- 1) and the vascular cell adhesion molecule-1 (VCAM-1). All inflammatory markers are expressed rapidly during the inflammatory process and have been reported to correlate to a certain degree to the stage of inflammation. Both passive and active targeting has been explored for imaging inflammation using molecular ultrasound. Passive targeting does not provide molecular-level information, but rather enables qualitative detection of inflammation as microbubbles attach to (and eventually phagocytosed by) activated leukocytes. This leads to damping, but the microbubbles remain acoustically responsive and still produce an acoustic signal [74]. One strategy for passive targeting is to incorporate negatively charged phosphatidylserine into the shell, which promotes microbubble attachment to activated leukocytes. This technique was used to image tissue necrosis factor-␣-treated inflamed cremaster muscle in mice and to assess the severity and extent of postischemic myocardial inflammation in dogs [74]. For active targeting, monoclonal antibodies that bind inflammatory markers like E- and P-selectin, ICAM-1, and VCAM-1 have been outfitted onto microbubbles for molecular ultrasound imaging of inflammation [8]. P-selectin, which is expressed immediately after an ischemic stimulus, was used to assess postischemic injury in the mouse kidney [63]. P-selectin targeted microbubbles were also used to identify recently ischemic myocardium in a myocardial ischemia reperfusion model in mice [64]. A similar study used sialyl Lewis (a P-selectin ligand) conjugated microbubbles in a rat model for myocardial ischemia reperfusion to visualize “the ischemic memory” noninvasively using molecular ultrasound imaging [65], highlightening the potential of molecular ultrasound imaging as a rapid and straight forward bedside test for screening patients with atypical chest pain for a recent ischemic event. The gut-specific marker “mucosal addressin cellular adhesion molecule-1” (MAdCAM1) was used as a target to image a mouse model for inflammatory bowel disease [66]. Bachmann et al. [66], using transabdominal ultrasound, demonstrated a significant accumulation of MAdCAM-1 targeted microbubbles as compared to nonspecific ones in focal areas of ileal inflammation, thus producing stronger acoustic signals. VCAM-1 antibodyconjugated microbubbles were utilized by Kaufmann et al. [67] to demonstrate a correlation between different stages of arteriosclerosis and the retention of microbubbles. The inflammation disease model system used was ApoE deficient mice, which form inflammatory plaques in the aorta when fed a high-cholesterol diet. Two studies have been reported to investigate transplant rejection in vivo in animal models. The first study employed passively leukocyte-targeted microbubbles with phosphatidylserine to assess acute allograft rejection

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in a rat cardiac transplantation model [68]. The degree of rejection in transplanted hearts could be directly revealed by the magnitude of intramyocardial infiltration of macrophages and T lymphocytes due to accumulation of targeted microbubbles at the site of inflammation. The second study used microbubbles conjugated to ICAM-1 antibody for selective imaging of cardiac transplant rejection in rats; approximately a tenfold increase in contrast intensity was observed in rejecting myocardium [33].

Molecular Ultrasound Imaging of Thrombus Formation A thrombus or blood clot is achieved via the aggregation of platelets that form a platelet plug, and the activation of the humoral coagulation system. A thrombus is normal in cases of injury, but pathologic in instances of thrombosis. Research in detection of thrombus formation by ultrasound imaging has been ongoing to develop contrast agents that could detect diseases such as stroke, myocardial infarct, and deep vein thrombosis. Various molecular markers conjugated to microbubbles, PFC nanoparticles, and liposomes were developed for animal studies: for platelet targeting, antibodies against GPIIb/IIIa receptors expressed on activated platelets or RGD peptides recognized by the active binding site of GPIIb/IIIa were developed, and various monoclonal antibodies were generated against markers associated with atheroma development such as fibrin, ICAM-1, VCAM-1, and tissue factor. Microbubble binding studies on arteriolar and venular clots in a mouse cremaster muscle model were conducted by Schumann et al. [76]. They confirmed binding of targeted microbubbles in both venules and arterioles. Hamilton et al. [71] used these targeted liposomes for intravascular ultrasound imaging of injured vessels of miniswine as a model system used to create various stages of atheroma. Different types of targeted liposome were generated with conjugation of anti-ICAM-1, anti-VCAM-1, anti-fibrin, and anti-tissue factor antibodies. These targeted liposomes demonstrated targeted enhancement in the vessel walls 5 min after intravenous administration. Similarly, Demos et al. [53] injected echogenic liposome targeted to atherosclerotic plaque created in the Yucatan miniswine animal model. The liposomes were conjugated with anti-fibrinogen or anti-ICAM-1 antibody. The liposomes attached to thrombi and to atherosclerotic arterial wall. Mean acoustic intensities of blood alone and blood with nontargeted and targeted agents showed that targeted liposomes dramatically increased the echogenicity of blood. Lanza et al. [38] employed targeted PFC nanoparticles for intravascular ultrasound and MRI contrast imaging. The PFC nanoparticles were outfitted with biotin–avidin coupling system along with anti-fibrin monoclonal antibodies. Frequency of thrombus detection increased from 2% to 83% after single administration and increased to 96% after second administration of targeted PFC nanoparticles with increased acoustic reflectivity of carotid thrombi of dogs. In a related study, a stretch induced tissue factor was imaged in balloon stretched pig carotid arteries by administering a tissue factor targeted antibody on biotinylated PFC nanoparticles. After administration of the targeted nanoparticles, the arteries were imaged with a 20-MHz intravascular ultrasound system, and the targeted contrast more than doubled the gray scale intensity of the injured areas [72]. Unger et al. [69] have developed microbubbles labeled with RGD analog that enhanced the echogenicity of induced thrombi in dogs. Molecular Ultrasound Imaging of Lymph Nodes The feasibility of using targeted microbubbles to image peripheral lymph nodes under normal conditions in animal models of mice and dogs was tested by Hauff et al. [70] by the use of stimulated acoustic emission (SAE). SAE involves color or power Doppler imaging with the transmission power set high enough to ensure microbubble disruption on the first pulse. This causes a transient high-

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amplitude response. Since an ultrasound Doppler system correlates the signal backscattered from a target within a number of successive pulses, the loss of signal correlation caused by the transient bubble collapse is interpreted by the machine as a random Doppler shift, resulting in a mosaic of colors at the location of microbubbles even without flow. SAE works particularly well with the air-filled microbubbles, as used by Hauff and colleagues. It has proved a powerful method of studying passively targeted agents [77] but has not previously been used to image active targeting. Microbubbles were conjugated to L-selectin specific ligand MECA-79. L-selectin is expressed on circulating lymphocytes and is involved in homing of the lymphocytes to lymph nodes. In these experiments, mice were sacrificed after intravenous administration of L-selectin targeted microbubbles. The lymph nodes were removed and examined by using harmonic color Doppler ultrasound in a tank containing degassed water. The lymph nodes of all the mice showed enhanced acoustic signal due to accumulation of L-selectin targeted microbubbles. In another experiment, anesthetized dogs were scanned with ultrasound in vivo after administration of L-selectin targeted microbubbles intravenously. The targeted microbubbles accumulated significantly in healthy lymph nodes. Thus L-selectin ligand-specific US contrast agent could be a candidate for an indirect method of lymphography for the safe and less invasive US identification of lymph nodes—for example, when performing ultrasound-guided biopsy [70].

11.4 CHALLENGES AND FUTURE DIRECTIONS OF MOLECULAR ULTRASOUND IMAGING Ultrasound imaging lacked effective contrast agent to render it a molecular ultrasound imaging tool until recently. With the introduction of microbubble contrast agents, great progress has been made in the field of contrast enhanced molecular ultrasound imaging. Furthermore, the development of targeted contrast agents has made it possible to image the pathophysiology of many diseases at the molecular and cellular levels using molecular ultrasound imaging and has opened up new avenues for using molecular imaging for therapy. A lot of investigative research has been done to detect intravascular events that play a role in cardiovascular and cerebrovascular diseases, including inflammatory responses, angiogenesis, and thrombus formation by targeted contrast-enhanced ultrasound in the last few years. Yet the field currently is restricted to animal models of human disease. Humanizing antibodies used for target detection using targeted microbubbles can be very expensive. Furthermore, the current biotin–streptavidin system used for preclinical trials to attach ligands onto the shell of ultrasound contrast agents is not suitable for clinical use, due to the immunogenic and allergic risk that streptavidin presents to humans, especially in the case of repeated use. Investigation of synthetic novel peptides that recognize a variety of molecular markers are under way and may provide a cost-effective alternative, circumventing the problem of allergic reactions in humans [14, 15]. Recently, it has been shown that a novel, clinically translatable contrast microbubble targeted to human KDR (which corresponds to VEGFR2 in mice) were designed using a small peptide directly integrated into the shell of contrast microbubbles (thereby avoiding binding chemistry that would preclude a clinical translation into patients) [15]. Molecular ultrasound imaging signal using this human KDR-targeted contrast agent showed significantly higher imaging signal in human colon cancer xenografts in mice than control microbubbles and allowed monitoring of antiangiogenic therapy in vivo [15]. This novel contrast agent was the first that was designed specifically for a clinical translation into patients; first in

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human clinical trials are expected for the near future assessing the potential of molecular ultrasound imaging in a clinical environment [15]. Several new microbubble constructs such as the described human KDR-targeted microbubbles, comprehensive scanner systems, and imaging protocols are currently under development to expand this technology to the clinical arena. Targeted contrast-enhanced ultrasound imaging could potentially be used in a wide range of applications, including diagnostics, image-guided biopsies, and treatment follow-up. Lastly, microbubbles today are still restricted to intravascular space; research into nanoparticles would likely help to overcome this aspect. Finally, extensive research is being directed toward development of the next generation of microbubbles, which are capable of encapsulating therapeutic agents and releasing them when exposed to high MI ultrasound waves. Microbubbles in combination with a therapeutic agent provide the vehicle for targeting molecular events and thus combining imaging with pathophysiology and ultimately therapy. Therapeutic agents could include genes, thrombolytics, and oncological drugs, and this technique has the clinical potential to increase therapeutic efficacy while decreasing systemic side effects. Clinical use of microbubbles has faced problems with regard to safety issues. In 2007 the United States Food and Drug Administration (FDA) ordered black box warnings for two contrast agents—Definity, formerly sold by Bristol-Myers Squibb, and Optison, sold by GE Healthcare—following reports of 11 deaths allegedly associated with the two agents. Optison consists of perflutren protein-type A microspheres, whereas Definity (available as Luminity in Europe and Australia) is a preparation of liposome-encapsulated microspheres containing perflutren. Both contrast agents were initially approved for clinical use in patients for echocardiography to enhance suboptimal echocardiograms to opacify the left ventricular chamber and to improve the delineation of the left ventricular endocardial border. The related ban nearly entirely halted the use of Definity in the United States. In 2008 the FDA eased off the stringent black box warnings that greatly restricted use of Definity and Optison. The FDA announcement followed publication of a retrospective study in the Journal of the American College of Cardiology (2007; 1704–1706). The review, involving 18,671 consecutive patients at St. Luke’s Mid-America Health Institute in Kansas City, Missouri, found no increased mortality risk for patients who received Definity-enhanced echocardiography compared with patients who were imaged without the agent. Some FDA-mandated labeling restrictions remain: Definity is now contraindicated for patients with known or suspected right-to-left, bidirectional, or transient right-to-left cardiac shunts and hypersensitivity to perflutren, an ingredient of the microspherical agents. Intra-arterial injection is still banned. So far, microbubble ultrasound contrast media are not approved for noncardiac applications in the United States [78]. However, there is an ongoing Phase III clinical trial in thr USA testing the diagnostic accuracy of nontargeted contrast-enhanced ultrasound imaging (using the microbubble SonoVue) for liver lesion characterization compared to CT, MRI or histology as the golden standard. Following successful completion of this clinical trial, US FDA approval of nontargeted contrast-enhanced ultrasound imaging for radiological indications is expected. Contrast-enhanced ultrasound imaging with nontargeted microbubbles is already widely used in Canada, Europe, and Asia. and microbubbles have been shown to be nontoxic with an extremely low adverse event rate as low as 0.13% (29 per 23,988 examinations) [79]. In conclusion, molecular ultrasound is an emerging molecular imaging approach finding its niche among other molecular imaging modalities. High spatial and temporal resolution, real-time imaging, noninvasiveness, relatively low costs, lack of ionizing irradiation and wide availability among the imaging community throughout the world are important advantages that will further define the role of this novel imaging technique both in preclinical

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research and clinical applications. In addition, ongoing improvements in ultrasound technology and sophisticated contrast agent design with novel high-affinity targeting ligands using clinically translatable binding chemistry, and further improvements in biodistribution of ultrasound contrast agents beyond the vasculature will further expand the clinical role of molecular ultrasound for imaging diseases at the molecular level in medicine [80, 81, 82].

REFERENCES 1. Hoskins, P. R.; Thrush, A.; Martin, K.; Whittingham, T. A. Diagnostic Ultrasound. Greenwich Medical Media Ltd, London, 2003. 2. Hauff, P.; Foster, S. Ultrasound Basics. Springer-Verlag: Berlin, 2008. 3. Dogra, V.; Rubens, D. J. Ultrasound Secrets. Hanley & Belfus Medical Publishers: Philadelphia, PA, 2003. 4. Aljarboua, Z. Ultrasound effect on biological tissue. Open Educational Resource. 2008. 5. Willmann, J. K.; van Bruggen, N.; et al. Molecular imaging in drug development. Nat. Rev. Drug Discov. 2008, 7(7), 591–607. 6. Rumack, C. M.; Wilson, S. R.; Charboneau, J. W.; Johnson, J. Diagnostic Ultrasound. Mosby: New York, 1998. 7. Leighton, G. The Acoustic Bubble. Academic Press, Inc: London, UK, 1997, 1–117. 8. Lindner, J. R. Molecular imaging with contrast ultrasound and targeted microbubbles. J. Nucl. Cardiol. 2004, 11(2), 215–221. 9. Soye, J. A.; Mullan, C. P.; et al. The use of contrast-enhanced ultrasound in the characterisation of focal liver lesions. Ulster Med. J. 2007, 76(1), 22–25. 10. Stanley, B. Safe use of ultrasound contrast agents. Ultrasound Med. Biol. 2007, 33(2), 171–172. 11. Klibanov, A. L. Ligand-carrying gas-filled microbubbles: ultrasound contrast agents for targeted molecular imaging. Bioconjug. Chem. 2005, 16(1), 9–17. 12. Junru, Wu, W. L. N. Ultrasound, cavitation bubbles and their interaction with cells. Adv. Drug Deliv. Rev. 2008, 60, 1103–1116. 13. Klibanov, A. L. Molecular imaging with targeted ultrasound contrast microbubbles. Ernst Schering Res. Found. Workshop 2005, 49, 171–191. 14. Willmann, J. K.; Kimura, R. H.; Deshpande, N.; Lutz, A. M.; Rosenberg, J.; Cochran, J. R.; Gambhir, S. S. Novel integrin-binding ultrasound contrast agent for molecular ultrasound imaging of tumor angiogenesis. J. Nucl. Med. 2010, 51, 433–440. 15. Pysz, M. A.; Foygel, K.; Rosenberg, J.; Gambhir, S. S.; Schneider, M.; Willmann, J. K. Antiangiogenic cancer therapy: monitoring with molecular US and a clinically translatable contrast agent (BR55). Radiology 2010, 256(2), 519–527. 16. Pochon, S.; Tardy, I.; Bussat, P.; Bettinger, T.; Brochot, J.; von Wronski, M.; Passantino, L.; Schneider, M. BR55: a lipopeptide-based VEGFR2-targeted ultrasound contrast agent for molecular imaging of angiogenesis. Invest Radiol. 2010, 45(2), 89–95. 17. Shah, P. M.; Gramiak, R.; et al. Determinants of atrial (S4) and ventricular (S3) gallop sounds in primary myocardial disease. N. Engl. J. Med. 1968, 278(14), 753–758. 18. Burns, M. Marketing challenges continue for ultrasound contrast media. Biotech Systems Inc. 2004, http://www.biotechsystems.com 19. Lanza, G. M.; Wickline, S. A. Targeted ultrasonic contrast agents for molecular imaging and therapy. Curr. Probl. Cardiol. 2003, 28(12), 625–653. 20. Lindner, J. R.; Song, J.; et al. Noninvasive ultrasound imaging of inflammation using microbubbles targeted to activated leukocytes. Circulation 2000, 102(22), 2745–2750.

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21. Fisher Science Education, Hanover Park, IL. www.fishersci.com 22. Oussoren, C.; Zuidema, J.; Crommelin, D. J.; Storm, G. Lymphatic uptake and biodistribution of liposomes after subcutaneous injection. II. Influence of liposomal size, lipid compostion and lipid dose. Biochm. Biophys. Acta. 1997, 1328, 261–272. 23. Klibanov, A. L. Targeted delivery of gas-filled microspheres, contrast agents for ultrasound imaging. Adv. Drug Deliv. Rev. 1999, 37, 139–157. 24. Klibanov, A. L.; Rasche, P. T.; et al. Detection of individual microbubbles of ultrasound contrast agents: imaging of free-floating and targeted bubbles. Invest. Radiol. 2004, 39(3), 187–195. 25. McCulloch, M.; Gresser, C.; et al. Ultrasound contrast physics: a series on contrast echocardiography, article 3. J. Am. Soc. Echocardiogr. 2000, 13(10), 959–967. 26. Leong-Poi, H.; Song, J.; et al. Influence of microbubble shell properties on ultrasound signal: implications for low-power perfusion imaging. J. Am. Soc. Echocardiogr. 2002, 15(10 Pt 2), 1269–1276. 27. Chomas, J. E.; Pollard, R. E.; et al. Contrast-enhanced US of microcirculation of superficially implanted tumors in rats. Radiology 2003, 229(2), 439–446. 28. Pollard, R. E.; Sadlowski, A. R.; et al. Contrast-assisted destruction-replenishment ultrasound for the assessment of tumor microvasculature in a rat model. Technol. Cancer Res. Treat. 2002, 1(6), 459–470. 29. Willmann, J. K.; Paulmurugan, R.; et al. US imaging of tumor angiogenesis with microbubbles targeted to vascular endothelial growth factor receptor type 2 in mice. Radiology 2008, 246(2), 508–518. 30. Walday, P.; Tolleshaug, H.; et al. Biodistributions of air-filled albumin microspheres in rats and pigs. Biochem. J. 1994, 299(Pt 2), 437–443. 31. Perkins, A. C.; Frier, M.; et al. Human biodistribution of an ultrasound contrast agent (Quantison) by radiolabelling and gamma scintigraphy. Br. J. Radiol. 1997, 70(834), 603–611. 32. Willmann, J. K.; Cheng, Z.; et al. Targeted microbubbles for imaging tumor angiogenesis: assessment of whole-body biodistribution with dynamic micro-PET in mice. Radiology 2008, 249(1), 212–219. 33. Weller, G. E.; Lu, E.; et al. Ultrasound imaging of acute cardiac transplant rejection with microbubbles targeted to intercellular adhesion molecule—1. Circulation 2003, 108(2), 218–224. 34. Korpanty, G.; Carbon, J. G.; et al. Monitoring response to anticancer therapy by targeting microbubbles to tumor vasculature. Clin. Cancer Res. 2007, 13(1), 323–330. 35. Palmowski, M.; Huppert, J.; et al. Molecular profiling of angiogenesis with targeted ultrasound imaging: early assessment of antiangiogenic therapy effects. Mol. Cancer Ther. 2008, 7(1), 101–109. 36. Gref, R.; Domb, P.; Quellec, P.; Blunk, T.; Muller, R. H.; Verbavatz, J. M.; Langer, R. The controlled intravenous delivery of drugs using PEG-coated sterically stabilized nanospheres. Adv. Drug Deliv. Rev. 1995, 16, 215–233. 37. Brigger, I.; Dubernet, C.; et al. Nanoparticles in cancer therapy and diagnosis. Adv. Drug. Deliv. Rev. 2002, 54(5), 631–651. 38. Lanza, G. M.; Wallace, K. D.; et al. High-frequency ultrasonic detection of thrombi with a targeted contrast system. Ultrasound Med. Biol. 1997, 23(6), 863–870. 39. Pana, D.; Wickline, S. A.; Caruthers, S. D. Nanomedicine: perspective and promises with liganddirected molecular imaging. Eur. J. Radiol. 2009, 70(2), 274–285. 40. Lanza, G. M.; Trousil, R. M.; Wallace, K. D.; Scott M. J.; Miller J. G.; Eisenberg P. R.; Gaffney P. J.; Wickline S. A. In vitro characterization of a novel, tissue-targeted ultrasonic contrast system with acoustic microscopy. J. Acoust. Soc. Am. 1998, 104, 3665–3672. 41. Kripfgans, O. D.; Fabiilli, M. L.; Cardon, P. L.; et al. On the acoustic vaporization of micrometersized droplets. J. Acoust. Soc. Am. 2004, 116(1), 272–281.

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64. Kaufmann, B. A.; Lewis, C.; Xie, A.; et al. Detection of recent myocardial ischaemia by molecular imaging of P-selectin with targeted contrast echocardiography. Eur. Heart J. 2007, 28(16), 2011. 65. Villanueva, F. S.; Lu, E.; et al. Myocardial ischemic memory imaging with molecular echocardiography. Circulation 2007, 115(3), 345–352. 66. Bachmann, C.; Klibanov, A. L.; et al. Targeting mucosal addressin cellular adhesion molecule (MAdCAM)-1 to noninvasively image experimental Crohn’s disease. Gastroenterology 2006, 130(1), 8–16. 67. Kaufmann, B. A.; Sanders, J. M.; et al. Molecular imaging of inflammation in atherosclerosis with targeted ultrasound detection of vascular cell adhesion molecule—1. Circulation 2007, 116(3), 276–284. 68. Kondo, I.; Ohmori, K.; et al. Leukocyte-targeted myocardial contrast echocardiography can assess the degree of acute allograft rejection in a rat cardiac transplantation model. Circulation 2004, 109(8), 1056–1061. 69. Unger, E. C.; McCreery, T. P.; et al. In vitro studies of a new thrombus-specific ultrasound contrast agent. Am. J. Cardiol. 1998, 81(12A), 58G–61G. 70. Hauff, P.; Reinhardt, M.; et al. Molecular targeting of lymph nodes with L-selectin ligand-specific US contrast agent: a feasibility study in mice and dogs. Radiology 2004, 231(3), 667–673. 71. Hamilton, A. J.; Huang, S. L.; et al. Intravascular ultrasound molecular imaging of atheroma components in vivo. J. Am. Coll. Cardiol. 2004, 43(3), 453–460. 72. Lanza, G. M.; Abendschein, D. R.; et al. In vivo molecular imaging of stretch-induced tissue factor in carotid arteries with ligand-targeted nanoparticles. J. Am. Soc. Echocardiogr. 2000, 13(6), 608–614. 73. Bloch, S. H.; Dayton, P. A.; et al. Targeted imaging using ultrasound contrast agents. Progess and opportunities for clinical and research applications. IEEE Eng. Med. Biol. Mag. 2004, 23(5), 18–29. 74. Kiessling, F.; Huppert, J.; et al. Functional and molecular ultrasound imaging: concepts and contrast agents. Curr. Med. Chem. 2009, 16(5), 627–642. 75. Lindner, J. R.; Dayton, P. A.; Coggins, M. P.; Ley, K.; Song, J.; Ferrara, K.; Kaul, S. Noninvasive imaging of inflammation by ultrasound detection of phagocytosed microbubbles. Circulation 2000, 102, 531–538. 76. Schumann, P. A.; Christiansen, J. P.; et al. Targeted-microbubble binding selectively to GPIIb IIIa receptors of platelet thrombi. Invest. Radiol. 2002, 37(11), 587–593. 77. Blomley, M. J.; Albrecht, T.; et al. Improved imaging of liver metastases with stimulated acoustic emission in the late phase of enhancement with the US contrast agent SH U 508A: early experience. Radiology 1999, 210(2), 409–416. 78. Brice, J. “FDA eases black box restrictions for ultrasound contrast media.” Diagnostic Imaging, 2008, http://www.diagnosticimaging.com. 79. Kim, T. K.; M.; Jang, H.-J.; Wilson, S. R. Microbubble contrast agents for ultrasound imaging—safety and efficacy in abdominal and vascular imaging. US Radiology 2008, 1(1), 54–57. 80. Deshpande, N.; Needles, A.; Willmann, J. K. Molecular ultrasound imaging: current status and future directions. Clin. Radiol. 2010, 65, 567–581. 81. Pysz, M.; Gambhir, S. S.; Willmann, J. K. Molecular Imaging: current status and emerging strategies. Clin. Radiol. 2010, 65, 500–515. 82. Deshpande, N.; Pysz, M.; Willmann, J. K. Molecular ultrasound assessment of tumor angiogenesis. Angiogenesis 2010 (Epub ahead of print).

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CHAPTER 12

Ultrasound-Based Molecular Imaging Using Nanoagents SRIVALLEESHA MALLIDI, MOHAMMAD MEHRMOHAMMADI, KIMBERLY HOMAN, BO WANG, MIN QU, and TIMOTHY LARSON Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA

KONSTANTIN SOKOLOV Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, and Department of Medical Physics, University of Texas M.D. Anderson Cancer Center, Houston, Texas, USA

STANISLAV EMELIANOV Department of Biomedical Engineering, University of Texas at Austin, Austin, Texas, USA

12.1 INTRODUCTION The need to understand the anatomical and functional aspects of the human body has led to the development of various noninvasive imaging modalities such as X-ray computed tomography (CT), ultrasonography (US), and magnetic resonance imaging (MRI). These imaging modalities are used extensively to diagnose pathologies such as cancer. In particular, ultrasonography, or ultrasound imaging, has gained popularity due to its excellent temporal resolution, reasonable penetration depth, portability, and low cost. Furthermore, ultrasound imaging is a nonionizing technique with no known adverse bioeffects. Currently, ultrasound imaging is being used in various medical fields ranging from obstetric medicine to cardiovascular applications. Obtaining an ultrasound image can be explained as a three-step process. First, the ultrasonic transducer generates pulses of ultrasound waves that are sent through a patient’s body. These waves then interact with organ boundaries and complex tissues, producing echoes due to reflection or scattering. Finally, the backscattered echoes are detected by the same transducer used for transmission of ultrasound waves. Digital postprocessing of these ultrasound signals results in formation of a gray scale image of the body or tissue cross section. Thus contrast in ultrasound images is due to the difference in the acoustic impedance of the tissues being imaged.

Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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Advancements in the field of molecular biology (i.e., discovery of various diseasespecific biomarkers) coupled with development of ultrasound biomicroscopy [1] (imaging performed at the acoustic frequency range of 40–100 MHz) has prompted development of site-specific ultrasound contrast agents that can provide molecular information in the context of a high-resolution anatomical map of the body. Indeed, microbubbles. For example, microbubbles have enhanced ultrasound contrast due to their acoustic impedance difference with tissue. When targeted to various biomarkers in the vascular lumen, these microbubbles can provide some molecular specific ultrasound imaging contrast. However, the gas-filled microbubbles have short lifespans (on the order of minutes) in the body, and due to their large size (∼0.5–500 ␮m diameter) are mostly used as intravascular tracers [2–7]. Other types of ultrasound contrast agents such as liposomes have been developed for molecular imaging [4, 8]. Liposomes have longer circulation time compared to microbubbles and they can also readily be conjugated to various biomarkers. However, passage of liposomes (∼800 nm in diameter) through endothelial gap junctions in the leaky vascular of pathologies such as cancer is size prohibitive. The vasculature of most cancers have endothelial gap junctions from 300 to 800 nm in size, and therefore liposomes would likely not reach the tumor interstitial space [9]. Recently developed ultrasound nanoparticle contrast agents such as perfluorocarbon nanoparticles [4] and silica nanoparticles [10] could extravasate through the leaky vasculature of a tumor into the interstitial space but are less echogenic compared to microbubbles. Metal based nanoparticles (∼5–100 nm in diameter) are being used extensively as molecular specific contrast agents for various imaging modalities such as optical imaging and MRI. Many researchers have demonstrated that metal nanoparticles extravasate and accumulate in tumors due to the enhanced permeability and retention (EPR) effect [11, 12]. This effect is caused by the leaky nature of tumor vessels. Thus by an injection of correctly sized metal nanoparticles, passive accumulation in tumors can be achieved. Furthermore, the metal nanoparticles can be made pathology-specific by bioconjugating them with monoclonal antibodies or antibody fragments. These bioconjugates can be attached either directly to the metal or covalently bound via linker segments. Metal nanoparticles (NPs) such as gold NPs have well-known bioconjugation protocols. Metal NPs are at the same size scale as large protein complexes and so can be targeted to subcellular structures. However, these metal nanoparticles cannot be used as ultrasound contrast agents because their size is much below the resolution of clinically available ultrasound imaging systems. On the other hand, molecular imaging with the metal NPs is possible using ultrasound based imaging modalities, namely, photoacoustic imaging [13, 14] and magneto-motive ultrasound imaging (MMUS) [15, 16]; that is, molecular and functional information could be obtained in the context of the anatomical map of the tissue. For example, gold nanoparticles exhibiting surface plasmon resonance properties and superparamagnetic iron oxide (SPIO) nanoparticles are used as photoacoustic and magnetoacoustic contrast agents, respectively. The synergy among the ultrasound, photoacoustic, and MMUS imaging modalities is schematically represented in Figure 12.1. The ultrasound transducer and the receiver electronics of an ultrasound imaging system are common for both photoacoustic and magnetoacoustic imaging. In photoacoustic imaging, the tissue is illuminated by nanosecond pulsed laser light and the subsequently emitted photoacoustic waves are detected by the ultrasound transducer (Fig. 12.1). The use of photoabsorbers such as gold nanoparticles enhances the contrast in photoacoustic imaging. In MMUS imaging, ultrasound images of the tissue labeled with magnetic nanoparticles are acquired under magnetic field excitation.

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Contrast mechanism RF pulse transmit

Amplitude of received RF echoes

Ultrasound imaging

Amplitude of tissue displacement

Magneto-motive ultrasound imaging Amplitude of photoacoustic transients

Tissue Acoustic Impedance Receive echo

RF pulse transmit

Receive echo

Photoacoustic transients

Motion induced in tissue labeled with magnetic NPs

Harmonic or Pulsed magnetic field

Tissue Optical Absorption

Pulsed Laser

Photoacoustic imaging

FIGURE 12.1 Schematic representation of contrast mechanism and image display in ultrasound, magneto-motive ultrasound, and photoacoustic imaging modalities.

The ultrasound images are then digitally postprocessed to observe the magnetically induced tissue motion (Fig. 12.1). In this chapter, we briefly review the fundamentals of photoacoustic and magnetoacoustic imaging modalities and also demonstrate their molecular imaging capabilities using metal nanoparticles. Furthermore, the feasibility of combining the photoacoustic and magnetoacoustic imaging techniques will be discussed. Finally, the advantages, limitations, and future prospects of ultrasound based photoacoustic and magnetoacoustic imaging modalities are illustrated.

12.2 PHOTOACOUSTIC IMAGING Photoacoustic imaging, also known as optoacoustic and thermoacoustic imaging, involves three steps [17–19]: (1) short laser pulses irradiate the tissue, (2) the tissue absorbs the laser energy; undergoes thermoelastic expansion, and subsequently generates a photoacoustic pressure wave; and (3) the generated pressure wave is detected by an ultrasound transducer. Note that the thermoelastic expansion occurs because the laser pulse duration is shorter than the thermal relaxation time of the tissue (thermal confinement condition). The acoustic pressure P(z) generated at a certain depth z using laser illumination of wavelength ␭ can be expressed as [17, 18]
P(z) =

␤cs2 CP

 ␮a (␭)F(z, ␭)

(12.1)

where ␤ is the thermal expansion coefficient, cs is the speed of sound, CP is the heat capacity at constant pressure, ␮a is the optical absorption coefficient, and F(z) is the laser fluence

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at depth z. The photoacoustic pressure wave is a broadband radiofrequency (RF) signal and can be detected using an ultrasound transducer that determines the resolution of the photoacoustic imaging. A resolution on the order of tens of micrometers can be achieved using a high-frequency ultrasound transducer. The detected photoacoustic RF signals or transients are then processed and displayed as a photoacoustic image. The expression (␤cs2 /CP ) in Eq. (12.1) is called the Gruneisen coefficient, a temperaturedependent parameter. For a constant temperature and laser fluence irradiating the tissue, the photoacoustic signal strength is proportional to the optical absorption coefficient of the tissue. Indeed, the distribution of optical absorption properties of the tissues can be obtained by analyzing photoacoustic images captured at multiple wavelengths. For example, functional information such as the oxygenation of blood or the presence of atherosclerotic plaques may be extracted from multiwavelength photoacoustic images [20, 21]. Photoacoustic imaging exploits both the high contrast associated with optical imaging techniques (due to optical absorption properties) and the spatial resolution of ultrasound imaging (due to detection by the ultrasonic transducer). Unlike optical imaging modalities, where the penetration depth is limited by optical backscattering from the tissue, photoacoustic imaging can image deeper since it detects sound versus light. Moreover, greater penetration depth in tissue can be achieved using near-infrared (NIR) wavelengths because endogenous chromophores such as melanin and blood absorb less light in the NIR range. Similar to optical imaging techniques, the use of exogenous contrast agents or photoabsorbers with higher optical absorption in the optical NIR window could facilitate the detection of pathologies in photoacoustic imaging. Metal nanoparticles, which have greater absorbance compared to conventional dyes such as indocyanine green, qualify as contrast agents in photoacoustic imaging. A variety of shapes and sizes of metal nanoparticles including gold or silver nanospheres, rods, and shells can be used as photoabsorbers [14, 22–25]. It’s apparent that by varying the shape and aspect ratio of nanostructures, particles can be manufactured to absorb light at a desired wavelength across a wide spectrum including the near-infrared spectrum, where the absorption of light by tissue is minimal. Moreover, the change in the optical absorption properties due to the plasmon coupling effect of closely spaced nanoparticle assemblies can also be detected using photoacoustic imaging [14]. 12.2.1 Description of Ultrasound Based Photoacoustic Imaging System A block diagram of the combined ultrasound and photoacoustic imaging system is presented in Figure 12.2. The imaging system primarily involves an integrated probe consisting of a ultrasonic and laser light delivery [13]. The light is delivered to the tissue through a fiberoptic bundle as shown in photographs of an integrated probe with a single element high-frequency transducer or an array transducer, respectively. The laser light delivery could also be done through a combination of various optical elements such as an axicon and prisms [18, 26, 27]. The integrated probe could be handheld or controlled using a positioning stage for mechanical scanning. For special applications such as intravascular photoacoustic imaging, an intravascular ultrasound transducer can be integrated with side fire fiber as shown in Figure 12.2c. The combined imaging system can work either in the ultrasound or photoacoustic imaging mode. In the ultrasound mode, the transducer transmits the rf pulse and receives the backscattered echo from the tissue. In photoacoustic mode, the short laser pulse excites the tissue and the generated acoustic transients are received by the ultrasonic transducer. The

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Trigger Ultrasound Pulser Ultrasound imaging Mode selection

Positioning stage or handheld

Integrated probe Receive Electronics

Transducer

Digital processing of RF data

Image Display

Fiber bundle Photoacoustic imaging (a)

Trigger

Laser

Transducer

(b)

(c)

Optical Fiber

(d)

FIGURE 12.2 (a) Block diagram of combined ultrasound and photoacoustic imaging system. (b–d) Photographs of an integrated probe consisting of a high-frequency single element transducer (b), an intravascular ultrasound transducer (c), and a linear array transducer (d) with optical delivery system.

ultrasound and photoacoustic rf data can be acquired sequentially from the same imaging plane using a single element transducer or an array transducer. In either case, spatially coregistered ultrasound and photoacoustic images can be obtained [13, 28].

12.2.2 Enhancement of Photoacoustic Contrast Using Silver Nanocages In most works to date, gold nanoparticles have been used, almost exclusively, to enhance photoacoustic contrast. Here we present an alternative: silver nanoagents. Being in the noble metal category, both silver and gold exhibit plasmonic resonance in the visible to NIR spectrum of light. Silver, however, has slightly better light absorption and therefore in theory should be a stronger photoacoustic contrast agent [29, 30]. In our attempts to explore this theory, silver nanocages built around a silica core were developed with core sizes ranging from 180 to 520 nm (Fig. 12.3a). These nanocages absorb light broadly across NIR wavelengths. To test their contrast properties in tissue, these nanocages were injected into an ex vivo porcine pancreas and imaged photoacoustically. The result of the photoacoustic and ultrasound imaging performed is shown in Figure 12.3b. Specifically, the combined ultrasound and photoacoustic system with a 7.5-MHz linear array ultrasound transducer and 800-nm laser illumination was employed to image silver nanocages injected into an ex vivo porcine pancreas. The 183-nm silica core, silver outer cage particles (100 ␮L

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(a) Ultrasound

Photoacoustic

Combined

Image

Image

Image

(b)

(c)

(d)

FIGURE 12.3 (a) SEM image of silver nanocages. (b) Ultrasound, (c) photoacoustic, and (d) combined images of ex vivo pancreas tissue injected with silver nanocages. Adapted with permission from Ref. 55.

of 1010 particles/mL) were injected via syringe, approximately 8–10 mm below the pancreas surface. Spatially coregistered ultrasound and photoacoustic rf data was captured using a Cortex ultrasound imaging system (Winprobe Corporation, North Palm Beach, FL, USA). The ultrasound image (Fig. 12.3b) defines the pancreas area, the photoacoustic image (Fig. 12.3c) shows the signal received from the nanocages (white inset), and the combined image (Fig. 12.3d) clearly depicts the location of the nanocages against the background ultrasonic image of the organ. Thus if these nanocages were injected systematically and accumulated in a cancerous area, then combined ultrasound and photoacoustic images could be used to locate them inside the tissue, helping clinicians to better define diseased areas. In this example silver nanoparticles were used instead of the well-characterized gold nanoparticles. Silver nanoparticles have both advantages and disadvantages. One known advantage is silver’s ability to absorb and scatter light better than gold [29]. One disadvantage is silver’s reactivity. Gold is relatively inert, owing to its biocompatibility. Under certain conditions, silver is more reactive and has been shown to be cytotoxic in some in vitro studies [31]. However, silver is well known for its antimicrobial properties and has been used for decades in the treatment of burns [32]. In recent years, silver is being used to line catheters to decrease infection levels [33, 34]. Thus, there is a resurgence in the use of silver in biomedical applications and future studies will be needed to determine its true efficacy and biocompatibility.

12.2.3 Monitoring Accumulation of Gold Nanoparticles in Tumor Using Photoacoustic Microscopy Photoacoustic imaging can also be used to monitor accumulation of nanoparticles in the tumor site over time. A tumor xenograft with A431 cells (human epithelial cancer cell line) was performed in nude mice. After the tumor reached approximately 10 mm in diameter, gold nanoparticles (40 nm in diameter) specifically targeted to the epidermal growth factor receptor [35] (EGFR) are imaged using single element transducer (25 MHz central frequency) and 532-nm wavelength illumination. The combined ultrasound and photoacoustic images of the tumor before and after the injection of gold nanoparticles are shown in Figure 12.4. Clearly, the accumulation of gold NPs over time is evident (Fig. 12.4f–h).

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After injection of Au NPs 0 hrs

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FIGURE 12.4 (a) Monitoring the accumulation of 40-nm gold nanoparticles in tumor using combined ultrasound (a–d) and photoacoustic (e–h) imaging at 532-nm wavelength illumination. The images represent a 13-mm × 11.25-mm field of view.

12.3 MAGNETO-MOTIVE ULTRASOUND IMAGING Magnetoacoustic imaging is a technique that employs high-frequency ultrasound to visualize structural and physiological properties of magnetically labeled tissues. The magneto-acoustic imaging technique capitalizes on significant contrast between magnetic susceptibility of normal tissue constituents and magnetic nanoparticles as well as on deep penetration of magnetic fields into human tissue. Specifically, the magnetoacoustic imaging procedure involves two steps: (1) magnetic nanoparticles such as superparamagnetic iron oxide (SPIO) nanoparticles are used to specifically label pathological tissues, and (2) a pulsed or time-varying magnetic field excites the magnetic nanoparticles to induce motion in the magnetically labeled tissues, along with simultaneous detection of the motion using noninvasive, deep penetrating ultrasound imaging techniques [16, 24, 36–38]. In the last few decades, much research has been devoted to the synthesis of magnetic nanoparticles. Several types of commercial or custom designed nanoparticles such as the superparamagnetic iron oxide (SPIO) nanoparticles were approved by the FDA and has been widely used in clinical applications as MRI contrast agent [36, 39]. Magnetic nanoparticles have been synthesized with a number of different compositions and phases, including iron oxides (Fe3 O4 and ␥ -Fe2 O3 ), pure metals (Fe,Co), spinel-type ferromagnets (MgFe2 O4 , MnFe2 O4 , and CoFe2 O4 ), and alloys (CoPt3 and FePt). A review of the synthesis and functionalization of magnetic nanoparticles has been provided elsewhere [40, 41]. The displacement occurring in the magnetically labeled tissue is dependent on several parameters such as the applied magnetic field and the susceptibility of the magnetic NPs. To understand and estimate the induced motion in the magnetically labeled tissue, the characteristics of the nanoparticles and the magneto-motive force acting on the magnetic nanoparticles need to be analyzed. The magneto-motive force acting on a particle can be expressed as Fm = (m • ∇)B,

(12.2)

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where m is the magnetic moment. Considering a z-directional magnetic flux B = (0,0,Bz ) and the magnetic moment m of paramagnetic particles as (0,0,mz ), the magneto-motive force F m can then be expressed as Fm = (m z • ∇)Bz ,

(12.3)

The magnetic moment mz experienced by a magnetic nanoparticle located in a weakly diamagnetic medium like tissue can be written m z = Vm Mz

(12.4)

where V m is the volume of the magnetic portion of the nanoparticle and can be described as Vm = Vnp × f m , where V np is the total size of the nanoparticle and f m is dimensionless factor called the fraction of magnetite and represents the volumetric ratio of magnetic material in a nanoparticle. The volumetric magnetization M can be written as Mz = (␹ np − ␹ medium )Hz . The magnetic susceptibility of the medium, that is, human tissue, is on order of 10−6 (−11 × 10−6 ≤ ␹ Tissue ≥ −7 × 10−6 ) while the susceptibility of magnetite nanoparticles (Fe3 O4 ) is 70 and it is 250 and 600 for other nickel and cobalt magnetic agents [42, 43]. Hence ␹ medium is assumed to be negligible. Another assumption is that the magnetic flux density (B) doesn’t change significantly over the nanoparticle due to its small size; the volumetric magnetization M can be expressed as Mz = ␹ np

Bz ␮0

(12.5)

Therefore from Eqs. (12.3) and (12.4), Fm =

Vnp f m ␹ np (Bz • ∇)Bz . ␮0

(12.6)

Since (Bz • ∇)Bz = 12 ∇(Bz • Bz ) = Bz

∂ Bz , ∂z

the magneto-motive force acting on a nanoparticles F m can be expressed as Fm =

Vnp f m ␹ np ∂ Bz Bz ␮0 ∂z

(12.7)

Clearly, from Eq. (12.7) it can be seen that there is a significant contrast in the force experienced by the normal and iron-laden tissue. The magneto-motive force is proportional to size (V np f m ) and the susceptibility (␹ np ) of nanoparticles. Increasing the size of the nanoparticles has some limitations for specific labeling or cellular uptake as larger nanoparticles are harder to diffuse out of the leaky vessels. Therefore the magnetic susceptibility (␹ np ) of nanoparticles plays an important role in determining the sensitivity of the magnetoacoustic imaging technique. Iron oxide nanoparticles have high magnetization up to 70 emu/gram; Fe, but other magnetic materials like nickel and cobalt or some alloys can have higher magnetization [44–46]. However, safety and toxicity issues of these materials are still subject to more investigation. Encapsulation of the toxic magnetic alloy core into

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a carbon or polymer shell could be one of the solutions to reduce the cytotoxicity of the metal nanoparticles [47, 48]. The excitation magnetic field can be either continuous-time (harmonic) or pulsed mode. Pulsed excitation has several advantages over harmonic excitation such as shorter operation time and increased magnetic flux density, thus allowing the imaging of deeper tissue structures [15]. Pulsed excitation also has less severe thermal management constraints. On the other hand, a harmonic system creates motion at a predetermined frequency and possibly can be used to identify and filter out the sources of tissue motion due to cardiac and respiratory systems. 12.3.1 Description of Magneto-motive Ultrasound Imaging System The block diagram of the magnetoacoustic imaging system is shown in Figure 12.5. The system consists of two major parts: (1) the magnetic field generation and (2) the ultrasound imaging and data acquisition. According to Ampere’s law, the magnetic field of a solenoid coil, B, has the same time characteristics as the supplied current. Hence a controllable current amplifier or a flash circuit could be used to generate a harmonic or pulsed magnetic field. A conical iron core can also be incorporated into the solenoid to maximize and localize the magnetic field strength applied to tissue specimens [16, 36]. The displacement or the induced motion in the tissue can be monitored using an ultrasound imaging system equipped with an ultrasonic array transducer or with a single element high-frequency transducer. In magnetoacoustic imaging, it is important to capture ultrasound RF data before, during and after the application before, during and after the application of the magnetic excitation to extract the relative motion with respect to the stationary reference. The induced motion can

Trigger generator

Ultrasound Transducer

Controllable delay Positioning Control

Control Unit

Ultrasound Pulser/Receiver

Sample

Stepper Motor Magnetic coil Iron Core

Current Amplifier

Waveform Generator

Signal Acquisition A/D converter Speckle tracking module US/pMMUS image formation

Processing Unit

Cooling casing

FIGURE 12.5 Block diagram of the combined ultrasound and magneto-motive ultrasound imaging system.

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be detected using various quantitative and qualitative techniques including phase-shift or block-matching speckle tracking, Doppler ultrasound, and color/power Doppler. 12.3.2 Demonstration of Magneto-motive Ultrasound Imaging Using Magnetic Iron Oxide Nanoparticles (MIONs) The ability of magnetoacoustic imaging was demonstrated using tissue phantoms fabricated with the J774A.1 cell line (mouse monocytes–macrophages). Like most cells of macrophage phenotype, the J774A.1 cells rapidly uptake dextran-coated nanoparticles [49–51]. Two tissue phantoms were prepared for the study: (1) a control phantom with macrophage cells only and (2) a labeled phantom prepared with MION loaded macrophage cells. Briefly, the cells were cultured in Dulbecco’s Modified Eagle Media (DMEM), supplemented with 5% fetal bovine serum (FBS) at 37 ◦ C in 5% CO2 . To label cells with MIONs, they were incubated with a suspension of nanoparticles in culturing media. After 24 h of incubation with the nanoparticles, the cells were harvested and suspended in a warm (35 ◦ C) gelatin solution (8% w/v) containing 0.5% silica particles (30 ␮m in diameter). The concentration of the control (unloaded) cells in gelatin was approximately equal to those in the loaded cells. The gelatin suspension with cells (control or labeled) was then pipetted into rubber spacers placed in a petri dish. The gelatin solution was allowed to harden at room temperature for approximately 10 min. In addition, a 1–2-mm thick pure gelatin layer was placed on top of the gelatin layer with cells. Finally, to facilitate imaging, the petri dish was filled with 1× PBS solution (phosphate buffered saline) to maintain the appropriate pH in the medium surrounding the tissue phantoms. The cells from each tissue phantom were imaged optically using a Leica DM 6000 upright microscope in epi-illuminated darkfield mode (Fig. 12.6a,b). Images were collected through a 20×, 0.5 NA darkfield objective and detected using an ultrasensitive 12-bit CCD camera. The unlabeled macrophage cells appear bluish white due to their intrinsic light scattering properties (Fig. 12.6a). The macrophage cells labeled with iron nanoparticles appear as orange regions due to light scattering of iron oxide nanoparticles (Fig. 12.6b). The B-scan and the color Doppler images of the tissue phantoms were obtained using a commercially available ultrasound imaging system (Sonic RP by Ultrasonix, Inc.) equipped with a linear array transducer (38-mm aperture, 5-MHz central frequency). The B-scan images

(a)

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(d)

(f)

20 m

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FIGURE 12.6 Optical dark-field images of (a) nonlabeled and (b) magnetically labeled macrophages. B-scan ultrasound images (c,d) and color Doppler images (e,f) of nonlabeled and labeled macrophages embedded in 8% gelatin background.

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cannot directly differentiate between the samples as the difference in image contrast is only derived from nonspecific variations in the phantom materials. The color Doppler image of the phantom with labeled cells clearly indicates the presence of tissue motion and is different compared to that of the control phantom (no motion). The results confirm that magnetoacoustic imaging can readily detect the presence of MIONs in the cells and tissue [16].

12.4 COMBINED PHOTOACOUSTIC AND MAGNETO-MOTIVE ULTRASOUND IMAGING Currently, in the field of molecular imaging, emphasis is being laid on multimodality nano contrast agents that can provide complementary functional information regarding pathologies such as cancer. For example, core–satellite structured dual functional nanoparticles comprised of a dye-doped silica “core” and multiple “satellites” of magnetic nanoparticles were utilized for optical and MR imaging of neuroblastoma cells [52]. Hybrid gold coated iron oxide nanoparticles were also used for combined optical and MR imaging of cancer cells [53]. A combination of MR and optical imaging techniques provides both anatomical and functional information on pathology. However, obtaining the images using the two techniques at the same spatial cross section could be challenging as the imaging equipments required vary significantly. On the other hand, photoacoustic and magnetomotive ultrasound imaging modalities can utilize these hybrid nanostructures possessing both optical absorption properties and magnetic properties as contrast agents. Moreover, photoacoustic and magneto-motive ultrasound imaging techniques utilize the same receiver electronics as ultrasound imaging (Fig. 12.1), and hence can be transparently integrated to provide complementary functional and morphological information at the same imaging cross section. Tissue mimicking samples made with polyvinyl alcohol (PVA) were used to demonstrate the feasibility of combined ultrasound, photoacoustic, and magneto-motive ultrasound imaging. The procedure to make PVA inclusions has been described elsewhere [28]. Briefly, the first inclusion (Fig. 12.7a) was made with 8% PVA and contained 15 micro-meter silica powder (0.5% w/v) for ultrasound contrast, 40-nm diameter spherical gold nanoparticles for optical contrast, and 20-nm (8-nm magnetic core) diameter iron oxide nanoparticles

FIGURE 12.7 Schematic representation (left) of phantom used for combined ultrasound, photoacoustic, and magneto-motive imaging. The phantom consisted of PVA inclusions with (a) mixture of Au and Fe3 O4 iron nanoparticles and (b) no nanoparticles. (c) Ultrasound, (d) photoacoustic, and (e) magneto-motive images of the PVA inclusions. The images represent a 2.5 mm × 3.5 mm field of view. Adapted with permission from Ref. 56.

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for magnetic contrast. The second inclusion (Fig. 12.7b) was also fabricated using 8% PVA, but it contained only silica powder and no nanoparticles—that is, the inclusion had only ultrasound scatters. The inclusions were placed in a water tank attached to a threedimensional (3D) positioning stage to facilitate mechanical scanning. A schematic representation of the experimental setup used for combined ultrasound photoacoustic and MMUS imaging is shown in Figure 12.7 (left panel). The integrated probe (ultrasound transducer and optical fiber delivery system) and the magnetic solenoid are placed on opposite sides of the sample. A single element focused transducer operating at 48-MHz central frequency was used to detect ultrasound and photoacoustic signals. A Q-switched Nd:YAG laser operating at 532-nm wavelength was used to generate photoacoustic transients. A magnetic pulse excitation of approximately 0.5 tesla was delivered to the sample via the conical iron core incorporated in a solenoid. At a particular spatial location, ultrasonic A-lines (with and without magnetic excitation) followed by photoacoustic A-lines were acquired. The sample was mechanically moved to the next spatial location using the 3D positioning stage to acquire the next set of ultrasonic and photoacoustic A-lines. The step size of the mechanical scan was determined by the lateral resolution of the ultrasound transducer. During offline processing of the acquired RF data, a digital bandpass filter was applied to increase the signal-to-noise ratio. The ultrasound and photoacoustic signals were extracted from the A-line records. The analytic signals were obtained by applying a Hilbert transform on the filtered ultrasound and photoacoustic signals and the images were displayed after spatial interpolation. To obtain a MMUS image, the displacement of the inclusions due to pulsed magnetic excitation was determined by correlating ultrasonic A-line records acquired with and without magnetic excitation. The procedure was repeated for each A-line record acquired at different spatial locations and the displacements obtained were plotted as an image after spatial interpolation. The ultrasound, photoacoustic, and MMUS images of the inclusions are shown in Figure 12.7c–e, respectively. The spatial location and the anatomical shape of the two inclusions are clearly indicated in the ultrasound image. Note how, in the photoacoustic image, only one of the inclusions produced photoacoustic signals. This observation was expected since the first inclusion contained gold nanoparticles while no photoacoustic absorbers were present in the second inclusion. The displacement of the inclusions in the MMUS image is displayed on a color map, where white indicates maximum displacement of 100 ␮m and black indicates zero displacement. Clearly, the inclusion with iron nanoparticles displaced more under magnetic excitation than the second inclusion with no iron nanoparticles. The results in Figure 12.7 clearly indicate the feasibility of combining ultrasonic, photoacoustic, and MMUS imaging. In the current experimental setup, the magnetic coil and the integrated probe are placed on opposite sides of the sample. This configuration might not be feasible for in vivo applications. A transducer with an optical fiber in the center for laser light delivery [54] and magnetic coil surrounding it (Fig. 12.5) could make an effective probe for the combined ultrasound, photoacoustic, and MMUS imaging.

12.5 CONCLUSIONS The photoacoustic and magnetoacoustic imaging modalities can be integrated with the widely available ultrasound imaging system with minimum effort. The excellent features of the combined ultrasound, photoacoustic, and magnetoacoustic imaging

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system—noninvasive, nonionizing, descent spatial resolution, and extended penetration depth—can all be obtained at low cost. Moreover, these complementary imaging modalities can simultaneously provide anatomical, optical, and biomechanical properties of the tissue. Given the availability of complex nanostructures with high magnetic susceptibility and optical absorption properties, we anticipate the combination of ultrasound guided photoacoustic and magnetoacoustic imaging techniques could be a valuable tool for early detection of pathologies such as cancer. Further studies are required to evaluate this molecular imaging technique in vivo.

ACKNOWLEDGMENTS Partial support from the National Institutes of Health under grants EB008101 and EB008821 is gratefully acknowledged. The authors would like to thank Dr. Salavat Aglyamov and Dr. Andrei Karpiouk of the University of Texas at Austin for their valuable inputs regarding the development of combined imaging system.

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of ultrasmall superparamagnetic particles of iron oxide in human atherosclerotic plaques can be detected by in vivo magnetic resonance imaging. Circulation 2003, 107, 2453–2458. Lee, J.-H., Jun, Y.-W., Yeon, S.-I., Shin, J.-S.; Cheon, J. Dual-mode nanoparticle probes for high-performance magnetic resonance and fluorescence imaging of neuroblastoma13. Angew. Chem. Int. Ed. 2006, 45, 8160–8162. Larson, T. A.; Bankson, J.; Aaron, J.; Sokolov, K. Hybrid plasmonic magnetic nanoparticles as molecular specific agents for MRI/optical imaging and photothermal therapy of cancer cells. Nanotechnology 2007, 325101. Roy, G. M. K.; Erwin, H.; Wiendelt, S.; Ton, G. v. L.; Frits, F. M. d. M. Photoacoustic imaging of blood vessels with a double-ring sensor featuring a narrow angular aperture. J. Biomed. Optics 2004, 9, 1327–1335. Homan, K.; Shah, J.; Gomez, S.; Gensler, H.; Karpiouk, A.; Brannon-Peppas, L.; Emelianov, S. Silver nanosystems for photoacoustic imaging and image-guided therapy. J. Biomed. Optics 2010, 15, 021316. Qu, M.; Mallidi, S.; Mehrmohammadi, M.; Ma, L. L.; Johnston, K.; Sokolov, K.; Emelianov, S. Y. Phantom study with combined photoacoustic and magneto-acoustic imaging technique, Proceedings of the 31st Annual International IEEE EMBS Conference, 2009, 4763–4766.

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CHAPTER 13

MRI Contrast Agents Based on Inorganic Nanoparticles HYON BIN NA and TAEGHWAN HYEON National Creative Research Initiative Center for Oxide Nanocrystalline Materials, and School of Chemical and Biological Engineering, Seoul National University, Seoul, South Korea

13.1 INTRODUCTION Magnetic resonance imaging (MRI) is currently one of the most powerful diagnosis tools in medical science [1]. MRI produces images through monitoring the relaxation processes of water protons under a magnetic field. With this technique, it is possible to obtain realtime images of the internal anatomy and physiology of living organisms in a noninvasive manner. Since it can give anatomic images of soft tissue with high resolution, it has been the preferred tool for imaging the brain and the central nervous system, for assessing cardiac function, and for detecting tumors. Although MRI itself gives detailed images, making a diagnosis based purely on the resulting images may not be accurate since normal tissues and lesions often show small differences in relaxation time. There are several strategies to obtain high-resolution MR images such as the use of high magnetic field and the design of coil systems. In economical and practical terms, it is more feasible to develop supplements that can maximize the ability of imaging tools. One of the most effective supplements is a chemical compound known as a contrast agent that is introduced to a living body for the improvement of visibility in the image. In particular, biological and functional information can be obtained in image form as a result of the interrelation of the contrast agent and the biological system. Therefore a MRI contrast agent is an essential research field in biological and medical sciences to supply a vision for the analysis of biological information and the diagnosis of diseases. Most of the presently available MRI contrast agents are paramagnetic complexes, usually gadolinium (Gd3+ ) chelates [2]. Among them, Gd-DTPA has been the most widely used. Its main clinical applications are focused on detecting the breakage of the blood–brain barrier (BBB) and on changes in vascularity, flow dynamics, and perfusion. Twenty years ago, a different class of contrast agent, superparamagnetic iron oxide (SPIO), was developed, and it has received great attention as a liver contrasting agent [3]. It was the first nanoparticulate MRI contrast agent and is still used clinically. Gd-based contrast agents Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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enhance the signal in T1-weighted images [2]. On the other hand, SPIO provides a very strong contrast effect in T2-weighted images due to its different contrasting mechanism [3]. Furthermore, its nanoparticulate properties represented by the nanosized dimension and shape allow different biodistribution and opportunity beyond the conventional imaging of chemical agents. The recent development of molecular and cellular imaging, which enables visualization of the disease-specific biomarkers at the molecular and cellular levels, has led to increased recognition of nanoparticles as MRI contrast agents, where nanoparticulate iron oxide has been the prevailing and the only clinically used nanoparticulate agent. As a result of the tremendous progress in nanotechnology, many researchers have developed new nanoparticulate MRI contrast agents that have further improved contrasting abilities and have extra functions. In the following sections, we review the progress in inorganic nanoparticles as MRI contrast agents [4]. In particular, this chapter is focused on the core nanoparticles related to their contrast mechanisms. First, we discuss the T2 contrast agent based on the superparamagnetic property which is the main part in nanoparticular MRI contrast agents. Newly developed nanoparticulate T1 contrast agents are introduced in the latter part of the chapter. 13.2 BASIC PRINCIPLES AND CLASSES OF MRI CONTRAST AGENTS “Contrast” refers to the signal differences between adjacent regions in images, and when the target of the image is the living body they could be “tissue and tissue,” “tissue and vessel,” and “tissue and bone.” Contrast agents make an enhancement of contrast around those interests. Contrast agents for X-ray and CT show contrasting effects according to the electron density difference, and they make direct contrast effects on their locations. However, the contrast mechanism is more complicated for MRI where the contrast enhancement occurs as a result of the interaction between the contrast agents and neighboring water protons, which can be affected by many intrinsic and extrinsic factors such as proton density and MRI pulse sequences. The basic principle of MRI is based on nuclear magnetic resonance (NMR) together with the relaxation of proton spins in a magnetic field [1]. When the nuclei of protons are exposed to a strong magnetic field, their spins align themselves either parallel or antiparallel to the magnetic field. During their alignment, the spins precess under a specified frequency known as the Larmor frequency (␻0 ) (see Fig. 13.1a). When the “resonance” frequency in the radiofrequency (rf) range is introduced to the nuclei, the protons absorb energy and are excited to the antiparallel state. After the disappearance of the rf pulse, the excited nuclei relax to their initial, lower-energy state (Fig. 13.1b). There are two different relaxation pathways. The first, called longitudinal relaxation or T1 relaxation, involves the decreased net magnetization (Mz ) recovering to the initial state (Fig. 13.1c). The second, called transverse relaxation or T2 relaxation, involves the induced magnetization on the perpendicular plane (Mxy ) disappearing by the dephasing of the spins (Fig. 13.1d). Based on their relaxation processes, the contrast agents are classified as T1 contrast agents and T2 contrast agents. Commercially available T1 contrast agents are usually paramagnetic complexes while T2 contrast agents are based on iron oxide nanoparticles, which are the most representative nanoparticulate agents. 13.3 T2 NANOPARTICULATE MRI CONTRAST AGENTS Most of reported nanoparticualte MRI contrast agents were T2 contrast agents based on their superparamagnetic properties. Thus there has been much research on their contrast

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(a)

(b) z ω0

p rf

ul

se ω

ω0 = γB0

B0

Mz

ω0: Larmor frequency

γ : gyromagnetic ratio B0 : Magnetic field

281

y

=

ω0

Mz Mxy

x (c)

(d)

Mz

Mxy

M0 M0(1 – e–1)

Mxye–1

T1’ T1

t

T2’ T2

t

FIGURE 13.1 Basic principle of magnetic resonance imaging. (a) Spins align parallel or antiparallel to the magnetic field and precess under resonance frequency (Larmor frequency, ␻0 ). (b) After induction of a rf pulse, the magnetization of spins changes. Excited spins take the relaxation process of (c) T1 relaxation and (d) T2 relaxation.

effects and clinical trials, and we first introduce T2 contrast agents. The main enhancement of T2 contrast agents is the acceleration on the T2 relaxation process, spin–spin relaxation. When the rf pulse is applied to spins, transverse magnetization on the xy plane (Mxy ) perpendicular to the direction of the static magnetic field is generated (Fig. 13.1b). Net magnetization M, as a vector, has the components Mz and Mxy , which make the interrelated process of spins. The change in Mz is due to energy transfer, whereas that in Mxy is due to the process of spin dephasing, that is, the randomization of the magnetization of excited spins with the same phase coherence immediately after the application of the rf pulse. Their phase coherence in the xy plane disappears due to the difference of magnetic field experienced by the protons. The magnetic field difference is produced by the system performance in shimming and the magnetic properties of imaging objects. Although the inhomogeneity of the static magnetic field by the system imperfection can be reduced by a variety of tools including the shimming coils and shimming algorithms and the usage of the spin echo sequence to reverse this effect, it affects the decay of transverse magnetization. As other source of field inhomogeneity, the magnetic properties of imaging objects can cause phase incoherence. The spin–spin interaction between the hydrogen nuclei or electrons causes a loss of transverse coherence, which makes the true and characteristic T2 relaxation of tissues. For example, the proton interaction of macromolecules in tissue can induce a local magnetic field, as well as a change in the actual magnetic field in their vicinity. Furthermore, the local magnetic field gradient can be induced from the differences in the magnetic susceptibility between the adjacent and different tissues or by contrast agents. Therefore transverse relaxation is affected by inhomogeneous magnetic field produced from tissue-inherent factors or external sources and the total relaxation time, T2∗ , is

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described by 1 1 = + ␥ BS T2∗ T2

(13.1)

where ␥ BS represents the relaxation by the field inhomogeneities and is called susceptibility effects. The magnetization of paramagnetic materials, such as gadolinium complexes, is directly dependent on the number of ions and they have no magnetization in the absence of an external magnetic field. However, ferromagnetic iron oxide has a very large magnetic susceptibility, which can persist even upon removal of the external magnetic field. nanosized iron oxide particles are superparamagnetic, losing their magnetization in the absence of an external magnetic field. However, when an external magnetic field is applied, they exhibit a strong magnetization, which can cause microscopic field inhomogeneity and activate the dephasing of protons. Therefore iron oxide nanoparticles shorten T 2 and T2∗ relaxation times of the neighboring regions and produce a decreased signal intensity in T2- and T2*-weighted MR images. 13.3.1 Dextran-Coated Iron Oxide Nanoparticles Since their first use as MRI contrast agents 20 years ago, iron oxide nanoparticles (usually magnetite (Fe3 O4 ) or maghemite (␥ -Fe2 O3 )) have become extremely popular due to their dramatic ability to shorten T2∗ relaxation times in the liver, spleen, and bone marrow, by selective uptake and accumulation in the cells of the reticuloendothelial system (RES) [3]. With their high magnetization, their selective signal loss allowed for a new class of MRI contrast agents in the world of dominant T1 contrast agents based on ionic complexes. Since the magnetic property of the nanoparticles and their biological distribution are directly dependent on their size, they have been classified by size as follows: (1) micrometer-sized paramagnetic iron oxide (MPIO; several micrometers), (2) superparamagnetic iron oxide (SPIO; hundreds of nanometers); and ultimately, (3) ultrasmall superparamagnetic iron oxide (USPIO: less than 50 nm) [4]. Among them, two classes are widely used in MRI: SPIO and USPIO. There are several approved products in the SPIO family: Feridex® (Berlex) in the United States or Endrem® (Guerbet) in Europe, and Resovist® (Schering) in Europe and Asia. Smaller nanoparticles, USPIO, have similar composition and originate from SPIO. At the beginning they were prepared through size fraction of an SPIO mixture. Nowadays, uniform USPIOs are produced with the improvement of a synthetic technique (Combidex® (Advanced Magnetics) in the United States and Sinerem® (Guerbet) in Europe). Products with USPIO are under consideration for clinical uses by the U.S. Food and Drug Administration (FDA). The most representative and traditional method to prepare SPIOs and USPIOs is the reduction and coprecipitation reaction of a mixture of ferrous and ferric salts by addition of an alkaline solution under vigorous stirring or sonication [5]. This precipitation reaction is performed in the presence of stabilizers such as hydrophilic polymers, dextran derivatives. Figure 13.2 shows the representative synthetic scheme of SPIO and USPIO, and an electron microscopic image of ferumoxides (Feridex® or Endrem® ) and MION. Because there were insufficient controls in the synthetic process, resulting particles were formed with a broad range of sizes. Furthermore, these particles consisted of multiple iron oxide cores within a dextran stabilization shell. The size of particles is the main factor that controls their

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FIGURE 13.2 (a) Scheme of the preparation of SPIO and USPIO. (b,c) Transmission electron microscope (TEM) images of (b) ferumoxides [5a] and (c) MION [7] GFC*: gel filtration chromatography. (Reproduced with permission of Elsevier Inc. and the Massachusetts Medical Society.)

biological characteristics such as blood half-life and biodistribution. Because SPIO has relatively large overall size and related opsonization by phagocytic cells located in the RES, it shows fast clearance from the body and short lifetime. Small nanoparticles have a longer plasma circulation time due to their slow excretion by the liver. Therefore iron oxide nanoparticles for molecular imaging are usually in the class of USPIO (<50 nm). Weissleder and co-workers first demonstrated MRI applications of USPIO [6, 7]. As mentioned before, early USPIOs were prepared via size fraction of an SPIO mixture by size-exclusion gel chromatography. The progress of synthetic methods has brought various USPIO products such as Combidex® , MION (monocrystalline iron oxide nanoparticles), and CLIO (crosslinked iron oxide). In particular, the amination of dextran laid down the foundations that allowed USPIO to become a functionalized and smart imaging platform for molecular imaging. As shown in Figure 13.2, USPIO is much smaller than SPIO with single iron oxide core. This decrease in size results in lower magnetism due to a loss of ferromagnetic coupling and an increase of quadrupole-coupled paramagnetic doublets [8]. Therefore the relaxation properties of SPIO and USPIO are very different. The efficiency of the relaxation property, relaxivity, shows the dependency on the size of nanoparticles because the enhancement of relaxation is related to the magnetic property, especially in spin–spin relaxation (detailed explanation is given in the later sections). The relaxivity coefficient r2 of SPIO is around 150 mmol−1 s−1 , whereas that of USPIO is a much smaller value of around 60 mmol−1 s−1 . However, it is hard to say that SPIO is a more efficient contrast agent than USPIO based only on the relaxivity value because their biological behaviors are very different due to their overall sizes. Clinical uses of SPIO and USPIO are often directed by size-dependent distribution in the body rather than their relaxivities. SPIO is selectively taken up by the Kupffer cells in the liver, spleen, and bone marrow, and the usual clinical targets of SPIO have been liver diseases [3]. If the normal liver architecture is destroyed by a hepatic disease, such as a primary liver tumor or liver metastasis, the region will have a lack of Kupffer cells. Due to the negligible uptake by the abnormal part, the SPIO presents a strong contrast between normal and abnormal tissue, thereby enabling clear detection of the abnormal tissue. However, SPIO’s rapid uptake by

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(a)

(b)

(c)

FIGURE 13.3 (a) Mechanism of action of ultrasmall superparamagnetic iron oxide nanoparticles. (b,c) MR images of a metastatic lymph node (b) before and (c) after use of USPIO [7]. (Reproduced with permission of the Massachusetts Medical Society.)

the liver leads to its fast excretion from blood plasma, which induces a limitation in the SPIO’s ability to depict the molecular and biological functions of other organs and legions. A typical clinical application of USPIOs is lymph node imaging [7]. The detection of lymph nodes is critical for accurate tumor staging and subsequent therapeutic planning. Since USPIOs are very small, they can be extravasated from the blood vessels into interstitial spaces and they are transported to lymph nodes through lymphatic vessels (Fig. 13.3a). As nodes that have malignant cells cannot undergo phagocytosis by nodal macrophages, nanoparticles are taken up only by normal nodes. Their accumulation within normal nodes produces significant susceptibility effects, and the specific uptake allows more sensitive detection between normal and malignant cells. Therefore they are expected to improve the diagnosis of metastatic tumors, and they are currently under investigation for human applications. Weissleder and co-workers reported the successful detection of lymph node metastases in patients with prostate cancer using USPIO (Fig. 13.3b) [7]. USPIOs have

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also been tested as a blood pool agent because they are readily distributed in intravascular extracellular space. For example, many clinicians have tried to use USPIO for angiography, functional MRI (fMRI), and passive targeted imaging of tumors. However, the aforementioned applications are based mainly on the biological behavior of iron oxide nanoparticles. In the view of molecular imaging, recent interest has focused on active imaging with maximization of nanoparticulate features. As mentioned in the Introduction, nanoparticles have advantages over paramagnetic complex agents. Nanosized particles can easily be taken up by both macrophage cells and nonphagocytic cells. Moreover, targeted imaging is possible because nanoparticles have a large surface area that can be conjugated with biological and targeting probes such as antibodies, oligonucleotides, aptamers, and other imaging probes. Iron oxide nanoparticles have shown outstanding results in the field of target-specific in vivo or in vitro imaging, in particular, monitoring the migration and tracking of cells, and disease targeted imaging. Cell therapy is a personalized treatment that has been used for more than 50 years. In particular, recent research on stem cells gives hope to those suffering from incurable diseases. For successful cell therapy and the monitoring of cell tracking and differentiation in vivo, it is critical to develop methods for the noninvasive assessment of the fate and distribution of cells. Because the magnetic nanoparticles’ large susceptibility provides highly sensitive tracking with very low detection limits, their use enables the labeling of target cells. MRI anatomic imaging using SPIO and USPIO has been employed for phagocytic cells, such as neutrophils, macrophages, and monocytes, and for nonphagocytic cells, such as lymphocytes, glioma cells, and stem cells [4]. Initial demonstrations of cellular imaging were macrophage imaging including hepatic imaging and lymph node imaging through intravenous injection of SPIO or USPIO mentioned earlier. Because monocytes and macrophages can phagocytize nanosized particles, those phagocytic cells and related diseases are targets of in vivo labeling and imaging. For instance, experimental autoimmune encephalomyelitis (EAE) shows the development of macrophages around disrupted blood–brain barrier (BBB), and transplant rejection induces inflammation and macrophages. USPIO has been applied in these in vivo labeling with its long lifetime [4]. Recent research has focused on in vivo MRI imaging of ex vivo labeled cells with SPIO and USPIO for highly efficient and selective labeling on targets. Moore and co-workers prepared labeled pancreatic islet cells through the incubation islets with UPIO stabilized by a modified dextran, which carries fluorochromes for near-infrared optical imaging [9]. They demonstrated in vivo tracking of transplanted pancreatic islet cells using MRI in diabetic mice in real time. In the initial studies, those cells were labeled by simple incubation with a high concentration of native nanoparticles. Because simple incubation has the limitations of long incubation time and low labeling efficiency in the case of the nonphagocytic cells, various chemical, mechanical, and biological approaches were developed to deliver the nanoparticles efficiently into the cells [10–13]. First, the use of transfection agents and electroporation were adapted from the methods used to introduce DNA or plasmids [11]. Bulte and co-workers reported the efficient labeling of stem cells with magnetic iron oxide through electromagnetoporation, which combines the electromechanical permeability changes of the cellular membrane and the accelerated movements of iron oxide by the magnetic field [12]. Another approach is the modification of the nanoparticles by the attachment of biomaterials that facilitate particle binding to the anionic cellular membrane, followed by internalization. Viruses, peptides, proteins, and antibodies have been conjugated or coated with USPIO and SPIO [13]. In particular, dextran-stabilized MIONs were popularly conjugated with biomaterials because MIONs can readily be functionalized

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with amine groups by epichlorohydrin treatment. Josephson et al. [13a] reported that crosslinked iron oxide nanoparticles conjugated with HIV-Tat proteins (CLIO-Tat) showed efficient nonphagocytic cell labeling through activated macropinocytosis. The facile process of functionalizing dextran-stabilized iron oxide nanoparticles with amine groups (CLIO) established a platform for active molecular imaging. Many kinds of biomolecules including antibodies, proteins, peptides, polysaccharides, and aptamers were covalently bound on the surface of the iron oxide nanoparticles for their site-specific accumulation at the targets of interest [14–16]. This functionalization of nanoparticles with bioactive materials moved MRI into the next stage of active targeted imaging. Kang et al. [14] reported antibody-conjugated CLIO and in vitro targeted MR imaging of E-selectin in endothelial cells. Furthermore, amine groups can be conjugated with small molecules, such as optical dyes for multimodal imaging, and synthetic ligands to develop targeted imaging and therapeutic materials with rapid screening and imaging [15, 16]. Weissleder and co-workers developed biocompatible magnetic relaxation switches (MRSs) to detect molecular interactions, such as DNA–DNA, protein–protein, protein–small molecule, and enzyme reactions [16], based on the observation that self-assembled magnetic nanoparticles exhibited enhanced spin–spin relaxation times (T 2 ) compared to the individual magnetic nanoparticle dispersions. 13.3.2 New T2 MRI Contrast Agents For last 20 years, dextran-stabilized iron oxide nanoparticles (SPIO, USPIO) have been applied as MRI contrast agents for conventional clinical imaging and molecular imaging. As mentioned before, they are not uniform particles and even have multiple cores inside a particle due to the limitations of their synthetic processes. For molecular and cellular studies, imaging with high resolution and high sensitivity are required to fulfill the quantization and interpretation issues. The T2 contrasting effect of agents is dependent on their magnetic properties, and those of nanosized materials are strongly dependent on the quality of the nanoparticle. However, a coprecipitation reaction in water at a relatively low temperature produces the formation of polydisperse and poorly crystalline nanoparticles with inferior magnetic properties. With the progress of nanotechnology, various nanoparticles have been prepared with improved uniformity and crystallinity. Among them, magnetic nanoparticles have received lots of attention from many scientists because of their potential for various application fields including ferrofluid magnetic refrigeration systems and magnetic carriers for drugs and catalysis [17]. Solvothermal or thermal decomposition processes have been reported as synthetic strategies for high-quality nanoparticles. However, they need to be modified to fulfill specific requirements for biomedical imaging. Of course, nanoparticles have to possess suitable properties for the contrasting mechanism. Second, nanoparticles should be compatible with biological systems. Finally, nanoparticles should have reactive moieties on the surface for conjugation with biological active materials. In general, the contrasting ability originates from the core materials, whereas the biocompatibility and conjugation capability are related to the surface properties. Recently, numerous studies have been conducted to develop new MRI contrast agents based on magnetic nanoparticles with core materials of improved magnetic properties and suitable surface characteristics.

Synthesis of Uniform Nanoparticles Various magnetic nanoparticles have been researched as signal-enhancing cores for T2 contrast effect, and iron oxide and related ferrite

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nanoparticles are the most popular class for T2 agents. Interestingly, they have attracted many scientists not only in the imaging area but in various fields of fundamental and applied sciences because of their unique magnetic properties, which cannot be achieved by their bulk counterparts [17, 18]. Therefore many material scientists have done research on their syntheses and properties. There are several key issues in the synthesis of nanoparticles including magnetic nanoparticles: (1) particle size distribution (uniformity), (2) particle size control, (3) crystallinity and crystal structure, and (4) large-scale production for applications [18a]. That is, the final goal is to produce size-tunable nanoparticles in large quantities. For the last few decades, many scientists prepared nanoparticles through two different approaches: the “top–down” approach, which is molecular-scale lithography via physical methods, and the “bottom–up” approach, which employs solution-phase colloidal chemistry [17, 18]. The advantage of the physical methods is the production of a large quantity of nanoparticles, but it is very difficult to control the size and distribution of nanoparticles. In contrast, colloidal chemical synthetic methods can be used to synthesize uniform nanoparticles with controlled particle size and shape, although generally only subgram quantities are produced. Very recently, colloidal chemical methods were developed to synthesize a large quantity of uniform nanoparticles [19]. In this chapter, we briefly discuss these colloidal chemical methods for nanoparticles. The synthesis of metal and metal oxide nanoparticles, which are major candidates for the core parts of MRI contrast agents, has been well documented in several review articles [18]. The mechanism studies revealed that uniform-sized nanocrystals can be produced when the burst of nucleation, which enables the separation of nucleation and growth processes, is combined with the subsequent diffusion controlled growth process through which the size focusing works [18a,c]. Several chemical synthetic methods have been used to synthesize uniform nanoparticles of metals and metal oxides. The “hot-injection” method, which involves the injection of organometallic precursors in surfactant solutions at high temperature followed by aging at slightly lower temperature, has been used extensively to synthesize uniform nanocrystals of CdSe, Co, and Fe. However, incomplete separation between nucleation and growth processes often resulted in a mixture of nanoparticles with various sizes, and a size-selection procedure is necessary to prepare uniform nanoparticles (Fig. 13.4a). Moreover, for the burst of nucleation process, highly reactive precursors, which are very unstable and even unavailable for many materials, are required. This hot-injection process is performed at high temperature and under extremely vigorous conditions, which can become very dangerous. Consequently, this hot-injection method cannot be used for large-scale synthesis of uniformly sized nanoparticles. The Hyeon group developed a “heat-up” process to synthesize various uniform oxide nanoparticles, in which reaction mixtures composed of metal precursors, surfactants, and solvents are slowly heated from room temperature to high temperature followed by aging [18c]. Kinetics studies showed that the overall process of the size distribution control consists of three steps: the accumulation of monomers from the thermal decomposition of metal precursors, burst nucleation, and diffusion controlled growth for size focusing to produce uniform nanoparticles (Fig. 13.4b). The Hyeon group reported the synthesis of highly uniform iron and iron oxide nanoparticles without a size-selection process via the heat-up method [20a]. In this report, an iron precursor (Fe(CO)5 ) was mixed with a surfactant (oleic acid) in solvent (octylether) and they were slowly heated to reflux. Sun and co-workers synthesized monodisperse magnetite nanoparticles by a similar heatup process using iron acetylacetonate as a precursor and 1,2-hexadecandiol as a mild

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(a)

(b)

FIGURE 13.4 Schematic illustration of (a) hot-injection method [18a] and (b) heat-up method. (Reproduced with permission of the Royal Society of Chemistry.)

reductant in the presence of oleic acid and oleylamine [20b]. They also synthesized metal ferrite nanoparticles such as MnFe2 O4 and CoFe2 O4 [20b]. In 2004, the Hyeon group reported large-scale synthesis of uniform metal oxide nanocrystals via the heat-up method using metal–oleate complexes as the precursors [19a]. Metal–oleate complexes are easily prepared from the reaction between metal chlorides and sodium oleate. This modified heat-up method has been applied to the synthesis of uniform nanocrystals of various materials such as hexagonal-cone-shaped ZnO and pencil-shaped CoO [20c,d]. Especially, iron oxide nanocrystals synthesized via this method show remarkable size uniformity. Because of its simplicity, the heat-up method using a metal–oleate precursor is highly advantageous for scale-up, as demonstrated by the production of iron oxide nanocrystals in tenths of grams from a single batch reaction without the deterioration of the size uniformity.

Size-Dependent Magnetic Properties and T2 Contrast Effects Magnetic nanoparticles have mainly been applied as T2 contrast agents. Acceleration of spin–spin relaxation (T2 shortening) by magnetic nanoparticles results from the dephasing of the magnetic moments due to the magnetic field gradients created by the small magnetized particles. In this process, the major relaxation mechanism is the dipolar outer-sphere interaction between the water proton spins and the magnetic moment of the magnetic nanoparticles. Therefore, as shown in the model suggested by Koenig and Keller, spin–spin relaxation is

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dependent on the magnetic moment of the nanoparticles (␮) [21]: R2 =

1 a ␥ 2 ␮2 CNP J (␻, ␶D ) = T2 dNP D

(13.2)

where a is a constant, dNP the diameter of the nanoparticle, D is the diffusion coefficient, ␮ is the magnetic moment of the nanoparticles, ␥ is the gyromagnetic ratio of the water proton, CNP is the concentration of the nanoparticles, and J(␻, ␶ D ) is the spectral density function. In other words, to be efficient T2 contrast agents, nanoparticles should possess a large magnetization. Although magnetism is an intrinsic property of bulk materials, the magnetic properties of nanoparticles are strongly dependent on their size, shape, and surface property [22]. A dramatic increase in surface area induces preponderant surface canting effects and consequently increases the magnetization, as described by the following equation [23]: m S = MS [(r − d)/r ]3

(13.3)

where mS is the saturation magnetization of the nanoparticle, M S is the saturation magnetization of the bulk material, r is the size of the nanoparticle, and d is the thickness of the disordered surface layer. To make efficient T2 contrast agents, we have to control the magnetic properties of the nanoparticles through the designed control of (1) the intrinsic material properties such as material composition and crystal structure, and (2) extrinsic factors such as size and shape [23c]. Most T2 contrast agents are based on iron oxide nanoparticles and are composed of aggregated nanoparticles with multiple iron oxide (4–5 nm) cores and dextran coating. USPIOs with a single core (MION, CLIO) have low relaxivities due to their small core size. As described earlier, uniform and highly crystalline iron oxide nanoparticles have been synthesized by various synthetic methods in the last decade [18–20]. Recently, some of these uniform-sized iron oxide nanoparticles have been used for MRI contrast agents. Cheon and co-workers systematically studied the relationships among size, magnetism, and relaxivity of uniform-sized iron oxide nanoparticles (see Fig. 13.5a) [22b, 23c]. Larger iron oxide nanoparticles (MEIO) possess higher magnetization values, exhibit stronger T2 contrast effects, and possess higher magnetization values. The r2 relaxivity coefficient is 78 mM−1 s−1 for 4-nm MEIO and this gradually increases 106, 130, and to 218 mM−1 s−1 for 6-, 9-, and 12-nm size MEIO nanoparticles, respectively. Large iron oxide nanoparticles have a large magnetization and high r2 relaxivity, which makes it possible to increase the sensitivity in T2/T2*-weighted images and to ease the toxicity concern by decreasing the agent dose. Because magnetism is an intrinsic property of materials, magnetic nanoparticles with high magnetism have recently been synthesized to develop new MRI contrast agents with improved relaxation properties and biocompatibility. Alloy materials can be candidates of more efficient T2 contrast agents. Various bimetallic ferrite nanoparticles including CoFe2 O4 , MnFe2 O4 , and NiFe2 O4 have been tested as T2 contrast agents (Fig. 13.5b and Table 13.1) [23c, 24]. In particular, the MnFe2 O4 nanoparticles (Mn-MEIO) have been found to have a very high magnetization values of 110 emu/g and large r2 relaxivity of 358 mM−1 s−1 [24]. This relaxivity is about six times higher than that of CLIO nanoparticles.

290 Dextran Carboxydextran Dextran Dextran DMSA[b] DMSA[b]

DMSA[b] DMSA[b] PEG PEG

Fe3 O4 , ␥ -Fe2 O3 Fe3 O4 Fe3 O4 Fe3 O4 Fe3 O4 MnFe2 O4

CoFe2 O4 NiFe2 O4 Fe3 O4 FePt (fcc)

Ferumoxides (Feridex) Ferucarbotran (Resovist) Ferumoxtran (Combidex) CLIO-Tat

WSIO (MEIO)

MnMEIO

CoMEIO NiMEIO Au-Fe3 O4 Au-FePt

4.96 4 5.85 5 4 6 9 12 6 9 12 12 12 20 6

Diameter of Core [nm] 160 60 35 30

Hydrodynamic Diameter [nm]

61 60 25 43 80 101 68 98 110 99 85

45

Magnetization [emu/g][a]

120 186 65 62 78 106 130 218 208 265 358 172 152 114 59

r2 [mM-1 s-1 ]

1.5 1.5 3.0 3.0

1.5

1.5

1.5 1.5 1.5 1.5

B0 [T]

5a 5b 5b 13a 22b, 23b 22b, 23b 22b, 23b 22b, 23b 23b, 24 23b, 24 23b, 24 23b, 24 23b, 24 41a 41b

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Surface

Core Material

Name

TABLE 13.1 Properties of T2 Contrast Agents

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(b) 6 nm

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250 MEIO

Co-MEIO

Ni-MEIO

200 150 100 50 0

150 5

size

10

(nm)

100

50 15

0

ms

)

Fe)

/g( emu

Relaxivity coefficient (mM–1s–1)

r 2(mM–1s–1)

Mn-MEIO

400 300 200 100 0 Mn-MEIO

MEIO

Co-MEIO

Ni-MEIO

(

FIGURE 13.5 (a) The size-dependent magnetic property and r2 relaxivity of monodisperse iron oxide nanoparticles. (b) The magnetic property and r2 relaxivity of several bimetallic ferrite nanoparticles [23b]. (Reproduced with permission of the American Chemical Society.)

Biocompatible Packing and Modification of T2 Nanoparticulate Contrast Agents For many biomedical applications, in particular, for in vivo imaging, nanoparticles should possess a good colloidal stability and low toxicity in a biological environment. Furthermore, surface modification is important for the nanoparticles to impart additional functions because bioactive materials are conjugated through the reactive groups on the surface. Dextran-stabilized iron oxide nanoparticles such as CLIO and MION were synthesized by coprecipitation methods, which prepare iron oxide nanoparticles with biocompatible dextran directly, and they did not need any additional step to stabilize in aqueous media. On the other hand, because conventional dextran is very stable and inert, it is very important to have reactive moieties for further functionalizations of nanoparticles. For example, amination has been investigated from the very early days of USPIO. A representative method is the crosslinking with epichlorohydrin and the resulting product is CLIO [6]. Although high-quality magnetic nanoparticles with uniform particle size and high crystallinity were synthesized from thermal decomposition of metal precursors in organic media, they are hydrophobic and consequently are not sufficiently stable in aqueous media for biomedical applications. Therefore modifying these nanoparticles is essential to endow them with hydrophilic properties so that they can be used extensively for various biomedical applications [25]. Polymer and silica derivatives are typical materials for the surface modification of nanoparticles for biomedical applications, to make inorganic nanoparticles compatible and stable, and what’s more, they are functionalized easily. Recently, various surface modification methods were developed and two representative strategies are ligand exchange with water dispersible ligands and encapsulation with biocompatible shells (Fig. 13.6). The ligand exchange method [26] has several advantages including simple procedure, thin passivating layer, and small overall size. Nanoparticles prepared in organic solvents

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Ligand exchange with hydrophilic ligands

Hydrophobic shell

Encapsulation with hydrophilic surface

FIGURE 13.6 Two strategies for biocompatible nanoparticles: (a) ligand exchange with hydrophilic ligands via substitution reaction of stabilizing ligands on nanoparticles reaction and (b) Encapsulation by hydrophilic matrix such as polymers and silica derivatives.

are stabilized by surfactants that have their hydrophilic head groups bound onto the surface of the nanoparticles and the hydrophobic hydrocarbon tails facing the solvent. Since the atoms on the surface of the nanoparticles have an affinity for the functional groups, surface functionalization with proper hydrophilic ligands allows the phase transfer of nanoparticles from organic media to aqueous media. Various kinds of nanoparticles can undergo substitution reactions with excess bifunctional ligands that consist of a strong binding moiety to the surface of the nanoparticles and relatively low binding hydrophilic group. Because metal oxide nanoparticles have a relatively unreactive surface compared to other kinds of nanoparticles such as chalcogenide-based quantum dots, there are very limited kinds of functional groups available. For example, metal oxide nanoparticles do not bind strongly to alkanethiols that serve as excellent ligands for nanoparticles of chalcogenides and metals. Consequently, there has been intensive recent research on the development of new ligand systems that can bind strongly to metal oxide nanoparticles, in particular, iron oxide nanoparticles. Dopamine was used as a ligand to immobilize functional molecules on iron oxide nanoparticles [26a,b], whereas Peng and co-workers took advantage of the interaction between hydroxamic acid and iron oxide [26c]. Cheon and co-workers were able to form waterdispersible iron oxide nanoparticles using a small organic molecule, 2,3-dimercaptosuccinic acid (DMSA), which has a bidentate carboxyl group [22b]. The resulting iron oxide nanoparticles were dispersible in water and they were successfully used for in vivo MRI. To improved the biocompatibility of the nanoparticles, the backbones of the ligands were designed to contain biocompatible polymers such as polyethylene glycol (PEG). Sun and co-workers demonstrated that a dopamine derivatized PEG (DPA-PEG) was an efficient ligand to stabilize 9-nm Fe3 O4 nanoparticles. The resulting nanoparticles were stable in biological media and showed very low uptake by macrophage cells [26b]. PEG-derivatized phosphine oxide ligands (PO-PEGs), consisting of a biocompatible PEG tail group and surface coordinating phosphine oxide head group, could stabilize several kinds of oxide nanoparticles in aqueous media [26d]. They were prepared by simple reaction between

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POCl3 and PEGs. Polymeric or dendrimeric ligands have also been used for the stabilization of iron oxide nanoparticles in aqueous media derived from their multiple functional groups and bulky structure [26e]. These ligands interact through a number of bonds on the nanoparticles by ligand exchange. The second strategy for surface modification is the formation of biocompatible shells or matrices around the nanoparticles. The advantages of this strategy include easy modification and expansion in those shell structures. Although overall sizes of products are relatively larger than those from the ligand exchange method, they can contain other kinds of material together and form multifunctional structures. There are a variety of biocompatible shell formation methods; they can broadly be classified according to the shell materials and the encapsulation processes. Typical shell materials are silica and polymers. Crosslinked amorphous silica shell can stabilize oxide nanoparticles. Although the toxicity of nanosized silica is still in debate, there have been promising in vitro and in vivo reports on silica-encapsulated nanoparticles [27–29]. Ying and co-workers reported silica-coated nanoparticles produced by base-catalyzed silica formation in a reverse microemulsion [28a,b]. Although the silica layer is relatively thick (10–30 nm), the encapsulation process is simple and various kinds of nanoparticles can be encapsulated. It also allows the encapsulation of various kinds of nanoparticles simultaneously within one shell. Fluorescent dye molecules can be incorporated in silica shells during the condensation reaction of silane reagents. Lee and co-workers reported magnetic fluorescent core–shell nanoparticles (Fig. 13.7a) [28c]. Cobalt ferrite (CoFe2 O4 ) nanoparticles were synthesized by coprecipitation in aqueous media and a dye-doped silica shell was formed around a nanoparticle via the modified St¨ober method. The surface of the silica shells was further functionalized with PEG and conjugated with an antibody to target cancer cells. Resulting core–shell nanoparticles exhibited T2 contrast enhancing properties in the phantom test and the confocal images of the cancer cells treated with the nanoparticles showed strong fluorescence. Huang and co-workers also synthesized core–shell nanoparticles composed of iron oxide nanoparticles and FITC-doped silica shells [28d]. They synthesized the iron oxide nanoparticles by the thermal decomposition method and stabilized with surfactants such as oleic acid and oleylamine. Because resulting iron oxide nanoparticles were hydrophobic, they were coated with dye-doped silica shells via a reverse microemulsion system. The authors demonstrated human stem cell tracking using these multi-imaging nanoparticles. The probes were easily internalized into human mesenchymal stem cells (hMSCs), and they were injected subcutaneously into the dorsal flank of a mouse and could be visualized in a clinical 1.5-T MRI imager. The Hyeon group reported the fabrication of magnetic fluorescent mesoporous silica nanoparticles composed of mesoporous silica matrix embedded with iron oxide nanoparticles and quantum dots [29a]. Hydrophobic magnetic and fluorescent nanoparticles were transferred into the aqueous phase using cetyl trimethylammonium bromide (CTAB) as a secondary surfactant. Because CTAB also acted as a structure directing agent of the mesostructures, the subsequent silica sol-gel reaction resulted in the formation of mesoporous silica spheres embedded with magnetic and fluorescent nanoparticles. Their high surface area and large pore volumes allowed for the loading of chemical drugs and controlled release. There are several other reports on composite materials of magnetic nanoparticles and mesoporous silica materials [29b–f]. However, they could not be used for in vivo MR imaging because of their large overall size of >100 nm. Very recently, the Hyeon group also synthesized discrete and monodisperse core–shell mesoporous silica NPs smaller than 100 nm by using single Fe3 O4 nanoparticles as cores [29g]. The uniform 3-nm mesoporous

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FIGURE 13.7 Silica-encapsulated iron oxide nanoparticles. (a) TEM images and schematic illustration of CoFe2 O4 –silica (core–shell) functionalized with organic dyes and antibodies [28c]. (b) The synthesis of magnetite nanoparticle/mesoporous silica core–shell and their in vivo multimodal imaging (MRI and optical imaging) [29g]. (Reproduced with permission of Wiley-VCH Verlag GmbH & Co. KGaA.)

shell can contain fluorescent dye and anticancer drug and the resulting composite magnetic silica nanoparticles were applied to simultaneous in vivo magnetic resonance (MR) and fluorescence imaging, and drug delivery vehicle (Fig. 13.7b). Various natural and synthetic polymers have been used extensively in the preparation of biocompatible nanoparticles. The immobilization of polymers can reduce the safety and toxicology concerns of nanoparticles for the clinical applications. Because the silica stabilized nanoparticles often experience precipitation and gel formation, additional biocompatible polymers (e.g., PEG) have been immobilized on the surface of the silica shell to improve the colloidal stability [28, 29]. Furthermore these polymer–nanoparticle hybrids can be utilized for multifunctional biomedical applications with simultaneous drug delivery and imaging capability [29]. Polyesters, such as poly(d,l-lactide-co-glycolide) (PLGA), poly(d,l-lactide) (PLA), and poly(glycolide) (PGA), have been most popularly used for these applications [30]. Gao and co-workers employed amphiphilic block copolymers of methoxy- and maleimideterminated poly(ethylene glycol)-block-poly(d,l-lactide) (PEG-PLA) to fabricate polymer

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micelles [30b]. Hydrophobic iron oxide nanoparticles and hydrophobic drug (doxorubicin, DOX) were spontaneously incorporated into the hydrophobic PLA core part and hydrophilic PEG was exposed to the aqueous environment. The resulting particles were relatively large (>50 nm) because several iron oxide nanoparticles were clustered in the hydrophobic core part. They were functionalized with RGD peptide, and they targeted cancer cells in in vitro T2 MRI and showed therapeutic effects via the release of loaded doxorubicin. Kim et al. [30c] fabricated multifunctional PLGA nanoparticles with particle size of 100–200 nm by simultaneously immobilizing magnetite nanoparticles, quantum dots, and anticancer drug (doxorubicin) in PLGA matrix via a conventional oil-in-water emulsion–evaporation process [30c]. Using the multifunctional polymer nanoparticles, they demonstrated simultaneous cancer-targeted MR imaging and optical imaging, as well as drug delivery. In addition, the loaded magnetite nanoparticles facilitated the magnetic guiding of the polymer particles, thereby increasing the synergetic targeting efficiency. Jiang and co-workers fabricated hollow Fe3 O4 –polymer hybrid nanospheres by the addition of Fe3 O4 nanoparticles to an aqueous solution of polymer–monomer pairs composed of the cationic chitosan polymer and the anionic acrylic acid monomer, followed by polymerization of acrylic acid and selective crosslinking of chitosan at the end of polymerization [30d,e]. The phantom test of magnetic resonance imaging showed that the synthesized hybrid hollow nanospheres had a significant magnetic resonance signal enhancement in T2-weighted image [30e]. PEG, a representative biocompatible polymer, has received great attention due to its nonfouling property, which supports a resistance to protein adsorption and an ability to bypass the RES and natural barriers such as the nasal mucosa [31]. PEGs have been used extensively as stabilizing materials for many nanoparticles in biomedical applications, in particular, in long circulating in vivo imaging systems. Because PEG itself is very inert, surface-anchorable materials such as copolymers, phospholipids, and silica are combined with PEGs to encapsulate the nanoparticles. Dubertret et al. [32a] reported that PEGphospholipid block copolymers could form a stable micelle structure on quantum dots via the hydrophobic interaction between hydrophobic tail groups of the surfactants and phospholipid parts. The outer surface of the nanoparticles is comprised of a dense PEG layer, which is stable in biological media. This process is highly reliable and can generally be applicable to many other kinds of nanoparticles. Using a very similar strategy, various water-dispersible metal oxide nanoparticles, including iron oxide nanoparticles, were generated [32b]. Because these PEG-phospholipid block copolymers are very expensive, they cannot be applied for the large-scale preparation. Amphiphilic di- and triblock copolymers have also been used as stabilizing shell materials for water-dispersible nanoparticles. Their hydrophobic blocks can strongly interact with the hydrophobic surface of the nanoparticles, whereas the outer hydrophilic blocks can make the nanoparticles dispersible in water [33]. These block copolymers are relatively inexpensive and can be derivatized with other functional groups for additional functionalization. One disadvantage of using these block copolymers is that their large shell thickness is derived from the high molecular weight, which limits their many potential applications. Oligomeric and dendritic molecules were also used as the shell materials for water-dispersible nanoparticles because they form thinner shells while preserving their stability in aqueous media. For example, Yang and co-workers demonstrated that the cyclic oligosaccharides that have hydrophobic cavities and hydrophilic rims can transfer the nanoparticles from an organic to an aqueous phase [34]. Weller and co-workers showed that the relaxivities of magnetic nanoparticles were dependent not only on the size of the core nanoparticles but also on the types of shells [35].

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(b)

(a)

900 40 nm

+ encapsulation

15 nm

relaxivity [mM-1s-1]

ge

an

ch

ex

r2 exchange r2* micelles r2 micelles r2* exchange r2* encapsulated r2 encapsulated r2* Resovist r2 Resovist

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O mi + ce lle s

0 2 250 nm

4

6

8

10

12

14

16

18

20

d [nm]

FIGURE 13.8 (a) Three kinds of methods to prepare water-dispersed manganese ferrite nanoparticles. (b) The r2 relaxities of water-dispersed manganese ferrite nanoparticles with three kinds of shell systems [35]. (Reproduced with permission of the American Chemical Society.)

The MnFe2 O4 nanoparticles synthesized in high boiling ether solvent were transferred to water using three different approaches, including ligand exchange to form a watersoluble polymer shell, embedding into an amphiphilic polymer shell, and encapsulating in large micelles formed by lipids. In the first two polymer-based systems, nanoparticles are individually dispersed to form homogeneous dispersion with a hydrodynamic radius of 30–40 nm, whereas aggregated nanoparticles were randomly distributed inside the micelles with a hydrodynamic radius of 250 nm. Interestingly, the relaxivity, r2*, is much higher for the micellar system than for the polymer-stabilized particles using the same-sized nanoparticles (Fig. 13.8). Nanosized and biocompatible iron oxide nanoparticles have been applied extensively to the diagnosis of cancers. They can be accumulated spontaneously in tumor sites via the enhanced permeability and retention (EPR) effect, which is the enhanced accumulation of macromolecular species including nanoparticles in tumor tissues that have abnormal blood vessels [36]. Consequently, iron oxide nanoparticles were successfully used to image tumors without any targeting probes, which is called passive targeting [37]. For more efficient targeted imaging, the surface of the iron oxide nanoparticles needs to be conjugated with active targeting probes such as antibodies and proteins. Magnetic nanoparticles have been conjugated with various bioactive materials such as antibodies, oligonucleotides, peptides, and proteins. Cheon and co-workers prepared Herceptin-conjugated iron oxide nanoparticles, and they were delivered selectively and imaged tumor cells by interactions with the human epidermal growth factor receptor (Her2/neu), which is usually overexpressed in breast cancers [22b]. The Cheon group also demonstrated that Herceptin-conjugated manganese ferrite (MnMEIO) nanoparticles with high magnetization showed more sensitive in vivo cancer targeted imaging with large r2 relaxivity (Figure 13.9) [24]. This result demonstrated that advanced MRI contrast agents consisting of high magnetic moment nanoparticles and appropriate targeting agents could enable the ultrasensitive detection of various types of cancer in T2/T2*-weighted MRI. As another demonstration of targeted MRI, Gao and co-workers conjugated a cancer-targeting antibody, anticarcinoembryonic antigen (CEA) monoclonal antibody rch 24 onto uniform PEG-coated

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FIGURE 13.9 In vivo MR detection of cancer after administration of magnetic nanoparticles– Herceptin conjugates. Manganese ferrite nanoparticles (MnMEIO) (a–c) show higher signal enhancement than crosslinked iron oxide (CLIO) (d–f) [24]. (Reproduced with permission of Nature Publishing Group.)

iron oxide nanoparticles and successfully performed targeted MR imaging for human colon carcinoma tumors [38]. Zhang and co-workers demonstrated that PEG-coated iron oxide nanoparticles conjugated with targeting peptide (chlorotoxin) were preferentially accumulated within gliomas and exhibited highly contrast-enhanced MR imaging [39].

T2 Nanoparticulate Contrast Agents of Unique Structures Recently, new multifunctional nanomedical platforms have been fabricated by combining various nanostructured materials with different functions, making it possible to accomplish multimodal imaging and simultaneous diagnosis and therapy [40]. Heterodimers of magnetic nanoparticles and other nanoparticles can serve as multimodal imaging agents such as MRI contrast agents and optical probes [41]. Dumbbell-like nanoparticles composed of magnetic nanoparticles and gold nanoparticles have been used for dual MR and optical imaging. The Sun group reported the dual modal imaging properties of Fe3 O4 –Au dumbbell nanoparticles, which were prepared by the growth of Fe3 O4 on as-prepared Au nanoparticles in the presence of oleic acid and oleylamine (Fig. 13.10) [41a]. Fe3 O4 and Au components had different surface properties, and they were modified by dopamine and thiol groups, respectively. The epidermal growth factor receptor antibody (EGFRA) was conjugated on the surface of Fe3 O4 , and the resulting heterostructured nanoparticles were successfully applied to the cancer-targeted MR imaging. Furthermore, the unique surface plasmon property of the Au component enabled reflection imaging. The Cheon group demonstrated dual modal imaging using heterodimeric FePt–Au nanoparticles, which were synthesized by the catalytic growth of Au on the surface of the FePt nanoparticles [41b]. Antibody-conjugated FePt–Au nanoparticles were used as both T2 MRI contrast agent and biosensor, and neutravidin-conjugated nanoparticles acted as detecting probes on the biotin patterned

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FIGURE 13.10 (a) TEM image of Au–Fe3 O4 nanoparticles. (b) T2-weighted MR image of A431 cells labeled with Au–Fe3 O4 nanoparticles. (c) Reflection image of A431 cells labeled with Au–Fe3 O4 nanoparticles [41a]. (Reproduced with permission of Wiley-VCH Verlag GmbH & Co. KGaA.)

biochip. Ying and co-workers reported Fe3 O4 –CdSe heterodimer nanoparticles by growing CdSe on the surface of the as-synthesized Fe3 O4 nanoparticles [41c]. Although there are no MRI demonstrations in their paper, it might be a candidate as an MRI contrast agent. The Hyeon group also reported that heterostructured nanoparticles composed of various combinations of a metal (Au, Ag, Pt, or Ni) and an oxide (Fe3 O4 or MnO) were readily synthesized from thermal decomposition of mixtures of metal–oleate complexes and metal–oleylamine complexes [41d]. Nanoparticles of Au–MnO and Au–Fe3 O4 have the potential to serve in many multifunctional biomedical applications, such as multimodal imaging or detection probes. The Hyeon group fabricated biocompatible hollow iron oxide nanocapsules and demonstrated their in vitro T2 MRI and drug delivery capabilities to cancer cells. They were fabricated from akagenite (␤-FeOOH) nanorods via the wrap–bake–peel process, which involves silica coating, heat treatment, and finally the removal of the silica layer. During heat treatment at 500 ◦ C, first in air and then in 10% hydrogen atmosphere the ␥ -FeOOH nanorods were transformed into magnetite (Fe3 O4 ) capsules. Removing the silica shells resulted in the formation of water-dispersible hollow Fe3 O4 nanocapsules. Large hollow pores could be loaded with chemical drugs that were released in cancer cells, and the magnetite shells were used as the T2 MRI contrast agent (Fig. 13.11) [42].

13.3.3 T1 Nanoparticulate MRI Contrast Agents Over the last 20 years, most nanoparticulate contrast agents have been T2 contrast agents using iron oxide nanoparticles. However, these magnetic nanoparticle-based T2 contrast agents have several disadvantages that limit their extensive clinical applications. First, they are negative contrast agents, which give a signal decreasing effect. The resulting dark signal could be confused with other pathogenic conditions and makes images of lower contrast than T1 contrasted images. Moreover, the high susceptibility of the T2 contrast agents induces distortion of the magnetic field on neighboring normal tissues. This distortion of the background is called the susceptibility artifact or “blooming effect,” which generates obscure images and demolishes the background around the lesions [4]. Because of the

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(a) β-FeOOH

β-FeOOH/sillica

Haematite/sillica Heat treatment

Sillica coating

Haematite

Removal of sillica using NaOH

100 nm

Reduction

Removal of sillica using NaOH

Magnetite

Magnetite/sillica

(b)

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120 PEG-MNC-D0X Fe-DOX

0

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viability (x)

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80 60 40 20

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62

27

11

0

0.001

0.010

0.100

1.000

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DOX (µM)

FIGURE 13.11 (a) Schematic illustration of the procedure for the synthesis of uniform and waterdispersible iron oxide nanocapsules and their TEM images. (b) T2-weighted MR images of the magnetite nanocapsules. (c) In vitro cytotoxicity of free DOX and DOX loading nanocapsules against SKBR-3 cells [42]. (Reproduced with permission of the Nature Publishing Group.)

limitations of these T2 agents, most extensively and clinically used MRI contrast agents are based on gadolinium complex-based T1 agents. T1 relaxation is the process of equilibration of the net magnetization (Mz ) after the introduction of an rf pulse. This change of Mz is a consequence of energy transfer between the proton spin system and the nearby matrix of molecules. All biological systems are composed of various molecules and organisms, and they have different relaxation behaviors and different T 1 relaxation times. The presence of paramagnetic ions near the tissue enhances its relaxation and shortens the T1 relaxation time. In particular, transition and lanthanide metal ions with a large number of unpaired electrons, such as Gd3+ , Mn2+ , and Fe3+ , show very effective relaxation [2]. T1 contrast agents enhance T1 relaxation, which makes a signal enhancement on images. Compared to T2 contrast agents, the major advantage of T1 contrast agents is positive contrast imaging by signal enhancement, which can maximize the forte of MRI (i.e., anatomic imaging with high spatial resolution). Furthermore, their bright signal can be distinguished clearly from other pathogenic or biological conditions. As T1 contrasting agents are basically paramagnetic, they do not disrupt the magnetic homogeneity over the large dimension, which can disturb other anatomic backgrounds. Since Gd3+ has seven unpaired electrons with a large magnetic moment, most T1 contrast agents are Gd3+ -based agents. However, due to the toxicity of heavy metal ions, the conventional contrast agents are in the form of ionic complexes with chelating ligands,

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which are thermodynamically and kinetically stable and less toxic. There is, however, no biochemistry based on gadolinium(III) ion in natural human system. In spite of their fewer unpaired electrons and lower magnetic moments, manganese(II), iron(III), and copper(II) ions could be alternative candidates. Manganese(II) ion, in particular, plays various important roles in many biological processes such as cofactors of enzymes and a release controller of neurotransmitters. Although there are some manganese(II) complex contrast agents, Mn2+ itself, in the form of MnCl2 solution, has been most frequently used. It has shown very prominent contrasting effects that can reveal detailed physiological and biological information, and it constitutes a new imaging category known as manganese-enhanced MRI (MEMRI). In particular, MEMRI can visualize the anatomic structure of the brain and its neuronal activity [43], which cannot be obtained with any of the gadolinium(II)-based contrast agents. Unfortunately, however, MEMRI can only be applied in animal studies because Mn2+ ions cause hepatic failure and have cardiac toxicity. As shown above, the present T1 contrast agents are based on paramagnetic ions and are used in the form of ion complexes. They have short life spans in the body and work in a nonspecific manner. Most T1 contrast agents reside within the extracellular space and usually interact with the blood so that they have some limitations as molecular probes for longer time tracking. As shown in recent studies of iron oxide nanoparticles, nanoparticulate agents are very promising for molecular and cellular imaging, which aims to visualize the disease-specific biomarkers at the molecular and cellular levels, respectively. However, the negative contrasting effect and magnetic susceptibility artifacts of iron oxide nanoparticles can be significant drawbacks when using iron oxide nanoparticles. This is because the resulting dark signal can mislead the clinical diagnosis in T2-weighted MRI as the signal is often confused with the signals from bleeding, calcification, or metal deposits, and the susceptibility artifacts distort the background image [4]. Recently, intensive research has been devoted to developing new T1 contrast agents that overcome the above-mentioned drawbacks of Gd3+ ion- and Mn2+ ion-based T1 contrast agents and SPIO-based T2 contrast agents. Briefly, these new classes of contrast agents should satisfy the following characteristics: (1) positive (T1) contrast ability, (2) easy intracellular uptake and accumulation for imaging cellular distribution and functions, (3) a nanoparticulate form for easy surface modification and efficient labeling with biological and bioactive materials, and (4) favorable pharmacokinetics and dynamics for easy delivery and efficient distribution to the biomarkers with minimal side effects. The first class of particulate T1 contrast agents is based on nanostructured frames that have many anchoring sites for paramagnetic ions [44]. Those particles can carry a large number of paramagnetic payloads and produce strong T1 contrast. Various platforms, such as silicas, dendrimers, perfluorocarbons, emulsions, and nanotubes, have been used. Lanza and co-workers used perfluorocarbon nanoparticle incorporated paramagnetic Gd–DTPA complexes for the sensitive detection of fibrin and for the molecular detection of angiogenesis [44a,b]. Lin and co-workers developed silica nanoparticles [44c], metal–organic framework [44d], and mesoporous silica (Fig. 13.12) [44e] to contain a large number of Gd3+ ions and to develop multifunctional properties. Carbon nanotubes could act as the framework that holds Gd3+ ions, either on the surface [44f] or in the structural defect sites [44g]. In these materials, many metal ions are concentrated in a defined volume and their biological behavior and relaxivities are different from those of the complex agents. Basically, this type of contrast agent is an extension of the paramagnetic complex agents so that the maximum number of ions is limited by the density of anchoring groups on the surface. Furthermore, the synthetic procedures are generally very complicated and expensive. The most significant limitation is their large overall size of >100 nm, which is much larger than

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FIGURE 13.12 (a) Schematic illustration of Gd–hybrid mesoporous silica nanospheres. (b) Precontrast (left) and postcontrast (right) T1-weighted mouse MR images [44e]. (Reproduced with permission of the American Chemical Society.)

the inorganic nanoparticles, such as USPIO. To prevent easy excretion by the RES and to compete with the small-sized T2 contrast agents, their overall size should be <100 nm, which is satisfied by inorganic nanoparticles. Very recently, paramagnetic inorganic nanoparticles were developed as new T1 contrast agents. To be effective T1 contrast agents, the ratio between the transverse and longitudinal relaxivities (expressed as r2/r1), which is a defining parameter indicating whether the contrast agent can be employed as a positive or negative agent, has to be low [21]. That is, nanoparticles have a large paramagnetic property (large r1) with negligible magnetic anisotropy (small r2). Nanoparticles of the many compounds of transition metals and lanthanide metals could be good candidates for T1 MRI contrast agents because the surface of the nanoparticles contains a large amount of metal ions with high magnetic moments. The most obvious nanoparticles for T1 agents are gadolinium-based nanoparticles because many gadolinium complexes have been used extensively as clinical MRI contrast agents. Nanoparticles of gadolinium oxide (Gd2 O3 ) [45], gadolinium fluoride (GdF3 ) [46], and gadolinium phosphate (GdPO4 ) [47] have been investigated as MRI contrast agents. They exhibited signal-enhancing contrast in T1-weighted images with low r2/r1 values. The Gd2 O3 nanoparticle-based agents are generally composed of small core nanoparticles of <5 nm and stabilizing shells of dextran [45a], PEG [45b], and silica [45c]. Bridot et al. [45c] reported the use of biocompatible gadolinium oxide nanoparticles with polysiloxane shells containing a fluorescent dye (GadoSiPEG2C) for in vivo dual imaging of magnetic resonance and fluorescence imaging. Water-dispersible GdF3 (or GdF3 /LaF3 ) nanoparticles were prepared with either positively charged surface by conjugation with 2-aminoethyl phosphate groups (GdF3 /LaF3 :AEP) or negatively charged surface by coating with citrate groups (GdF3 :cit) [46]. Dextran-coated GdPO4 nanoparticles (PGP/dextranK01) were synthesized by a hydrothermal process in the presence of dextran [47]. These Gd-based nanoparticles had significantly low r2/r1 values and they showed strong positive contrast effects. However, uniformly sized nanoparticles of gadolinium or related lanthanide compounds have not yet been demonstrated. Very recently, MnO nanoparticles were developed as a new T1 MRI contrast agent for body organs such as the brain, liver, and kidney [48]. In particular, clear T1-weighted MR images of the brain structures, depicting fine anatomic structures, were obtained (Fig. 13.13). This clear anatomic imaging of the various brain structures can be utilized

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FIGURE 13.13 (a,b) Typical coronal (top), axial (middle), and sagittal (bottom) views of a T1weighted MONEMRI before (a) and after (b) the administration of the MnO nanoparticles. (c) Breast cancer cells were selectively enhanced in T1-weighted MRI by the Herceptin-functionalized MnO nanoparticles [48]. (Reproduced with permission of Wiley-VCH Verlag GmbH & Co. KGaA.)

for many applications, such as basic neuroscience research and the diagnosis of clinical neurological diseases (e.g., neurodegenerative diseases). Furthermore, functionalized MnO nanoparticles, by conjugation with a tumor-specific antibody, were used for the selective imaging of the breast cancer cells in a metastatic brain tumor (Fig. 13.13c). Cancer cells were selectively enhanced in T1-weighted MRI because the functionalized MnO nanoparticles with tumor specific-antibody were delivered to and accumulated in the cancer cells with a clear marginal detectability without destroying the anatomic background. This clear marginal detectability while preserving the anatomic background images is a marked superiority of using a T1 contrast agent over a T2 contrast agent.

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Gilad et al. [49] reported the labeling and tracking of rat glioma cells with either MnO nanoparticles or SPIO (Figure 13.14). Cells labeled with MnO nanoparticles and SPIO were each transplanted in contralateral brain hemispheres of a rat and tracked by in vivo MRI. SPIO-labeled cells showed a strong negative contrast in T2/T2*-weighted MRI. However, it is difficult to distinguish them from blood/hemosiderin-associated hypointense regions. MnO-labeled transplanted cells had a higher R1 than the brain tissue surrounding the transplanted cells. Therefore MnO-labeled cells could be successfully detected and distinguished with positive contrast for in vivo T1-weighted MRI. The authors showed a MR “double labeling” technique using opposite contrasts simultaneously on two sites, where one is labeled with MnO and the other with SPIO nanoparticles. This labeling method can be applied to the tracking of cell populations that are injected in different locations or at different time points by complementing and amplifying two kinds of contrasting effect. Very recently, hollow structured manganese oxide nanoparticles were reported. The Hyeon group synthesized various hollow oxide nanoparticles via nanoscale acid etching using MnO as the starting material and alkylphosphonic acid impurity as the etchant [50a]. These hollow nanoparticles were paramagnetic at room temperature, and when dispersed in water they showed spin relaxation enhancement effect. Hollow manganese oxide nanoparticles were synthesized by selective removal of the core of nanoparticles in an

FIGURE 13.14 In vivo MR images of labeled rat glioma cells 24 h after transplantation in the striata of rat brain. (a–c): Spin echo images. (d–f): R1 maps. (g–i): R2 maps. (j–l): R1/R2 merged maps. MnO- and SPIO-labeled cells (a,b), and one with SPIO-labeled cells and unlabeled cells (c; control) [49]. (Reproduced with permission of Wiley-Liss, Inc.)

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acidic buffer and these hollow manganese oxide nanoparticles were applied simultaneously as MRI contrast agent and drug delivery vehicle [50a].

13.4 CONCLUSION In this chapter, we report the progress for MRI contrast agents based on inorganic nanoparticles. Over the last 20 years, various inorganic nanoparticles have been investigated as MRI contrast agents through the collaborative efforts of scientists working in material science, biology, and medicine. In particular, magnetic iron oxide nanoparticles have been used extensively as T2 MRI contrast agents due to their strong ability to shorten T2* relaxation times. Dextran-stabilized SPIOs have been used clinically for the diagnosis of liver diseases because SPIOs are selectively taken up by the Kupffer cells. Smaller iron oxide nanoparticles, USPIOs, have been applied for lymph node imaging and are under consideration for clinical uses. More recently, uniform nanoparticles with high crystallinity have been synthesized from the thermal decomposition of metal precursors at high temperature in the presence of surfactants in organic media. After proper surface modification to impart biocompatibility, these uniform ferrite nanoparticles have been employed successfully as new T2 MRI contrast agents with improved relaxation effects and multifunctional properties. Furthermore, the functionalization with bioactive materials enabled these nanoparticlebased contrast agents to be applicable to targeted imaging via the site-specific accumulation of nanoparticles at the targets of interest. Recently, extensive research has been conducted to develop nanoparticle-based T1 contrast agents to overcome the drawbacks of iron oxide nanoparticle-based negative T2 contrast agents. Gadolinium complexes immobilized in various nanostructured materials, including nanoporous silicas, dendrimers, perfluorocarbon nanoparticles, and nanotubes, have been investigated as new T1 MRI contrast agents due to their ability to carry a large number of paramagnetic payloads and produce strong T1 contrast. More recently, inorganic nanoparticles such as Gd2 O3 , GdF3 , and GdPO4 have been investigated as signal-generating cores of T1 relaxation. Uniformly sized MnO nanoparticles were developed as a new T1 MRI contrast agent for body organs such as the brain, liver, and kidney. In particular, clear T1-weighted MR images of the brain structures, depicting fine anatomic structures, were obtained. The development of multifunctional platforms for either multimodal imaging or simultaneous imaging and therapy has been intensively pursued through the combinations of various nanomaterials. Most of these newly developed nanoparticle-based MRI contrast agents are still in the stage of in vitro testing or preliminary animal studies. There are many issues to be clearly addressed prior to their extensive use in clinical diagnosis, including toxicological effects, long-term stability, and pharmacokinetics. Intensive interdisciplinary collaborative research is essential to achieve the ultimate goal of using nanoparticle-based MRI contrast agents for molecular and active imaging that can be applied to personalized diagnosis.

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CHAPTER 14

Cellular Magnetic Labeling with Iron Oxide Nanoparticles ´ SEBASTIEN BOUTRY, SOPHIE LAURENT, LUCE VANDER ELST, and ROBERT N. MULLER Department of General, Organic and Biomedical Chemistry, NMR and Molecular Imaging Laboratory, University of Mons, Mons, Belgium

14.1 INTRODUCTION For molecular magnetic resonance imaging (MRI), an amplification of the molecular labeling is required. This can be achieved, for example, by avidin–biotin systems in the extracellular compartment [1]. Cellular internalization of the contrast agent is a way to increase the labeling of targeted cells. The challenge for cellular MRI is obtaining a sufficient uptake of MRI contrast agents in cells, especially nonphagocytic cells, in order to render them distinct from the surrounding tissue. Although this approach gives less specific results than molecular MRI, it offers suitable systems for the noninvasive and repeated in vivo monitoring (identification and tracking) of cells by MRI. Magnetic labeling can be realized in vitro or in vivo through specific (involving a cell surface receptor) or nonspecific cellular internalization pathways with contrast agents considered as positive (complexes of paramagnetic ions such as gadolinium(III) or manganese(II); shortening the T1 of their adjacent hydrogen nuclei as a predominant effect and brightening the regions taking them up on a T1-weighted MR image), and often with contrast agents considered as negative (superparamagnetic particles based on a monocrystalline or polycrystalline iron oxide core, shortening the T2 and T2* much more than the T1 of hydrogen nuclei in their neighborhood and generating a signal darkening on T2 (T2* )-weighted MR images). Many studies have indeed been performed in vitro and in vivo with iron oxide nanoparticles [2, 3]. Another mandatory point about cellular MRI is that the loading of cells with contrast agents does not have to affect normal cell function [3]. This chapter mainly focuses on the use of native (unmodified) or unspecifically cell-targeted iron oxides as cell markers. Some examples of cellular labeling performed with other MRI contrast agents, or obtained via specific internalization pathways, are also given.

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14.2 SPECIFIC MAGNETIC LABELING OF CELLS This type of magnetic labeling occurs via receptor-mediated endocytosis (Fig. 14.1), which is a highly specific and efficient mechanism of internalization. The transferrin receptor (TfR) pathway has been used to label cells in several studies. Transferrin (Tf), an iron chelating protein, interacts with TfR and enters cells by receptor-mediated endocytosis, supplying them with iron from stores in the liver. As iron is an essential element involved in cell proliferation, a high number of TfRs is expressed in proliferating cells [4]. Thus, this mechanism of iron uptake appears as an interesting marker for tumor cells, or an interesting way to label progenitor cells. Furthermore, the high turnover rate of the TfR avoids the phenomenon of receptor saturation [5]. Oligodendrocyte progenitors (CG-4) have been magnetically labeled in culture with monocrystalline iron oxide nanocompound (MION) covalently linked to an antibody (OX-26) specific for TfR. The binding of OX-26 to TfR induced the internalization of the nanoparticles. Labeled cells were tracked by MRI after transplantation into the spinal cord of myelin-deficient rats and were found to synthesize myelin [6]. In another study, the in vitro receptor-mediated endocytosis of Tf covalently linked to ultrasmall superparamagnetic iron oxide (USPIO) (Tf-USPIO) by human epidermoid carcinoma A431 was observed. Tf-USPIO was also injected intravenously (200 ␮mol Fe/kg) into rats bearing implanted mammary carcinoma. Despite the fact that most of the injected dose was captured by the liver and the spleen, MR imaging results showed a signal reduction in the tumor only with Tf-USPIO and not with the parent compound or a nonspecific albumin-bound USPIO, suggesting a transferrin-mediated endocytosis of USPIO by the tumor [5]. In another approach, rat 9L gliosarcoma cells were transfected with a modified gene coding for the human TfR, to constitutively express high levels of the receptor protein. The engineered cells were implanted in mice flank and a strong negative contrast was detected in the resulting tumors after intravenous injection of MION conjugated to human holo-transferrin (Tf-MION) (3 mg Fe/animal). This result suggested an efficient uptake of Tf-MION by tumor cells and provided the possibility to image transgene expression [7]. The approach, consisting of the manipulation of iron metabolism genes to provide cellular iron accumulation and subsequent MRI contrast, was further developed in a study where coexpression of transgenic human TfR and human ferritin H-subunit was induced in a stably transfected neural stem cell line (C17) [8]. Ferritin (FT) is composed of H- and L-subunits. Twenty-four subunits form the apoferritin, which is able to incorporate up to 4500 iron atoms. The H-subunit has a ferroxidase activity that promotes iron oxidation and incorporation [9, 10]. Thus FT-bound iron has a semicrystalline structure and acts as an endogenous cellular negative contrast agent. These studies demonstrated that transfected cells grown in iron-rich medium incorporated significantly more iron than control cells, inducing a contrast effect in MRI that was preserved after transplantation

Extracellular compartment Cell membrane

Intracellular compartment CA Endocytic vesicle formation

CA Specific ligand of the cell surface receptor Cell surface receptor

FIGURE 14.1 Schematic representation of the receptor-mediated endocytosis of a contrast agent (CA).

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of cells into the mouse brain. After uptake through TfR, iron was stored by FT without affecting cell viability [8]. The usefulness of FT as an endogenous contrast agent has been reported also after inducing the overexpression of this protein by C6 rat glioma cells. Indeed, intracellular iron is redistributed in excess FT, which generates negative contrast. As iron chelation induces a transient decrease of intracellular iron concentration, TfR is expressed in compensation to increase the iron uptake and restore iron homeostasis. This newly endocytosed iron will also be sequestered by overexpressed FT, leading to an enhancement of the contrast effect [11]. The internalization pathway of a vitamin has also been used often to shuttle iron oxides in cells. The folate (or vitamin B9 ) receptor has been targeted with folic acid-linked MRI contrast agents in order to induce an uptake of the cell tag by cancer cells overexpressing the receptor [12]. The internalization of superparamagnetic nanoparticles bearing chlorotoxin in the cytoplasm of 9L glioma cells has been reported. This 36 amino acid peptide indeed renders the nanoprobe capable of targeting tumor cells expressing membrane-bound matrix metalloproteinase-2 [13]. Prostate cancer cells could be labeled via receptor-mediated endocytosis with iron oxides conjugated to an antibody specific for the prostate-specific membrane antigen (PSMA) [14]. Intracellular accumulation of iron oxide nanoparticles in cells has been achieved also by targeting hormone receptors. The single chain anti-epidermal growth factor receptor (EGFR) antibody has been bound to iron oxide nanoparticles to serve as an intracellular marker of EGFR-expressing cancer cells [15]. Luteinizing hormone releasing hormone (LHRH) and luteinizing hormone/chorionic gonadotropin (LH/CG)conjugated iron oxide nanoprobes have been used to label breast cancer cells [16].

14.3 NONSPECIFIC MAGNETIC LABELING OF CELLS This type of magnetic labeling can occur via natural mechanisms of endocytosis, which are not receptor dependent. In this context of nonspecific magnetic labeling of cells, most of the studies concern the in vitro labeling of nonphagocytic cells potentially useful for cell therapy. Among these cells of interest we find mesenchymal cells, embryonic stem cells, neural stem cells, oligodendrocyte progenitor cells, endothelial progenitor cells, and muscle progenitor cells. In other words, undifferentiated or less differentiated cells that, after their reintroduction in a living organism, could hopefully reach the damaged area (ischemic lesion) of an organ and differentiate to replace dead cells or, more specifically for endothelial precursors, reach a tumor were angiogenesis is occurring. Many in vitro cellular magnetic labeling studies are performed also on CD34+ hematopoietic stem cells or immune cells (lymphocytes) because the study of their in vivo trafficking could help us to better understand the immune response and, more particularly, to improve cell-based cancer therapies [2, 3, 17]. Thus visualization of the in vivo behavior of these cells is crucial and depends on their prelabeling in culture. Furthermore, the study of contrast agent accumulation in phagocytes and tumor cells can become a tool for the biological characterization of pathologies [18, 19]. 14.3.1 Fluid-Phase Endocytosis of Contrast Agents With or Without the Help of a Transfection Agent: Influence of the Contrast Agent Surface Charge Gadolinium-based contrast agents have been developed for the in vitro labeling of cells. Gd chelates coupled to fluorescent rhodamine dextran (GRID), or the combined use of Gd

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COO-

OOC N

N Eu3+

N

N

CH3

-

OOC OH

FIGURE 14.2 Structure of Eu-HP-DO3A. R and europium (Eu, fluorescent) complexed by the chelate HP-DO3A (ProHance
, Bracco Diagnostics, Milan, Italy) (Fig. 14.2), constitute bimodal agents. Indeed, the detection of reimplanted labeled cells can be achieved by in vivo MRI and corroborated by fluorescent histology [3, 20, 21]. Lanthanide ions such as Eu(III) or terbium(III) (Tb(III)) have been used to build complexes called paramagnetic chemical exchange saturation transfer (PARACEST) agents. When the mobile (exchangeable) protons of these complexes are irradiated at their absorption (resonance) frequency, they transfer saturated magnetization to the surrounding bulk water, acting so as negative contrast agents. Thus contrast alteration on MR images is based on a phenomenon called chemical exchange dependent saturation transfer (CEST). As the resonance frequency of the exchangeable protons of the complex is influenced differently according to the type of paramagnetic ion used, PARACEST agents can generate their own specific contrast when irradiated at the specific absorption frequency of their mobile protons. In other words, cells labeled in vitro with different PARACEST agents, by internalization of the complex through pinocytosis, could be distinguished and separately tracked in vivo (Fig. 14.3a) [22]. Several attempts have been made to improve the in vitro cellular uptake of Gd chelates. Gd-DTPA has been combined with transfection agents such as calcium phosphate or the lipofection agent LipofectamineTM , usually facilitating the delivery of functional DNA into the cell [23]. Unmodified dextran-coated iron oxide nanoparticles can be used to label cells in vitro. Lymphocytes or T cells in culture have been loaded with this type of contrast agent through fluid-phase endocytosis [24, 25]. Rat bone marrow mesenchymal cells and mouse

Cell penetrating peptide Cell membrane

Extracellular compartment d

a

CA

Intracellular compartment

CA c Endocytic vesicle formation

CA b

CA

Cell transfection reagent

FIGURE 14.3 Schematic representation of the uptake of contrast agents (CAs) by cells through nonspecific (receptor-independent) endocytic processes: (a) fluid-phase endocytosis (pinocytosis) of the native CA, (b) endocytosis after adsorption of the CA mixed with a cell transfection reagent to the cell surface, (c) endocytosis after adsorption of the native CA to the cell surface, and (d) endocytosis after adsorption of a “cell penetrating peptide”-conjugated CA conjugated to the cell surface.

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embryonic stem cells were simply incubated with superparamagnetic iron oxide (SPIO) R ) at a relatively high concentration (around 100 ␮g Fe/mL in the cell culture (Endorem
medium for 48–72 h of incubation) [26–29]. In both cases, the in vivo fate of labeled cells was successfully MR imaged in vivo. However, in many studies, the cellular uptake R ) is optimized by mixing the superparamagnetic contrast agent with a of SPIO (Feridex
cationic transfection agent such as poly-l-lysine. It is hypothesized that these dextran-coated SPIO nanoparticles can be coated with the positively charged transfection agent through R R (or Endorem
) electrostatic interactions. The citrate-containing formulation of Feridex
may allow for the presence of a few carboxyl groups at the particle surface, which was indeed R ) [30, 31]. Thanks to found to be negatively charged (zeta potential of -41 mV for Feridex
the positive charges of poly-l-lysine, the complex rapidly binds to the negatively charged cell membrane and is shuttled into cells through endosomes, producing a cellular magnetic labeling equivalent to the loading obtained by simply adding SPIO to the culture medium of cells (10–20 pg Fe/cell). However, the iron incubation concentration is lower (25 ␮g Fe/mL (vs. ∼100 ␮g Fe/mL)) and the magnetic labeling duration can potentially be shorter (2–48 h, depending on the cell type (vs. 48–72 h)). In the case of human mesenchymal stem cells, an iron content of 16 pg/cell has been found after a 2-h incubation [3, 32, 33]. The same approach has been developed with other commercial transfection agents, which have been combined with iron oxide nanoparticles in order to form complexes possessing a net positive charge that can associate with the negatively charged cell surface [34]. The resulting increase of contrast agent uptake has allowed for an improvement of the magnetic labeling of cultured cells. Another polycationic amine, protamine sulfate (Food and Drug Administration apR ) proved), has been used instead of poly-l-lysine to form complexes with SPIO (Feridex
+ TM and label human mesenchymal stem cells or CD34 cells [35]. Superfect consists of activated dendrimers with a spherical architecture and branches radiating from a central core and terminated by positively charged amino groups [36]. This commercially available transfection reagent has also been suitable for the magnetic labeling of mammalian R ). An iron content of 30 pg Fe/cell has been obtained for hucells with SPIO (Feridex
R complex man mesenchymal stem cells incubated for 2 h with the SuperfectTM –Feridex

R (25 ␮g Fe/mL), a value twice higher than with poly-l-lysine–Feridex complex under the same conditions [33, 35]. The mechanisms initiating the endocytosis of dendrimers and polyamine-complexed iron oxide nanoparticles are thought to be of an electrostatic origin. Indeed, the adsorption of the positively charged complex on the negatively charged cell surface is supposed to induce membrane bending and membrane disruption. These phenomena induce invaginations of the cell membrane, leading to the encapsulation of the complexes in endosomes (Fig. 14.3b) [3, 32]. A carboxylated dendrimer has also been used as a coating for iron oxide nanoparticles. The carboxylate groups gave an anionic surface to the dendrimers. The resulting magnetodendrimer, called MD-100 (Fig. 14.4), has successfully served as a magnetic marker for different cultured cell types, originating from mouse, rat, and human, including stem cells. The uptake mechanism of MD-100 is also believed to result from electrostatic interactions between the dendrimer and the cell surface. The highly charged polymers bind on multiple sites on the cell membrane, inducing bending and disruption of the plasma membrane. An amount of 9 pg Fe/cell has been reached for CG-4 rat oligodendrocyte progenitors after 48 hours of incubation with MD-100 (25 ␮g Fe/mL) [3, 37, 38]. A binding of nanoparticles bearing negative charges to the cell surface, which is also mainly negatively charged, appears as contradictory, and the reason for such an event is unknown. Nevertheless, a high level of

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γ-Fe2O3 magnetodendrimers

FIGURE 14.4 Structure of magnetodendrimers (adapted from [32]).

cell internalization, comparable to the one obtained with the magnetodendrimer MD-100, has been obtained also with anionic iron oxide nanoparticles. The negative surface charge of these nanoparticles, due to the unbound carboxylate groups of their dimercaptosuccinic acid coating, has been hypothesized to allow them to interact strongly and nonspecifically with cationic sites of the plasma membrane, which are much more scarcely distributed than anionic areas. This mechanism involving an adsorption of the nanoparticles on the cell surface through electrostatic interactions, preceding their internalization, is called adsorptive endocytosis (Fig. 14.3c) [39]. Commercially available anionic nanoparticles (carboxydextran-coated SPIO R R )) have been found to be more efficient than dextran-coated SPIO (Feridex
) to (Resovist
label cultured cells. Furthermore, the effectiveness of the intracellular uptake of nanoparticles has been found to be dependent on the density of carboxyl groups at the nanoparticle surface (with there being an optimum density), as well as on the cell type. The uptake of anionic nanoparticles has indeed been observed to be more pronounced in mesenchymal stem cells than in human cervical carcinoma (HeLa) cells, and a lower density of carboxyl groups was necessary to obtain an effective uptake of nanoparticles by mesenchymal stem R was captured in a higher amount cells compared to the latter cancer cells. Indeed, Feridex
by mesenchymal stem cells than by HeLa cells. It was hypothesized that the citrate used R might allow for the presence of a few carboxyl groups at in the preparation of Feridex
R the surface of nanoparticles, which would be sufficient to trigger an uptake of Feridex
by mesenchymal stem cells, but not by HeLa cells [30]. The magnetic labeling of cultured nonphagocytic adherent cells with iron oxide nanoparticles has been found to be dependent on the relative concentration of the magnetic tag and of the cells in culture, on the nanoparticle hydrodynamic diameter, and on the coating charge (Fig. 14.5) [40]. 14.3.2 Cellular Delivery of Contrast Agents by Cationic Liposomes The same approach as for dendrimers and polyamines has been developed with liposomal transfection reagents [34]. Several types of superparamagnetic iron oxide nanoparticles, R R R , Endorem
, Sinerem
(dextran-coated), or P7228 (covered by an including Feridex
anionic dextran derivative), have been incorporated into cationic transfection liposomes (such as LipofectamineTM or FuGENETM ). The lipid complex, carrying a net positive charge, binds to the negatively charged cell surface (Fig. 14.3b). The delivery of the contrast agent to the cytoplasm then depends on the liposome–cell membrane fusion. It has been demonstrated that this fusion only occurred after uptake of the liposome into the endocytic

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FIGURE 14.5 T2-weighted MR images (TR/TE: 3000/15 ms, 16th echo (240 ms), 5 × 105 cells/mL in 2% gelatin) of 3T6 fibroblasts, after 48 h of SPIO labeling with 50 ␮g Fe/mL of R R or Resovist
(upper row samples). Samples imaged on the bottom row (from left to Endorem

R R at 0.01 mM Fe, and 2% right: Resovist nanoparticles at 0.01 mM Fe, unlabeled cells, Endorem
gelatin) are for comparison (from [40]).

pathway [41]. These cationic liposome–iron oxide nanoparticle complexes have been used to magnetically label several types of mammalian cells including mouse embryonic stem cells, human mesenchymal stem cells, rat oligodendrocyte progenitor CG-4 cells, rabbit skeletal myoblasts, human hematopoietic progenitor cells, mouse lymphocytes, or HeLa cells [29, 33, 42, 43]. The iron content of human mesenchymal stem cells was 7.6 pg/cell R complex. In other after a 2-h incubation with 25 ␮g Fe/mL of LipofectamineTM –Feridex
R complex used terms, this loading was twice lower than with the poly-l-lysine-Feridex
under the same conditions [33]. 14.3.3 Cellular Delivery of Contrast Agents Through a Cell Penetrating Peptide Another nonspecific way of cellular internalization exists, since some natural proteins, such as the human immunodeficiency virus (HIV) Tat (transactivator of transcription) protein, have the ability to penetrate cell membranes directly. The basic domain of the HIV Tat protein, formed of an arginine-rich sequence, has been identified as the domain responsible for the translocation of this protein. The highly cationic peptide derived from this basic domain is called a cell-penetrating peptide because of its ability to cross the plasma membrane and consequently to drag the rest of the protein with it [44, 45]. The mechanism by which this task is performed is based on an interaction of the peptide with glycosaminoglycans attached to cell surface heparan sulfate proteoglycans, initiating an adsorptive endocytic process (Fig. 14.3d) [46, 47]. Gd chelates and iron oxide nanoparticles have been conjugated to the Tat basic domain peptide (or more simply, Tat peptide) in order to efficiently transport the contrast agents into cells. Mammalian cells have been labeled successfully with Tat peptide-derivatized Gd-DOTA,[48, 49] or superparamagnetic contrast agents linked to the Tat peptide [50–52]. Human CD34+ hematopoietic stem cells labeled for 1 h with the Tat peptide-conjugated iron oxide nanoparticles (100 ␮g Fe/mL) contained from 10 up to 30 pg Fe/cell [50].

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14.3.4 Instant Cellular Delivery of Contrast Agents by Electroporation This method, called magnetoelectroporation (MEP), circumvents a prolonged incubation of cells and the use of a transfection agent, and allows achieving an instant magnetic labeling of cells. Mesenchymal stem cells and neural stem cells have been labeled with R SPIO (Feridex
) by the MEP method. The nanoparticles were taken up in endosomes and the iron content of cells (picogram range/cell) was comparable to values obtained with transfection agents. Furthermore, when used under properly calibrated conditions (a pulse of 130 V for 17 ms), MEP is claimed to have no effect on cell viability, proliferation, or differentiation [53]. 14.3.5 Fluid-Phase Endocytosis of Contrast Agents with Micron-Sized Particles Some advantages have been found in the use of micrometer-sized particles of iron oxide (MPIO) to label cells, as compared to the dextran-coated iron oxide nanoparticles SPIO or USPIO. For the latter, the endocytosis of millions of nanometer-sized particles is required to achieve sufficient contrast, and cell division can dilute the label beyond the detection threshold. MPIO (Bangs Laboratories, Fishers, Indiana, USA) are magnetite cores encapsulated with styrene/divinyl benzene. Various sizes exist (from 0.76 to 5.80 ␮m in diameter). The smallest ones (0.76–1.63 nm) also contain a fluorescein-5-isothiocyanate analog (Dragon Green). Human CD34+ hematopoietic progenitor cells and porcine mesenchymal stem cells have been reported to efficiently internalize 0.9-␮m MPIO through endocytosis under incubation conditions (concentration and duration) comparable to those used with dextrancoated nanoparticles (e.g., 18 h with 45.6 ␮g Fe/mL), and to lead to an iron load of hundreds of picograms per cell [54]. The styrene/divinyl benzene coating is thought to be inert inside cells and cannot be degraded within the cell, which is not the case with the dextran coating [55]. As the T2* effect of an iron oxide particle increases with its size, these MPIO have a larger T2* effect than SPIO in MRI, making possible the detection of single particles by in vitro MRI with a 50-␮m resolution in agarose phantoms or in cultured cells [56]. The possibility to MR image labeled cells by detecting single endocytosed particles overcomes the problems related to the dilution of the magnetic label induced by cell divisions [57]. Furthermore, it has been shown that mouse hepatocytes, prelabeled with 1.63-␮m MPIO, were detectable by MRI as single cells in the liver of recipient mice 1 month after their transplantation into the spleen [58].

14.4 MAGNETIC LABELING OF PHAGOCYTES R ) undergo a higher It has been demonstrated that dextran-coated SPIO (Endorem
R ) in vitro, macrophage uptake than the smaller-sized dextran-coated USPIO (Sinerem
confirming their respective in vivo biodistribution; SPIO are rapidly cleared from the circulation as a result of their capture mainly by liver Kuppfer cells, and USPIO have a longer blood half-life, allowing them to reach macrophages situated in less accessible territories, such as lymph nodes [59]. Scavenger receptors (SRs) are relatively nonspecific receptors mainly expressed by macrophages. These receptors bind to a broad range of negatively charged macromolecules (class A) or modified low-density lipoprotein (LDL) (class A and class B). The class A SR

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Extracellular compartment Opsonin-independent receptor

Phagocyte membrane extension

CA a

Cell membrane

CA Intracellular compartment

b

Opsonin-dependent receptor Opsonin

FIGURE 14.6 Schematic representation of the uptake of MRI contrast agents (CAs) by phagocytic cells: (a) endocytosis initiated by a receptor mediating an opsonin-independent recognition of the CA as a phagocytic target, and (b) endocytosis initiated by a receptor mediating a recognition of the CA as a phagocytic target through opsonins adsorbed to its surface.

(SR-A) types I and II bind anionic proteins, lipids, polynucleotides, and polysaccharides including dextran sulfate or fucose sulfate (fucoidan) and bacterial lipopolysaccharides [59]. The SR-A types mediate an opsonin-independent recognition of microorganisms or apoptotic cells before their phagocytosis (Fig. 14.6a) [60]. Competition experiments have suggested that the endocytosis of SPIO by cultured macrophages was mediated by the SR-A types. Indeed, ligands of the SR-A (fucose sulfate and polyinosinic acid) were effective inhibitors of the SPIO capture by cultured mouse peritoneal macrophages [59]. Phagocytosis concerns the uptake by macrophages of large material, such as bacteria (∼1 ␮m) or senescent eukaryotic cells (∼10 ␮m), which are much larger than the hydrodynamic diameter of iron oxide nanoparticles. The larger-sized dextran-coated SPIO (80–150 nm) can nevertheless initiate their macrophage endocytosis by interacting with the SR-A [59]. The macrophage uptake of dextran-coated USPIO (20–40 nm) occurs primarily through a pinocytosis mechanism and leads to a weaker internalization of nanoparticles than with dextran-coated SPIO [59, 61]. However, it has been demonstrated that citrate-coated very small superparamagnetic iron oxide nanoparticles with a hydrodynamic diameter of 8 nm (VSOP-C125) were efficiently taken up by macrophages through a mechanism involving actin, which is the cytoskeleton component determinant for phagocytosis. This result suggests that the pronounced negative surface charge of the nanoparticles could help their macrophage endocytosis by allowing an interaction with a receptor initiating phagocytosis. Another important point in this context is that the clustering of the nanoparticles, leading to the formation of larger-sized particulate aggregates, can also result in different uptake stimuli for macrophages (i.e., phagocytosis rather than pinocytosis) [61]. Particles can also bind to nonspecific scavenger receptors after adsorption to their surface of proteins from the heat-inactivated fetal bovine serum supplement added in culture medium [62]. In fresh plasma, MION have been covered by vitronectin, by fibronectin, and by fragments of the complement fraction C3 (C3b and its fragment C3dg) bound to IgG heavy or light chain [18]. This opsonization induced the recognition of MION by specific receptors on the macrophages (Fig. 14.6b) [63]. Indeed, IgG heavy chain-complexed dimeric forms of C3b and C3dg are ligands for the complement receptors CR1 and CR2, respectively [64]. The vitronectin receptor and the ␤2 integrins, such as the complement receptors CR3 and CR4, usually bind the opsonin iC3b (fragment of C3b). They can also mediate the nonopsonic recognition of phagocytic targets [60, 63]. Iron oxide nanoparticles have been targeted to a ␤2 integrin for the in vitro magnetic labeling of dendritic cells before their reimplantation and in vivo MRI monitoring. By conjugation to an antibody, iron oxide nanoparticles

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have been targeted to CD11c, the alpha subunit of the ␤2 integrin CR4 (CD11c/CD18), and allowed to obtain an efficient intracellular labeling in vitro through receptor-mediated endocytosis [64]. Gd-based contrast agents (Gd-DTPA) have been linked to long fatty acid chains in order to form 0.5–1-␮m insoluble particles. These foreign particles were captured by dendritic cells in vitro through phagocytosis and intracellular lipases cleaved the long aliphatic side chains, restoring the solubility of Gd-DTPA and consequently its contrast agent activity inside the cells [65].

14.5 TOXICITY From the toxicity point of view, iron oxides are generally well tolerated by cells. No significant short-term toxicity of iron oxide nanoparticles has been reported [3, 26, 33, 37, 42]. Several studies have shown that the endosomal incorporation of iron oxide nanoparticles R complexed did not affect cell viability or proliferation, but it was found that Feridex
to poly-l-lysine had an effect on the human mesenchymal stem cells’ differentiation into chondrocytes [66]. Furthermore, the same complex significantly affected the viability of marrow stromal cells as compared to unlabeled control cells when incubations were done for long periods (24 or 72 h) with relatively high concentrations (100 and 250 ␮g Fe/mL). R –poly-l-lysine complexes at 25 ␮g Fe/mL did On the contrary, a labeling with Feridex
not induce any significant mortality of mesenchymal stem cells, even after a 72-h incubation period, and long-term toxicity studies performed after a 16–18-h incubation period of mesenchymal stem cells with the same amount of complex did not show any significant cell death as compared to unlabeled control cells until 15 days postlabeling [67]. The metabolic fate of internalized iron has not been well determined in current studies. It is assumed that iron atoms are integrated in the normal iron turnover inside the cell [3]. Nonetheless, incubation with iron oxide at high concentrations (1 mgFe/mL for 24 h) has been found to cause the generation of free radicals, a decrease in proliferation, and cell death [43]. VSOP were massively taken up by macrophages and were reported to induce a transient oxidative stress R in these cells [68]. Human neural precursor cells that were labeled for 72 h with Sinerem
(800 ␮gFe/mL) did not show any loss of short-term viability. However, a decrease in their subsequent survival and a reduction in their neurosphere-forming capacity were observed when cells were dissociated and plated in culture under standard growth conditions after incubation with the USPIO (more than 80% of cell death at the first subculturing passage postlabeling). This phenomenon was avoided with lower incubation concentrations (e.g., 400 ␮gFe/mL) and shorter incubation times (e.g., 48 h). Poly-l-lysine has been found to R R and Endorem
without impairing cell improve the labeling efficiency of both Sinerem
survival, even surprisingly increasing the short-term survival of the human neural precurR , and 200 and 400 sor cells at high iron incubation doses (100 ␮gFe/mL for Endorem

R
R ␮gFe/mL for Sinerem ) [69]. Higher Sinerem incubation concentrations (up to 22 mg Fe/mL) and a shorter incubation time (12 h) have been used to perform a cationic liposome R )-helped labeling of D3 embryonic stem cells [70]. Toxicity was mainly in(Metafectene
R /mL). duced by the transfection agent (50% of immediate cell death at 25 ␮L Metafectene

R A small influence of Sinerem alone on cell proliferation was only observed at the highest iron incubation concentration. In the same study, comparisons between different cell lines (D3 embryonic stem cells, C17.2 neuronal progenitor cells, and dendritic cells) were also performed: just as the uptake of a given superparamagnetic cell tag is cell line specific, the tolerance to the transfection agent was also observed to be dependent on the cell type [70].

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14.6 APPLICATIONS OF CELLULAR MAGNETIC LABELING WITH UNMODIFIED OR UNSPECIFICALLY CELL-TARGETED MRI CONTRAST AGENTS The main application of cellular magnetic labeling is the in vivo tracking of cells by the noninvasive MRI technique, principally in the context of cell therapies. This method provides the possibility to visualize the migration of cells such as lymphocytes during the immune response. It also allows for the monitoring of the in vivo fate and behavior of cells (e.g., stem or progenitor cells) transfused in a diseased living organism, or directly transplanted in a damaged organ. Cellular imaging of phagocytes infiltrated in inflammatory areas has also been reported, as well as the magnetic labeling of tumors. In many studies, iron oxide nanoparticles have been chosen for cell labeling in vitro or in vivo and have been effective as magnetic tags to follow the cells in vivo by MRI [3]. 14.6.1 MRI Monitoring of Cells Transplanted or Transfused In Vivo After In Vitro Magnetic Labeling The in vivo distribution of magnetically labeled cells from the hematopoietic lineage has been studied by MRI. Hybridomas (a fusion between B lymphocytes and myeloma cells) internalized anionic nanoparticles in vitro and were intraperitoneally injected in the mouse. MRI suggested a homing of the labeled cells in the spleen 24 h postinjection. These cells were also found to proliferate in the spleen [71]. Magnetically labeled human hematopoietic progenitor cells were intravenously injected in athymic mice. MRI monitoring showed signal attenuations in spleen, liver, and bone marrow. Histology confirmed the homing of the labeled cells in these recipient organs [72]. Tat peptide-derivatized iron oxide nanoparticles have also been applied to image cell trafficking in vivo after intravenous injection of CD34+ cells into immunodeficient mice. These cells, prelabeled with superparamagnetic–Tat peptide conjugates, were detected in the bone marrow by MRI [50]. The MRI tracking of prelabeled lymphocytes (T cells) has also been performed in the context of the mouse experimental autoimmune encephalomyelitis (EAE) model, allowing researchers to collect information about the pathogenesis of this dysmyelinating disease [73]. Because of its high relevance for immune cell-based anticancer therapies, the in vivo trafficking of tumor-targeted lymphocytes has been studied in several models. Iron oxide nanoparticles conjugated to a large number of Tat peptides were used to label CD8+ cytotoxic T lymphocytes stimulated with the ovalbumin-derived immunogenic peptide. Ovalbumin-transfected B16 melanoma cells were subcutaneously implanted in mice and intraperitoneally injected with the labeled cells. The tumoral recruitment of these cells was visualized by MRI and confirmed by histological analysis [74]. In a similar type of animal model, ovalbumin-specific splenocytes, labeled with iron oxide nanoparticles, were injected into mice with ovalbumin-expressing tumors. MRI results suggested that cells went first to the spleen (homing, 24 h postinjection) before being recruited by the tumor (48–72 h postinjection) [75]. The in vivo MRI monitoring of the accumulation of immune cells in tumors was also reported in a study performed with human natural killer cells directed against HER2/neu receptors, expressed by mammary tumors. Cells were magnetically labeled with iron oxide nanoparticles and intravenously injected into mice bearing HER2/neu-positive mammary tumors. The tumoral signal decay in MRI, confirmed by histology results, demonstrated the presence of the labeled natural killer cells in the tumors [76]. These different reports have shown the usefulness of MRI as

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a noninvasive method for the monitoring of cell-based immunotherapies. Furthermore, a human approach to therapeutic cell tracking appeared as clinically safe. Dendritic cells, which can play the role of an antitumor cellular vaccine since they are antigen-presenting cells, were SPIO-labeled and detected by MRI after injection in the lymph nodes of patients suffering from melanoma. MRI showed how accurate the nodal delivery was of the iron oxide-loaded autologous dendritic cells [77]. Many attempts of stem cell transplantation have been performed with the aim of repairing central nervous system (brain or spinal cord) disorders, and promoting anatomical and functional recovery. The magnetic prelabeling of these cells allows visualization of their presence and their migration by MRI. In order to provide a strategy of cell therapy to reverse dysmyelinating diseases, stem cell-derived oligodendrocyte progenitors (CG-4) were magnetically labeled and were observed to migrate after transplantation in the spinal cord of 7-day-old myelin deficient rats. Furthermore, histological analysis identified iron-loaded cells as oligodendrocytes, and detected newly formed myelin in areas corresponding to the localization of labeled cells on MR images [6]. Oligodendroglial progenitors derived from neural stem cells were also magnetically labeled in vitro and transplanted intraventricularly in the brain of neonatal rats for a successful in vivo MRI tracking [37]. The possibility to repair spinal cord injuries has been investigated also by studying the migration and behavior of transplanted labeled stem cells. Marrow stromal cells (multipotent progenitor R ) and intravenously cells extracted from bone marrow) were labeled with SPIO (Endorem
injected into rats 1 week after induction of a spinal cord injury. It was observed that the paraplegic rats showed good scores of recovery from the locomotor and the hind limb sensitivity points of view during the 5 weeks following marrow stromal cell injection. Hypointense areas observed in ex vivo MRI suggested a migration of the SPIO-labeled cells to the lesion site, and histology confirmed the presence of iron-loaded cells in the lesion. Furthermore, a reduction of the lesion size was noticed in marrow stromal cell-injected animals, suggesting tissue regeneration due to these cells. As the regeneration of axons can be affected by the formation of scar tissue in the injured region, biocompatible macroporous hydrogels (pore size: 10–50 ␮m) have been used to anatomically replace a removed part of the spinal cord and prevent scarring. A block of hydrogel seeded with SPIO-labeled marrow stromal cells was inserted in the place of the right half of a spinal cord segment and was visible by MRI. Hydrogels create an environment where intrinsic growth factors can diffuse and promote marrow stromal cell migration. Indeed, 6 weeks after implantation, the hydrogel was adherent to the spinal segments and neurofilaments were stained inside by immunohistochemistry, demonstrating the growth of axons into the hydrogel [28]. Another central nervous system injury requiring a replacement of lost cells is cerebral stroke. In the context of brain lesions, the fate (migration and differentiation) of in vivo implanted stem cells, after magnetic labeling in culture, has been monitored by the noninvasive MRI technique. The behavior of marrow stromal cells, embryonic stem cells (ESCs), and CD34+ hematopoietic progenitor cells in the injured brain has been studied using a model of cortical photochemical lesion. On T2-weighted MR images, SPIO-labeled marrow stromal cells were visible as a hypointense signal at the level of their implantation site, in the hemisphere contralateral to the lesion. A signal darkening was noticed within 1 and 3 weeks in the damaged cortical tissue, around the necrotic part of the lesion, suggesting a migration of labeled cells. Immunohistochemistry revealed that a small proportion (less than 3%) of the large population of iron-loaded cells present in the lesion expressed the neuronal marker NeuN. Intravenously injected SPIO-labeled marrow stromal cells infiltrated the damaged hemisphere at the border of the lesion. The presence of labeled cells was visible by MRI

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and confirmed by histology [26, 28]. The same type of experiment was performed with ESCs. However, to avoid the risk of forming embryonic tumors once grafted in vivo, ESCs were differentiated into neural stem cells before implantation. These ESCs were SPIOlabeled and intravenously or intracerebrally (contralateral to the lesion) injected into rats 1 week after the induction of a photochemical cortical lesion. In both cases, MRI revealed a hypointense signal in the lesion 1–2 weeks after injection or implantation. As ESCs were modified to express a fluorescent marker (green fluorescent protein, GFP), fluorescence microscopy performed on histological slices confirmed the presence of labeled cells in this injured site. MRI and histology results also suggested that intracerebrally implanted cells were migrating across the corpus callosum to reach the lesion. Iron-loaded ESCs found in the injured region after intravenous injection were observed to achieve a final differentiation and were identified as mostly astrocytes (70%), but also oligodendrocytes (1%) and neurons (5%) [27, 28]. Under the same experimental conditions, magnetically labeled human CD34+ cells were observed to reach the lesioned target 24 h after contralateral R ) implantation [28]. In other studies, ESCs magnetically labeled with USPIO (Sinerem
combined to lipofection agent have been implanted in the brain of rats suffering from a focal cerebral ischemia. Two weeks after this induced stroke, labeled cells were implanted in the contralateral hemisphere and observed by MRI. A line of hypointense signal was visualized along the corpus callosum 6 days after grafting of the cells, suggesting a migration of the cells toward the lesion. Indeed, 2 days later, the ischemic hemisphere became darker, suggesting that labeled cells had moved from the injection site and accumulated in the damaged area. Once again, other techniques, such as immunocytochemistry, confirmed the presence of implanted cells in the areas of MR signal darkening. Furthermore, changes in the labeled cells’ shape suggested that their differentiation occurred during their migration. Neuron-like cells were observed in the ischemic area [42]. This phenomenon of transplanted stem cells attracted by an ischemic site in the brain has also been observed with neural stem cells labeled using a bimodal contrast agent, the GRID, which allowed for the detection of cells by MRI and fluorescent microscopy [78]. The migration of iron oxide-labeled marrow stromal cells and neural progenitor cells toward another type of brain injury, a gliosarcoma, also has been observed by MRI after the injection of cells in a tail vein or in a subarachnoid space: the cisterna magna [79]. Stem cell transplantation can also be important as a therapy for neurodegenerative diseases. GRID-labeled neural stem cells were intracerebrally transplanted in a rat model of Huntington disease and observed to infiltrate the damaged hemisphere, migrating from their contralateral transplantation site [3]. Another possibility to track magnetically labeled stem cells in the context of ischemic injury has been tested in a model of swine myocardial infarction. Mesenchymal stem cells labeled with fluorescent micrometer-sized particles of iron oxide (MPIO) were percutaneously injected into the intact or infarcted myocardium. Signal darkening was detected in the beating normal or infarcted heart using cardiac MRI (fast gradient echo (FGE) or steady-state free precession (SSFP)) until 3 weeks after the cells’ implantation (confirmed ex vivo by MRI and histology). Fluorescence microscopy performed on histological slices showed that implanted mesenchymal stem cells were elongated and aligned with the host myocardium, suggesting their in vivo differentiation [80]. Using a similar model, SPIO-labeled mesenchymal stem cells were injected in the infarct and could be visualized by in vivo MRI until 1 week postimplantation. However, MRI also suggested that the size of the lesion increased and that the negative contrast decreased. Possible explanations for the loss of contrast were division of mesenchymal stem cells inducing a SPIO dilution, a migration of labeled cells, or eventually both [81]. In the same myocardial infarction

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model, SPIO-labeled myogenic precursor cells were injected in the infarcted area of pig’s heart. The necrotic region was identified in the left ventricle through endoventricular electromechanical mapping, as the part of the myocardium with the most reduced electrical and mechanical activity. Implantation sites of labeled myogenic precursor cells were detected in vivo by cardiac MRI and ex vivo by iron staining on histological slices [82]. Other experiments have been performed in rats or mice. Rat smooth muscle cells labeled with dimercaptosuccinic acid-coated particles (anionic particles) were implanted in rat hearts with ischemic injury. MR imaging of implantation sites was performed ex vivo on the heart excised from 2 to 48 h after grafting of cells. An anatomical correlation between the hypointense signal in myocardium (in T2- or T2* -weighted MRI) and the labeled cells detected in histology or immunohistochemistry was observed [83]. SPIO-labeled ESCs have also been visualized in vivo by T2-weighted MRI of the left ventricle of the intact rat heart [84]. Magnetically labeled cardiac progenitor cells have been transplanted into the healthy or injured myocardium of mice. High-resolution MRI allowed the precise determination of the location of labeled cells in the mouse’s heart. It was observed that a cardiac ischemic lesion appeared as a hyposignal area, similarly sized on both T2* -weighted and proton density-weighted MR images. However, after transplantation of iron oxide-loaded cardiac progenitor cells to the infarcted mouse myocardium, the signal attenuation area noticed on proton density MR images reflected only the lesion size, while it was larger on T2* -weighted MR images because of the higher detection sensitivity of magnetically labeled cells allowed by the latter parameter weighting [85]. In order to target other potentially damaged organs with stem cells, the fate of SPIOlabeled mesenchymal stem cells injected into the portal veins of rats treated with carbon tetrachloride to induce centrolobular liver necrosis, as well as into the renal arteries of healthy rats, has been studied by MRI and correlated to the histological observation of the iron-loaded cells’ distribution in the liver or kidney [86]. However, further experiments have shown that SPIO-labeled mesenchymal stem cells systemically injected (i.e., intravenously and not through the feeding artery) in rats with kidney mesangiolysis homed to the injured kidney but were below the in vivo MRI detection threshold. Only ex vivo images of the kidneys showed areas of lower intensity that were correlated to iron staining by histology [87]. Type 1 diabetes mellitus has been treated experimentally with another kind of cell therapy. Instead of stem cells, pancreatic islets containing ␤-cells (producing insulin) were transplanted into the liver of diabetic rats through injection in the portal vein. The pancreatic R ) and their hepatic location was islets were magnetically labeled with SPIO (Resovist
visualized by MRI. The labeled pancreatic islets, situated in the liver sinusoids, were observed as hypointense areas homogeneously distributed in the liver. Furthermore, a normal glycemia was restored in previously diabetic rats 1 week after transplantation of the islets [88]. 14.6.2 MRI Monitoring of Cells After Magnetic Labeling In Vivo Visualizing cells in vivo by MRI without any in vitro magnetic prelabeling requires a sufficient cellular uptake of the MRI contrast agent provided by the blood circulation after intravenous administration.

Inflammation Hepatic MRI using iron oxides is based on the capture of nanoparticles by the resident macrophages of the liver: the Kupffer cells. As inflammatory diseases are associated with monocyte/macrophage activity, this macrophage targeting approach has been

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further developed to visualize inflammatory lesions in the body. Thus following intravenous injection of iron oxide nanoparticles, inflammatory foci can be visualized by MRI. This has been achieved in experimental models of brain inflammatory diseases. After adminisR ) at a concentration of 300 ␮mol Fe/kg, macrophage activity tration of USPIO (Sinerem
was imaged in rats with experimental autoimmune encephalomyelitis (EAE), thanks to the capture of the nanoparticles by peripheral macrophages (blood monocytes or lymph node macrophages) subsequently invading the inflamed brain. However, the injected iron dose was about seven times higher than the appropriate dose for human use (45 ␮mol Fe/kg (2.6 mg/kg)). This was explained by the fact that a high iron dose is necessary to saturate the uptake capability of liver, spleen, or bone marrow macrophages. Furthermore, R in rats is only 2 h, while it is 20–24 h at the normal dose, the blood half-life of Sinerem
in humans [89]. In rat models of cerebral stroke, infiltration of macrophages in ischemic lesions has also been visualized by MRI following USPIO uptake. The localization of these hypointense areas on T2-weighted MR images was correlated to the histological detection of iron and to the immunohistological labeling of macrophages postmortem [90, 91]. In this context of brain damages associated with inflammatory events, macrophage activity has been studied by MRI after USPIO infusion also in human patients suffering from stroke [92]. In vivo cellular MR imaging of USPIO-labeled macrophages has found successful applications in other animal models of disease, such as in kidney injuries (nephropathies) induced by a nephrotoxic drug (puromycin aminonucleoside) or in organ graft rejection (e.g., after renal allograft transplantation without any immunosuppressive drug) [3, 93, 94]. The accumulation of long-circulating USPIO in monocytes and macrophages recruited to atherosclerotic plaques has been demonstrated also in hyperlipidemic rabbits [95, 96]. Furthermore, USPIO-enhanced MRI has also been used for in vivo detection of macrophages in human plaques. The strongest signal decreases, due to iron-loaded macrophage accumulation, were observed in ruptured and rupture-prone human atherosclerotic lesions rather than in stable lesions, allowing for a differentiation between low-risk and high-risk plaques on MR images [3, 97]. MPIO also have been intravenously injected into rats with the aim of labeling macrophage infiltrates in the rejection site of cardiac graft, allowing for the noninvasive monitoring of organ transplant rejection by MRI [98].

Tumors The in vivo tumoral accumulation of iron oxide nanoparticles has been investigated given the great interest of detecting tumors by MRI. In a rat model of intracerebrally implanted glioma, MION were intravenously injected (10 mg Fe/kg) and induced a signal increase of the tumor on T1-weighted MR images and a signal decrease on T2-weighted MR images, particularly at the tumoral periphery, 24 h after injection. Microscopic and histological examinations showed that the nanoparticles were intracellularly and extracellularly distributed in the tumor, with a higher level of accumulation in the periphery. Indeed, the highest microvascular density of a tumor is found in this area, which is consequently a region of preferential extravasation. Tumor-associated macrophages and vascular endothelial cells were observed to take up the nanoparticles, but the majority of nanoparticle-containing cells were tumor cells. Furthermore, the cellular uptake of the nanoparticles seems to be directly related to the tumor growth—in other words, directly related to the cellular cycle [19]. The same type of MION distribution in the tumor cells adjacent to the hyperpermeable tumor–brain interface has been reported in another study [99]. USPIO were injected intravenously into rats bearing intracerebral tumors (LX-1 small-cell lung carcinoma) and did not show the same distribution pattern as MION. Within 24 h postinjection, T1-weighted MR images showed an enhancement of the signal at the periphery of the tumor, surrounding a

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brighter region corresponding to a necrotic core. Histological analysis allowed for the detection of iron in and around necrotic areas, and at the tumor margin, but only in macrophages, reactive astrocytes, or gliosis (scar made of a dense fibrous network of neuroglia). Regarding LX-1 small-cell lung carcinoma, other tumor types exhibited a weaker T1-enhancement pattern with USPIO and consequently showed a sparser iron staining in histology, but confirmed the distribution of the nanoparticles in cells infiltrated at the brain–tumor interface R ), rapidly and in areas of necrosis. It was noticed that bigger nanoparticles (SPIO, Feridex
captured by spleen or liver macrophages, were ineffective for imaging brain tumors [100]. MION and USPIO are long-circulating iron oxide nanoparticles, able to passively target tumors by extravasation through the leaky tumoral capillaries. This interstitial accumulation of USPIO has also been suggested by a tumoral signal enhancement in T1-weighted MRI, 12 h after intravenous injection of the nanoparticles in mice bearing nonnecrotic subcutaneously implanted tumors (human prostatic adenocarcinoma). However, a signal decrease of the tumor was generated in T2-weighted MRI by dimercaptosuccinic acidcoated anionic iron oxide nanoparticles, 24 h postinjection. It suggested an intracellular tumoral distribution of these nanoparticles [101]. Magnetic labeling of tumors has also been achieved by magnetically targeting liposomes loaded with iron oxide nanoparticles to human prostatic adenocarcinoma implanted in mice. A preferential accumulation of these polyethylene glycol-coated long-circulating magnetoliposomes (200 nm in diameter) was induced by placing a magnet on the skin above the tumor during the circulation of the intravenously injected magnetoliposomes. MRI showed a strong signal darkening of the tumor exposed for 24 h to the magnet, compared to nonexposed tumors. It was observed that, due to a higher global magnetic moment, the magnetoliposomes could be magnetically driven to a tumor more efficiently than nonencapsulated USPIO. The mechanism of tumoral accumulation of magnetoliposomes is diffusion to the interstitium, through the leaky vasculature. Histological analysis revealed the presence of magnetoliposomes mainly in the most highly vascularized zones: in capillaries, interstitium, and cells (macrophages, fibroblasts, endothelial cells), but not in tumor cells. Iron was also found in liver and spleen [102]. Paramagnetic contrast agents have also been used as tumor-specific contrast agents in vivo. Gd-based metalloporphyrins, such as Gadophrin-2 (Bayer Schering Pharma AG), can selectively accumulate in tumors because of their high affinity to necrotic tissues. It has been suggested that their ability to bind plasma albumin is responsible for their slow extravasation and subsequent accumulation into the necrotic tumor interstitium [103, 104].

14.7 CONCLUSION AND PERSPECTIVES Iron oxide particles, made of polymer- or monomer-coated iron oxide crystals, are well known in clinical applications as intravenously injected negative MRI contrast agents. In vivo, they are sooner (for larger ones) or later (for smaller ones) taken up by cells of the mononuclear phagocyte system. Iron oxide particles have been presented as an interesting tool for cellular magnetic labeling because of their greater effect on the MR signal as compared to paramagnetic ions, allowing their MRI detection at relatively low concentrations (nanomolar) in tissues and thus opening the field of cellular MRI research. The ability of iron oxide particles to be taken up by phagocytic cells after intravenous injection has been further exploited to image macrophage-invaded areas in the context of inflammatory pathologies or tumors. MRI monitoring of nonphagocytic cells is often possible by a magnetic labeling of these cells performed in vitro, prior to their in vivo implantation. In

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other terms, iron oxide particles must be internalized by cultured cells in an optimal way. In many cases, particles are used unmodified (polymer- or monomer-coated iron oxide crystals) or they can be unspecifically cell-targeted (mixed with DNA transfection reagents, conjugated to a cell-penetrating peptide), triggering their endocytosis via nonspecific pathways. Iron oxide nanoparticles can also be linked to a ligand specific for a certain receptor at the cell surface, allowing their receptor-mediated uptake. In brief, the magnetic labeling of cells with MRI contrast agents, in particular, with iron oxide nanoparticles, is receiving increasing attention since in vivo applications exist in different pathological contexts. Since R R R (Endorem
), Resovist
) were they are taken up by phagocytic cells, SPIO (Feridex
originally developed and approved for clinical MRI to increase the lesion-to-liver contrast in malignancies where Kupffer cells are absent. It is worth mentioning that these SPIO were recently taken off the market due to lack of sales in that context. Nevertheless, they were used for cellular magnetic labeling studies in patients, suggesting that a great deal of future research and development will be driven in the field of cellular magnetic labeling and cellular MRI [105]. REFERENCES 1. Artemov, D.; Mori, N.; Okollie, B.; Bhujwalla, Z. M. MR molecular imaging of the Her-2/neu receptor in breast cancer cells using targeted iron oxide nanoparticles, Magn. Reson. Med. 2003, 49, 403–408. 2. Bulte, J. W. M.; Modo, M. M. J. Nanoparticles in Biomedical Imaging, Emerging Technologies and Applications. Springer: New York, 2007. 3. Modo, M.; Hoehn, M.; Bulte, J. W. Cellular MR imaging, Mol. Imaging 2005, 4, 143–164. 4. Hopkins, C. R.; Trowbridge, I. S. Internalization and processing of transferrin and the transferrin receptor in human carcinoma A431 cells. J. Cell Biol. 1983, 97, 508–521. 5. Kresse, M.; Wagner, S.; Pfefferer, D.; Lawaczeck, R.; Elste, V.; Semmler, W. Targeting of ultrasmall superparamagnetic iron oxide (USPIO) particles to tumor cells in vivo by using transferrin receptor pathways. Magn. Reson. Med. 1998, 40, 236–242. 6. Bulte, J. W.; Zhang, S.; van Gelderen, P.; Herynek, V.; Jordan, E. K.; Duncan, I. D.; Frank, J. A. Neurotransplantation of magnetically labeled oligodendrocyte progenitors: magnetic resonance tracking of cell migration and myelination. Proc. Natl. Acad. Sci. U.S.A. 1999, 96, 15256–15261. 7. Weissleder, R.; Moore, A.; Mahmood, U.; Bhorade, R.; Benveniste, H.; Chiocca, E. A.; Basilion, J. P. In vivo magnetic resonance imaging of transgene expression. Nat. Med. 2000, 6, 351–355. 8. Deans, A. E.; Wadghiri, Y. Z.; Bernas, L. M.; Yu, X.; Rutt, B. K.; Turnbull, D. H. Cellular MRI contrast via coexpression of transferrin receptor and ferritin. Magn. Reson. Med. 2006, 56, 51–59. 9. Harrison, P. M.; Arosio, P. The ferritins: molecular properties, iron storage function and cellular regulation. Biochim. Biophys. Acta 1996, 1275, 161–203. 10. Treffry, A.; Zhao, Z.; Quail, M. A.; Guest, J. R.; Harrison, P. M. Dinuclear center of ferritin: studies of iron binding and oxidation show differences in the two iron sites. Biochemistry. 1997, 36, 432–441. 11. Cohen, B.; Dafni, H.; Meir, G.; Harmelin, A.; Neeman, M. Ferritin as an endogenous MRI reporter for noninvasive imaging of gene expression in C6 glioma tumors. Neoplasia. 2005, 7, 109–117. 12. Wang, Z. J.; Boddington, S.; Wendland, M.; Meier, R.; Corot, C.; Daldrup-Link, H. MR imaging of ovarian tumors using folate-receptor-targeted contrast agents. Pediatr. Radiol. 2008, 38, 529–537.

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47. Iozzo, R. V. Heparan sulfate proteoglycans: intricate molecules with intriguing functions. J. Clin. Invest. 2001,108, 165–167. 48. Bhorade, R.; Weissleder, R.; Nakakoshi, T.; Moore, A.; Tung, C. H. Macrocyclic chelators with paramagnetic cations are internalized into mammalian cells via a HIV-tat derived membrane translocation peptide. Bioconjug. Chem. 2000, 11, 301–305. 49. Prantner, A. M.; Sharma, V.; Garbow, J. R.; Piwnica-Worms, D. Synthesis and characterization of a Gd-DOTA-D-permeation peptide for magnetic resonance relaxation enhancement of intracellular targets. Mol. Imaging 2003, 2, 333–341. 50. Lewin, M.; Carlesso, N.; Tung, C. H.; Tang, X. W.; Cory, D.; Scadden, D. T.; Weissleder, R. Tat peptide-derivatized magnetic nanoparticles allow in vivo tracking and recovery of progenitor cells. Nat. Biotechnol. 2000, 18, 410–414. 51. Josephson, L.; Tung, C. H.; Moore, A.; Weissleder, R. High-efficiency intracellular magnetic labeling with novel superparamagnetic-Tat peptide conjugates. Bioconjug. Chem. 1999, 10, 186–191. 52. Tung, C. H.; Weissleder, R. Arginine containing peptides as delivery vectors. Adv. Drug. Deliv. Rev. 2003, 55, 281–294. 53. Walczak, P.; Kedziorek, D. A.; Gilad, A. A.; Lin, S.; Bulte, J. W. Instant MR labeling of stem cells using magnetoelectroporation. Magn. Reson. Med. 2005, 54, 769–774. 54. Hinds, K. A.; Hill, J. M.; Shapiro, E. M.; Laukkanen, M. O.; Silva, A. C.; Combs, C. A.; Varnet, T. R.; Balaban, R. S.; Koretsky, A. P.; Dunbar, C. E. Highly efficient endosomal labeling of progenitor and stem cells with large magnetic particles allows magnetic resonance imaging of single cells. Blood 2003, 102, 867–872. 55. Shapiro, E. M.; Skrtic, S.; Koretsky, A. P. Sizing it up: cellular MRI using micron-sized iron oxide particles. Magn. Reson. Med. 2005, 53, 329–338. 56. Weisskoff, R. M.; Zuo, C. S.; Boxerman, J. L.; Rosen, B. R. Microscopic susceptibility variation and transverse relaxation: theory and experiment. Magn. Reson. Med. 1994, 31, 601– 610. 57. Shapiro, E. M.; Skrtic, S.; Sharer, K.; Hill, J. M.; Dunbar, C. E.; Koretsky, A. P. MRI detection of single particles for cellular imaging. Proc. Natl. Acad. Sci. U.S.A. 2004, 101, 10901–10906. 58. Shapiro, E. M.; Sharer, K.; Skrtic, S.; Koretsky, A. P. In vivo detection of single cells by MRI. Magn. Reson. Med. 2006, 55, 242–249. 59. Raynal, I.; Prigent, P.; Peyramaure, S.; Najid, A.; Rebuzzi, C.; Corot, C. Macrophage endocytosis of superparamagnetic iron oxide nanoparticles: mechanisms and comparison of ferumoxides and ferumoxtran-10. Invest. Radiol. 2004, 39, 56–63. 60. Palecanda, A.; Paulauskis, J.; Al-Mutairi, E.; Imrich, A.; Qin, G.; Suzuki, H.; Kodama, T.; Tryggvason, K.; Koziel, H.; Kobzik, L. Role of the scavenger receptor MARCO in alveolar macrophage binding of unopsonized environmental particles. J. Exp. Med. 1999, 189, 1497–1506. 61. Fleige, G.; Seeberger, F.; Laux, D.; Kresse, M.; Taupitz, M.; Pilgrimm, H.; Zimmer, C. In vitro characterization of two different ultrasmall iron oxide particles for magnetic resonance cell tracking. Invest. Radiol. 2002, 37, 482–488. 62. Champion, J. A.; Mitragotri, S. Role of target geometry in phagocytosis. Proc. Natl. Acad. Sci. U.S.A. 2006, 103, 4930–4934. 63. Jelezarova, E.; Luginbuehl, A.; Lutz, H. U. C3b2-IgG complexes retain dimeric C3 fragments at all levels of inactivation. J. Biol. Chem. 2003, 278, 51806–51812. 64. Ahrens, E. T.; Feili-Hariri, M.; Xu, H.; Genove, G.; Morel, P. A. Receptor-mediated endocytosis of iron-oxide particles provides efficient labeling of dendritic cells for in vivo MR imaging. Magn. Reson. Med. 2003, 49, 1006–1013.

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79. Zhang, Z.; Jiang, Q.; Jiang, F.; Ding, G.; Zhang, R.; Wang, L.; Zhang, L.; Robin, A. M.; Katakowski, M.; Chopp, M. In vivo magnetic resonance imaging tracks adult neural progenitor cell targeting of brain tumor. Neuroimage 2004, 23, 281–287. 80. Hill, J. M.; Dick, A. J.; Raman, V. K.; Thompson, R. B.; Yu, Z. X.; Hinds, K. A.; Pessanha, B. S.; Guttman, M. A.; Varney, T. R.; Martin, B. J.; Dunbar, C. E.; McVeigh, E. R.; Lederman, R. J. Serial cardiac magnetic resonance imaging of injected mesenchymal stem cells. Circulation 2003, 108, 1009–1014. 81. Kraitchman, D. L.; Heldman, A. W.; Atalar, E.; Amado, L. C.; Martin, B. J.; Pittenger, M. F.; Hare, J. M.; Bulte, J. W. In vivo magnetic resonance imaging of mesenchymal stem cells in myocardial infarction. Circulation 2003, 107, 2290–2293. 82. Garot, J.; Unterseeh, T.; Teiger, E.; Champagne, S.; Chazaud, B.; Gherardi, R.; Hittinger, L.; Gueret, P.; Rahmouni, A. Magnetic resonance imaging of targeted catheter-based implantation of myogenic precursor cells into infarcted left ventricular myocardium. J. Am. Coll. Cardiol. 2003, 41, 1841–1846. 83. Rivi`ere, C.; Boudghene, F. P.; Gazeau, F.; Roger, J.; Pons, J. N.; Laissy, J. P.; Allaire, E.; Michel, J. B.; Letourneur, D.; Deux, J. F. Iron oxide nanoparticle-labeled rat smooth muscle cells: cardiac MR imaging for cell graft monitoring and quantitation. Radiology 2005, 235, 959–967. 84. Tallheden, T.; Nanmark, U.; Lorentzon, M.; Rakotonirainy, O.; Soussi, B.; Waagstein, F.; Jeppsson, A.; Sjogren-Jansson, E.; Lindhal, A.; Omerovic, E. In vivo MR imaging of magnetically labeled human embryonic stem cells. Life Sci. 2006, 79, 999–1006. 85. Kustermann, E.; Roell, W.; Breitbach, M.; Wecker, S.; Wiedermann, D.; Buehrlr, C.; Welz, A.; Hescheler, J.; Fleischmann, B. K.; Hoehn, M. Stem cell implantation in ischemic mouse heart: a high-resolution magnetic resonance imaging investigation. NMR Biomed. 2005, 18, 362–370. 86. Bos, C.; Delmas, Y.; Desmouliere, A.; Solanilla, A.; Hauger, O.; Grosset, C.; Dubus, I.; Ivanovic, Z.; Rosenbaum, J.; Charbord, P.; Combe, C.; Bulte, J. W.; Moonen, C. T.; Ripoche, J.; Grenier, N. In vivo MR imaging of intravascularly injected magnetically labeled mesenchymal stem cells in rat kidney and liver. Radiology 2004, 233, 781–789. 87. Hauger, O.; Frost, E. E.; van Heeswijk, R.; Deminiere, C.; Xue, R.; Delmas, Y.; Combe, C.; Moonen, C. T.; Grenier, N.; Bulte, J. W. MR evaluation of the glomerular homing of magnetically labeled mesenchymal stem cells in a rat model of nephropathy. Radiology 2006, 238, 200–210. 88. Jirak, D.; Kriz, J.; Herynek, V.; Andersson, B.; Girman, P.; Burian, M.; Saudek, F.; Hajek, M. MRI of transplanted pancreatic islets. Magn. Reson. Med. 2004, 52, 1228–1233. 89. Dousset, V.; Doche, B.; Petry, K. G.; Brochet, B.; Delalande, C.; Caille, J. M. Correlation between clinical status and macrophage activity imaging in the central nervous system of rats. Acad. Radiol. 2002, 9 (Suppl 1), S156–S159. 90. Rausch, M.; Sauter, A.; Frohlich, J.; Neubacher, U.; Radu, E. W.; Rudin, M. Dynamic patterns of USPIO enhancement can be observed in macrophages after ischemic brain damage. Magn. Reson. Med. 2001, 46, 1018–1022. 91. Schroeter, M.; Saleh, A.; Wiedermann, D.; Hoehn, M.; Jander, S. Histochemical detection of ultrasmall superparamagnetic iron oxide (USPIO) contrast medium uptake in experimental brain ischemia. Magn. Reson. Med. 2004, 52, 403–406. 92. Saleh, A.; Schroeter, M.; Jonkmanns, C.; Hartung, H. P.; Modder, U.; Jander, S. In vivo MRI of brain inflammation in human ischaemic stroke. Brain 2004, 127, 1670–1677. 93. Hauger, O.; Delalande, C.; Trillaud, H.; Deminiere, C.; Quesson, B.; Kahn, H.; Cambar, J.; Combe, C.; Grenier, N. MR imaging of intrarenal macrophage infiltration in an experimental model of nephrotic syndrome. Magn. Reson. Med. 1999, 41, 156–162. 94. Beckmann, N.; Cannet, C.; Fringeli-Tanner, M.; Baumann, D.; Pally, C.; Bruns, C.; Zerwes, H. G.; Andriambeloson, E.; Bigaud, M. Macrophage labeling by SPIO as an early marker of

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CHAPTER 15

Nanoparticles Containing Rare Earth Ions: A Tunable Tool for MRI ` C. RIVIERE Laboratoire de Physique de la Matiere ` Condensee ´ et Nanostructures, Universite´ de Lyon, Lyon, France

S. ROUX Laboratoire de Physico-Chimie des Materiaux Luminescents, Universite´ de Lyon, Lyon, France ´

R. BAZZI Laboratoire Physico-Chimie des Electrolytes, Collo¨ıdes et Sciences Analytiques, Universite´ Pierre et Marie Curie, Paris, France

J.-L. BRIDOT Service de Chimie Gen Organique et Biomedicale, Laboratoire de RMN et d’Imagerie ´ erale, ´ ´ Moleculaire, Universite´ de Mons-Hainaut, Mons, Belgium ´

C. BILLOTEY Laboratoire CREATIS–Animage, Universite´ Claude Bernard, Lyon, France

P. PERRIAT Groupe d’Etudes de Metallurgie Physique et de Physique des Materiaux, Universite´ Claude ´ ´ Bernard, Lyon, France

O. TILLEMENT Laboratoire de Physico-Chimie des Materiaux Luminescents, Universite´ de Lyon, Lyon, France ´

15.1 INTRODUCTION Thanks to its highly resolved three-dimensional (3D) noninvasive imaging capabilities, magnetic resonance imaging (MRI) appears as one of the most powerful clinical diagnostic techniques. MRI can not only give precise anatomical information, but it can also provide functional measurements. MRI takes advantage of the magnetic resonance of water molecules that compose our bodies. The contrast obtained in MR images is due to the differences in the density and relaxation time of water protons within the defined

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volume of investigation. While the density of water protons will depend on the level of hydration of the different structures of the body, the relaxation time will depend on the environment of water molecules and the degree of mobility of the water molecules within this environment. MRI contrast is thus dependent on chemical, physical, and biological properties of the investigated tissue. Any changes in the tissue (protein density, edema, tumor, etc.) will affect the magnetic interactions between water protons, leading to changes in MRI contrast properties of that tissue. The relaxation of water protons is also dependent on the strength of the static magnetic field and on the pulse sequence of radiofrequency (rf) waves. Thanks to the development of various types of sequences, it is possible to get images mainly reflecting water proton density, transverse relaxation time (T2 imaging), or longitudinal relaxation time (T1 imaging). The signal tends to increase with decreasing relaxation time T1 and decrease with decreasing relaxation time T2. The interpretation of the resulting images therefore leads to the delineation and identification of most tissues. To increase MRI sensitivity, contrast agents have been developed that modify the relaxation time of water protons in their vicinity, hence modifying the resulting signal intensity detected in this region. Contrast agents are divided into two groups, the negative and the positive ones. Negative contrast agents are mainly superparamagnetic iron oxide nanoparticles, which induce a large shortening of the transverse relaxation time T2 of surrounding water protons, resulting in a decrease of tissue signal intensity, and hence darkening of the MR images. Positive contrast agents are mainly paramagnetic chelates, which induce a large shortening of the longitudinal relaxation time T1 of surrounding water protons, resulting in an increase of tissue signal intensity, and hence brightening of the MR images. Even if quantitative results can be obtained in theory for both negative and positive contrast agents, it requires development of important data treatment and physical interpretation for negative contrast agents, whereas it is much more straightforward for positive contrast agents. Among positive contrast agents, gadolinium(III), a rare earth compound, is the most widely used paramagnetic element, due to its seven unpaired electrons and relatively long electronic relaxation. Contrast properties are dependent on two key features: the water-exchange rate between bulk water and water bound to the Gd ions and the rotational correlation time of the Gd containing entities. As free gadolinium is extremely toxic, gadolinium ions are caged within chelates ® structures, such as DTPA (diethylenetriaminepentaacetic acid, Magnevist ), and DOTA ® (1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid, Dotarem ) [1]. As many nanostructures presented in this chapter are based on the use of such complexes, the structure of gadolinium chelates is described in detail in Section 15.2. This caging limits the number of coordinated water molecules associated with each Gd ion, and hence limits the modification of signal intensity that could be induced with such a paramagnetic element. Another limitation is due to the small size (<1 nm) of clinically used Gd complexes, which leads to short rotational correlation time, hence limiting the modification of proton relaxivity. Their small size also reduces the blood half-life of Gd complexes, and blood-pool imaging is only accessible during the first few minutes after intravenous injection of the contrast agent. Moreover, as it is quickly eliminated from the body, both passive (by taking advantage of the enhanced permeability and retention (EPR) effect in tumoral tissues) and active targeting (by grafting relevant target-specific ligand on the surface of the contrast agents) are limited. In that context, many nanostructures have been developed in the past few decades (1) to increase MRI sensitivity while keeping a positive contrast enhancement

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to get quantitative results more easily; and (2) to tune the biodistribution of the contrast agent for tissue-targeting purpose. Moreover, these nanostructures can often support additional attractive features to combine imaging and therapeutic functionalities and/or to induce a contrast in different complementary imaging techniques (optic, MRI, PET, ultrasound, etc.). This chapter specifically focuses on the different nanostructure-based strategies employed to increase the sensitivity and specificity of the positive contrast agents. To increase the sensitivity of rare earth based contrast agents, three parameters are usually considered: (1) increasing the number of Gd per contrast agent, (2) increasing the rotational correlation time by increasing molecular weight and size, and (3) increasing the number of coordinated water molecules. The sensitivity of a given contrast agent is classically quantified by measurements of the longitudinal and transverse relaxivity values, r1 and r2 , which refer to the amount of increase in 1/T1 and 1/T2, respectively, per millimolar of agent (often given as per mM of Gd). The ratio r2 /r1 is used to classify the contrast agents as positive T1 agents, which usually have r2 /r1 ratios of 1–2, or as negative T2 agents, which have r2 /r1 ratios as high as 10 or more. For nanostructure elements, it is not only the relaxivity per Gd that defines the efficiency of the contrast agent but also the number of chelated Gd per nano-object. It is then also interesting to compute the relaxivity values per mM of nano-objects, instead of per mM of Gd. Section 15.3 is dedicated to a review of different strategies to entrap the rare earth chelates in 3D nanostructures, while Section 15.4 is devoted to the use of nanostructures with rare earth surface functionalization, including the development of crystalline nanoparticles based on inorganic or organometallic compounds of rare earth elements. In Section 15.5, we introduce the use of such MR contrast agents for both imaging and therapy.

15.2 STRUCTURE OF GADOLINIUM CHELATES AND CHEMICAL MODIFICATION The structure of gadolinium chelates was extensively described in several reviews, which are taken as the authority in the MRI field [1, 2]. Generally, gadolinium chelates are composed of a gadolinium(III) ion buried in a ligand, which is designed for occupying at most eight binding sites at the metal center. The remaining coordination sites are occupied by solvent water molecules. The presence of the water molecules in the coordination sphere of Gd(III) ion and its exchange with the water molecules of the surrounding solution are essential for affecting the relaxation time of solvent protons. The hydration of the gadolinium(III) complex is an important parameter since the higher the number of water molecules in the coordination sphere, the more pronounced are the effects on the relaxation and therefore on the contrast of the MR images. However, the gain in relaxivity is often depreciated by the loss of stability of the complex, which leads to the undesirable liberation of toxic Gd3+ in the body. The majority of clinically approved contrast agents that are commonly used for medical imaging are gadolinium(III) chelates whose ligand is based on a linear or cyclic polyaminocarboxylate motif [1, 2]. Among them, linear DTPA and cyclic DOTA are well suited for the immobilization of gadolinium(III) onto the nanoparticles (Fig. 15.1). First, Gd-DTPA and Gd-DOTA (actually Gd(DTPA)2− and Gd(DOTA)− ) form

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O

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OH O

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FIGURE 15.1 The structures of commercially available ligands DTPA (left) and DOTA (right).

stable octacoordinated chelates. The presence of several carboxylic acid moieties in these ligands can be exploited for the covalent immobilization of gadolinium(III) chelates. Their grafting can be performed directly onto the surface or indirectly by using a crosslinker or by modification of the ligand. It will depend on the chemical composition of the nanoparticle surface. Since an amide group is, regarding hydrolysis, more resistant than an ester function, the immobilization of gadolinium(III) chelates, whatever the grafting mode, requires a condensation reaction between COOH groups of the ligand and an amine function (Fig. 15.2). This reaction is generally promoted by a water-soluble carbodiimide (EDC) and N-hydroxysuccinimide (NHS), which are added to the carboxylic acid containing molecules before the reaction with amine [3]. The conversion of one or two acid groups into amide is accompanied by a lowering of the stability constant of the resulting gadolinium chelate because the coordination strength of amide is lower than the one of carboxylate. Despite this, their use for imaging application is possible because, in most cases, the decrease of the stability constant is of the same magnitude as the one observed for the clinically approved Gd-DTPA-bis(amide) (DTPA-BMA:Gd, K = 1016.85 ). In order to circumvent the decrease of stability constant, the carboxylic acid moieties must therefore be preserved. This implies that the carboxylic groups of the ligands cannot be used for the covalent grafting of the gadolinium complexes. In this case, the immobilization of gadolinium chelates requires the presence of a reactive group on the diethylenetriamine backbone besides the carboxylic acid groups. Brechbiel and co-workers developed the synthesis of DTPA and DOTA derivatives comprised of a grafting group (amine, isothiocyanate) carried by the ethylene chain (Fig. 15.3) [4, 5]. Despite the advantages that these functionalized ligands can afford for the covalent grafting of gadolinium(III) chelates, they are rarely applied, in contrast to DTPA and DOTA, because of their time-consuming multistep synthesis.

15.3 NANOSTRUCTURES AS 3D SCAFFOLD FOR RARE EARTH CONTRAST AGENT DELIVERY Various strategies have been developed to concentrate paramagnetic contrast agent within a 3D scaffold and thereby to increase MRI sensitivity per contrast agent. One approach is to encapsulate contrast agents within lipid-based nanostructures such as liposomes, or within porous mesoporous nanostructures. More recently, carbon nanotubes and metallofullerenes have also been studied as potentially attractive scaffolds for rare earth contrast agents. Those three approaches will be developed in the following sections.

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+ HOOC

HOOC

COOH COOH

HOOC SCN

COOH

S

+

N

N

N

N COOH

FIGURE 15.2 Reaction of DTPA, DOTA, and DTPA and DOTA-analog with amine groups.

15.3.1 Lipid-Based Structures In “lipid-based structures,” nano-objects possessing an outer membrane and/or an inner cavity will be described. This refers mainly to liposome-based and lipoprotein-based contrast agents. Liposomes are particularly attractive, being phospholipid-based biocompatible structures of adjustable size distribution. They can coencapsulate hydrophobic (within the phospholipid bilayer), hydrophilic (inside the vesicle), and amphiphilic molecules, such as therapeutic chemicals, genes, and contrast agents [6–13]. In order to avoid opsonization and recognition by the reticulo endothelium system (RES), liposomes are usually coated with polyethylene glycol (PEG). The furtiveness

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X

O HO HO

O

O N

N O

O

O

O OH

OH

N O

N

O

N

N

O

O

OH

OH

O

X= NCS, NH2

NCS or NH2-Bn-DTPA

DTPA-BA

COOH

HOOC X

N

N

N

N COOH

HOOC X= NCS, NH2

NCS or NH2-Bn-DOTA

FIGURE 15.3 The structure of the DTPA and DOTA analog bearing a highly reactive isothiocyanate (NCS), amino (NH2 ), or anhydride group.

of such long-circulating liposomes makes them promising passive drug delivery devices ® [13–16]. Stealth liposomes are used commercially to deliver chemotherapeutic drugs such as doxorubicin and cisplatin (Doxil). The use of liposomes as paramagnetic ion carriers has been investigated since the early 1990s. Two issues are to be addressed in the design and synthesis of these biomimeticbased structures: the rare earth payload and the signal intensity enhancement induced by the liposome-associated paramagnetic ions. Even if the relaxivities are improved with such structures, contrast enhancement is usually not as high as the one predicted by the number of Gd complexes that can be trapped in or within these biomimetic nano-objects (Table 15.1). In all these studies, the moderate increase in contrast properties is attributed to the moderate degree of freedom of the rare earth ions, which cannot fully interact with the water protons in the environment. Several attempts to address these two specific issues (payload and signal enhancement) have been carried out over the past two decades. As far as relativities are concerned, liposomes bearing rare earth ions must be divided into two classes, depending on the location of the paramagnetic entities: (1) encapsulated in the liposome cavity and (2) in the membrane.

Paramagnetic Complexes Encapsulated in the Liposome Cavity The use of liposomes as T1 contrast agent was first investigated by encapsulating Gd-DTPA complexes within the liposomes cavity [17]. Even if the relaxivity per contrast agent object was increased with this strategy, the liposome encapsulation significantly lowered the T1 relaxivity (in mM−1 · s−1 of Gd) compared to free Gd complexes. The relaxivity obtained for liposomes varied linearly with the surface-to-volume ratio: the smaller the liposomes, the greater the relaxivity. Even the smallest liposomes (70 nm) have a relaxivity nearly twofold lower than the one obtained with the free complex. This has been attributed to the limited access of Gd atoms to the bulk water molecules because of the liposomal bilayer,

®

10.3

3.4 3.4 9.2 3.9 6.8

3.4 3.4 3.9 6.8

4.3 7

9.1

4.8 3.8

Clinical Gd Complexes

9.1

45,000 43,000 1800 510 15–20 27,000 ?

125 117 130 40 9 100 20–80 nm/ 3–10 nm

22 170

39 10.4

7.2 1 17

41 ?

49,7 ?

? ? 26

1.3 ? 1.9 ?

2.5 × 104 ? 1.1 × 106 ?

2 × 104 150–200 5.9 × 105 ?

? ? 1.5

1.1 1

1.4 1.1

1.1

? ? 4.7 × 104

4.3 7

4.8 3.8

10.3

3.2 × 105 4,3 × 104 3.1 × 104

Nanostructures as 3D Scaffolds for Rare Earth Contrast Agent Delivery

1 1 1 1 1

1

r1 r2 r1 r2 Nb (Gd) (mM−1 · s−1 ) (mM−1 · s−1 ) (mM−1 · s−1 ) (mM−1 · s−1 ) r1 /r2

Relaxivities per Nanostructure

0.47 T 1.5 T

0.47 T 1.5 T

1.4 T 0.47 T 1.5 T

1T 1T 3T 1T 3T

1T

Field (T)

(Continued)

48 39

117 28

19 18 23

157 157 144 157 51

156

Reference

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Porous polymerosomes Gd-encapsulated Stealth liposomes DPPC and Gd-DTPA-bisamide derivative Shell-crosslinked nanoparticles Lipoprotein micelles Gd in the lipid membrane Nanorods Gadonanotubes

Gd-DOTA (Dotarem ) ® Gd-DTPA (Magnevist ) Gd-DTPA ® Gadodiamide (Omniscan ) Gd-Si-DTTA complex

Gd3+ free

Size (nm)

Relaxivities per Gd

TABLE 15.1 Comparative Relaxivity Values of the Different Rare Earth Containing Nanostructures

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340

Gd-DOTA-QD Gd-DOTA-QD Au@DTDTPA-Gd150 Au@DTDTPA-Gd50 Au-Gd Gd2 O3 @PEG

Silica nanoparticles Silica nanoparticles Gd-Si-DTTA Silica nanoparticles Gd-Si-DTPA Silica nanoparticles Gd-Si-DTTA LbL technique

Gadofullerenes Gd@C82

67 31

1 1

?

28.8 10.2 23

24,300

79 131

65.5 110.8 34 67 31

7 × 10 2.48 × 105 ? 5

79 131

1.6 × 10 2.7 × 106 ?

16,000 10,200 63,200 10,200 28,000 ? 45 150 50 ?

100 37 40 37 44 10 8 2.4 2.4 75 4.3

19 ? 23 3.9 3.9 23.7 12

9 19.7 7.8 19 55 ? 54 ? ? 89.5

116 60 12.3 55

1.9 × 106 6.1 × 105 7.8 × 105 5.6 × 105 1.6 × 106 6779 2438 ? ? ?

1.4 × 105 2 × 105 4.9 × 105 1.9 × 105 5.3 × 105 365 1019 585 195 ?

6

2.9 19 2.3 ? ? 3.8

13 3 1.6 2.9

1.2 4.2

2.3 11 1.5

r1 r2 r1 r2 Nb (Gd) (mM−1 · s−1 ) (mM−1 · s−1 ) (mM−1 · s−1 ) (mM−1 · s−1 ) r1 /r2

Relaxivities per Nanostructure

Nanostructures with Rare Earth Surface Functionalization for Multimodal Imaging

27 nm Pore size: 2,4 nm 120 nm Pore size: 2 nm 1 nm

Size (nm)

Relaxivities per Gd

3T ? 1.4 T 7T 7T 1.5 T 1.5 T

4.7 T 3T 3T 3T

0.47 T 4.7 T

3T 9.4 T 9.4 T

Field (T)

52 87 88 101,102 101,102 105 140

158 51 51 52

34 34

49

47

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Mesoporous silica nanoparticles

Mesoporous silica nanospheres

TABLE 15.1 (Continued)

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Hollow sphere gadolinium oxide nanoparticles: shell thickness 21 nm, interior cavity 91 nm Porous sphere gadolinium oxide nanoparticles: pores 3–12 nm GdF3 nanoparticles@cit GdF3 nanoparticles@AEP Gadolinium carbonate (Gd2 O(CO3 )2 · H2 O) particles—spheres

Gd2 O3 @DMSA-PEG Gd2 O3 @CA-PEG Gd2 O3 core/silica shell nanoparticles: shell thickness 2 nm Gd2 O3 core/silica shell nanoparticles: shell thickness 2 nm Gd2 O3 core/silica shell nanoparticles: shell thickness 2 nm Dysprosium oxide nanoparticles core/silica shell

16.8 3.17 2.71 16.5

2110 8900 ? ? ?

? 6.3 × 106 3.25 × 105 ?

3.8 4.6 1 nm core/70 total 1 nm core/60 total 130

187 130 51 500

17.7

?

?

4.4

8.8

420

14.2 7.8 8.8

4.3 4.3 2.2

? ? 210.7

2.6 × 105

3.9 × 104

2.107 8.8 × 105 ?

?

?

?

? ? ?

?

?

675

190

6.1 × 104

1.9 × 104

?

4.8 × 103

3.7 × 103

? ? 13

1.9

1.5

?

?

6.6

3.3

1.3

14 T 14 T 3T

3T

3T

17.6 T

7T

7T

7T

124 124 126

144

144

142

142

131

131

140 140 131

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31.4

26.6

?

?

28.9

28.8

11.4

1.5 T 1.5 T

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which lowers the rate of water exchange between the bulk and the interior of the liposome. The presence of cholesterol in the liposomal bilayer was found to decrease even further the relaxivity by reducing the permeability coefficient of water molecules. Another disadvantage of such an approach was the possible leakage of low molecular weight compounds from the vesicles in vivo. More recently, PEGylated Gd-encapsulated liposomes (117 nm) presenting longcirculating properties and high Gd content were investigated [18]. The decrease in relaxivity compared to the free Gd complexes was overcome by an efficient payload (43,000 Gd per liposomes on average), leading to efficient and long-lasting contrast properties, which are extremely attractive properties for blood-pool imaging and MR angiography. Other interesting alternatives are also under investigation, such as dendrimer-attached Gd complexes encapsulated in so-called porous polymerosomes [19]. Porous polymerosomes are 125-nm structures composed of an outer porous membrane that enables better water exchange across the bilayer, stabilized by a polymer shell (an amphiphilic diblock copolymer). High T1 relaxivity obtained with such a nanostructure (r1 (60 MHz, 1.41 T, 40 ◦ C) = 7.2 mM−1 · s−1 versus 3.1 mM−1 · s−1 for free dendrimer-attached Gd complexes and 3.9 mM−1 · s−1 for free Gd-DTPA in the same condition) is assumed to be due to high Gd payload (approximately 45,000 Gd/porous polymerosome) and faster flux of water across the bilayer together with slower rotational correlation lifetime of the encapsulated Gd dendrimers.

Paramagnetic Complexes in the Membrane To overcome the decrease in water accessibility of encapsulated Gd complexes, a second strategy consists of entrapping paramagnetic complexes within the membrane. This strategy has been used successfully with chelating moieties modified with different hydrophobic groups such as alkyl chains of varying lengths attached via amide linkages, Gd-DTPA-stearylamine (Gd-DTPA-SA), GdDTPA-phosphatidyl ethanolamine (DTPA-PE) [20], or more recently Gd-DTPA-polylysine linked to a polychelating amphiphilic polymer (Gd-DTPA-PLL- PAP) [21, 22]. Laurent and co-workers [23] have studied in detail the effects of chain type and length of amphiphilic complexes bearing gadolinium complexes within the phospholipid membrane. The relaxivity of the studied paramagnetic unilamellar DPPC (1,2-sn-glycero-3phosphatidylcholine) liposomes was found to be mainly related to the matching of the carbon chain length and of the chemical structure of the phospholipid and the Gd complex. The best efficiency as an MRI constrast agent was obtained when the membrane is less structured, that is, when the immobilization of the paramagnetic complex within the membrane was lowered, enabling efficient water exchange rate. Fortunately, the incorporation of PEG moieties on the outer surface of the membrane, classically used to increase the contrast agent’s stability and blood half-life, also contributes to an increase in contrast properties, as water molecules are tightly associated with hydrophilic PEG molecule [24]. Sterically stabilized ∼100-nm liposomes with Gd complexes coupled to a fatty acid chain (dimyristoyl-sn-glycero-3-phosphoethanolamineN-diethylene-triaminepentaacetate—DMPE-DTPA) exhibit threefold higher relaxivity at 1.5 T compared to conventional paramagnetic complexes and a prolonged blood halflife (120 min), leading to a progressive signal enhancement in the tumor region (with a maximum obtained 20 h after liposome injection) [25] . Limitations of this approach include the low loading capacity and the high quantity of the lipid in the carrier. One strategy to increase the number of complexes per liposome is the use of amphiphilic polychelating polymers. The advantage of this approach is that one

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phospholipid can carry a polymer molecule with multiple paramagnetic complexes on the liposome surface. It was found that liposome modification with Gd-DTPA-polylysine can increase the rare earth load per vesicle 5- to 15-fold [21], while decreasing the lipid-to-Gd ratio. Even if this approach does not increase the relaxivity per mole of gadolinium, the relaxivity per nano-objects is increased (Table 15.1). Most recent studies have focused on the use of PEG-liposomal membrane-incorporating polychelating amphiphilic polymer heavily loaded with gadolinium and functionalized with ligand specifically targeting cancer cells for sensitive and specific tumor detection [22].

Stable Lipoprotein Mimetic Micelles A promising alternative approach is the use of lipoprotein-based structures. These structures can be considered as attractive biomimetic nanostructures to target vulnerable plaques (in atherosclerosis), tumor, or other inflammatory related disease. To our best knowledge, the design of such objects is among the most advanced techniques concerning in vivo targeting and imaging of the early stages of atherosclerosis. High-density lipoprotein (HDL) is a native structure that naturally interacts with atherosclerotic plaques [26]. In vivo, HDL helps remove excess cholesterol from macrophages in atherosclerotic plaques. It is composed of a hydrophobic core of triglycerides and cholesterol esters covered in a monolayer of phospholipids into which apolipoprotein A-I (apoA-I) is embedded [27]. After extraction, the apoliproteins are reconstituted with phospholipids, including a phospholipid-based contrast agent such as GdDTPA-DMPE (1,2-dimyristoyl-sn-glycero-3-phosphoethanolamine diethylenetriamine pentaacetic acid) [28] or Gd-DTPA-BSA (bovine serum albumin [29]). Depending on the dialysis method, both spherical [28] and disks-like [29] nano-objects can be obtained, bearing about 20 molecules of Gd. As the protein is endogeneous, it is not recognized by the reticuloendothelial system, resulting in a low level of immunogenic host reaction. In addition, the small size of the final nano-object obtained (7–12 nm in diameter) is much smaller than the one obtained with liposomes and is thus more adapted to penetrate the plaques. As far as contrast properties are concerned, due to the reduced rotation of the Gd complexes, compared with free Gd-DTPA, the obtained relaxivity values are relatively important (two- to threefold increase compared to Gd-DTPA, with r1 = 10.4 · mM−1 · s−1 at 65 MHz [28]). This type of lipoprotein-based contrast agent is readily internalized by macrophage cell lines in vitro [27], resulting in high contrast enhancement of labeled cell pellets. Using such a contrast agent in a mouse model of atherosclerosis (hyperlipidemic mice, Apo-E knockout), in vivo signal enhancement in atherosclerotic plaques rich in macrophages has been reported 24 h post intravenous injection, confirming the targeting efficiency of this contrast agent [28, 29]. Alternative solutions are currently under investigation to optimize this type of structure: (1) the use of synthetic peptides instead of HDL [26] to avoid the lengthy separation procedure required to isolate HDL and also to reduce immunogenicity; and (2) inclusion of an organic core (gold, iron oxide, or quantum dots) for multimodal imaging [27]. In addition, by postfunctionalization with a molecule known to be expressed in large amount in a given state of the development of atherosclerosis or tumor progression (such as VCAM-1, E-selectin, or ␣v ␤3 integrin), even more specific targeting could be achieved with such a contrast agent [27].

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pH-Sensitive Liposomes In order to detect pathologies associated with pH modifications, pH-sensitive liposomes have been developed recently [30–32]. Basically, the idea is to modify the liposome structure as a function of pH values. This modification will then induce a selective modification in the contrast enhancement, providing a direct visualization of in vivo pH variations. Classically, the system consists of gadolinium diethylenetriamine pentaacetic acid bismethylamide (gadodiamide, GdDTPA-BMA) encapsulated within liposomes. The modification of contrast enhancement can be induced by a destabilization of the liposome’s membrane, hence liberating entrapped Gd complexes. As stated above, the relaxation properties of free Gd complexes are usually higher than the one of Gd complexes entrapped within the liposome cavity. The liberation of Gd complexes due to pH variation will hence increase the contrast enhancement in that particular area. The difficult part resides in the control of the destabilization process, as the liposomes should remain stable within the bloodstream, before reaching the pathological area. Many formulations are currently under investigation. One of the most promising is the dipalmitoylphosphatidylethanolamine/ dipalmitoylglycerosuccinate system (DPPE/DPSG) [31], which remains stable in blood while displaying large pH response. 15.3.2 Carbon Nanotubes and Metallofullerenes Different carbon structures have shown recently their potential as MRI T1 contrast agents. Two types of structures have been shown to have extremely high relaxivity values at clinical magnetic field strength and are currently under investigation: (1) metallofulleres and (2) carbon nanotubes.

Metallofullerenes Metallofullerenes consist of endohedral cages of about 1 nm that encapsulate atoms, clusters, or small molecules. Gadolinium containing metallofullerenes (Gd@C60 or Gd@C82 mainly), so-called gadofullerenes, are emerging as novel caging structures concurrently replacing clinically used gadolinium chelates. They exhibit high relaxivity properties while preventing Gd3+ dissociation in vivo thanks to the confinement of the paramagnetic ion within the fullerene cage. The metallofullerenes are typically obtained thanks to an arc discharge method of metal/graphite composite rods. The resulting soot must then be extracted with carbon disulfide. The soluble fullerene M@C82 is then purified and isolated by two-stage high performance liquid chromatography (HPLC). The obtained toluene solution of M@C82 is then mixed with NaOH solution containing a few drops of tetrabutylammonium hydroxide (TBAH) as catalyst. This step produces a reaction mixture that can easily be isolated by decantation. Aqueous solution of M@C82 (OH)n (gadofullerenols in the case M = Gd) is then precipitated with methanol and the precipitate is further rinsed with methanol to ensure the removal of the catalyst and NaOH. The rinsed precipitate is finally dissolved in water and further purified by dialysis against water. Because this synthesis is labor intensive, has low yield (∼1% wt), and only 4% of sublimable Gd endohedrals are of the soluble Gd@C82 variety, other routes of synthesis have recently been investigated to produce soluble M@C60 , which represents the major portion of all metallofullerenes produced in the traditional arc synthesis method. Indeed, the first generation of M@C60 (M being any metal ion) materials were insoluble in water and air sensitive, which prevents their use in biomedical applications. New techniques of derivatization, solubilization, and stabilization of these species have overcome these limitations [33], by forming soluble carboxylate gadofullerenes, Gd@C60[C(COOH)].

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The obtained Gd@C82 and Gd@C60 were investigated as potential MRI contrast agents. They both show high water proton relaxivity with r1 , respectively, 20 and 10 times higher than clinically used Gd-DTPA both in vitro (r1 (Gd@C82 ) = 81 mM−1 · s−1 at 1 T, 25 ◦ C [34] and r1 (Gd@C60 ) = 40 mM−1 · s−1 [35], against 4 mM−1 . s−1 for Gd-DTPA in the same condition) and in vivo (enhancement detected after intravenous (IV) injection of a dose 20 times lower than the one currently used for Gd-DTPA [34]). It is difficult to understand the mechanisms underlying this contrast enhancement, as the encapsulated Gd ion is unable to interact directly with water protons. Even if the precise electronic and magnetic properties of gadofullerenols are not known at present, the relaxivity enhancement is attributed to electronic interaction of cages of Gd ions with bulk water protons through the fullerene cage walls and their functionalized groups [35]. Gd@C82 has indeed paramagnetic properties due to the three electrons transferred from Gd to the C82 fullerene cage, leading to seven unpaired electrons in the Gd orbitals inside the C82 cage and one unpaired electron on the cage [36]. The enhanced relaxation of the surrounding water protons is thus likely due to dipolar interactions between unpaired electrons and the large number of water molecules on the surface of the fullerenes [1], together with a decrease in the overall molecular rotational motion. As opposed to Gd chelates, whose relaxation mechanism is mainly based on an “innersphere” mechanism (direct interaction of water protons with Gd ions, that rapidly exchange with bulk water molecules—fast rotational correlation motion), the mechanism for gado–fullerene is entirely “outer-sphere” (related to diffusion of water molecules in the vicinity of the contrast agents) [35]. Although there is no experimental evidence on the dynamic structure of the surrounding water molecules, the presence of hydrogen bonds at the surface of the fullerene, together with their negative surface charge, may lower the rotation of the water molecules in direct contact with the surface of the fullerene and increase the effective hydrodynamic diameter, inducing substantial decrease in the overall molecular rotational motion [37]. Moreover, relaxivity measurements as a function of field strength (NMRD profile) show high-field peaks, typical of slowly rotating objects [35]. NMRD measurements were best fitted with a rotational correlation time two to three orders of magnitude higher than the one obtained with Gd chelates (on the order of nanoseconds for gadofullerenols versus picoseconds for Gd chelates), and a rate of water exchange two to three times greater than the rate of water exchange from the inner sphere water to bulk water for Gd3+ . The relaxivity enhancement has also been related to an increase in hydrodynamic diameters of gadofullerenes (hence a decrease in its tumbling rate) by aggregation processes as a function of pH [38]. Taking advantage of the controlled aggregation process, gadofullerenes appear as interesting pH-sensitive agents. The strong pH dependency of the aggregation (increasing hydrodynamic diameter with decreasing pH), and hence the strong pH dependency of the relaxivities (increasing contrast enhancement with decreasing pH) can be used to induce contrast enhancement in regions with low pH values, such as in tumors [35, 38]. A biodistribution study in a mouse model indicated that, in vivo, the Gd@C82 (OH) aggregates are recognized by the reticuloendothelial system (RES), with strong signal enhancement and Gd concentration in lung, liver, spleen, and kidney after IV administration [34]. In contrast, biodistribution of the Gd@C60 [C(COOH)2 ]10 in a rat model shows a low level of recognition by the RES for that type of gadofullerenols, which is rapidly appearing in the kidneys, with only minimal uptake into the liver and excretion via the bladder within 1 h of injection [33].

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Carbon Nanotubes Since other chapters of this book are dedicated to carbon nanotubes for optical imaging and multimodal imaging, this section is restricted to a short description of interesting results concerning the use of carbon nanotubes as a versatile structure to get a high load of Gd complexes. The so-called gadonanotubes developed by Sitharaman and co-workers [39] is certainly one of the most advanced attempts to use carbon nanotubes as T1 MRI contrast agents. In their study, they used short nanotubes (20–100 nm). Such single-walled carbon nanotubes (SWNTs) have recently been shown to shuttle various cargoes across cellular membrane without cytotoxicity [40–42]. Such gadonanotubes exhibit relaxivity values at clinical field strength 40-fold higher than the current free Gd complexes with r1 ∼ 170 mM−1 · s−1 (at 1.5 T and 40 ◦ C) [39]. Compared to gadofullerenes, the interesting point is the good transport properties of water and protons [38, 43], enabling access of water molecules to Gd complexes within the carbon nanotubes. Proton relaxivity measurements as a function of the magnetic field strength (NMRD profile) of such systems show a profile remarkably different from other paramagnetic agents, notably with constant high relaxivity values at high field. As MRI resolution increases with the magnetic field strength, the use of such a system could circumvent the actual limitation of paramagnetic complexes, which usually show a decrease in relaxivity above 1.5 T. More recently, pH-sensitive gadonanotubes have been developed, which exhibit a twofold change in relaxivity between pH 7.4 and 6 [44]. Such a system is particularly attractive to achieve smart MRI detection. Cancerous and normal tissues do not have the same pH, with lower pH usually associated with cancerous areas. Detecting such pH variation is thus an interesting approach for the detection of cancer. 15.3.3 Mesoporous Nanoparticles Mesoporous nanoparticles are mainly based on silica nanoparticles. Such rigid structure has a large pore volume, with tunable pore size and surface functionalization. The biocompatibility of such a system is now established [45, 46] and mesoporous silica nanoparticles are emerging as ideals agents for biomedical application. Superparamagnetic nanoparticles and paramagnetic complexes were successfully incorporated in these structures for their use as MRI contrast agents. As far as paramagnetic molecules are concerned, different Gd complexes have been used: Gd-DTTA [47] (diethylenetriaminetetraacetate), Gd-DTPA [48, 49], or Gd-DOTA [50]. The synthesis is based on a sol-gel procedure using surfactant such as CTAB (cetyltrimethylammonium bromide) as template. A first route consists of cocondensation under basic conditions of tetraethyl orthosilicate (TEOS) with chelating agent (such as DTPA) conjugated with silane. The obtained solution of mesoporous nanostructure is then mixed with GdCl3 , which enables caging of Gd ions within chelates [48, 49]. A second route consists of condensation of tetraethyl orthosilicate (TEOS) under basic conditions. The obtained mesoporous structures are then coated with Gd chelates conjugated with silane complex via a siloxane linkage [47]. Mesoporous nano-objects of tunable form (nanospheres [47, 49, 50], nanorods of various aspect ratio [48]), size (from 60 to 1100 nm), and pore size (from 2 to 10 nm) can be obtained [47]. As far as MRI contrast properties are concerned, as the accessibility of paramagnetic entities to water molecules is one of the most important features in obtaining high contrast

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agent relaxivity properties [51, 52], the large surface-to-volume ratio of mesoporous nanoobjects appears as an interesting property for the design of sensitive contrast agents. In fact, Gd containing mesoporous nano-objects present a remarkable relaxivity increase, compared to free Gd complexes or Gd entrapped in plain silica nanoparticles. This enhancement in contrast properties is due to a concomitant effect of the porosity of the structure, which enables free movement of water molecules in and out of the structure, and of the interaction of Gd complexes with the silica structure, which impedes the rotational movement of the lanthanide complexes. This decrease in tumbling rate of the paramagnetic unit consequently increases the relaxivities of water molecules in the vicinity of the Gd complexes. Five- to tenfold increase in longitudinal and transverse relaxivities were reported for mesoporous nanorods [48] of 100 nm with a 3/5 aspect ratio and a pore size of 2–3 nm, as well as on spherical mesoporous nanoparticles of 20–50 nm with pore size of 3 · 5 nm [50]. Lower enhancement on a per-Gd basis was reported for larger nanospheres (600–700 nm with pore size of 8 nm) [50]. The difference in relaxivities between various mesoporous nanostructures is attributed to the different accessibility of the water molecules to the Gd complexes, localized inside or outside the silica pores [50]. It is hypothesized that only Gd complexes grafted outside the nanopores contribute to the observed relaxivity, whereas the complexes grafted inside the nanopores do not modify the relaxation of water molecules. This may be due to a hindered or limited (slow) access of the water molecules to the Gd complexes inside the pores. As the localization of the Gd complex on the inner or outer surfaces of the mesoporous silica particles strongly depends on the pore size, the corresponding relaxivity is similarly affected by the pore size. Gd load per nano-object is also a key issue to increase the relaxivity on a per contrastobject basis. However, the relation between Gd load and relaxivity enhancement is not linear. Lin and co-workers [53] reported higher relaxivities for Gd load ranging from 1.6% to 3.1% wt. The r1 value reaches its maximum for 1.6% Gd load (23.6 mM−1 · s−1 at 9.4 T and 23 ◦ C) and decreases abruptly as the Gd loading increases from 2.3% to 3.1%, falling to relaxivity values close to free Gd-DTPA at 6.8% (4.4 versus 3.9 mM−1 · s−1 for GdDTPA at 9.4 T and 23 ◦ C). At such high Gd loading, the dipole–dipole interaction among the paramagnetic ions is presumed to be more significant, hence shortening the electronic relaxation time. Gd-containing mesoporous nanoparticles have been found to have potential application as blood-pool contrast agents [47], but also as tissue-targeting contrast agents and cell labeling agents. As the use of mesoporous structures as drug carriers has been demonstrated [54, 55], potential application combining drug and sensitive contrast agent targeting are envisaged. Postgrafting of the ligand of interest is possible thanks to a large number of amine groups on the surface of the mesoporous nanoparticles at the end of the synthesis. Tissue targeting is thus possible in theory with such structures; however, no reports of such application have been made to our knowledge. The presence of amine groups has mainly been used for the design of multifunctional mesoporous nanoparticles combining optical and magnetic properties. The fluorescent dye can be incorporated in the mesoporous nanostructures during synthesis, by adding a small amount of organic dyes conjugated to aminopropyltriethoxysilane [47]. Detailed description of this approach is given in Section 15.4.2. Another interesting approach that is developed (Section 15.4.2) is to combine the incorporation of Gd with other lanthanides such as Eu and Tb [53], which possess excellent

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luminescent properties in the visible range. This approach enables the design of multimodal contrast agents that combine high resolution and sensitivity of optical setup, with 3D localization provided by MRI. The use of mesoporous nanoparticles as cell markers has also been studied: these structures are readily internalized by many types of cells [46–48, 56], without the need of any transfection agent. Hsiao and colleagues have recently reported the use of hybrid mesoporous nanoparticles with Gd for stem cell tracking and were able to track as few as 4000 stem cells for 14 days using 1.5-T clinical MRI [49]. However, precise study on the effect of internalization on contrast properties has to be performed for long-term cell trafficking follow-up [57, 58]. 15.4 MULTIFUNCTIONAL NANOSTRUCTURES CONTAINING RARE EARTH ELEMENTS FOR MULTIMODAL IMAGING 15.4.1 Introduction The derivatization of the nanoparticle surface by gadolinium chelates constitutes an attractive route for the elaboration of highly sensitive MRI contrast agents. As already stated in the previous section, this strategy is expected to improve the enhancement of the positive contrast of MR images in comparison to isolated gadolinium chelate because of the large amount of paramagnetic centers per probe and of the reduction of the rotational motion. Since the great attention that nanoparticles received rests on their possibility to gather a large range of properties in a reduced volume, the immobilization of gadolinium chelates on nanoparticles increases the multifunctional character of the nanoparticles. Many works focused on the combination of MRI modality and fluorescent nanoparticles. The combination of fluorescence and magnetic resonance imaging in the same nano-object provides probes that are characterized by a high sensitivity (due to the fluorescence properties) and high resolution (owing to the specifications of MRI). Fluorescence of these paramagnetic nanoparticles can arise from the encapsulation of organic dyes (extrinsic fluorescence) or from intrinsic properties of the nanoparticles. Among the latter, quantum dots are probably the most widely known. Other combinations of imaging techniques can also be envisaged by functionalizing gold nanoparticles with gadolinium chelates since gold nanoparticles are able to induce a contrast enhancement in CT imaging. The association of several imaging modalities is very important because it ensures better reliability of observations and also allows monitoring of the contrast agents at various scales (from the subcellular scale to the whole organism) by different techniques [59–62]. Moreover, an increasing number of works have recently emphasized the therapeutic activity of the nanoparticles, which can be remotely controlled by an external physical stimulus ( e.g., magnetic field, visible and NIR light, X-ray photons, and neutron beam). In other words, harmless nanoparticles can be rendered cytotoxic after energy absorption [63, 64]. For optimal effect, follow-up of these therapeutic agents is required from administration to elimination from the body. For instance, the activation of these nanoparticles must be performed only when they are accumulated preferentially in the pathological zone. The possibility to image the fate of the nanoparticles is consequently a crucial issue. Since most of these nanoparticles are not detected by MRI, the functionalization of their surface by gadolinium chelates is highly expected. However, it depends on the chemical properties of the nanoparticle’s surface. For this reason, the composition of the ligand and/or the surface must be tuned for ensuring an efficient anchoring of the gadolinium chelates.

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15.4.2 Functionalization of Nanoparticles by Gadolinium(III) Chelates The functionalization of the nanoparticles aims at the synthesis of highly sensitive MRI contrast agents and/or the combination of several imaging techniques and/or the follow-up of therapeutic agents. Most of the particles developed for eventual biomedical applications are essentially composed of silica (actually polysiloxane), quantum dots, and gold nanoparticles. The functionalization of nanoparticles by gadolinium chelate depends on the chemical composition of the surface, which is in most cases imposed by the nature of the nanoparticles.

Fluorescent Silica Nanoparticles and Other Oxide Nanoparticles Silica nanoparticles can easily be produced by applying either the protocol developed by St¨ober or microemulsion techniques. St¨ober’s protocol consists of the alkaline hydrolysis– condensation of tetraethyl orthosilicate (TEOS) poured in a large volume of ethanol [65]. The high dilution and the aqueous NH4 OH to TEOS ratio are key parameters for yielding monodisperse silica spheres in a large range of sizes (from 20 to 800 nm). Control of the size, when microemulsion techniques are applied, is ensured by the volume of micelles, which behave as a nanoreactor [66]. Moreover, these synthesis protocols can be exploited for the encapsulation of metal oxide nanoparticles because of the great affinity between the surface of these particles and polysiloxane shell precursors. Besides their simplicity, both methods are very attractive because multifunctional nanoparticles can be obtained by minor variations of the composition of the polysiloxane precursors and/or by postfunctionalization. Indeed, a large variety of polysiloxane precursors are available for envisaging accurate control of the composition of the silica nanoparticles and for anticipating a functionalization before and/or after the hydrolysis–condensation. The production of multifunctional silica nanoparticles can be performed in several steps for the elaboration of multilayered nanostructures [67]. Each layer and the surface after derivatization contribute to the multifunctional character of the nanoparticles since they can each be designed for expressing a peculiar feature. For instance, aminopropyltriethoxysilane (APTES) was widely used in mixture with TEOS for introducing in silica nanoparticles, in each layer or on the surface, amino groups, which can act as a grafting site for the immobilization of organic dyes, hydrophilic molecules for improving the colloidal stability in aqueous biological media, and biotargeting groups for increasing the accumulation of the nanoparticles in the zone of interest within the organism. Van Blaaderen and co-workers demonstrated that the conjugation of fluorescein isothiocyanate to APTES before the hydrolysis–condensation allows the formation of fluorescent silica nanoparticles in which the organic dyes are distributed in the whole nanoparticles or in the whole thickness of the polysiloxane shell, whereas the postfunctionalization (i.e., after the hydrolysis–condensation) of the amino groups led to the tethering of the organic dyes onto the nanoparticles [67]. Wiesner and co-workers developed the synthesis of highly fluorescent nanoparticles, which are composed of an organic dye (rhodamine B isothiocyanate (RBITC), cyanine 5 NHS ester (Cy5-NHS)) loaded core and a protective shell [68]. The fluorescent cores were obtained by adapting St¨ober’s protocol: the hydrolysis–condensation was performed on a mixture of TEOS and organic dyes conjugated to APTES. In a second step, the protective shell was obtained by hydrolysis–condensation of a mixture of TEOS and APTES whose amino group makes possible the covalent immobilization of neutral polyethylene glycol chains. Such a postfunctionalization favors renal clearance of the nanoparticles and significantly reduces the liver uptake after in vivo injection of these

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highly fluorescent nanoprobes [69]. The microemulsion technique can also be exploited for the synthesis of fluorescent silica nanoprobes. The research group of W. Tan determined the experimental conditions for the production of fluorescent nanoparticles in water in oil (W/O) microemulsion (reverse microemulsion). Water droplets (micelles) are dispersed in the oil (cyclohexane) phase and stabilized by a surfactant layer (Triton X-100) for preventing the coalescence of droplets. The alkaline hydrolysis–condensation of TEOS and organic dye conjugated APTES takes place inside the micelles. This group demonstrated the versatility of this strategy since various organic and inorganic dyes can be covalently or not encapsulated in the amorphous silica matrix [66, 70]. They succeeded in tuning the color of the resulting nanoprobes and therefore in covering the visible spectral domain by embedding in the same nanoparticles several distinct organic dyes but with various ratios [71]. Moreover, these nanoparticles can be postfunctionalized. In other words, the surface of these fluorescent nanoparticles can be derivatized by the covalent immobilization of oligonucleotide for lowering the detection threshold of hybrization detection [70] or by the condensation of negatively charged polysiloxane precursors for improving colloidal stability [66]. This brief overview, which is naturally not exhaustive, aims at drawing attention to the possible ways of providing multifunctional silica nanoparticles. They were exploited for the synthesis of fluorescent and paramagnetic silica nanoparticles. Fluorescent, radio-opaque, and paramagnetic silica nanoparticles were synthesized by applying reverse microemulsion (water-in-oil system) techniques. Soluble ruthenium(II) based fluorophores are encapsulated in a silica core, which is further embedded in paramagnetic and aminated shells inside water droplets of the reverse microemulsion. The paramagnetic character of these nanoparticles was conferred by the capture of gadolinium(III) ions since the fluorescent core is covalently functionalized by N-(trimethoxysilylpropyl) ethyldiamine (TSPETE) [72]. The water solubility of these particles is ensured by the formation of an outer shell obtained by polymerization of TEOS, APTES, and 3-(trihydroxysilyl)propyl methylphosphonate (THPMP). The presence of negatively charged THPMP compensates the positive charge of APTES and generates the colloidal stability in water because the fucntionalization by THPMP provides a high negative charge to the surface of the particles. Moreover, the presence of amine functions, resulting from APTES, can be used for bioconjugation of the particles. Due to their composition, these particles behave as in vitro contrast agents for fluorescence imaging (Ru(bpy) dyes), MRI (TSPETE-Gd), and computed tomography imaging (Ru and Gd elements). However, these particles showed a lower radio-opacity than Omnipaque (iodine-based organic compound), a widely used contrast agent for clinical CT because Ru and Gd exhibit atomic numbers that are smaller than the atomic number of iodine. On the other hand, these particles exert a higher contrast enhancement for MRI than in the case of the commercially available MRI contrast agent Gadoteridol, (Gd-HP-DO3A. Unfortunately, no data about the stability of the complex TSPETE-Gd3+ was available although it is suspected to be relatively low because TSPETE molecules have only five reactive coordination sites. Despite these imperfections, the development of these particles is important because it opens the door to multimodal imaging applications based on gadolinium(III) chelate coated nanoparticles. Lin and colleagues performed the postfunctionalization of fluorescent silica nanoparticles by two different gadolinium chelates, which carry one (Gd-Si-DTTA) or two (Gd-Si-DTPA) propyltrimethoxysilyl moieties [51]. After encapsulation of a luminescent [Ru(bpy)3 ]Cl2 core (bpy = 2,2 -bypyridine) in an amorphous silica matrix, the surface of the resulting nanoparticles was derivatized by condensation of silanol (Si OH) groups present on the nanoparticles and the methoxysilyl groups of the gadolinium chelates. The authors

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noticed that the functionalization with Gd-Si-DTTA led to the formation of a monolayer of gadolinium chelates, whereas a thick paramagnetic multilayered shell was provided when Gd-Si-DTPA is used for the postfunctionalization of the fluorescent nanoparticles. The final size of the nanoparticles is around 37 and 40 nm, respectively. The covalent immobilization of these gadolinium chelates confers to the nanoparticles the ability to enhance the positive contrast of the MR images since they exhibit higher longitudinal relaxivity values per mM Gd3+ ions than the one for unbound gadolinium complexes. Moreover, each Gd-Si-DTTA chelate bound to the fluorescent silica nanoparticles exhibits higher relaxivities r1 and r2 than that of Gd-Si-DTPA grafted to similar nanoparticles (r1 = 19.7 s−1 · mM−1 Gd3+ and r2 = 60 s−1 · mM−1 Gd3+ vs. r1 = 7.8 s−1 · mM−1 Gd3+ and r2 = 12.3 mM−1 · s−1 Gd3+ ). This can be reasonably explained by a higher water exchange rate since gadolinium chelates Gd-Si-DTTA coordinate two water molecules whereas Gd-Si-DTPA chelates coordinate only one water molecule. In addition, another feature of nanoparticles coated by Gd-Si-DTTA plays in favor of a higher longitudinal relaxivity. Since the surface functionalization of the fluorescent silica nanoparticles by Gd-Si-DTTA yields a monolayered shell in contrast to the derivatization of the nanoparticles by Gd-Si-DTPA, gadolinium(III) ions are therefore more accessible. But the silica nanoparticles coated by Gd-Si-DTPA exhibit a higher longitudinal relaxivity r1 when the comparison is based on the amount of particulate contrast agents since each particle coated by Gd-Si-DTPA contains more gadolinium(III) ions. It must be pointed out that these nanoparticles have an impressive longitudinal relaxivity whatever the coating (r1 = 2.0 ×105 s−1 · mM−1 and r2 = 6.1 × 105 s−1 · mM−1 vs. r1 = 4.9 × 105 s−1 · mM−1 and r2 = 7.8 × 105 s−1 · mM−1 ). The great potential of these luminescent and paramagnetic nanoparticles was confirmed by in vitro experiments, carried out on monocyte cells with Gd-Si-DTTA coated nanoparticles. Their uptake by these cells was successfully and efficiently monitored by fluorescence imaging, flow cytometric experiments and MRI (T1 and T2-weighted images). This study reveals that the internalization of the nanoparticles in the monocyte cells does not alter the viability and that the fluorescent and paramagnetic nanoparticles allow multimodal in vitro imaging. These authors explored another strategy, for developing particulate contrast agents for imaging cells by fluorescence and MRI. From fluorescent nanoparticles coated by Gd-SiDTTA (see above), they applied a layer-by-layer (LbL) techniques for immobilizing a higher amount of gadolinium chelates [52]. Since Gd-Si-DTTA coated nanoparticles are highly anionic, a layer of cationic Gd-DOTA oligomers was deposited onto the negatively charged surface via electrostatic interactions. As a result, the particles exhibit a positive charge that allows the further deposition of anionic polystyrenesulfonate (PSS). The LbL assembly strategy, which consists in successive depositions of cationic Gd-DOTA oligomers and anionic PSS, yields luminescent and paramagnetic nanoparticles with alternate multilayer coatings of positively charged Gd-DOTA oligomers and negatively charged PSS. It was demonstrated that the size of the multilayered nanoparticles linearly increases with the number of depositions (from 37 nm (for a single layer) to 43 nm (for seven layers)). A similar evolution is observed for the relaxivities r1 and r2 of the particulate contrast agents because the alternate deposition of an additional cationic layer led to an increase in the amount of gadolinium(III) chelates for each nanoparticle. The per particle r1 values increase from 1.94×105 mM−1 · s−1 for nanoparticles coated by a single Gd-Si-DTTA layer to 5.34 × 105 mM−1 · s−1 for Gd-Si-DTTA coated nanoparticles after the deposition of seven layers (four of cationic Gd-DOTA oligomers and three of anionic PSS), whereas the per particle r2 values increase from 5.61×105 mM−1 · s−1 to 1.55×106 mM−1 · s−1 . The LbL self-assembly thus offers an interesting strategy for increasing nanoparticle MR relaxivities. It also affords the

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opportunity to design these particulate contrast agents for the targeting of cancerous cells, which is highly recommended for the early detection of tumors. Since LbL self-assembled nanoparticles are terminated with anionic PSS polymers, a biotargeting peptide sequence containing arginine–glycine–aspartate (RGD) and seven consecutive lysine (K) residues (K7RGD) can be immobilized onto the surface through electrostatic interaction between positively charged lysine residues of the K7RGD sequence and negatively charged PSS layer. The RGD sequence is well known for its strong interaction with ␣v ␤3 integrins that are upregulated in growing endothelial and angiogenic cancer cells. Such a functionalization led to the efficient targeting of human colon cancer (HT-29) cells and calf pulmonary artery endothelial (CPAE) cells as revealed by in vitro experiments monitored by fluorescence imaging and MRI. Although these results are very promising, the in vivo application of these contrast agents is not yet validated. Thanks to the trimethoxysilyl group, which allows its grafting onto the oxide surface, Si-Gd-DTTA can be used successfully for rendering mesoporous silica nanoparticles paramagnetic [47]. As described Section 15.3.2, such a postfunctionalization confers to the mesoporous nanoparticles very high longitudinal and transversal relaxivities r1 and r2 , which amounts to 7.0 × 105 mM−1 · s−1 (3 T)/2.48 × 105 mM−1 · s−1 (9.4 T) and 1.6 × 106 mM−1 · s−1 (3 T)/2.7 × 106 mM−1 · s−1 (9.4 T) on a per millimolar particle basis, respectively. These relaxivity values are much larger than the solid silica nanoparticles that are coated with multilayers of the Gd-DTPA derivative (see above [51]). The enhanced MR relaxivity was attributed to the ready access of water molecules through the nanochannels of the mesoporous particles derivatized by gadolinium chelates Si-Gd-DTTA. When the functionalization is performed by a mixture of Si-Gd-DTTA and of rhodamine B conjugated to aminopropyltriethoxysilane (4 mol %, relative to Gd-Si-DTTA), the resulting mesoporous nanoparticles can behave as contrast agents for fluorescence and MR imaging. The uptake of these fluorescent and paramagnetic mesoporous nanoparticles by monocyte cells was revealed both by the fluorescence and the contrast enhancement MR images. Moreover, they appeared well suited for in vivo MRI after intravenous injection in the tail vein since aorta and liver are clearly visible on T1- and T2-weighted images, respectively. These nanoparticles are indeed able to induce a positive and a negative contrast. Even if hepatic uptake renders liver imaging possible thanks to the phagocytosis of the nanoparticles by the resident liver macrophage cells (Kupfer cells), it however leads to a quick clearance from the blood and limits the use of these nanoparticles as in vivo contrast agents. In addition to gadolinium, luminescent rare earth ions (Tb3+ , Eu3+ , etc.) are coordinated by the same ligands and therefore can be immobilized onto the nanoparticles. The postfunctionalization of a fluorescent silica core that contains Ru(bpy) as the fluorophore with silylated polyaminocarboxylate ligands allows the immobilization of both Tb3+ and Gd3+ ions. The presence of Tb3+ confers to the particles emission properties that are complementary to those of Ru2+ entrapped inside the core, whereas gadolinium ions are expected to render the nanoparticles paramagnetic [73]. The silylated gadolinium chelates can also be successfully applied for the functionalization of metal oxide nanoparticles since they are covered by OH groups, which act as anchoring sites. Frias and co-workers reported on the modification of boehmite (AlO(OH)) nanoparticles with silylated gadolinium(III) chelates (derivatives of DO3A) [74]. These particles are therefore able to enhance the positive contrast of MR images. As expected, these ligands, which are covalently bound to the nanoparticle surface through the condensation between methoxy groups of silylated ligands and OH groups of the surface, are able to capture luminescent lanthanide ions (Eu3+ and Tb3+ ).

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FIGURE 15.4 Schematic illustration of the procedure for the synthesis of the silica encapsulated oxide nanoparticles. The DTPA or DOTA was attached to the amine groups of the surface through an amide reaction.

The postfunctionalization of fluorescent silica nanoparticles by silylated gadolinium(III) chelates is very efficient, as demonstrated by the work of Lin’s research group. Since most of the clinically approved gadolinium(III) chelates are based on polyaminocarboxylate motif, they can easily be immobilized on aminated nanoparticles through the formation of amide linkage (Fig. 15.4). Amine functions in nanoparticles are generally introduced by the inorganic copolymerization of TEOS and aminated polysiloxane precursors (APP like APTES or APTMS) or by the derivatization of oxide nanoparticles by APP. Delville and co-workers succeeded in converting alumina and silica nanoparticles into fluorescent and paramagnetic contrast agents [75]. Such a tuning was performed after the coating of the bare nanoparticles with a polysiloxane shell yielded by the hydrolysis–condensation of APTMS. Such an encapsulation allows tethering organic dyes (rhodamine isothiocyanate) and polyaminocarboxylate ligand (DTPA derivative) for the immobilization of gadolinium(III) ions onto the surface of the nanoparticles. DTPA ligand was grafted to the surface either by the reaction between aminated surface and bis-anhydride derivative of DTPA or by the peptide reaction promoted by the carbodiimide EDC in the presence of NHS between aminated surface and DTPA. Despite a low amount of gadolinium chelates immobilized on the nanoparticles (4 gadolinium ions per nanoparticle), a strong positive and negative contrast enhancement is observed. This increase in relaxivity is consistent with the increase in rotational correlation time, which results from both the increase in molecular weight and the maximum compactness of the SiO2 structure as compared to organic molecules. Indeed, the rigid nature of the inorganic nanoparticle tends to avoid the Gd3+ chelate from freely rotating or even folding back. Since the propensity of nanoparticles to be internalized in cells is higher than in the case of molecules, the immobilization of the gadolinium chelates onto the nanoparticles is attractive because an efficient labeling of microglial cells by paramagnetic agents is expected. The uptake of these particles by the cells was confirmed by the strong contrast enhancement and fluorescence. Microglial cells are well known to be

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chemoattracted and strongly accumulate particularly in brain pathological situations, including tumors. These cells can therefore behave as cargo for freighting a large amount of fluorescent and paramagnetic nanoparticles in the zone of interest. As a result, labeled cells can be followed up from the injection site to the target, which will be clearly delineated thanks to the presence of a large amount of contrast agents. This type of detection will open the route for more efficient contrast agents for MR cell tracking. The aforementioned examples obviously show that the encapsulation of nanoparticles in a polysiloxane shell is an efficient route for rendering the nanoparticles paramagnetic through the immobilization of gadolinium ions either by the use of silylated polyaminocarboxylate ligands or by the conjugation of polyaminocarboxylate ligands with amine functions that are present on the surface of the polysiloxane shell if aminated precursors are used. However, this method requires strict control of the experimental conditions to avoid bulk polymerization of the polysiloxane precursors. The surface modification of transition metal oxide nanoparticles can also be performed by an alternative method that relies on the grafting of gadolinium(III) chelates onto naked nanoparticles. These ligands are terminated by an anchoring function, which ensures their immobilization onto the nanoparticles since they are designed for establishing a strong interaction with the surface. The selective reactivity of the surface of TiO2 nanoparticles to ortho-substituted enediols (such as dopamine) was exploited for grafting gadolinium(III) chelates [76]. Prior to the derivatization of the TiO2 nanoparticles, dopamine modified MR contrast agents (DOPA-DO3A) were synthesized by coupling commercially available succinimidyl ester activated 1,4,7,10-tetraazacyclododecane-1,4,7-tetraacetic acid to dopamine hydrochloride. The gadolinium(III) ions were captured by the resulting ligands. As expected, these particles are able to exert a contrast enhancement of MR images since they exhibit a longitudinal relaxivity of 3.5 mM−1 · s−1 per gadolinium ion, which is similar to the value of clinically used small molecule contrast agents. Since this method of functionalization leads to the immobilization of a single monolayer of gadolinium(III) chelates, the longitudinal relaxivity per particles remains moderate (61 mM−1 · s−1 ) in comparison to the shell obtained from silylated ligands (up to 7.0×105 mM−1 · s−1 , see above [47, 51]). In addition to dopamine modified contrast agents, these nanoparticles can also be functionalized by DNA oligonucleotides (DNA-DOPA-DO3A-Gd) for targeted imaging and therapy. DNADOPA-DO3A-Gd nanoparticles were successfully applied for the labelling PC3M cells, which were revealed by two-dimensional X-ray fluorescence maps and by T1-weighted MR images. Since phosphonic acid groups have been found to be very effective in anchoring organic molecules to various metal oxide surfaces, gadolinium(III) complexes of polyaminocarboxylate ligands containing a phosphonate group in a pendant arm were used for the surface modification of TiO2 nanoparticles [77].

Paramagnetic Quantum Dots (QDs) The functionalization of QDs by gadolinium chelates represents another route for the elaboration of fluorescent positive contrast agents combining fluorescence imaging and MRI. QDs refer to a crystallite cluster of II–VI semiconductors (CdS, CdSe, ZnS, CdSe@ZnS, etc.) exhibiting unique optical properties that arise from quantum confinement effects characteristic of these small dimensions (1–8 nm) [78–80]. Their attractive features (high photostability, high quantum yield, controllable and narrow emission band, broad excitation band) should contribute to a breakthrough in fluorescence imaging even if their routine application is still controversial (owing to the toxicity of Cd2+ ions) [81, 82]. Their relatively narrow emission band whose position

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is controlled either by the particle size or, for a given size, by tuning the composition of the alloyed semiconducting QD (CdSeTe), together with their broad absorption spectra offer a large palette of distinct color enabling multilabeling [83]. Since several synthetic routes have been developed for the preparation of water-soluble QDs with narrow size distribution and high quantum yields, the literature is overflowing with examples of biological materials labeled by QDs [84, 85]. Indeed, the most common synthetic routes of QDs lead to hydrophobic fluorescent nanoparticles since they are coated by trioctylphosphine/ trioctylphosphine oxide (TOP/TOPO). The biological application of QDs requires their functionalization by capping hydrophilic molecules. Besides the water solubility, the functionalization plays two supplementary important roles: capping ligands (1) can serve as anchoring sites for grafting molecules and (2) can ensure, by passivating the QD surface, their protection and consequently the preservation of optical properties. A representative list of caps and the QD dispersal strategies they use is provided in a review written by the research group of Mattoussi [86]. These strategies can be grouped into three major routes (Fig. 15.5). The first one consists in the exchange of a the hydrophobic monolayer of TOP/TOPO by hydrophilic thiol or phosphine terminated ligands (mercaptocarbonic acids, alkylthiol terminated DNA, thioalkylated oligo-ethyleneglycols). This functionalization is driven by mass action and is possible because sulfur or phosphorus atoms can establish strong dative bonds with the chalcogenide species present at the surface of QDs. However, it must be pointed out that the colloidal and chemical stability is considerably improved when hydrophilic ligands used for replacing TOP/TOPO carry at least two thiol or phosphine functions (dihydrolipoic acid derivatives, oligomeric phosphines). The second strategy described in the literature is based on the encapsulation of QDs in a silica shell. As mentioned earlier, silica shells afford multiple ways for further surface modification. The first step of the formation of the silica shell lies in the replacement of TOP/TOPO by mercaptopropyltrimethoxysilane (MPTMS) whose thiol end allows the immobilization onto the QD. The growth of the silica shell is therefore generated by hydrolysis–condensation of a polysiloxane precursor mixture (TEOS, APTES, MPTMS, etc.) from the alkoxysilyl groups of mercaptopropyltrimethoxysilane bound to the QD surface. The third method preserves the native TOP/TOPO on the QDs and uses variants of amphiphilic “diblock” and “triblock” copolymers and phospholipids to tightly interleave/interdigitate the alkylphosphine ligands through hydrophobic attraction, whereas the hydrophilic outer block permits aqueous dispersion and further derivitization. These three strategies (Fig. 15.5) were applied for derivatizing fluorescent QDs with gadolinium chelates. To add the MRI contrast ability, QDs can be functionalized with Gd3+ DOTA (1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid) complexes as shown by the study of Jin et al. [87]. For achieving this, the TOP/TOPO layer was first replaced by glutathione, which is composed of a thiol function for the immobilization of the ligand onto the NIR emitting QD, an amine group for further functionalization, and two carboxylic moieties for water compatibility and also for further functionalization. The addition of DOTA NHS ester to glutathione coated QDs led to the covalent grafting of DOTA through the formation of amide bond. In a last functionalization step, QDs are made paramagnetic by the capture of gadolinium(III) ions from DOTA (Gd3+ -DOTA-QD). The functionalization induces an increase of the hydrodynamic diameter from 7 to 10 nm. This strategy seems to be efficient since Gd3+ -DOTA-QD exhibits a relatively high longitudinal relaxivity r1 (365 mM−1 · s−1 ). Moreover, these paramagnetic QDs have a great potential for in vivo imaging since the implantation of a tube filled with Gd3+ -DOTA-QD into the abdomen of a mouse can be detected by NIR-fluorescence imaging and by MRI. Unfortunately, no data

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NH2

FIGURE 15.5 Schematic of generic QD functionalization.

are given about the intravenous injection of these nanoparticles and their biodistribution, which is a key parameter for the biomedical application of these QDs. Gerion et al. [88] demonstrated that the functionalization of QDs by gadolinium(III) chelates can be performed, as expected, after their encapsulation in a silica shell. The shell was obtained in two steps. After a priming step replacing trioctylphosphine oxide (TOPO) surfactants on the QD surface with mercaptopropyltrimethoxysilane (MPTMS), polymerization of siloxane was performed in methanol under slightly basic conditions. In a second step, addition of fresh MPTMS and polyethylene glycol(PEG)–propyltrimethoxysilane permitted the introduction of functional (SH) and stabilizing PEG groups on the surface of the SiO2 shell. As reported in the section devoted to fluorescent silica nanoparticles, ligands used for the immobilization of paramagnetic ions can easily be tethered to the polysiloxane shell provided that functional groups are present. In order to preserve the eight coordination sites of DOTA, an aromatic amine was grafted to one of the ethylene bridges of DOTA (NH2 -Bn-DOTA, Fig. 15.3). DOTA derivative with paramagnetic Gd3+ chelates are linked to the thiols of the silica shell embedding the QDs by using the sulfo-SMCC crosslinker. The maleimide group of sulfo-SMCC allows the grafting to the thiol function due to the formation of thioether linkage while the NHS ester group at the opposite end of the crosslinker favors the formation of amide linkage with the aromatic amine of the DOTA derivative. As compared to unbound Gd-DOTA, the Gd3+ ion relaxivity r1 for the gadolinium chelates attached to the silica-coated nanoparticles is increased approximately sixfold (23 mM−1 · s−1 vs. 3.7 mM−1 · s−1 ). Chemical analyses revealed that about 50 Gd-DOTA are grafted onto the nanoparticles. As a result, the paramagnetic QDs (8 nm) exhibit a relatively high longitudinal relaxivity (1019 mM−1 · s−1 , 1.4 T). Preliminary in vivo tests indicate that paramagnetic silica-coated

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nanoparticles after intravenous injection provide a contrast enhancement in MRI. The accumulation of paramagnetic nanoparticles in the bladder was evidenced by the appearance of brighter zones in the bladder 5 min after the injection. However, these nanoparticles do not accumulate in the other organs since no contrast enhancement was observed elsewhere. This indicates that the nanoparticles freely circulate without undesirable accumulation and most of them are quickly removed by renal excretion. Both previous works led to the development of paramagnetic QDs suitable for MRI applications. Their functionalization by gadolinium(III) chelates rests on replacement of the hydrophobic TOP/TOPO layer by a hydrophilic one. However, the TOP/TOPO layer can be exploited for a further functionalization of QDs as demonstrated by the studies performed by Bakalova and co-workers and by Mulder and co-workers [89–92]. The preservation of the TOP/TOPO layer limits the deleterious effect of polar species on the luminescence since the hydrophobic layer constitutes an efficient barrier. In contrast to the procedure of Gerion et al. [88], which requires the replacement of TOP/TOPO layer by MPTMS, Bakalova et al. [89] demonstrated that QDs can be encapsulated by a silica shell from the hydrophobic layer. Prior to the inorganic polymerization of the polysiloxane network, hydrophobic QDs were covered by vinyltrimethoxysilane. The hydrophobic tail of this precursor (i.e., the vinyl moiety) was entrapped in the hydrophobic TOP/TOPO layer, whereas the methoxysilyl end group pointed toward the outside. These methoxysilyl groups serve as anchoring sites for the polysiloxane shell yielded by hydrolysis–condensation of a mixture of TEOS and APTES. Owing to the amine function, which is a characteristic of the latter, the surface is ready for surface modification. The MRI contrast agent ability was conferred to the QD either by the entrapment of amphilic gadolinium(III) chelates (common name: Resolve Al–Gd) before the encapsulation in the silica shell or by the surface modification of the silica-coated QDs by hydrophilic gadolinium(III) chelates, which are DOTA derivatives (Gd-DOTA-Bn-NCS) (Fig. 15.3). The isothiocyanatobenzyl group ( CH2
N C S) of Gd-DOTA-Bn-NCS carried by one of the ethylene bridges of the macrocycle ensures the covalent grafting of the gadolinium chelates by condensation with the surface amine function leading to the formation of a thio-urea linkage (Fig. 15.2). Whatever method is used to render them paramagnetic, silica-coated QDs are able to induce the shortening of the longitudinal relaxation time T1. This was reflected by the enhancement of the contrast of T1-weighted images. In vitro and in vivo studies showed that they can be applied as probes for multimodal imaging. Indeed, the internalization of these particles can be monitored by both fluorescence imaging and MRI. It was demonstrated that silica-coated QDs do not alter the viability of labeled cells. Furthermore, labeled cells under continuous UV or laser irradiation remain viable, whereas it generates a cytotoxic effect when cells are loaded with uncoated QDs. The silica shell seems to suppress the photosensitization effect, which was previously reported for uncoated QDs. The low toxicity of these contrast agents is encouraging for in vivo applications. However, accumulation in the spleen must be solved since this reflects the uptake by macrophages, which is commonly observed when positively charged particles are injected intravenously. This capture should be overcome by PEGylation. The encapsulation of TOP/TOPO-coated QDs by the silica shell was also performed via a reverse microemulsion technique [90]. But in this case, the silica shell was formed between the QDs and the TOP/TOPO layer. The postfunctionalization of the resulting particles for rendering them paramagnetic and hydrophilic was realized according to the protocol developed by Mulder’s research group for TOP/TOPO-coated QDs [93].

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Hydrophobic TOP/TOPO-coated QDs can be rendered hydrophilic by coating them with amphiphilic PEGylated phospholipids (PEG lipids: PEG-DSPE (1,2-distearoyl-sn-glycero3-phosphoethanolamine-N-[methoxy-(polyethylene glycol)-2000]). These molecules are widely used to stabilize liposomes for pharmaceutical applications because PEG chains form a hydrophilic layer on the liposomal surface. The coating of hydrophobic QDs (or QDs encapsulated in a silica shell) with a mixture of PEG lipids and of paramagnetic lipids (i.e., gadolinium(III) chelate conjugated to hydrophobic stearylamide) yields watersoluble and fluorescent positive contrast agents for fluorescence imaging and MRI. The postfunctionalization of hydrophobic silica-coated QDs with PEG lipids markedly improves the biodistribution and the pharmacokinetics of these contrast agents [94]. Moreover, some PEG lipids are terminated by a maleimide moiety, which allows bioconjugation. The research group of Mulder succeeded in functionalizing these nanoprobes by RGD, and annexin A5 proteins [91,92]. The grafting of RGD, which is well known for targeting ␣v ␤5 and ␣v ␤3 integrins, favors the uptake of these paramagnetic QDs by HUVEC due to the overexpression of ␣v ␤3 integrins. The ability of these contrast agents to target ␣v ␤5 and ␣v ␤3 integrins opens the door to the detection of neoangiogenosis, which is associated with the development of tumors, whereas the conjugation of Annexin A5 to these nanoprobes makes them well suited for the detection of apoptosis by fluorescence imaging and MRI. The strategy of the Mulder research group, which is based on the synthesis of lipid-based nanoparticles, is very attractive because it leads to a large variety of contrast agents for multimodality imaging [93]. By a convenient choice of the building blocks (the lipids, the nanoparticles, the biotargeting groups, etc.) the behavior of these platforms can be accurately tuned for a definite application. Recently, they demonstrated that endogenous nanoparticles (high-density lipoproteins, HDLs) can also be modified for multimodality imaging by incorporating paramagnetic and/or fluorescent lipids but also by including within each HDL nanoparticle a hydrophobic inorganic core such as iron oxide, QDs, or gold [27]. Since it is known to migrate to atherosclerotic plaques and remove cholesterol from macrophages, HDL was exploited by Frias and co-workers to develop a probe for molecular imaging of macrophages [28, 29]. The inclusion of an inorganic core affords additional features for imaging purposes. For instance, the presence of a hydrophobic gold core in a paramagnetic and fluorescent HDL allows monitoring gold–HDL by MRI, fluorescence imaging, and X-ray imaging since the high atomic number of the gold element confers to the gold nanoparticles a great ability to absorb X-ray photons.

Paramagnetic Gold Nanoparticles One of the most promising applications of nanosized gold objects lies in their ability to combine imaging and therapy [95]. Two main strategies were developed for designing gold particles that can be used for imaging and destroying tumors. The first one results from the properties of the gold element, whereas the second one rests on the tunable optical properties of gold particles. Gold particles are indeed characterized by strong light diffusion and absorption cross sections, which are sensitive to the size, shape, and dielectric environment [96]. By tuning these parameters, gold-based nano-objects (gold-coated silica particles and gold nanorods) can be synthesized and exploited for NIR imaging and therapy as demonstrated by Halas and co-workers and El-Sayed and co-workers [97, 98]. Whatever the size, shape, and environment of the gold nano-objects, they strongly absorb X-ray photons. Hainfeld et al. [99, 100] demonstrated that ultrasmall hydrophilic gold nanoparticles (1.9 nm) can be monitored by X-ray imaging after intravenous injection to mice and can be applied for radiotherapy. However, X-ray imaging is not completely harmless for envisaging a long-term biodistribution

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study without interference with radiotherapy. Moreover, this technique suffers from a low sensitivity. In order to circumvent this problem, the use of gold nanoparticles designed for magnetic resonance imaging was proposed [101, 102]. Roux and co-workers succeeded in designing gold nanoparticles for MRI by reducing gold salt in the presence of a dithiolated ligand DTDTPA (Au@DTDTPA). As a result, 2.4-nm sized gold cores embedded in a DTDTPA shell were obtained. DTDTPA, is a dithiolated derivative of DTPA, which is commonly used as a gadolinium chelate (DTPA-Gd) for clinical MRI. The transformation of DTPA into DTDTPA was performed by the condensation of a DTPA bis-anhydride molecule with two aminoethanethiol molecules. It was demonstrated that each gold core is encapsulated in a multilayered shell, which is composed of about 150 DTDTPA molecules. The tethering and formation of this multilayered shell are actually ensured by the presence of two thiol functions per DTDTPA. Since the thiol function exhibits a great affinity for gold, some DTDTPA molecules present in the shell are immobilized onto gold nanoparticles through the strong interaction between gold and sulfur atoms of one thiol function. Thiol functions that were not involved in the grafting of the shell onto the nanoparticles were implied in the formation of disulfide bonds. These bonds make possible the formation of a multilayered shell around each gold core. Due to the presence of the DTDTPA shell, 150 Gd3+ ions can be immobilized on a gold nanoparticle (Au@DTDTPA-Gd150 ). However, a good compromise between colloidal stability and the amount of active ions for MRI was found when the number of gadolinium(III) ions per particle is fixed at 50 (Au@DTDTPAGd50 ). The behavior of Au@DTDTPA-Gd50 in the bloodstream after intravenous injection can easily be monitored by both X-ray imaging (in transmission and computed tomography modes) and MRI (r1 = 3.9 mM−1 · s−1 based on gadolinium concentration and r1 = 195 mM−1 · s−1 based on nanoparticle concentration). Biodistribution studies performed with both medical imaging techniques and with postmortem elemental analysis of the organs revealed that these paramagnetic gold nanoparticles freely circulate in the bloodstream and do not accumulate in liver, spleen, and lungs. In the case of tumor bearing rats (or mice), it must be pointed out, however, that accumulation is observed in the tumor due to the EPR effect. This phenomenon was exploited for radiotherapy [103]. It was demonstrated that the survival of diseased rats was lengthened only when they were treated by X-ray beam after intravenous injection of Au@DTDTPA-Gd50 . Owing to its imaging and therapeutic potentialities, Au@DTDTPA-Gd50 seems to be suited for in vivo medical applications. Gold nanoparticles coated by rare earth chelates were also produced for development of a fluorescent sensor since some rare earth ions are characterized by a very photostable luminescence, which is distinguished from organic dyes and QDs by narrower emission peaks and longer lifetime of the excited state. Gunlaugsson and co-workers demonstrated that the luminescence of europium buried in a heptadendate ligand (a tetraazamacrocycle) can be modulated by exploiting the deleterious quenching effect of water on rare earth luminescence [104]. The luminescence is switched off when two water molecules are present in the coordination sphere, whereas the displacement of water by a diketone derivative bearing an aromatic ring reestablishes the light, which is even enhanced due to the effect of the antenna from the aromatic ring. This principle developed for modulating fluorescence can be applied for “smart” MRI. When water is present in the coordination sphere, these gold particles can exert an enhancement of the positive contrast of MR images, whereas no enhancement is visible when water was displaced by analytes. MRI contrast agents that are sensitive to the presence of definite biomolecules can therefore be imagined.

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Gold nanoparticles can also be rendered paramagnetic by the covalent grafting of gadolinium(III) chelates to biomolecules, which are then immobilized on the particles for targeting them to the zone of interest. For instance, Chung and co-workers applied this strategy for preparing paramagnetic gold nanostructures, which are able to combine dual modality imaging (MRI and optical imaging) and targeted therapy [105]. First, hollow gold nanoparticles coated by thiol ended PEG and cystein-protein-G were synthesized. Each class of molecules grafted to the hollow nanoparticles ensures a specific role. PEG was used for improving the colloidal stability and therefore the particle blood circulation time and the efficiency of particle internalization, while protein G molecules were introduced to optimize the orientation of antibodies (anti-HER2) on the surface of the hollow gold nanoparticles. Meanwhile anti-HER2 antibodies were derivatized by gadolinium chelate thanks to the covalent coupling between the amino group of the antibody and the isothiocyanate group of p-SCN-Bn-DTPA (Figs. 15.2 and 15.3). After chelation of gadolinium(III) ions by DTPA conjugated to anti-HER2, the paramagnetic antibody (anti-HER2-Gd-DTPA) was immobilized onto hollow gold nanoparticles. The resulting particles were successfully applied as both MRI (r1 = 23.7 mM−1 · s−1 ) and optical imaging contrast enhancers of cancer cells since they can be observed by optical microscopy (due to the optical properties of hollow gold nanoparticles) and by MRI. The presence of anti-HER2 on the nanoparticles favors the internalization in SKBR3 cells (breast cancer cells that overexpress epidermal growth factor receptors targeted by anti-HER2). Labeled cells that can therefore be visualized by dark-field-based scattering images and MRI were destroyed after NIR irradiation due to the photothermal effect, which occurs when NIR light is absorbed by the gold nanostructures. The multifunctional hollow gold nanoparticles are also expected to be used for imaging specific cancer sites using MRI, and then for destroying cancer cells selectively via illumination with NIR light. However, their large size should impede in vivo application of these paramagnetic therapeutic agents. For this reason, data regarding their biodistribution and their pharmacokinetic parameters, which are not available, are required. 15.4.3 Gadolinium(III) Containing Crystalline Nanoparticles

Introduction The development of gadolinium(III)-based nanoparticles, which is discussed in this chapter, aims at affording positive contrast agents for MRI with a stronger ability of contrast enhancement (due to both the increase of gadolinium(III) ion number and the decrease of rotational motion), better pharmacokinetic parameters, a better control of biodistribution, and a larger palette of properties than molecular gadolinium complexes provide. To achieve this goal, Gd-DTPA or Gd-DOTA functionalized polymer [106, 107], selfassembled peptide amphiphiles [108,109] or viral capsid [110], Gd-DTPA terminated dendrimer [111, 112], gadolinium complex loaded liposomes [113], high-density lipoprotein nanoparticles [28, 29], micelles [114, 115] or polymeric nanoparticles [116, 117], gadolinium ions entrapped in zeolites [118], fullerenes [35], carbon nanotubes [39], clays [119] or mesoporous silica nanoparticles [53], and gadolinium chelates immobilized on quantum dots [91, 92], on lipid particles [120], and on gold nanoparticles [101, 102] were synthesized and studied. They are all characterized by an increase of the molecular weight and of the amount of Gd(III) ions per contrast agent. As a consequence of their structure, some of them were easily functionalized by biotargeting groups and/or fluorescent molecules, conferring on them additional attractive features [28, 29, 91, 92, 106, 110–113, 120]. Although the number of gadolinium atoms is very high even for the smaller size, the potential of crystalline nanoparticles based on inorganic gadolinium such as gadolinium oxide [121–123],

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gadolinium fluoride nanoparticles [124], and gadolinium carbonate [125, 126] has only recently been evaluated. Pioneering studies devoted to gadolinium oxide nanoparticles were performed by Roberts and Watkins. They revealed the potential of gadolinium oxide nanoparticles to be applied as contrast agents for MRI. However, these preliminary works suffered from the availability of well-defined hydrophilic gadolinium oxide nanoparticles. The exploitation of the attractive characteristics of gadolinium oxide nanoparticles for MRI became conceivable only when a reproducible and efficient synthesis protocol of naked gadolinium oxide nanoparticles and their functionalization were developed.

Synthesis and Functionalization of Gadolinium Oxide Nanoparticles The research activity of Tillement’s group was decisive in this field. First, a reproducible and reliable synthesis of gadolinium oxide nanoparticles was reported by this group [127, 128]. These particles were synthesized by applying, with some modifications, the polyol routes, which opened the door to the synthesis of a large range of inorganic crystalline nanoparticles. The synthesis of gadolinium oxide nanoparticles is based on the alkaline hydrolysis of gadolinium chloride in diethylene glycol (DEG), a high boiling point diol. Diethylene glycol plays an important role. It allows carrying out the synthesis at 180 ◦ C and to accurately control the size of gadolinium oxide nanoparticles (between 1 and 5 nm according to the experimental conditions) because the high viscosity of DEG and its adsorption on growing particles limit their growth and avoid agglomeration. The alkaline hydrolysis of a mixture of gadolinium chloride and of chloride salt of luminescent rare earth ions (Tb3+ , Eu3+ , Nd3+ , etc.) in boiling DEG provides rare earth (RE) doped gadolinium oxide nanoparticles whose Gd to RE molar ratio is very similar to the initial ratio [127, 128]. Since gadolinium oxide is transparent in the visible range of light, it was demonstrated that these nanoparticles are fluorescent when rare earth ions are present in the crystalline nanoparticles. Despite the remarkable optical properties of luminescent rare earth ions (high photostability, narrow emission peaks, long lifetime of excited state), the in vivo application of luminescent doped gadolinium oxide nanoparticles is impeded by the low quantum yield and by the energy of the excitation wavelength (UV range). However, the in vivo application of gadolinium oxide nanoparticles remains very promising for MRI provided that the surface of the gadolinium oxide core is adapted to a biological environment. The second important contribution of Tillement’s research group lies in the functionalization of these cores. The strategy developed by this group rests on the encapsulation of the gadolinium oxide cores in a polysiloxane shell [129]. This functionalization mode provides several important advantages. Besides a protective effect of the core, which was validated by the preservation of the luminescence in acidic media when the cores are doped by Tb3+ ions, the polysiloxane shell exerts a positive influence on the luminescence of rare earth ions present in the cores since its presence induces the replacement of RE OH bonds, which behaves as quenchers for the luminescence by the RE O Si bonds and an energy transfer from the polysiloxane shell to the luminescent centers [130]. The encapsulation of luminescent rare earth doped gadolinium oxide cores therefore yielded nano-objects that are more fluorescent than the uncoated gadolinium oxide cores are. But the most striking feature of the polysiloxane shell results from its chemical nature, which can provide new properties to the particles. The polysiloxane shell was indeed obtained by hydrolysis–condensation of tetraethyl orthosilicate (TEOS) and by aminopropyltriethoxysilane (APTES). The use of APTES allows the sequential functionalization of the polysiloxane shell: before and after the hydrolysis–condensation step, that is, before and after the formation of the polysiloxane shell because each APTES

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carries an amino (NH2 ) function. Before the hydrolysis–condensation, APTES can be conjugated to organic dyes thanks to the coupling between the amino group of APTES and the isothiocyanate or NHS ester group of the organic fluorophores, while the polysiloxane shell can be postfunctionalized by the covalent grafting of hydrophilic PEG chains [131, 132] and/or of biotargeting groups [129, 133]. The encapsulation of gadolinium oxide cores enlarges the range of properties of these particles because the inner part of the polysiloxane shell is functionalized by organic dyes and the outer part is derivatized by hydrophilic molecules. As expected, these gadolinium oxide nanoparticles encapsulated in a fluorescent and PEGylated polysiloxane shell (GadoSiPEG) can be followed up after intravenous injection to mice and rats by fluorescence imaging (due to cyanine 5 covalently bound to the polysiloxane network) and by MRI (thanks to the presence of gadolinium(III) in the core of each particle). The in vivo imaging experiments revealed that the particles exhibit attractive biodistribution and pharmacokinetics parameters that are accurately controlled by the length of the PEG chains and the nature of the end group [131, 132]. They freely circulate in the bloodstream without undesirable nonspecific accumulation in liver, spleen, and lungs. Moreover, these nanoparticles are rather quickly removed from the body essentially by renal excretion. All these observations denote a safe behavior of GadoSiPEG nanoparticles when they are intravenously injected into small animals. However, the accumulation of these nanoparticles was observed after intravenous injection into gliosarcoma (9L) bearing rats. This was revealed by the obvious delineation of tumor due to the enhancement of the positive contrast induced by the presence of the particles in the tumor region. This accumulation, which can be explained by the EPR effect, was exploited for radiotherapy since the atomic number of the gadolinium element is sufficiently high for generating a dose enhancement of the X-ray beam. The treatment of brain tumor bearing rats by radiotherapy led to a longer survival only when hybrid gadolinium oxide nanoparticles were administered by intravenous injection before exposure to a therapeutic X-ray beam [134, 135]. The therapeutic activity of gadolinium oxide nanoparticles is not limited to radiotherapy. As widely mentioned in the literature, objects containing gadolinium could possibly replace the boron based compounds for neutron capture therapy. The internalization of GadoSiPEG nanoparticles in bioluminescent murine lymphoma cells (EL4-luc), which was monitored by fluorescence (rhodamine B isothiocyanate was covalently bound to the polysiloxane shell), MRI, and elemental analysis (ICP), generates no cytotoxicity and no alteration of the cell proliferation until [Gd]incubation = 0.3 mM. Unloaded cells are also not affected by exposure to a thermal neutron beam if the delivered dose is equal to 3 Gy (or less). But the irradiation of EL4-luc cells after incubation in the presence of GadoSiPEG ([Gd]incubation = 0.05 mM) by harmless thermal neutron beam (3 Gy) generates a great killing effect since all cells are destroyed [136]. Gadolinium oxide cores embedded in a fluorescent and PEGylated polysiloxane shell are very attractive multifunctional nanostructures for in vivo application and, in particular, for imaging guided therapy since they combine fluorescence imaging (organic dyes conjugated to the inner part of the polysiloxane shell), MRI (gadolinium(III) ions in the crystalline core), and therapeutic activity, which exploits some properties of the gadolinium element (radiotherapy and neutron therapy). The potential of gadolinium oxide cores for MRI was confirmed thanks to studies performed by a research group from Link¨oping University (Sweden) [121]. Their works is very interesting and complementary to the work done by Tillement’s group. Indeed, they propose an alternative way for derivatizing the gadolinium oxide cores and therefore for exploiting their paramagnetic character in MRI experiments. Their strategy consists in the capping of

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gadolinium oxide nanoparticles, which were synthesized according to the polyol route [127, 128], by organic acids or by the grafting of PEG-silane molecules. After their synthesis in DEG, the resulting gadolinium oxide cores are coated by DEG molecules, which can be replaced in hot DEG by citric acid or dimercaptosuccinic acid [137, 138]. This layer of organic acid molecules acts as a primer for further functionalization and, in particular, for the PEGylation, which is required for biological applications. The grafting of nonimmunogenic and nonantigenic PEG is expected to render the gadolinium oxide cores more robust, to improve the colloidal stability in biological media, to favor the uptake by cells, and to enhance blood retention. Citric-acid-coated gadolinium oxide cores were derivatized by PEG chains ended by a thiol (SH) group and an amino (NH2 ) group (Gd2 O3 @CA-PEG). The amino group of PEG permits condensation with carboxylic acid groups of citric acid, which were activated by EDC/NHS mixture prior to the coupling reaction. The thiol group at the opposite extremity can be used for the grafting of molecules bearing a maleimide group [3]. Since each dimercaptosuccinimide (DMSA) molecule grafted to gadolinium oxide cores carries two thiol functions, Mal-PEG-NHS (Mal and NHS for maleimide and N-hydroxysuccinimide groups, respectively) was used for the PEGylation of DMSA capped gadolinium oxide cores. The covalent coupling of PEG is ensured by the formation of a thioether linkage resulting from the condensation of the thiol group of DMSA with the maleimide moiety of Mal-PEG-NHS (Gd2 O3 @DMSA-PEG). The presence of the NHS group affords the possibility to graft aminated molecules (typically biomolecules that are rich in amine functions) onto the PEGylated nanoparticles. The PEGylation can also be performed using PEG-silane [139] since the alkoxysilyl end group of PEG-silane favors its grafting on oxide nanoparticles through condensation between the alkoxysilyl groups and hydroxyl (OH) groups present on the surface of the oxide materials (Gd2 O3 @PEG). These PEGylated oxide nanoparticles, which have been prepared by the research group from Link¨oping University, exhibit, as expected, relatively high longitudinal relaxivity r1 . The relaxivity r1 values of Gd2 O3 @PEG, Gd2 O3 @DMSA-PEG, and Gd2 O3 @CAPEG (based on the amount of gadolinium(III) ions in each sample) were 12.0, 14.2, and 7.8 mM−1 · s−1 , respectively [140]. Except for Gd2 O3 @CA-PEG, these PEGylated gadolinium oxide nanoparticles exhibit a higher relaxivity than that of GadoSiPEG developed by Tillement’s research group (see above). The longitudinal relaxivities of GadoSiPEG nanoparticles containing the smallest gadolinium oxide cores (2.2 nm and 3.8 nm) were 8.8 mM−1 · s−1 whatever the core size whereas r1 dropped to 4.4 mM−1 · s−1 when the size of the core amounted to 4.6 nm [131]. This difference can be explained by a better accessibility of water molecules to the surface of the oxide cores in the case of Gd2 O3 @PEG and Gd2 O3 @DMSA-PEG because the cores are encapsulated in a monolayer, whereas the gadolinium oxide cores of GadoSiPEG are embedded in a relatively thick polysiloxane layer. However, the cause of this difference is probably not so obvious since Gd2 O3 @CAPEG particles exhibit a lower relaxivity than GadoSiPEG particles do despite the thin PEG shell of Gd2 O3 @CA-PEG. It must be pointed out, however, that two different strategies for the functionalization of gadolinium oxide cores are available. They provide similar performance in terms of positive contrast enhancement for MRI. Unfortunately, their behavior after intravenous injection in small animals (rats, mice) cannot be compared because no in vivo imaging experiment was reported with the PEGylated organic acid capped gadolinium oxide cores in contrast to GadoSiPEG. Although the encapsulation of the gadolinium oxide cores by polysiloxane shell seems more difficult to control, it nevertheless allows enlarging more easily the palette of properties of the contrast agents since postfunctionalization of gadolinium oxide cores can be realized sequentially, in contrast to the encapsulation by

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organic acids. Besides MR imaging, which is efficient whatever the functionalization mode, GadoSiPEG nanoparticles can indeed be easily followed up by fluorescence imaging after their uptake by cells or after their intravenous injection into small animals [131, 132, 136, 141], whereas the detection of PEGylated organic acid capped gadolinium oxide nanoparticles by fluorescence imaging rests only on the doping of the core by rare earth ions [140]. Despite its interesting properties, the fluorescence imaging based on the luminescence of rare earth ions (Tb3+ ) is not very well suited for biological applications because the light emission of Tb3+ in gadolinium oxide is induced by the absorption of UV photons, which are harmful for living matter. Moreover, the UV excitation is not suited for fluorescence imaging of small animals since the UV irradiation is almost totally absorbed by the skin. Because of the great chemical similarities of the rare earth elements, the synthesis of gadolinium oxide nanoparticles by applying the polyol route was easily adapted to the synthesis of dysprosium oxide nanoparticles (diameter: 1 nm). It was demonstrated that these particles exhibit a high transverse relaxivity r2 that is dependent on the size [142]. The optimal size for the highest r2 value is between 60 and 70 nm: r2 varies between 190 mM−1 · s−1 (70 nm at 7 T) and 675 mM−1 · s−1 (60 nm at 17.6 T). Other routes for the preparation of gadolinium oxide based contrast agents have recently been developed. Gadolinium oxide nanoparticles in carbon nanotubes were synthesized for MRI application [143]. They were obtained after thermal treatment of gadolinium acetate entrapped in single-wall carbon nanohorns (SWNHs) with holes (NHox). Gd2 O3 -NHox dispersed in agarose gel induces a significant positive contrast enhancement as revealed by T1-weighted images of phantoms. However, no r1 value was given, probably because the variation of 1/T1 as a function of gadolinium concentration is not linear. Although the enhancement induced by Gd2 O3 -NHox appears sufficient for MRI, their application as an in vivo contrast agent seems to be limited due to the lack of colloidal stability in biological fluid and the toxicity of carbon nanotubes. Yeh and co-workers succeeded in synthesizing porous and hollow gadolinium oxide particles [144]. They demonstrated that hollow spheres produced better positive contrast enhancement (r1 = 17.7 mM−1 · s−1 and r2 /r1 = 1.50) than porous particles did (r1 = 16.8 mM−1 · s−1 and r2 /r1 = 1.87). From a chemical point of view the elaboration of such materials is very interesting; however, their application as contrast agents for MRI can be impeded by the too large size of these particles. The sizes of hollow Gd2 O3 spheres (∼130 nm) and of porous Gd2 O3 particles (∼187 nm) are obviously larger than the hydrodynamic diameter recommended for the renal excretion of the nanoparticles in the bloodstream (Dh < 5 nm) according to the studies by Frangioni and co-workers [84].

Synthesis of Gadolinium Containing Crystalline Nanoparticles At the same time, several routes were explored for obtaining alternative crystalline nanomaterials rich in gadolinium(III) by wet chemistry techniques. The potential of these gadolinium containing crystalline nanoparticles for MRI was naturally evaluated since there is a real need for efficient multifunctional contrast agents. Among the variety of materials, fluoride and carbonate particles catch the attention. Fluoride Nanoparticles Containing Gadolinium(III) Van Veggel and colleagues acquired great experience in the synthesis of lanthanide fluoride nanoparticles for luminescence application. They paid particular attention to the synthesis and characterization of gadolinium fluoride. Synthesis of the nanoparticles was carried out for 3–4 hours in water at moderate temperature (75 ◦ C) and rests on the reaction of ammonium hydroxide

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(NH4 OH) with sodium fluoride and gadolinium(III) nitrate (Gd(NO3 )3 · 6H2 O) in the presence of a hydrophilic stabilizing agent (citric acid or 2-aminoethylphosphate (AEP)). As a result, relatively large gadolinium fluoride nanoparticles capped by citric acid or AEP (GdF3 @cit (mean diameter 129.3 nm) and GdF3 @AEP (mean diameter 51.5 nm), respectively) were collected [124]. The presence of hydrophilic agents on the surface of the particles facilitates their dispersion in aqueous solution and prevents agglomeration. Moreover, citric acid and especially AEP allow the postfunctionalization of the nanoparticles by a biotargeting group (biotin) and by PEG chains [145]. Whatever the capping ligands, GdF3 nanoparticles exhibit moderate longitudinal relaxivities (3.17 and 2.71 mM−1 · s−1 per Gd3+ for GdF3 @cit and GdF3 @AEP, respectively) when it was expressed in terms of the number of gadolinium ions in the nanoparticles. But the relaxivities per mM nanoparticles reach high values that amount to 2.0 × 107 and 8.8 × 105 s−1 · mM−1 for the GdF3 @cit and GdF3 @AEP nanoparticles, respectively. This study suggests that Gd3+ ions located beneath the outer shell of Gd3+ are also contribute to the relaxation. This was confirmed by work focused on GadoSiPEG, which showed that the effect on r1 not only results from the Gd3+ ions present at the surface of the crystalline core but also partially from those inside the core. Indeed, the relaxivities r1 are identical for 2.2- and 3.8-nm Gd2 O3 cores, whereas the smallest particles possess the highest amount of surface Gd3+ ions for the same gadolinium concentration. It means that in both samples all Gd3+ ions have the same influence whatever their location in the crystalline structure and that the range of the influence of Gd3+ ions is at least 1.9 nm. NaYF4 nanocrystals constitute another interesting class of fluoride materials since they can easily be doped by various rare earth ions. These fluoride nanocrystals received much attention for photoluminescence applications because they are considered the most efficient host materials for upconversion phosphors. They exhibit visible emission upon IR excitation when doped with Yb3+ and Er3+ /Tm3+ ions. By an appropriate choice of rare earth ions incorporated in the fluoride matrix, the emission wavelength can be red-shifted relative to the excitation wavelength. Downconversion is therefore achieved and can be exploited for biological imaging. For instance, Chatterjee et al. [146] have recently demonstrated the use of upconversion nanophosphors (NaYF4 : Er3+ , Yb3+ ) for in vitro imaging of cancer cells and in vivo imaging in tissues. It is claimed that the advantages of these nanoprobes include absence of photodamage to living organisms, very low autofluorescence, high detection sensitivity, and high light penetration depth in biological tissues. If many studies devoted to the photoluminescence of rare earth doped NaYF4 nanocrystals were performed, data about the doping of these crystals by gadolinium remains rare despite the great potential of contrast agents with multimodality imaging. Prasad and co-workers contributed to the development of paramagnetic fluoride nanophosphors of size <50 nm. They reported the synthesis of functionalized fluoride nanocrystals, which form a stable aqueous dispersion that can be used for both optical imaging and MRI. Down- and upconverted luminescence from the rare earth ions (Eu3+ , Er3+ , and Yb3+ ) doped into nanometer-sized fluoride matrices allows optical imaging modality, and co-doping with Gd3+ confers to the resulting nanoparticles the ability to enhance the positive contrast of MR images. By a convenient surface derivatization (immobilization of antibodies (anti-claudin and anti-mesothelin) via the covalent coupling with mercaptopropionic acid anchored onto the nanoparticles), these nanoparticles can be selectively taken up by pancreatic cancer cells. The selective internalization in cells was demonstrated only by the light emitted by the nanoparticles (using a confocal microscope). Since the longitudinal relaxivity r1 value of these Gd3+ co-doped fluoride nanocrystals is very low (r1 = 0.14 mM−1 · s−1 and r2 /r1 = 62  1), in comparison to gadolinium oxide

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or gadolinium fluoride (GdF3 ) nanoparticles, the uptake of these nanoprobes cannot be monitored by MRI. The low value of r1 observed in the case of Gd3+ co-doped NaYF4 can be explained by the amount of gadolinium(III) ions in the crystalline fluoride matrix, which is largely lower than in the case of gadolinium oxide or gadolinium fluoride (GdF3 ) nanoparticles. If a small amount of luminescent rare earth ions is sufficient for detection by optical imaging techniques, this study showed clearly that a greater amount of paramagnetic ions is required for inducing positive contrast enhancement in MRI because this medical imaging is less sensitive. Nevertheless, MRI remains a more attractive in vivo imaging technique than fluorescence imaging techniques because the resolution of MRI is better. Moreover, a three-dimensional cartography of living bodies can be achieved by MRI, whereas it is impossible with fluorescence imaging techniques owing to the high absorption of biological tissues.

Carbonate Particles Containing Gadolinium(III) By refluxing aqueous solution containing urea and gadolinium chloride, amorphous gadolinium carbonate (Gd2 O(CO3 )2 · H2 O) particles can be obtained [126]. The shape (spherical, rhombus-, or rice-shaped nanoparticles) and the size can be tuned by varying the urea to gadolinium chloride molar ratio. Relaxation rate measurements revealed that the relaxivities depend on the size and on the shape of the particles. In the case of 500-nm sized spherical Gd2 O(CO3 )2 · H2 O particles, r1 is equal to the 16.5 mM−1 · s−1 (r2 /r1 = 12,7, 3 T). After injection, the behavior of these large particles can be monitored by MRI. As expected for large particles, accumulation in the liver was observed. These particles can be functionalized by an aminated polysiloxane shell thanks to the hydrolysis–condensation of the aforementioned mixture of TEOS and APTES. This polysiloxane shell permits a further functionalization. Due to the presence of the amino groups on the nanoparticles, gold nanoparticles can be immobilized since nitrogen atoms have a great affinity for gold atoms. These gold nanoparticles act as seeds for the growth of a gold layer. Gd2 O(CO3 )2 · H2 O particles can also be encapsulated by a gold shell whose thickness can be accurately controlled [125]. The encapsulation of these gadolinium carbonate particles by a gold shell induces a decrease of the longitudinal relaxivity with the increase of the thickness of the gold shell. However, the gold shell renders the resulting nanoparticles suitable for photothermal therapy [97]. GdO(CO3 )2 · H2 O particles encapsulated in a gold shell are able to combine MRI and therapy as revealed by in vitro experiments. However, the exploitation of this attractive feature requires the reduction of the size of this class of materials.

15.5 CONCLUSION AND OUTCOME Although many outstanding results have been achieved recently, the use of nanoparticles containing rare earth ion as MRI contrast agents is still in its infancy. Many of the studies are currently dealing with in vitro analysis or in vivo “proof-of-principles.” Deeper in vivo studies on relevant animal models and clinical trials will certainly give rise to many optimizations in the design, stability, and contrast enhancement of this new class of contrast agents. In addition to active and passive targeting for molecular imaging, many rare earth ion containing nanoparticles combining both imaging and therapeutic capacities are under

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development. These multifunctional nanoparticles open new possibilities for “personalized nanomedicine,” where both diagnosis and treatment can be best fitted to every single patient. As the size range of currently designed nano-objects is in the window of the vascular system (>20 nm), they are mainly targeting vascular related modifications such as angiogenesis associated with tumor progression or atherosclerosis. With the development of smaller nanoparticles containing rare earth ions (<20 nm) that can extravasate from the endothelium to the interstitial space, tumor cell targeting can also be envisioned. Different strategies are currently under investigation to combine active targeting with therapy: (1) entrapping pharmaceutical compounds within the nano-objects to enable selective drug delivery, (2) taking advantage of unique physical properties of these nano-objects to perform selective therapy, or (3) combining both. As the use of lanthanide-based nanoparticles is relatively recent, their use in multifunctional nanoplatforms has undergone little investigation so far. Nevertheless, recently reported work is highly promising. For example, successful attempts to administer drugs locally have been reported for atherosclerosis treatment in a rabbit model using perfluorocarbon nanoparticles (175–220 nm) loaded with Gd ions as the contrast agent and fumagillin as the drug [148]. Other highly promising strategies with minimum side effects rely on the exploitation of the unique physical properties of these nanosized carriers for efficient and selective radiotherapy, phototherapy, or neutron therapy. Zielhuis and colleagues [149] have successfully designed liposomes (130 nm) combining paramagnetic gadolinium ions and radionuclide rare earth ions holmium-166 (beta- and gamma-emitter with highly paramagnetic properties) for both imaging and radiotherapy. Several recent works have also demonstrated the possibility to combine imaging and phototherapy, using rare earth containing nanoparticles with absorbing nanomaterials such as gold [105, 125] or to a lesser extent carbon [150]. The principle relies in the conversion of absorbed light into local heat strong enough to damage nanoparticle containing cells. This is possible thanks to the strong light-absorption properties of nanomaterials (in particular, gold nanomaterials), known as surface plasmon resonance (SPR). The nano-objects are classically designed to tune the SPR peak in the near-infrared wavelength region, where the depth penetration into tissues is higher. The advantages of such a technique are minimum invasivity and an increase in selectivity, as damages are only done in the vicinity of the nano-objects, thus reducing severe injury of healthy tissue. So far, reported studies have used relatively large nanoparticles (50–450 nm [105, 125, 150]), limiting their use for local injection. There is no doubt, however, that these pioneering works will lead to many innovative nano-objects more suitable for cell targeting after intravenous injection in clinical applications. Another attractive feature of gold lies in its high atomic number. As a result, gold absorbs very strongly the X-ray photons. Gold nanoparticles therefore appear promising for use in radiotherapy since their presence in tumor permits the absorption of almost all X-ray photons. Consequently, healthy tissue is more preserved while the tumoricide effect the of X-ray beam is enhanced [99, 100, 103]. Last but not least, neutron therapy appears as an interesting alternative technique that has hardly been exploited. The principle relies on the irradiation of compounds with a high neutron capture cross section (neutrophages) by a thermal neutron beam. The irradiation by a thermal neutron beam or the presence of neutrophage compounds is harmless by itself. Selective destruction of surrounding cells can be achieved by combining nanoparticle targeting with thermal neutron beam irradiation. Most therapeutic protocols are now based on

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molecular boron compounds [151]. However, different groups have reported the possibility to take advantage of ␥ -ray emission and Auger electrons of gadolinium isotopes (155 Gd and 157 Gd) upon interaction with a thermal neutron [136, 152–155]. In particular, the in vitro study of Bridot and colleagues showed the therapeutic potential of hybrid nanoparticles combining high contrast properties and long blood half-life [131] and a size (7 nm in diameter) compatible with partial extravasation from the vasculature and cell targeting. Multifunctional rare earth based nanoparticles are far from clinical use, but different routes to achieve that goal are under investigation by many multidisciplinary joint research programs, starting from chemical design and synthesis to in vitro and in vivo toxicity and therapeutic efficacy analysis. The aim is to be able to get a diagnosis as early as possible and to treat the pathology with high efficacy and limited side effects. Rare earth based multifunctional nanoplatforms are one possible route to achieve high contrast properties (and thus lower the detection sensitivity threshold) as well as high specificity toward the targeted pathological sites. The near future will certainly see the design of many innovative rare earth based multifunctional nanoplatforms for early diagnosis and treatment, most presumably in oncology and atherosclerosis diseases.

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CHAPTER 16

Microfabricated Multispectral MRI Contrast Agents GARY ZABOW and ALAN KORETSKY Laboratory of Functional and Molecular Imaging, National Institute of Neurological Disorders and Stroke, National Institutes of Health, Bethesda, Maryland, USA

16.1 INTRODUCTION Since its inception in 1973 [1, 2], magnetic resonance imaging (MRI) has rapidly become a widely used medical diagnostic imaging technique. Primary advantages of MRI include its nonionizing nature and its excellent soft tissue contrast and high spatial resolution compared to other radiological imaging methods [3]. The relatively low inherent sensitivity of MRI is offset by being able to use generally abundant water as the main molecule for detection. The ability of water to report on a wide range of normal and pathophysiological conditions due to alterations in MRI detectable properties has proved remarkable. These include changes in the relaxation times T1 and T2 that are modified by many processes such as changes in local water content and in endogenous paramagnetic metal content such as iron. Techniques that sensitize MRI to the diffusion of water [4–6], to blood flow [7], and to oxygenation of hemoglobin [8] have dramatically expanded MRI detection of tissue function. In addition to detecting such intrinsic properties, MRI contrast can be modulated by addition of exogenous contrast agents. Most important among the MRI contrast agents are the small molecule chelates of the paramagnetic ion Gd3+ . They are used in a wide variety of applications, such as in detecting leakage of the blood–brain barrier that enables, for example, ready detection of brain tumors [9] and of plaques associated with multiple sclerosis [10]. Another class of MRI contrast agent is made from nanosized dextran-coated iron-oxide particles that are potent relaxation agents and that have been used to measure pathology in liver [11]. Manipulation of endogenous contrast and use of added contrast agents has perhaps made MRI the most versatile radiological imaging tool. The success of MRI for imaging anatomical and physiological/functional changes has led to growing interest in extending MRI for cellular and molecular imaging. Although the low sensitivity of magnetic resonance would initially make MRI seem an unlikely candidate for molecular imaging, the ability of contrast agents to alter the MRI properties of large numbers of surrounding water molecules enables indirect detection via the abundant Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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water signal. Because the surrounding water is continually undergoing random diffusion, and because the magnetic field disturbances caused by the contrast agents can often extend appreciable distances from the contrast agent itself, the agents can interact with surrounding volumes of water many times greater than their own physical sizes, resulting in large signal amplifications. Even relatively small quantities of contrast agent are therefore detectable, rendering MRI well suited to molecular and cellular level imaging despite any inherent low NMR sensitivity. For example, concentrations of the paramagnetic ion Mn2+ as low as 2 ␮M from voxels as small as 100 ␮m has been detected via shortening of T1 of water in MRI [12]. This represents a total amount of Mn2+ of 2 femtomoles. MRI detection of individual micrometer-sized superparamagnetic particles has also been demonstrated [13–15] even at imaging resolutions that were 50–100 times larger than the sizes of the particles used. This is due to the large magnetic field disturbances that these superparamagnetic particles generate in the surrounding water. Such superparamagnetic particles have been used to detect single cells in vitro [16–18] and in vivo [19] and to monitor the movement of such cells [20, 21]. Over the past decade, the use of MRI for cell labeling and tracking has enjoyed considerable success [22, 23], as have other MRI-based molecular and cellular imaging approaches. Examples include the possibility of early stage disease detection through particular cell-type or epitope recognition [24–26] and real-time in vivo imaging of gene expression and enzyme activity [27–31]. Also possible with MR techniques is the detection of various in vivo molecular binding events and interactions [32], enabling, for example, real-time analyte concentration measurement [33] thanks to changes in magnetic relaxivities induced by the spatial clustering of magnetic nanoparticle contrast agents [34]. All of this development of molecular- and cellular-based MRI depends on altering MRI contrast with appropriate agents. Thanks to the great success of the Gd chelates for clinical MRI, much research has gone into the development of numerous contrast agents, be they small molecules, macromolecules, or nanoparticles, and many excellent reviews have already been written on their synthesis, characterization, and use [22, 35–45]. All MRI contrast agents developed to date, however, have been chemically synthesized. Recently, it has been demonstrated that top–down microfabrication techniques can be used to generate novel properties for MRI contrast agents [46–48]. This work relied on fabricating specifically shaped microstructures from magnetic materials to control the NMR resonance of water in the structures in ways that make them uniquely identifiable and that open the prospect of large-scale multiplexing with MRI. The goal of this chapter is to describe this new approach to MRI contrast agents. First, however, a brief description of the main classes of chemically synthesized MRI agents is presented to lend perspective to the microfabricated agents.

16.2 BOTTOM–UP CHEMICALLY SYNTHESIZED CONTRAST AGENTS Broadly classified, existing chemically synthesized MRI contrast agents can be thought of as either molecular or particulate agents. The molecular agents are generally based on chelates of some paramagnetic metal ion, often one from the lanthanide series with a large number of unpaired electrons including, most notably, the above-mentioned gadolinium, Gd3+ . These chelates are typically used as T1 agents, requiring direct contact with water in the first or second coordination sphere of the metal. At high concentrations, however, they can also cause appreciable T2 * effects, forming the basis of using Gd chelates to monitor blood flow in the brain [49]. Over the past decade, clever chemistry has been used to control

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water access to the Gd3+ ion by relevant biological processes such as changes in calcium concentration [50] or specific enzyme activity [51]. This work has opened the possibility of using Gd3+ chelates to report on molecular and cellular events. There has also been an increased interest in Mn2+ as an MRI agent. This interest has arisen because Mn2+ goes intracellular and can be used to make unique images of anatomy as well as to monitor a number of interesting cellular properties such as calcium influx and connectivity between neurons [52, 53]. The particulate agents are generally based on various magnetic materials including, most commonly, superparamagnetic iron oxide (SPIO). Unlike the molecular-sized lanthanide chelates, the particulate agents’ linear sizes span over three orders of magnitude. At the smaller end are individual nanocrystals a few to several nanometers in size that form the core of so-called monocrystalline iron-oxide nanoparticles (MIONs) [54] and ultrasmall SPIO (USPIO) particles [55]. These cores are coated, most commonly with dextran, to ensure solubility and biocompatibility. The total hydrodynamic diameters of core and coat are typically in the tens of nanometer range. Various agglomerations of these individual iron oxide nanocrystals have been produced. For example, they have been encapsulated within dendrimer structures that are again a few tens of nanometer in size and known as magnetodendrimers [56], or collected into larger composite SPIO nanoparticles that commonly range from about 50 nm to 300 nm in diameter [57, 58]. At the larger end of the size scale are micrometer-sized particles of iron oxide (MPIOs) [15] with diameters of around 1 ␮m to as large as about 10 ␮m. Although these iron-oxide-based nano- and microparticles can appreciably reduce both T1 and T2 relaxation times, they are most commonly used as potent T2 * contrast agents, which tend to lead to substantial localized image darkening or reduced signal intensity. T2 * effects are caused by dephasing of magnetization due to the particles’ inhomogeneous magnetic fields. Therefore the potency of these agents is based on the size of the magnetic moment generated, which depends on the size and composition of the particles. Since MRI can detect very small perturbations in magnetic field (fractions of a part per million) these particles affect MRI signal over ranges much larger than the size of the particles themselves. Because of this amplification effect, only on the order of a million superparamagnetic nanoparticles are required per voxel for MRI detection, while at the microscale, as mentioned above, even individual microparticles can be detected. This high sensitivity to such particles has enabled MRI to detect targeting of specific receptors on specific cells [59] and is behind the already mentioned large range of work using these particles to label cells for monitoring cell migration with MRI. This success of iron-oxide particles has also encouraged interest in generating nanoparticles of other compositions including manganese oxide [60] and gadolinium oxide [61]. The flexibility of MRI has also led to a number of interesting techniques that let the T2 * properties of particles give signal enhancement rather than signal loss [62, 63]. And the distinction between paramagnetic T1 agents and particulate T2 * agents is blurring even further with proposals to let particulate contrast agents degrade under cellular conditions to give rise to paramagnetic agents [64]. Besides T1 and T2 agents, an exciting new class of MRI contrast agents has been developed in recent years known collectively as chemical-exchange saturation transfer (CEST) agents [65, 66]. These agents are based on off-resonance frequency shifts that indirectly generate contrast through the exchange of free unshifted protons of the surrounding bulk water with various bound—and thus chemically shifted—water, amide, amine, or hydroxyl protons of the CEST molecules. The beauty of such off-resonant agents is that they may be selectively addressable or detectable through their different chemical shifts, allowing a form of multispectral or “color” contrast rather than the traditional “monochrome” brightening or

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darkening of T1 and T2 agents. Recently, it has been demonstrated that certain paramagnetic chelates, so called PARACEST agents, can also be used, with the paramagnetic ion leading to larger chemical shifts of the exchangeable protons, increasing the spectral bandwidth achievable [41, 67]. CEST principles have recently also been applied to hyperpolarized xenon agents [68] and to lipid-based nanoparticles [40], including liposome and micelle structures that each aggregate together many CEST molecules [69] to effectively increase contrast per composite particle.

16.3 TOP–DOWN FABRICATED MULTISPECTRAL CONTRAST AGENTS Compared to the above chemically synthesized forms of contrast agents, top–down microfabricated contrast agents [47] represent a departure, being designed around techniques more commonly associated with the microchip industry—see, for example, Madou [70]. Borrowing more from the microengineer’s, than from the traditional chemist’s toolkit, they are most obviously different in terms of the physical forms they can assume and the fabrication methods used in their production. The question then is whether microfabrication, or perhaps even some combination of microfabrication together with bottom–up chemical synthesis (e.g., such as in the design of nucleic acid chips), can offer new and useful properties for MRI contrast agents. Among the many differences between top–down microfabrication and bottom–up chemical synthesis approaches, two in particular bear mention here. First, having been relentlessly honed and enhanced over decades by the microelectronics industry, the micro- and nanoscale spatial patterning capabilities of top–down fabrication approaches allow greater control and flexibility in the design of structure geometries than do purely bottom–up syntheses. Second, being often more physical than chemical in nature, certain microfabrication processes can be relatively independent of material choice, allowing structure compositions that can be more accurately controlled and more easily interchanged. Consequently, top–down fabrication has the potential for creating structures boasting substantially greater degrees of consistency in size, shape, and material composition. Additionally, the increased flexibility in material choices may allow fabricated contrast agents to be more biologically compatible than, for example, those based on lanthanide ions whose toxicity demands chelation and associated loss in agent efficiency [71, 72]. Traded off against these advantages, however, are also relative disadvantages related to minimum structure sizes and fabrication efficiencies and costs. Although direct write serial patterning techniques such an electron beam and scanned probe writing overcome most size limitations by offering true nanoscale definition, they are less efficient than parallel photolithographic processes. And while the electronics industry continues to find new ways to overcome inherent optical diffraction limitations, such parallel top–down fabrication processes are, at least for the time being, still not well suited to structure fabrication at the very smallest of nanoscale sizes that are accessible to bulk phase synthesis. Additionally, even given a massively parallel patterning technique and the economies of scale that accompany such wafer-level processing, the two-dimensional planar nature of top–down fabrication means that such fabrication generally remains less efficient and offers lower throughput than bottom–up chemical synthesis routes. Finally, top–down fabrication may be more demanding in terms of process complexity and cost, often requiring specialized equipment that renders it less accessible to traditional biology/chemistry lab settings. To justify this added complexity and expense, therefore, the top–down fabricated

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micro- and nanostructures should offer some extra functionality not readily achievable through traditional chemical synthesis methods. The main subject of this chapter presents one example of that, namely, exploiting microfabrication’s control over structure geometry to encode distinct spectral signatures into MRI agents. Note that these multispectral agents are almost certainly not the only form of novel contrast agent that could potentially be microfabricated; they just happen to be the first demonstrated example of one potential new direction opened up by top–down micro- and nanofabricated structures. 16.3.1 Contrast Agent Geometry While the use of ferromagnetic particles as MRI contrast agents was first proposed more than two decades ago [73, 74], interest has always focused primarily on particle composition, size, and clustering, with little attention paid to the geometry of the individual micro- or nanoparticles. For two reasons this is not surprising. First, from a chemical synthesis/selfassembly/magnetic clustering point of view, full control over the structure geometry is generally harder to realize than is control simply over overall structure size. Second, with the exception of a few positive-contrast schemes [62, 63], such agents have thus far primarily been used as negative contrast T2 * agents, designed simply to dephase the surrounding transverse proton magnetization. This makes sense since, with typical MRI voxels measuring hundreds of micrometers or more on a side, virtually all of the dephased hypointense signal results from water lying within the contrast agent nanoparticles’ farfields, that is, at distances away from the particle that are large compared to the particle’s size. With these far-fields being always dipolar in nature, the spatial contrast signal then naturally depends only on overall particle magnetization (i.e., on particle volume and composition) and is essentially independent of structure geometry. Although particle geometry may not directly affect spatial contrast, it has recently been shown [46] that, under certain conditions, the geometry can provide distinct spectrally shifted NMR frequency signals. As such, top–down fabricated ferro- or superparamagnetic agents can be microengineered to simultaneously provide both regular SPIO-type T2 * spatial contrast, as well as distinguishing spectral contrast. Viewed another way, the implication of the above is that by suitably controlling the local distribution of exactly the same quantity of magnetic material that would otherwise have been indiscriminately aggregated together into a traditional chemically synthesized nanoparticle, top–down microfabricated agents’ near-fields can be exploited to provide additional differentiating spectral contrast while their far-fields continue to provide exactly the same spatial contrast as before. As an introductory example, the scanning electron micrograph in Figure 16.1 compares a typical chemically synthesized magnetic microparticle to that of a microparticle whose geometrical structure has been engineered using surface micromachining techniques to produce magnetic field profiles suitable for the above-mentioned spectrally shifted NMR signals. Positioned in front of a standard, commercially available 4.5-␮m diameter magnetic bead (commonly used in biotechnology research as a separating agent, and in MRI as an MPIO T2 * agent), the image shows a sample microfabricated particle, approximately 2 ␮m in size, formed from two slabs of magnetic material separated by a nonmagnetic post. This particular microstructure is, of course, not intended to represent the only geometry that can be microfabricated, nor the only geometry that can produce distinct NMR spectral shifts; it is intended only as a suggestive visual introduction to some of the differences between typical chemically synthesized structures and those that can be produced with the extra geometrical structuring that top–down fabrication techniques enable.

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FIGURE 16.1 Top–down microfabrication versus bottom–up chemical synthesis. Scanning electron micrograph (SEM) showing a sample r = 1-␮m microfabricated structure in front of a commercial chemically synthesized 4.5-␮m diameter magnetic particle to show relative size. The top and bottom material slabs of the microfabricated structure are made from magnetic material (nickel). The central support post is nonmagnetic (copper). (Reproduced with permission from Zabow et al. [46].)

Viewed as specialized geometrical variants of regular chemically synthesized magnetic nanoparticles, the microfabricated multispectral agents’ added frequency-shifting ability can be understood by first reconsidering how regular magnetic nanoparticles generate T2 * contrast. As with any coherence phenomenon, SPIO agents’ reduction of the transverse coherence time is directly coupled to an associated frequency broadening; in this case, a broadening of the water NMR line. Proton Larmor precession frequencies are directly proportional to magnetic field magnitude. Therefore the spatially changing magnetic fields surrounding all magnetic particles lead to spatially varying proton precession frequencies that (1) lead to rapid dephasing of the neighboring transverse magnetization and the agents characteristic resulting negative image contrast, and (2) yield NMR frequency signatures that integrate out into broadening of the water line. To achieve instead a distinct frequency shift of the water NMR signal, rather than solely a broadening, the water protons need to be exposed to a field magnitude that is homogeneous over some extended spatial region and that is of a different magnitude than any surrounding field. Given the continuous spatial field decay away from any magnetic particle, this is not possible in any particle’s far-field region, but for appropriate geometries the requisite homogeneous field profiles can be engineered within the particle near-fields. For example, it has been shown [46] that when suitably aligned and magnetized to saturation by the background MRI field, an appropriately spaced pair of magnetizable disks can give rise to a homogeneous field region that can uniformly shift the NMR frequency of the water passing between those disks. A schematic of this double-disk geometry is shown in Figure 16.2 together with the spatial profile of the total field magnitude resulting from the magnetized structure’s field added to the background MRI field. Although not shown in the figure, the magnetic disks are held at a predetermined fixed distance apart by nonmagnetic posts, either internal or external to the structure (see Fig. 16.4(a) and Fig. 16.4(b)). By varying the disk spacings, thicknesses, and radii, or by changing their material composition to increase magnetic moment, the homogeneity and the magnitude of the surrounding magnetic fields can be engineered to yield the desired spectral signature. This enables water NMR frequencies to be shifted almost arbitrarily, certainly over ranges that exceed any that chemical synthesis has yet to achieve. For the shift in NMR frequencies to be well defined, it is critical that the structures

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FIGURE 16.2 Magnetic structure and field diagrams. (a) Diagram of the field (small black arrows) from two parallel disks magnetized to saturation by the background MRI field (large gray arrows). (b) Calculated (negative) field magnitude in the midplane through a typical magnetized disk set, contrasting the homogeneous nature between the disks with the rapid external decay. (Reproduced with permission from Zabow et al. [46].)

be aligned parallel to the applied magnetic field. The structures’ built-in magnetic shape anisotropy automatically ensures that this is the case, causing the particles to self-align to the applied magnetic field. The double-disk structures are by no means the only self-aligning structures that can give a well-controlled frequency shift. As another example, it has recently been demonstrated [48] that hollow magnetic cylinders of suitable aspect ratios can likewise generate shifted NMR signals from the water within the cylindrical shells. These hollow magnetizable cylindrical structures operate, and self-align, according to the same principles as the doubledisk structures and, at least qualitatively, can be understood as each being comprised simply of a series of rotationally superposed double-disk structures [48]. Accordingly, in this case, instead of disk thickness, spacing, and radii, it is the cylinder wall thicknesses, radii, and lengths that control the resulting spectral shifts. For the simplified cases of “thin” disks or “thin-walled” cylinders made from a material with saturation magnetic polarization J S , the resulting NMR frequency shifts ␦␻ can be analytically approximated as
Double disk:


␻ ≈ −4␥ JS


Hollow cylinder:


␻ ≈ −4␥ JS

hr 2 (4r 2 + 4s 2 )3/2 L␳ t (L 2 + 4␳ 2 )3/2

 

where ␥ is the gyromagnetic ratio, and the geometrical parameters r, 2s, h, L, ␳ , and t denote, respectively, the disk radii, center-to-center separation and thicknesses, and

FIGURE 16.3 Schematic of double-disk and hollow cylindrical contrast agent microstructure geometries. The double-disk structure (left) has disks of radii r, center-to-center separation 2s, and thickness h. The hollow cylinder (right) has radius ␳ , length L, and wall thickness t.

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the hollow cylinder length, radius, and wall thickness as illustrated schematically in Figure 16.3. In the above, the designation “thin” assumes that h  r ≈ 2s and t  L ≈ 2␳ . The above equations have been arranged in similar form and the geometrical correspondence between the two systems becomes even more apparent by noting the natural associations between disk diameter and cylinder length 2r ∼ L, between disk separation and cylinder diameter 2s ∼ 2␳ , and between disk thicknesses and cylinder wall thickness h ∼ t. SEMs of arrays of different double-disk and hollow cylindrical magnetic structures of various configurations and sizes, all of which can discretely shift NMR spectral peaks, are collected together in Figure 16.4. Note that it is neither the disk nor the cylindrical shell families of geometrical structures that should necessarily be viewed as the critical component for generating distinct spectrally shifted NMR signals. In fact, these particular geometries were chosen primarily on the basis of their being, among others, compatible with microfabrication processing technology. Rather, the critical requirement is only that the total resulting field possess a water-accessible spatially extended region over which its magnitude is homogeneously shifted away from that of all surrounding fields. The ability to encode different spectral signatures into magnetic particles, that would otherwise appear identical in an MRI scan, offers the potential to track and differentiate between multiple different particles and the different biological markers that may be specifically bound to them. While the underlying physical principles behind their operation are different, in spectral terms the engineered magnetic structures can be thought of as radiofrequency (rf) analogs to optically probed quantum dots [75–77]. Just like quantum dots, they therefore have the potential to allow similar multiplexed studies except, being rf rather than optical probes, the obscuring optical scattering properties of biological tissues no longer matter and the probes can potentially be tracked in vivo. A fundamental difference between these magnetic structures and quantum dots, however, is that there is nothing quantized about the magnetic structure sizes. Instead, it is the magnetic structure’s geometrical aspect ratios, rather than quantum dot’s semiconductor crystal size, that is engineered to provide the desired “color” or rf spectral response. Indeed, as implied by the dimensionless groupings of geometrical variables in the above frequency-shift equations, provided all dimensions

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FIGURE 16.4 Scanning electron micrographs of various microfabricated contrast agent microstructures. Panels (a) and (b) show arrays of microfabricated double-disk structures with single internal, or three external, support posts, respectively. Panel (c) shows an array of microfabricated hollow cylindrical structures.

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are scaled proportionally to one another, the shifts in the surrounding water NMR signals are independent of overall structure size, allowing a wide range of particle sizes. Apart from permitting multiple different “colors,” their spectral signals also enable confirmation as to whether hypointense regions observed within an MR image indicate the locations of administered T2 * contrast agent or are instead spurious natural image darkenings or signal voids arising from various sources of magnetic field inhomogeneities. Figure 16.5 illustrates these concepts, showing regular T2 * -weighted gradient echo MR imaging of a series of double-disk structures together with chemical shift MR imaging of those same structures. The unique spectral signatures were shown to enable different particle types (in this case

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FIGURE 16.5 Multispectral MRI. (a–d) Chemical shift imaging of demonstration 1.25-mm diameter particles magnetized by the background MRI field. Particle frequency was varied by changing the thickness of electroplated nickel layers that formed the magnetizable disk pairs. As with normal SPIO detection, magnetic dephasing due to the particles’ external fields enables the spatial imaging shown in the gradient-echo MRI (a). However, comparison between (a) and the chemical shift images (b) shows that the additional spectral information both differentiates between particle types and improves particle localization. The particles are shown schematically (not to scale) in (c). With particle spectra ((d), to the right of the corresponding chemical shift images in (b)) shifted well clear of the water proton line, different planes in the chemical shift imaging map isolate different particle types for unambiguous color coding with minimal background interference ((b), bottom panel). Although still visible in the gradient-echo image, the top corner particle of the letter “B” was damaged, causing its shifted frequency peak to vanish. (Reproduced with permission from Zabow et al. [46].)

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all double-disk structures but made with disks of different thicknesses) to be distinguished from one another and from the background. Additionally, because the spectral signals came from water between the structures’ two disks rather than from external surrounding water, as do the dephased magnetization signals responsible for the SPIO-like spatial contrast, the spectral signals afforded improved spatial resolution in the chemical shift images. 16.3.2 Chemical Shifting Versus Magnetic Microstructure Shifting Spectral shifting of NMR frequencies, per se, is nothing new; after all, it underpins all of NMR chemical spectroscopy where the magnetic shielding effect of the surrounding electron clouds locally alters the fields seen by that molecule’s protons, giving rise to different chemical shifts or characteristic NMR spectral “fingerprints” for different molecules. Although chemical shifts may appear unrelated to the shifts produced by the microfabricated agents, noting the parallels that do exist between the two helps illuminate some of the microfabricated agents’ relative advantages. Just as chemical shifts can be understood in terms of the molecules’ surrounding chemical electronic current loops, the microfabricated structures’ field shifting can be regarded analogously as arising from the effective surface Amperian electronic current loops that surround all magnetic objects. The difference is that since the composition, geometrical structure, and relative positionings of these magnetic objects can be arbitrarily engineered, so too can the strengths, locations, and paths of those Amperian surface current loops, providing a degree of explicit field control that is harder to achieve through chemical molecular engineering. Also, since ferromagnetism is fundamentally a multibody phenomenon relying on mutual cooperation between multiple spins, the fabricated microstructures, which represent extended material objects, can readily be ferromagnetic in nature; by contrast, molecular complexes are typically paramagnetic or diamagnetic in nature. When magnetized by the background MRI field, therefore, the microstructures’ greater magnetic permeabilities induce larger equivalent electronic currents than those of dia- or paramagnetic molecules, and therefore enable larger frequency shifts that, as explained below, allow for greater signal amplification. Indeed, since the microfabricated magnetic structures are magnetically saturated already in applied fields well below a tesla, their field shifts can be large and essentially independent of background field for typical MRI fields. This can be seen in the sample NMR spectra acquired from a set of double-disk structures shown in Figure 16.6. The curves, so-called z-spectra [78], show the spectrally shifted NMR water peaks for three different fields strengths, 11.7 T, 7.0 T, and 4.7 T. Although signal amplitudes differed (due to differing transmit/receive coil geometries and field-dependent differences in T1 ), the frequency shifts were independent of MRI field strength. By comparison, given the much lower permeabilities of para- or diamagnetic species, chemical shifts are normally proportional to the background MRI field strength and large molecular shifts therefore require large MRI fields. For the microfabricated structures this then leads to the somewhat unusual situation where conventional chemical shift nomenclature, typically involving measurements represented in relative parts per million (ppm), is no longer helpful. The shifts in Figure 16.6, for example, are therefore given in absolute units, not in ppm. 16.3.3 Magnetization Transfer for Amplification of the Detection of Microfabricated Structures Chemical shifts can be exploited not only for NMR chemical spectroscopy or associated chemical shift MR imaging as seen in Figures 16.5 and 16.6, but also for generating

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FIGURE 16.6 Field-independent spectral shifting: H2 O z-spectra for an array of double-disk structures (r = 2.5 ␮m) taken at three different MRI field strengths all exhibit the same 275-kHz spectral shift, indicating that structures have been fully magnetized to saturation by background MRI fields. (Reproduced with permission from Zabow et al. [46].)

frequency-dependent MRI spatial contrast through off-resonance proton exchange magnetization transfer [79, 80]. For example, such is the mechanism behind CEST and PARACEST contrast agents [41, 65, 67], where off-resonance pulses saturate out the magnetization of particular exchangeable protons within the molecular complexes comprising the CEST/PARACEST agents. Once these off-resonance protons subsequently undergo chemical exchange with the surrounding bulk water protons, the total on-resonance magnetization, and hence the signal intensity acquired from the surrounding bulk water pool, is diminished. Compared to chemical shift imaging, this magnetization transfer scheme represents a more indirect method of imaging, but one that can afford substantial signal enhancement. Specifically, if the free and bound protons exchange at a rate that is faster than the longitudinal relaxation time, T1 , but that is not too fast (explained below), then prior to each on-resonance signal acquisition, several off-resonance saturating pulses can be applied to multiply the total magnetization saturation and effectively increase the contrast agents’ signal-to-noise ratio. Likewise, the microfabricated structures reviewed here can also benefit from similar signal amplification strategies. The difference is that for the microfabricated agents the magnetization transfer does not rely on chemical exchange processes and is therefore not limited by chemical exchange rates; instead, exchange is limited only by the rate at which water molecules naturally self-diffuse around the structures and in and out of the structures’ internal homogeneous field regions. Because this diffusion continually replenishes the water within the structures’ internal homogeneous field regions with “fresh” water that has diffused in from outside, the spectral signals generated can appear to come from volumes of water that are orders of magnitude larger than the internal signal-generating regions of the particles themselves. Since the time period required to diffuse any given distance scales with the square of that distance, as structure sizes shrink, the water exchange rates, and hence potential NMR signal amplification, increase quadratically. This renders the microfabricated structures well suited to miniaturization down to the small sizes required for cellular and molecular level imaging applications since for any given total volume of contrast agent material, the total signal from an ensemble of contrast agent micro- or nanostructures is greater the smaller are the structures comprising that ensemble.

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An important caveat to continually increasing the exchange rate, however, is that it should not become so fast that it frequency-broadens the spectral peak by more than its shift. In other words, ensuring that the shifted frequency peak remains clearly resolvable from the unshifted background water requires that the frequency of water exchange remains less than the magnitude of the water’s NMR frequency shift or, equivalently, is slow relative to the NMR time scales involved. A more in-depth discussion of this so-called slowexchange condition and its significance in contrast agent design can be found in Woods’ review of PARACEST agents [41]. Although presented in terms of chemical-exchange limited molecular agents, the arguments in that review carry over equally to the diffusionexchange limited microfabricated agents discussed here. To summarize those arguments, maximizing an exchange-based agent’s potential for signal amplification requires maximizing the exchange rate, which in turn requires maximizing the magnitude of that agent’s frequency shift. Fortunately, given their ferromagnetic nature, even at low MRI background fields the microfabricated structures can produce large field shifts, exceeding those achievable by molecular analogs. For example, a wide range of engineered frequency shifts from tens of kilohertz to substantially beyond a megahertz have already been experimentally demonstrated [46, 48]. In principle, shifts still an order of magnitude larger should also be achievable, potentially allowing structures to be scaled down to the 100-nm range [46], well within typical SPIO nanoparticle size ranges. Apart from enabling large signal amplification, large shifts also bring other benefits. By shifting the signal of interest as far away from the bulk water signal as possible, much of the main competing background noise can be eliminated. Constraints on the power and envelope shapes of the applied rf pulses are significantly reduced since for large offset shifts, there is little risk of any residual rf power overlapping with the bulk water line. Similarly, multiple sets of different frequency shifts can also be clearly distinguished from one another. Also, detection becomes more forgiving of any small system field inhomogeneities, whose resulting frequency shift errors become negligible in comparison to microstructure frequency shifts. Since the amplitude of the agents’ spectrally shifted signals depends on the total magnetization exchange, it depends not only on the chemical- or diffusion-driven exchange rate for each proton, but also on the total number of those protons whose magnetization can be simultaneously saturated at any one time. Not being dependent on any molecule’s particular individual exchangeable proton sites but instead being designed to generate uniform field offsets over spatially extended regions, the microfabricated magnetic structures can interact with larger numbers of protons simultaneously to give large signals per particle. Combined with the signal enhancement due to diffusional exchange, it has been shown [46] that such microfabricated structures should be detectable at low concentrations that may exceed those of certain molecular-based agents and that compare favorably to chemically synthesized SPIO nanoparticle contrast agent concentration requirements. 16.3.4 Combining Imaging and Sensing The microfabricated agents’ spectral signals are derived from the water that flows through the particle structures. This enables several possibilities for transforming the multispectral contrast agents into rf probes that could potentially report on the presence of various biomolecules or enzymes, or monitor various in vivo physiological conditions. By selectively blocking or unblocking water access to the particles’ interior regions, the spectrally

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FIGURE 16.7 Controlling diffusion to turn tags on or off. (Main panel) High tilt-angle SEM image showing a square array of r = 2.5-␮m double-disk particles. Except for a defined circular region, all particles have their interiors filled, blocking water diffusion. (Top left inset) A higher magnification SEM image of the boundary between open and filled particles. (Top right inset) The resulting background-subtracted chemical shift MRI showing transferred magnetization saturation from the particles’ shifted resonance. Signal is visible from those particles that have water diffusing through their open interior region (labeled “On”) but not from those particles that have their interiors filled (labeled “Off”). The bottom of the image shows a region that contains no particles (labeled “No tags”), providing a null background signal comparison. A scratch (seen at the lower right corner) removed ∼100 particles (about 10–20 per voxel). Its visibility in the magnetic resonance image suggests the potential for high-resolution imaging to spectrally distinguish individual such particles. (Reproduced with permission from Zabow et al. [46].)

shifted NMR signals can disappear or reappear, respectively, allowing the structures to act as switches. The idea is illustrated in Figure 16.7, showing an array of double-disk structures that are identical in every respect except that some of the structures have their internal regions blocked. Instead of a standard gradient-echo MRI that spatially locates structures based on their resulting T2 * hypointense contrast, the MRI shown is a difference image formed by subtracting from a first on-resonant image, a second image that was generated after first applying a preceding set of off-resonant pulses at the particles’ shift frequencies, as described above. In this way the locations of those double-disk structures with geometries that are resonant with the preceding off-resonant pulses appear highlighted in the difference image while all else disappears. However, only those double-disk structures that have their interior regions accessible generate any off-resonance contrast signal; those structures with their interior regions blocked do not yield any spectrally shifted signal and hence are not visible. The potential for the structures’ contrast signal switching to report on local conditions is then clear: if the material that fills the particles’ internal regions or that coats the particles

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was designed to dissolve under certain conditions, such as temperature or pH level, or to break down in the presence of specific enzymes, the presence or absence of the particles’ spectral signals could provide real-time information about in vivo physiological events. This idea is, for example, not unlike one first demonstrated with gadolinium complexes, where water access to the inner Gd3+ ion was compromised by a galactopyranose blocking group sensitive to cleavage by ␤-galactosidase, a common marker gene [27]. However, the difference is that because the particles’ spatial T2 * contrast is independent of whether the particle interior admits water or not, the spatial locations of the particles remain always visible, allowing them to be simultaneously continuously tracked. Additionally, because each particle’s signal can be engineered to occur at a different frequency, multiple different physiological conditions can potentially be simultaneously tested for and differentiated between, even when occurring within the same MRI voxel. The strategy of admitting or denying water access to the particle interior is applicable to both the double-disk and the hollow cylinder structures. However, with their additional separating nonmagnetic posts, the double disks also have the potential to monitor local conditions in a more continuous manner through condition-dependent variations in the magnetic resonant frequencies of the structures themselves. For example, if the separating posts were made of a material responsive to local conditions, then the resulting change in disk spacing could continuously shift the NMR resonances of the water passing between the disks. This change would be instantaneously reflected in the shift in the water resonance, opening the possibility of fast MRI sensors. 16.3.5 Contrast Agent Microfabrication From a scientific point of view, the ability to arbitrarily control NMR resonances increases the magnetic particles’ imaging and sensing potential; from an engineering point of view, however, fully realizing that potential is not without its challenges. With their multispectral nature the magnetic structures epitomize the proverbial double-edged sword: while the structures’ resonances can be intentionally manipulated through controllably changing their geometry, any unintentional geometrical variations from one structure to the next will lead to unintended variation in their resonance frequencies. For NMR signals acquired from individual structures, small fabrication inaccuracies do not necessarily matter: at worst, they diminish that individual structure’s field homogeneity, reducing the sharpness of the structure’s spectral peak. However, for NMR signals acquired from large ensembles of nanostructures, interparticle frequency shifts integrate out into a further broadening of the structures’ combined NMR spectral peak that can significantly reduce signal intensity. For example, the double-disk structures can potentially yield shifted spectral peaks with linewidths on the order of a hundred times narrower than their shifts; accordingly, anything more than a 1% variation in any of the structures’ geometrical parameters or in their saturation magnetization (or material composition) across the ensemble of structures can noticeably reduce their signal intensity [47]. Figure 16.8 shows an example of this rapid signal reduction as a function of the ratio of cross-structure frequency variation to average frequency shift. The plot shows calculated histograms that approximate the double-disk’s NMR signatures by measuring, throughout the space surrounding an ensemble of structures, the total volumes over which each field magnitude (or, equivalently, Larmor precession frequency) occurs. Therefore it is not only the structures’ geometrical forms but also the requirement of minimizing cross-structure variation that favors top–down fabrication techniques over

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FIGURE 16.8 Impact of cross-wafer manufacturing variation on spectral linewidths. The main plot shows numerical calculation of the broadening in peak widths when integrating expected signal from an array of double-disk structures that suffer from interparticle cross-wafer processing variation. Curve labels show the integration range of resulting structure frequency shifts about the central frequency. Inset plots the associated relative signal loss as cross-wafer variation increases. Calculations are for structures with r = 1 ␮m, 2s = 0.85 ␮m, h = 50 nm, and J S = 0.6 T (nickel). (Reproduced with permission from Zabow et al. [47].)

traditional chemical syntheses. For example, producing sharp spectral peaks with the double-disk system requires that for each structure the top and bottom disks be well centered on top of each other, the separation gap between those disks be well defined, the radii, thicknesses, and magnetic moments of the top and bottom disks be well controlled and equal to each other, and from one structure to the next there be minimal cross-wafer variation between any of the structures’ magnetic or geometric parameters. Even with the more direct control that top–down fabrication provides, attention must still be paid to various systematic errors and cross-wafer variations particular to each microfabrication processing step. Full discussion of such considerations is beyond the scope of this chapter but can be found detailed elsewhere [47]. Also to be found elsewhere [48] are microfabrication details of the hollow cylindrical structures whose geometry dictates less conventional microfabrication techniques. The double-disk structures, by contrast, are better suited to conventional planar surface micromachining processes and therefore provide a better introduction to some of those processes. Outlined here are illustrative examples of two different approaches for microfabricating double-disk structures: one based on electrochemical deposition and the other on thermal evaporation. The higher deposition rates of chemical electroplating are indispensible for the fabrication of relatively large microscale structures, while the reduced surface roughness and cross-wafer variation possible with thermal evaporation techniques are better suited to more accurate, smaller microscale or nanoscale structures. Whether electroplated or evaporated, a precursor to all double-disk structures is a deposited planar metal trilayer, photolithographically patterned into arrays of trilayer cylindrical stacks, that each consist of a base magnetic layer, a middle sacrificial nonmagnetic layer, and a final top magnetic layer. The sacrificial middle layer is used to define the spacing between what becomes the top and bottom magnetic disks and is either wholly or partially removed depending on the nature of the desired final spacer post(s) linking top and bottom disks. In addition to these three layers, extra layers may be intermingled that may

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serve as various protective coatings for the final magnetic layers. For example, interleaving thin-film titanium or gold layers can make the final structures more biologically inert or more amenable to certain bioconjugation protocols that may be required for their specific targeting. In the vertical direction, perpendicular to the processing substrate, accurate definition of the thicknesses of the deposited metal layers can be achieved relatively easily through control of metal deposition rates and times. In this way the stacked layer design allows the thicknesses of top and bottom disks as well as their separation gap to be accurately defined. Disk radii meanwhile, being dependent on lateral patterning of the metal trilayer, are subject to various spatial resolution limits. However, because the disks are circular, their patterning scales relatively well; when encroaching on optical diffraction limits, other shapes would scale less well, since preferential blurring of higher spatial frequencies would lead to shape distortion. Scalability is beneficial because (1) structure frequency shifts are independent of overall structure size and the range of potential applications is therefore increased if structures can be fabricated over a large range of sizes, (2) for any given total amount of material, smaller structures can interact with a larger total volume of water, and (3) being able to fabricate, in particular, very small micro- or nanostructures makes for a less intrusive agent that may afford increased biological compatibility. Spatial patterning of the circular stacks may be performed either before or after the metal deposition steps. In the former approach a masking layer of photoresist is first patterned with arrays of circular openings defining those locations on the substrate wafer, where, after subsequent metal film deposition and photoresist removal, metal will remain in the form of the required cylindrical stacks; in the latter, the metal trilayer is first uniformly deposited everywhere on the substrate and covered by an inverse photoresist mask consisting of arrays of circular disks (instead of circular holes) that protect the metal beneath those circles as the surrounding metal is etched away. Which approach is preferred depends on considerations of desired final structure sizes and associated metal deposition and removal processes, on considerations of the resulting structure accuracy and cross-wafer uniformity that is demanded, and on considerations of the overall process complexity and throughput. The last consideration of processing efficiency is particularly important when considering that the top–down fabrication approach is here replacing more traditional bottom–up chemical syntheses that can offer substantially higher throughput. Indeed, it may often be desirable to trade off some degree of structure accuracy in favor of increased top–down process simplicity and throughput [47]. Turning to specific fabrication details, the diagrams in Figure 16.9 briefly outline the main steps in a sample process flow for microfabricating electroplated double-disk structures. A 10-nm thick adhesion layer of titanium is thermally evaporated onto the wafer substrate followed by a 100-nm layer of gold that serves as the electroplating seed layer and provides the necessary electrical contact. A first layer of magnetic material (e.g., nickel), followed by a nonmagnetic sacrificial layer of copper, and then a second layer of magnetic material are sequentially electroplated on top of this gold-coated substrate. Electroplating times are determined by the desired final disk thicknesses and separations, which are governed by double-disk field homogeneity concerns and desired NMR frequency shifts. A layer of photoresist is then spin-coated on top of this metal stack, photolithographically patterned and developed, and then hardbaked to leave arrays of resilient protective photoresist disks as shown in the schematic cross section of Figure 16.9a. An argon ion beam directed normal to the surface is then used to ion-mill through the exposed top magnetic layer and partway through the underlying copper (Fig. 16.9b). A selective copper wet etch then exposes the

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FIGURE 16.9 Process flow diagram for electroplated double-disk structures: (a) electroplated nickel–copper–nickel trilayer atop gold–titanium seed layer with patterned photoresist disks on top, (b) ion milling through top nickel layer and part way through copper, (c) partial copper wet etch, (d) additional photoresist coating after flood exposure and development, (e) nickel wet etch and additional photoresist removal, and (f) left: timed copper wet etch, right: SU-8 photopatterned external posts and complete copper removal. (Reproduced with permission from Zabow et al. [47].)

base magnetic layer, but this etch is stopped short of laterally etching completely through the copper (Fig. 16.9c). A second layer of photoresist is then spin-coated over the wafer, flood exposed, and developed (Fig. 16.9d). This results in the removal of this second layer of resist everywhere except for directly beneath the circular disks of the top magnetic layer, which effectively double as an optical mask during the exposure. A selective wet etch of the base magnetic layer and an acetone removal of the second photoresist layer (Fig. 16.9e) therefore leaves the base magnetic disks patterned identically to, and automatically aligned directly beneath, the top magnetic disks. Finally, either the remaining copper is subjected to a timed wet etch to leave single central support posts, or external posts are first patterned using, for example, a biologically compatible photoepoxy [81, 82], before the remaining copper is completely wet etched away (Fig. 16.9f). As another example, the process flow in Figure 16.10 sketches out an alternative fabrication approach based on thermal evaporation. Beginning similarly with a gold–titaniumcoated substrate, a double layer of resist is spin-coated over the wafer. The top layer of this bilayer resist is photosensitive and lithographically patterned to yield an array of circular holes once developed. Meanwhile, the base layer of the resist, a so-called lift-off resist (LOR), is not photosensitive, but instead develops isotropically. Appropriately timed, the resist development then exposes the underlying gold layer and leaves the reentrant undercut profile shown in the cross-section schematic of Figure 16.10a. Sequential evaporation of a first base magnetic layer, a middle sacrificial copper layer, and a final magnetic layer results in the metal deposition profile shown in Figure 16.10b with evaporation times being used to determine layer thicknesses. Because of the undercut bilayer resist profile, the metal deposit on the base gold layer and on the resist layers is discontinuous, allowing clean

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FIGURE 16.10 Process flow diagram for evaporated double-disk structures using lift-off patterning: (a) patterned bilayer lift-off resist (undercut profile) atop a gold–titanium base layer, (b) nickel–copper–nickel trilayer evaporation, (c) lift-off resist stack removal and partial copper wet etch, and (d) left: timed copper wet etch, right: SU-8 photopatterned external posts and complete copper removal. (Reproduced with permission from Zabow et al. [47].)

removal of the resist bilayer together with its metal coatings. A subsequent partial copper wet etch then leaves the metal profiles seen in Figure 16.10c. Because of the manner in which the evaporated material accumulates on the resist mask, the exposed circular regions in this mask continually shrink during the metal evaporation leaving tapered metal profiles and top disks that are somewhat smaller than their underlying base disks. This does affect resulting double-disk field homogeneity but can be compensated for by making the top disk layers thinner than the base layers [47]. Finally, as before with the electroplating scheme, the remaining copper may either be partially etched away or, if external posts are used, completely removed (Fig. 16.10d). Note that while each of the above steps show just two representative structures, through the parallel nature of photolithographic processing, each one of those steps is compatible with the simultaneous patterning of large numbers of copies across the substrate wafer. For example, a standard 6-inch wafer can accommodate simultaneous fabrication of on the order of 1010 ␮m-sized structures. The above process flows represent only two possible protocols; many other variants are possible. Ultimately, which processing protocol and materials are used will depend on the contrast agents’ intended applications. Likewise, process protocols will depend on the acceptable levels of compromise between required geometrical accuracy and microfabrication throughput that the final structures’ uses allow.

16.4 FUTURE OUTLOOK While the multispectral contrast agents described in this chapter have been successfully microfabricated in various shapes and sizes and successfully tested in various MRI scanners, they still represent a new approach to MRI contrast agent development. At this point the top–down microfabricated multispectral structures may be regarded as demonstrations only of a new physical platform, a precursor only to any proven new biological imaging agent. At the time of this writing, much remains to be done in terms of biocompatibility studies, the development of efficient strategies for the administering or biological uptake of these

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structures, and the structure microengineering itself. For example, structure compositions must still be optimized for biological compatibility, and methods and functionalization strategies developed for the specific biological targeting of the agents. Work continues on further miniaturization of the structures, on the exploration of alternative fabrication materials, strategies, and geometries, and on the improvement of the efficiencies and yields of those microfabrication processes. With the wealth of microfabrication strategies available, the relative ease of switching between materials that those strategies permit, and the many already well-developed proven methods for the bioconjugation of other nanoparticle agents discussed elsewhere in this book, the outlook for microfabricated cellular and molecular imaging agents is an optimistic one. In particular, the range of fabrication techniques that have been developed by the microelectronics industry is substantial. So too is that community’s ability to continually improve upon existing techniques or to invent new ones to overcome new fabrication challenges. Leveraging off such a strong existing technology base, it seems likely that various additional fabrication strategies and even completely alternative types of agents may emerge that will enable additional new, and as yet unenvisioned, functionalities and novel forms of MRI contrast and molecular imaging agents.

ACKNOWLEDGMENTS This work was supported in part by the NINDS NIH Intramural Research Program.

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CHAPTER 17

Radiolabeled Nanoplatforms: Imaging Hot Bullets Hitting Their Target RAFFAELLA ROSSIN Department of Biomolecular Engineering, Philips Research Europe, Eindhoven, The Netherlands

Nanomedicine is a rapidly evolving field. Every year, new and more sophisticated nanoplatforms (NPs) designed to deliver drugs to specific targets or for molecular imaging of disease states make their appearance on the scientific scene. The assessment of the in vivo distribution of these NPs is the first step in order to fully understand their potential as candidate nanomedicines and to predict and prevent toxic side effects. To this aim, the incorporation of a radionuclide in the NP is of great advantage, as radiometric assays are rapid and quantitative. Furthermore, noninvasive nuclear imaging techniques are highly sensitive, achieve high tissue penetration and good spatial resolution, and can accelerate the translation of new NPs into the clinic. This chapter gives an overview of the strategies to label NPs with a variety of radionuclides (radiometals and radiohalogens), focusing on some of the hurdles that scientists must overcome (or take into consideration) when administering these radioconjugates in vivo. Furthermore, the latest achievements in molecular imaging of angiogenesis with radiolabeled NPs are reviewed. The most recent publications on nanoplatforms labeled with ␥ - and ␤+ -emitting radionuclides for SPECT and PET imaging are reviewed. Different strategies to label nanoplatforms are described, particularly those involving radiometals (64 Cu, 111 In, 67/68 Ga, 86 Y, 99m Tc, and 186/188 Re) and radiohalogens (123/124/125/131 I, 76 B, and 18 F). Some of the problems occurring when evaluating radiolabeled nanoplatforms in vivo are also covered. Finally, the use of radiolabeled nanoplatforms for angiogenesis imaging with PET and SPECT is discussed.

ABBREVIATIONS %ID/g, percent injected dose per gram tissue; 111 In-acac, 111 In-acetylacetonate; FpyME, 1-[3-(2-[18 F]fluoropyridin-3-yloxy)propyl]pyrrole-2,5-dione; 18 FDG, 2-deoxy2-[18 F]fluoro-d-glucose; BFC, bifunctional chelating agent; BMEDA, N,N-bis(2-mercaptoethyl)-N
,N
-diethyl-ethylenediamine; c(RGDyK), cyclo(Arg-Gly-Asp-d-Tyr-Lys); 18

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CAT, Chloramine-T; CB-DO2A, 4,10-bis(carboxymethyl)-1,4,7,10-tetraazabicyclo[5.5.2] tetradecane; CB-TE2A, 4,11-bis(carboxymethyl)-1,4,8,11-tetraazabicyclo[6.6.2]hexadecane; CNT, carbon nanotube; CT, computed tomography; DF, deferoxamine; DOTA, 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid; Doxil, PEGylated liposomal doxorubicin; DTPA, diethylenetriaminepentaacetic acid; EC, electron capture; ECM, extracellular matrix; EPR, enhanced permeability and retention; HMPAO, hexamethyl propyleneamine oxime; HPMA, 2-(hydroxypropyl)methacrylamide; HYNIC, hydrazino nicotinamide; IO, iron oxide particle; IT, isomeric transition; MB, microbubble; MPS, mononuclear phagocyte system; MRI, magnetic resonance imaging; MWNT, multiwall carbon nanotube; NIRF, near-infrared fluorescence; NP, nanoplatform; PAMAM, poly(amido amino); PE, phosphatidylethanolamine; PEG, polyethylene glycol; PET, positron emission tomography; PPI, poly(propylene imine); QD, quantum dot; RGD, Arg-Gly-Asp; SarAr, 1-N-(4-aminobenzyl)-3,6,10,13,16,19-hexaazabicyclo[6.6.6]eicosane-1,8-diamine; siRNA, small interfering RNA; SOD, superoxide dismutase; SPECT, single photon emission computed tomography; SPIO, superparamagnetic iron oxide particle; SWNT, single-wall carbon nanotube; SWNTol, hydoxylated single-wall carbon nanotube; TETA, 1,4,8,11-tetraazacyclotetradecane-1,4,8,11-tetraacetic acid; US, ultrasound; VCAM-1, vascular cell adhesion molecule-1; VEGF, vascular endothelial growth factor; VEGFR-2, vascular endothelial growth factor receptor-2; ␤+ , positron.

17.1 INTRODUCTION Recent breakthroughs revealing the mechanisms of disease that have emerged from genomics and proteomics prompted great advances in the design of new therapy strategies. It was Paul Ehrlich who, at the beginning of the twentieth century, conceived the idea of a “magic bullet” capable of selectively hitting the diseased cells [1]. In the clinic, however, this concept is hampered by the fact that most small molecule drugs distribute throughout the entire body and few are perfectly selective in terms of action. This results in intrinsic toxic side effects when the drug has a cytotoxic action and/or a suboptimal efficacy. Therefore the design of “smart” drug carriers has become an integral part of drug development [2–5]. Such vehicles should be biocompatible, encapsulate a therapeutic cargo, and release it only at the site of action in a controlled and timely fashion and, possibly, provide a feedback to the clinician. Composite NPs are ideal candidates for this multitask assignment as they can carry multiple functionalities (targeting ligands and detectable labels among others). In addition, some NPs have intrinsic cytotoxic properties [6,7] or can be chemically engineered to degrade and release a therapeutic payload when triggered by the microenvironment (e.g., pH, enzymes) or by external stimuli (e.g., temperature, pressure, light) [8–10]. Although the first liposomes were described in 1965 [11] and the first controlled release of a protein from a polymer was described in 1976 [12], the term “nanomedicine”—the application of nanotechnology to healthcare—was introduced only recently [13]. The properties of drug carriers need constant tuning to meet healthcare needs. For this reason, nowadays the development of new “magic nanobullets” entails the combined multidisciplinary efforts of chemists, material scientists, biologists, engineers, mathematicians, and clinicians. Some examples of advanced carrier NPs include late-generation dendrimers,

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core–shell nanoparticles, CNTs, polymersomes, bubbles, multilayered polyelectrolyte assemblies, and mesoporous silica nanoparticles. The literature relating to new and more sophisticated NPs increases year by year. Biological data is provided in many of these studies, but the vast majority of experiments are carried out in vitro while these candidate drug vectors must perform their multiple tasks (reaching a target tissue or organ, delivering their intact cargo, and clear) in a living organism. The body is well protected from external agents by layers of epithelial cells that line the skin, the airways, the gastrointestinal tract, and the genital mucosa. Low molecular weight agents (<500 Da) can penetrate these barriers but, in general, macromolecular drugs and NPs cannot pass through the intercellular junctions [14]. To bypass these hurdles, most drug delivery systems are injected parenterally (intravenously, intraperitoneally, or subcutaneously). However, the macrophages of the MPS are very efficient in capturing circulating “foreign objects” (i.e., senescent cells, macromolecules, and particulates) and often NPs are eliminated from the blood circulation within minutes from injection and end up in the liver, spleen, and lymph nodes. The blood clearance was shown to depend on charge, size, and surface chemistry of the NP [15]. Significant improvements in biological half-life were obtained by tampering with the opsonization and complement activation processes driving macrophage recognition, for instance, by covering the NP surface with PEG [15–19]. However, eventually all injected NPs are eliminated from the circulation and clear mainly through the liver and intestine or through the kidneys (the size limits for glomerular filtration are 70 kDa for linear polymers [20] and 3–5 nm in diameter for globular assemblies [21]). Knowing the bioavailability of a candidate drug carrier is crucial, especially for cancer therapy, as the active (e.g., receptor mediated) or passive targeting (driven by the EPR effect [22]) of a tumor may require time. Also, the NP distribution and residence time in nontarget tissues need to be evaluated to predict and minimize toxic side effects in sensitive organs. Finally, the evaluation of candidate NPs in living organisms gives the scientist vital information on the effects of different administration routes on the kinetics and specific pathways of NP distribution upon administration in vivo [23–26]. The localization of a NP at the site of interest is often accomplished by conjugating multiple targeting moieties (peptides, antibody fragments, receptor binding small molecules, etc.) to its surface. Information on the NP localization within a living system (from a living cell to the entire body) can be obtained with multiple imaging modalities by including detectable functions in the nanostructure or by exploiting intrinsic imageable characteristics of the NP. This adds diagnostic capability to the therapeutic action of the NP, leading to the emerging field of “theranostic” (therapeutic + diagnostic) agents [27]. To date, the specific imaging capabilities of targeting NPs have been utilized mainly to visualize vascular markers of disease (e.g., integrins and VCAM-1 on activated endothelial cells, macrophages in atherosclerotic plaques), which are easily accessible to blood-borne bulky structures without the need of extravasation from blood vessels. In fact, the accumulation of macromolecular structures within the interstitial space of solid tumors is mostly driven by passive diffusion (the EPR effect) and not by specific interaction with tumor cell targets [28–30]. The application of IOs for MRI, QDs for optical imaging micro- and nanobubbles for US imaging, perfluorocarbon nanoparticles for 19 F nuclear magnetic resonance spectroscopy, and more are discussed in detail in other chapters of this book. This chapter focuses on the achievements in the use of radioactive labels to track NPs in vivo.

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17.2 NUCLEAR TECHNIQUES AND RADIONUCLIDES FOR NANOPLATFORM IN VIVO TRACKING Classic “cut and count” radiometric assays have been used for decades to assess the biodistribution of polymers and NPs labeled with long-lived radionuclides such as 3 H, 14 C, and 125 I [25, 31, 32]. More recently, high-resolution (<100 ␮m) autoradiography equipment and preclinical PET and SPECT have entered research laboratories worldwide, making it easier to evaluate the distribution of radiolabeled compounds in normal animals and disease models. Nuclear imaging modalities hold a key position among the noninvasive imaging techniques because of their exquisite sensitivity [33]. Microdosing combined with good spatial resolution (from 0.5–1 mm in small animal scanners to 5–10 mm in clinical scanners [34]) and the ability to provide quantitative information on tracer accumulation in deep tissues allow a rapid translation of investigational tracers from mouse to human. Furthermore, the use of noninvasive nuclear imaging techniques for longitudinal studies speeds up the drug development process while keeping the costs to a reasonable level, especially when pricey disease models and transgenic animals are used. The choice of radionuclide for NP labeling depends on the imaging technique and on the timing for in vivo evaluation. The radionuclides used in PET imaging decay by emitting ␤+ particles, which travel in tissue until they collide with one electron generating two back-to-back 511-keV photons nearly 180◦ apart (annihilation). The detection of all pairs of photons hitting opposite detectors in coincidence allows the construction of an image defining the position of the annihilation event. On the contrary, SPECT uses single photons emitted by ␥ -emitting radionuclides. These photons travel through the tissue and hit the scanner detector. A collimator (usually a lead adsorbing plate with parallel holes or a pinhole collimator) selects the photons along defined axes and stops the others, therefore allowing the construction of a projection image defining the tracer distribution in the scanned object. Multiple projections, acquired sequentially, are then reconstructed in a three-dimensional (3D) tomographic image. The different physics underlying PET and SPECT imaging lead to intrinsic strengths and weaknesses of these two techniques. For instance, the availability of ␤+ -emitting radionuclides for PET imaging can be a limiting factor, especially for short-lived cyclotronproduced radionuclides, which must be produced in-house. On the contrary, most ␥ emitters are relatively long-lived and commercially available. Furthermore, the detection of 511-keV annihilation photons does not allow imaging different ␤+ -emitting radiotracers at the same time with PET while the simultaneous detection of photons with different energies can be performed with SPECT by choosing the appropriate energy window in the scanner. However, the need for a collimator reduces significantly (up to two orders of magnitude [34]) the sensitivity of SPECT scanners compared to PET scanners. Overall, the high detection sensitivity, deep tissue penetration, and quantitative accuracy of PET and SPECT make them powerful techniques for elucidating physiology, metabolic pathways, and molecular targets in vivo. Therefore great effort has been dedicated to integrating these techniques with those providing detailed structural information, such as CT and MRI [35]. As a result, SPECT/CT and PET/CT systems have become available over the past years and, recently, the scientific community has welcomed the introduction of the first integrated PET/MR scanners [36]. Functionalized nanocarriers are ideal platforms for multimodality imaging, as they can produce enough contrast at the target to effect a signal change detectable with CT, MRI, or US imaging while carrying also labels for nuclear or optical imaging. Consequently, the high resolution anatomical information delivered by CT,

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MRI, and US imaging can be used to localize the target, while the functional information provided by more sensitive techniques such as PET, SPECT, and optical imaging, can be used to detect the release of the therapeutic payload or to monitor the therapeutic effects.

17.3 RADIONUCLIDES AND STRATEGIES FOR NANOPLATFORM RADIOLABELING The radionuclides most commonly used to track NPs in living systems are listed in Table 17.1 (␥ -emitters) and Table 17.2 (␤+ -emitters). The selection of the right radionuclide for an in vivo assay depends on several factors such as the technique of choice, the radionuclide availability, the experimental time window, and the radiochemistry. The need to detect enough counts to differentiate the target tissue from the background noise, especially for imaging, limits the use of ␥ - and ␤+ -emitting radionuclides to approximately four decay half-lives. Therefore the physical life of the radionuclide must match the biological life of the NP and the time window of the biological process under scrutiny. Short-lived radionuclides such as 18 F and 68 Ga find an application with short circulating NPs or in dynamic imaging studies early after NP injection. On the contrary, the prolonged evaluation of long-circulating NPs is usually carried out with medium- to long-lived radionuclides such as 64 Cu, 111 In, and others. The energy of the radionuclide emission is also a key issue. High-energy ␤+ emitted by many “nonstandard” PET radionuclides (i.e., ␤+ -emitters other than 18 F, 11 C, 15 O, and 13 N) travel for a relatively long distance (a few millimeters) in the target tissue before annihilating with an electron. Some of these radionuclides (e.g., 66 Ga) also emit highenergy ␥ rays in cascade during the decay. These two factors reduce the quality of the PET images and complicated algorithms must be used for image reconstruction [37–39]. The emission energy is a limitation also for SPECT imaging. In fact, low-energy photons (<100 keV) produce too much scatter in tissues while ␥ energies greater than 200 keV are difficult to collimate, therefore posing a challenge to image resolution and quantitation. TABLE 17.1 Physical Dataa for Useful ␥-Emitting Radionuclides for NP Radiolabeling Decay Mode (%)

Main ␥ keV (%)

␤max MeV (␤ave )

140 (89.1) 137 (9.5)

1.069 (0.347)

17.0 h 13.2 h 59.4 days

IT ␤− (93) EC (7) ␤− EC EC

I

8.0 days

␤−

Ga

3.3 days

EC

2.8 days

EC

Radionuclide

Half-life

99m

Tc 186 Re

6.0 h 3.7 days

188

Re I 125 I 131

123

67

111

In

155 (15.6) 159 (83.3) 27 (74.0); 31 (13.2); 35 (6.7) 284 (6.1); 364 (81.5); 636 (7.2) 93 (39.2); 185 (21.4); 300 (16.6) 171 (90.7); 245 (94.1)

Production 99

2.120 (0.763)

0.807 (0.182)

Mo/99m Tc generator Re(n,␥ )186 Re 186 W(p,n)186 Re 188 W/188 Re generator 124 Te(p,2n)123 I 124 Xe(n,␥ )125 Xe → 125 I 185

130

68

Te(n,␥ )131 Te → 131 I

Zn(p,2n)67 Ga

111 112

Cd(p,n)111 In Cd(p,2n)111 In

a Some data are obtained from http://www.nndc.bnl.gov (searchable National Nuclear Data Center, Brookhaven National Laboratory) (accessed October 1, 2009).

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TABLE 17.2 Physical Dataa for Useful ␤+ -Emitting Radionuclides for NP Radiolabeling

Radionuclide Half-life 64

Cu

12.7 h

68

Ga

67.7 min

86

Y

14.7 h

124

I

76

Br

18

F

Decay Mode (%)

Main ␥ keV (%)

␤+ (18) 511 (35.2) EC (44) ␤− (38) ␤+ (89) 511 (178.3)

EC (11) ␤+ (33) EC (67) 4.2 days ␤+ (23) EC (77) 16.2 h ␤+ (55) EC (45) 109.8 min ␤+ (97) EC (3)

511 (64); 628 (32.6); 1077 (85.2); 1153 (30.5) 511 (45); 603 (62.9); 723 (10.4) 511 (109); 559 (74.0); 657 (15.9); 1854 (14.7) 511 (193.5)

␤max MeV (␤ave )

Mean ␤+ Range (mm)

Production

0.653 (0.278) 0.70

64

Ni(p,n)64 Cu

1.899 (0.829) 2.40

68

Ge/68 Ga generator

3.141 (0.660) 2.46

86

Sr(p,n)86 Y

2.138 (0.820) 3.24

124

Te(p,n)124 I Te(d,2n)124 I 76 Se(p,n)76 Br 76 Se(d,2n)76 Br 18 O(p,n)18 F 20 Ne(d,␣)18 F 124

3.941 (1.180) 5.07 0.633 (0.250) 0.69

a Some

data are obtained from http://www.nndc.bnl.gov (searchable National Nuclear Data Center, Brookhaven National Laboratory) (accessed on October 1, 2009). See also Nayak and Brechbiel [85].

The stability of the radiolabeled species is yet another aspect of paramount importance when evaluating its behavior in a living system. For instance, some radiometals bind with relatively high affinity to serum proteins (e.g., transferrin) and therefore they need to be stably anchored to the molecule under evaluation. The metabolic breakdown of radiolabeled species may also produce radiocatabolites with a different distribution with respect to that of the parent molecule. For these reasons, assessing the integrity of a radiolabeled species is crucial when using nuclear techniques to track its distribution in a living system, as these techniques do not discriminate between the intact species and the degradation products. For the same reason, extra care needs to be taken in order to administer in vivo NP solutions free of radioactive impurities such as free radionuclides, unbound radiometal chelates, and radioactive by-products. A comprehensive review of all the different methods to radiolabel NPs is beyond the scope of this chapter due to the variety of nanostructures in preclinical evaluation and the diversity of radiolabeling approaches used to this aim. Detailed methods to radiolabel liposomes [40] and general procedures for specific radionuclides (64 Cu and 76 Br) [41] were published recently. The following sections contain an overview of the most recent publications on NPs labeled with radionuclides that form coordinative bonds (radiometals) and covalent bonds (radiohalogens). The properties of each radionuclide and the radiochemistry are discussed briefly. The discussion focuses also on some specific problems that have occurred during radiolabeling and/or in vivo evaluation. 17.3.1 Copper Radiolabeling Among the ␤+ -emitting radionuclides, 64 Cu is becoming the radionuclide of choice for assessing the in vivo distribution and targeting capabilities of NPs with small animal PET imaging. Copper-64 is produced with high specific activity on medical cyclotrons [42] and decays by emitting ␤+ (18%) and ␤− particles (38%). Hence 64 Cu has potential for both molecular imaging with PET and radiotherapy. From the imaging point of view, the

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COOH

HOOC N

N N HOOC

405

N COOH

N

COOH

COOH

NOTA

DTPA

HOOC

HOOC

COOH N

N

N

N

N

HOOC

COOH

HOOC

COOH

COOH N

N

N

N

HOOC

COOH

TETA

DOTA

COOH

COOH N

N

N

N

N

N

N

N

HOOC

HOOC

CB-TE2A

CB-DO2A

NH H2N

HN

NH

HN

NH

HN

NH2

diamSar

FIGURE 17.1 Representative chelators for radiometal complexation.

low energy of the ␤+ emitted by 64 Cu produces images with superb spatial resolution (0.70-mm mean ␤+ range). In addition, due to the convenient half-life (12.7 hours), 64 Culabeled tracers can be tracked in vivo for 24–48 hours. Besides 64 Cu, the family of Cu radionuclides includes other isotopes currently under investigation for PET (60 Cu, 61 Cu, and 62 Cu) and radiotherapy (67 Cu) applications. For this reason, the chemistry of radio-Cu applied to the labeling of small molecules, peptides, and antibodies has been thoroughly investigated [43–45]. A variety of NPs were labeled with 64 Cu by using BFCs, such as derivatives of the tetraazamacrocyclic TETA [29, 46, 47] and DOTA [48–61] or the poly-azacarboxylate DTPA [62] (Figs. 17.1 and 17.2), that bind the metal with high affinity. Usually, the BFC is conjugated to the preformed NP before radiolabeling by using reactive functions available on its surface. When using negatively charged nanosystems, the length of the spacer between the chelator and the surface was shown to influence significantly the conjugation yield of the BFC. In fact, the use of short spacers affected the specific activity of the 64 Cu-labeled NP [59]. In this scenario, direct conjugation of the BFCs to the NP building blocks before

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HOOC

NCS

HOOC HOOC

NCS N N HOOC

N COOH

N

H N

COOH

COOH

HOOC

p-SCN-Bz-DTPA O

N

N

N

Br O

COOH

COOH

HOOC

COOH

BAT COOH

O COOH N

O N O

N

HOOC

N

H N

N COOH

HOOC N

N

CHX-A"-DTPA

HOOC N

COOH

N

COOH

HOOC

DOTA-NHS

N

SCN

N N

N

HOOC

COOH

COOH N COOH

p-SCN-Bz-DOTA

NODAGA

O NH

O

N OH HN O

H2N HO N

O

NH

HN

NH

HN

NH

HN

NH2 N H

O N

Deferoxamine

OH

SarAr

NH2

FIGURE 17.2 Representative BFCs for radiometal complexation.

NP assembly was shown to improve markedly the radiolabeling yields [57]. Nonetheless, with this strategy some chelators end up buried within the core of the NP, where they are not available to bind the radiometal [50, 54, 57]. Jarret and co-workers reported also the conjugation of preformed 64 Cu-DOTA groups onto NPs [51]. This approach, however, is not ideal for the in vivo administration of high specific activity 64 Cu-NPs due to the length of the labeling, conjugation, and purification procedures. Despite the attractive features of 64 Cu, one drawback for using 64 Cu labeling for NP evaluation is the partial instability of Cu complexes with tetraazamacrocyclic chelators. In fact, in vivo transchelation of Cu-DOTA and Cu-TETA has been reported both in rodents [63, 64] and in humans [65, 66]. The precise location and mechanism of copper release from the chelators are still under debate. However, it is known that the release is caused by the reduction of Cu(II) to Cu(I). In liver, this is followed by Cu binding to SOD [63, 64], while in blood the radiometal binds to ceruloplasmin, metallothioneins, and other serum proteins [64, 66]. To overcome this problem, a series of new macrocyclic ligands (the cross-bridged CB-DO2A and CB-TE2A, and diamSar, Fig. 17.1), that form Cu complexes with improved stability, were developed [45]. To date, there are no published data suggesting metabolism of 64 Cu-DOTA-NPs and formation of 64 Cu-SOD in liver. The in vitro stability of a series of 64 Cu-NPs was confirmed for as long as 48 hours in serum [47, 50, 53, 54]. Moreover, Liu and co-workers used the intrinsic optical properties of 64 Cu-DOTA labeled SWNTs and found a correlation between the amount of radioactivity and the amount of SWNTs in mouse liver by using Raman spectroscopy [53]. A thorough evaluation of the in vivo metabolism of a series of polymeric NPs is under way at Washington University in St. Louis. The results of this study will help us understand the fate of 64 Cu-DOTA-NPs in liver.

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Mechanisms other than Cu transchelation may also cause degradation of 64 Cu-NPs in vivo. For instance, when using small animal PET to evaluate the distribution of 64 CuDOTA-siRNA particles in mice, Bartlett and co-workers observed a high accumulation of radioactivity in kidneys and bladder shortly after injection [48]. This result was not expected due to the large size of the siRNA NPs (80–120 nm). Therefore the authors investigated the possible release of 64 Cu-DOTA-siRNA from the NP in vitro and observed approximately 50% NP dissociation in 0.4-M NaCl. Given these findings, they hypothesized a salt-mediated disruption of the NP within the kidney as the possible cause for the observed in vivo behavior. Recently, we investigated the possible use of PET to image the accumulation of 64 Cu-labeled immunobeads in the lung endothelium in mice [55]. We investigated the metabolism of the beads in this organ and observed significant dissociation of radioactivity (approximately 50% of the total) from the beads 1 hour after injection. The radioactivity elimination was not due to Cu transchelation; it was more likely due to the release of the noncovalently bound 64 Cu-DOTA-IgG from the surface of the beads and to degradation of the protein. The formation of a small amount of a hydrophilic radiocatabolite was observed also by Liu et al. [53] following the injection of 64 Cu-DOTA-CNTs in mice. In fact, virtually no CNT kidney uptake was observed based on Raman measurements, while some kidney activity was visible in the PET images.

17.3.2 Indium Radiolabeling The cyclotron-produced 111 In (2.8-day half-life) is another commonly used radiometal for NP evaluation in living systems, as it allows one to detect NP in vivo trafficking for several days. Indium-111 decays by EC with two major ␥ emissions (171 and 245 keV) and is widely available through commercial sources. Like 64 Cu, 111 In binds avidly to poly-azacarboxylate chelators such as DTPA and DOTA. For NP radiolabeling, BFCs have been chemically conjugated to the surface [67–70] or appended to hydrophobic domains constituting the nanostructure (i.e., the lipid bilayer of liposomes or the core of lipid NPs [19, 71]). Non-BFC-based methods to label NPs with 111 In were also developed. For instance, Oyewumi et al. [72] labeled Gd NPs by entrapping 111 In-acac (Fig. 17.3), a small lipophilic complex, in the particle core during synthesis by an oil-in-water microemulsion procedure. A similar approach is used to label liposomes after manufacturing (after-loading approach). In fact, 111 In-acac and 111 In-oxine (Fig. 17.3) freely cross the lipid bilayer of liposomes [40, 73]. The premature release of the label in vivo is avoided by including DTPA within the liposome aqueous core, which transchelates the radiometal forming a hydrophilic moiety [40, 74]. Recently, Cartier and co-workers [75] reported also the direct labeling of NPs with 111 InCl3 , 68 GaCl3 , and GdCl3 . Surprisingly, the trivalent metal cations bound quickly to latex nanoparticles in water without the need for a BFC, although with low specific activity (220 kBq/mg NP for both radiometals and 10 ␮g/mg NP for Gd) [75]. This finding was explained with the formation of strong coordination bonds between the metal ions and multiple carboxylic groups on the NP surface. However, the in vivo data reported by Cartier et al. [75] were limited to 1 hour post NP injection. In another interesting study, H¨afeli et al. [76] labeled magnetic microspheres prepared from metallic iron and activated carbon by directly mixing the particles with 111 InCl3 , 111 In-oxine, 111 In-DOTA, and 111 In-Abz-DOTA (an amino benzyl derivative of DOTA). Indium-111-oxine and 111 In-Abz-DOTA adsorbed onto the surface of the microspheres with high efficiency but only 111 In-Abz-DOTA was

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O

O

O N

O

O

M

O O

N

N O N

O

N

N

O

O

M

O

Tc

N

H

In/Ga-Acac

In/Ga-Oxine

S S

S

Tc-HMPAO

SH

N

M S

S

N O

S

S

M S

N

S N

Tc/Re-SSS

Tc/Re-BMEDA

FIGURE 17.3 Representative complexes used to radiolabel NPs comprising a lipophilic domain.

stably retained on the particle for at least 7 days, both in vitro and in vivo. The authors did not make hypotheses of the interaction mechanisms and did not study further the different release of radiolabels from the particles.

17.3.3 Gallium Radiolabeling Gallium radionuclides are also used to label NPs. The ␥ -emitting 67 Ga is one of the most widespread radiometals in nuclear imaging because of the 185-keV photon emission and relatively long half-life (3.3 days), which allows marketing. Gallium-67 is often used to radiolabel liposomes with convenient after-loading methods [40]. Like 111 In, 67 Ga can pass through lipid bilayers as 67 Ga-oxine (Fig. 17.3) and then it is trapped within the aqueous core of the NP preloaded with a chelator. To this aim, strong chelators such as DF [77] (Fig. 17.2) or DTPA [30, 78] are preferred to weaker ones such as nitriloacetic acid. In fact, in case of release of the liposome content in vivo, Ga-DF and Ga-DTPA undergo fast renal elimination while the nitriloacetate complex exchanges the radiometal with serum proteins [79, 80]. The ␤+ -emitter 66 Ga could be a good candidate for in vivo evaluation of NPs with PET because of the 9.5-hour half-life [81]. Unfortunately, 66 Ga has a suboptimal decay for imaging applications, as it emits high-energy ␤+ (4.153 MeV) and ␥ rays (1.039 and 2.752 MeV) that degrade the image quality (8.06-mm ␤+ mean range) and increase the patient dose [37]. For this reason, 68 Ga is the only gallium radionuclide used for PET imaging nowadays. This is due to its superb imaging characteristics and to the development of a user-friendly

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and long-lasting generator that delivers 68 Ga solutions on demand without the need for an in-house cyclotron [82]. However, the 67.7-min half-life of 68 Ga is a drawback for in vivo NP applications because it gives only a few hours of imageable signal. Few studies on 68 Ga-labeled NPs have appeared in the literature. In 1979, Wagner and Welch [83] used 68 Ga-oxine to label DTPA-conjugated human serum albumin microspheres. These radiolabeled particles were found to be very stable toward transchelation in vitro. After intravenous injection in dogs with one surgically ligated lung, the microspheres quickly accumulated in the perfused lung, allowing PET imaging at 45 min postinjection. Recently, as previously mentioned, Cartier et al. [75] directly labeled the surface of latex NPs with 68 Ga and other trivalent (radio)metals. The in vivo distribution of such NPs was studied in normal rats with PET (68 Ga), MRI (Gd), and cut-and-count techniques (111 In). In this study, the only time point used for PET imaging was 15 min postinjection. At the same time point, no significant MRI contrast was visible in any organ despite the administration of a double amount of Gd-NPs (20 mg per rat) compared to 68 Ga-NPs [75]. Reasonably, this was due to the low specific activity of the injected (radio)labeled NPs. Therefore in spite of the obvious convenience of direct labeling methods, BFC-based approaches appear to be more suited to in vivo molecular imaging with NPs.

17.3.4 Yttrium Radiolabeling The medium-lived ␤+ -emitter 86 Y (14.7-hours half-life) can be produced with high specific activity on a medical cyclotron [84] and is being considered as an imaging analog of the therapeutic radionuclide 90 Y (100% ␤− ) [85]. Unfortunately, the widespread use of 86 Y is held back by its limited availability and complicated decay scheme (emission of high-energy ␤+ and prompt ␥ rays in cascade). Despite these drawbacks, McDevitt and colleagues recently used 86 Y to study the distribution of CNTs in normal mice injected intravenously and intraperitoneally [24]. The DOTA-conjugated CNTs were 86 Y-labeled with high yield and specific activity (555 MBq/mg) in mild conditions (30-min incubation at 60 ◦ C). The prolonged retention of radioactivity in the peritoneal cavity suggested in vivo stability of these 86 Y-labeled NPs, confirmed also by the presence of intact 86 Y-CNTs in the urine 1 hour postinjection. Both small animal PET imaging and cut-and-count techniques showed fast blood clearance of CNTs and high uptake in kidney, liver, and spleen up to 24 hours postinjection [24].

17.3.5 Technetium and Rhenium Radiolabeling The establishment and evolution of diagnostic nuclear medicine can probably be attributed to the existence of 99m Tc [86]. Technetium-99m has been the most used ␥ -emitter since the introduction of ␥ -cameras in hospitals and today approximately 80% of the diagnostic imaging procedures are still carried out using 99m Tc radiotracers. The prominent role of 99m Tc is due to its optimal nuclear properties (>99% decay by IT with the emission of a 140-keV ␥ photon) and convenient half-life (6.0 hours) and to the existence of an affordable 99 Mo/99m Tc generator. In the hospital pharmacy, the radiometal is eluted daily from the generator in saline solution and added to commercial kits containing all the chemicals necessary for the tracer synthesis, thus making the production of 99m Tc tracers widely accessible and cost effective. Another Tc isotope with potential application in nuclear imaging is 94m Tc. Technetium-94m is a cyclotron-produced radionuclide with a 52.0-min

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half-life. It decays 72% by emission of 2.438-MeV ␤+ (2.74-mm mean ␤+ range), and it is a candidate PET tracer [87]. Based on the similarity of the chemical properties to those of Tc, in the past two decades considerable effort was dedicated also to the development of tracers containing Re radionuclides. Rhenium-186 and 188 Re decay by emitting ␤− particles and imageable ␥ photons. As a result, Re radionuclides can be tracked in vivo during internal radiotherapy procedures and can be used for dosimetry studies. Rhenium-186 can be produced in nuclear reactors (low specific activity) and cyclotrons (high specific activity) [88] while high specific activity 188 Re is available from the 188 W/188 Re generator [89]. For this reason, 188 Re is becoming the most popular Re radioisotope for labeling of a variety of molecules, including NPs. For labeling purposes, the radiometal contained in the generator eluate (pertechnetate or perrhenate) is reduced from the nonreactive heptavalent state to a lower oxidation state. Usually, this reaction is carried out in the presence of mono- or polydentate ligands or BFCs containing heterofunctions (e.g., thiol, amino, amido, hydrazino, or phosphino groups [90]) that stabilize the radiometal in its final oxidation state and avoid the formation of over reduced, inert Tc/Re-oxide species (usually referred to as Tc/Re radiocolloids). Despite this general approach, a variety of NPs were labeled with a direct method, that is, without the use of a metal chelator. In theory, this is possible when the NP contains multiple heteroatoms able to bind 99m Tc in one of its reduced states. However, the in vivo stability of these 99m Tc-NPs varies from case to case. For instance, Banerjee and co-workers studied the labeling of native chitosan NPs with 99m Tc in the presence of SnCl2 or NaBH4 as reducing agents [91]. In this study, the use of SnCl2 in the absence of a metal chelator led to the formation of radiocolloids, which could not be separated from the NP mixture. The formation of colloids was not observed with NaBH4 , but the reduced Tc was weakly bound on the NPs. For these reasons, the in vivo administration of the first batch (SnCl2 ) of 99m Tc-labeled chitosan NPs in normal rabbits resulted in high radioactivity uptake in lungs and liver (colloids) while the administration of the second batch (NaBH4 ) produced high radioactivity uptake in the stomach (99m Tc-pertechnetate) [91]. Presumably, partial radiolabel release in vivo occurred also when injecting glucosamine modified MWNTs directly labeled with 99m Tc and SnCl2 . In fact, in the biodistribution study published by Guo et al. [92], the uptake and release of radioactivity in the stomach in the first few hours after MWNT administration mirrors that of free 99m Tc-pertechnetate [93]. Such high stomach uptake was not observed with SWNTs carrying a variety of functionalities and radiolabels [24, 53, 68]. The release of pertechnetate from directly labeled NPs is not a general phenomenon. In fact, Agashe et al. [94] directly labeled PPI dendrimers with 99m Tc-pertechnetate and SnCl2 and proved that the label was stably attached to the NPs both in vitro and in vivo. Similarly, Park and co-workers did not observe any release of 99m Tc-pertechnetate after intratumor injection of directly 99m Tc-labeled pullulan NPs [95]. Solid lipid nanoparticles were also labeled with 99m Tc in aqueous solution though a direct method [96, 97]. Interestingly, despite the paucity of heteroatoms in the structure, these NPs were able to retain 99m Tc in vitro when challenged with excess DTPA and Cys, and did not release the label in vivo. A more traditional approach to label the lipid core of lipid nanocapsules with 99m Tc was reported by Cahouet et al. [98]. In an early study, 99m Tcoxine was encapsulated in the NP lipid core while 125 I was used to label PE in the shell of the capsule with the Bolton–Hunter method [98]. Unfortunately, shortly after intravenous injection the 125 I activity was found in plasma while the 99m Tc activity was associated with

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red blood cells, reasonably due to 99m Tc-oxine exchange with red blood cell membranes. Markedly different results were obtained when including a lipophilic 99m Tc/188 Re-SSS complex (Fig. 17.3) in the NP core [99]. This formulation proved to be very stable in vivo and no release of radiolabel was observed in normal rodents 24 hours after intravenous administration. In addition, NPs loaded with 188 Re with the same strategy were injected intratumor in a glioma model and only 10% of the injected dose was released as perrhenate from the injection site in 3 days [100]. Various direct labeling approaches were used also with magnetic NPs. Silica-coated magnetite and SPIOs were functionalized with an antibody targeting human hepatocarcinoma cells and directly labeled with 188 Re-perrhenate after reduction of the protein disulfide bridges to sulfhydryl groups [101, 102]. With this approach, the radiometal incorporation was efficient but approximately 25% of the NP-bound radioactivity was released in vitro in 24 hours. On the contrary, when the surface of the magnetite/silica NPs was functionalized with a chelator (His) and reacted with the [188 Re(CO)3 (H2 O)3 ]+ synthon, less than 20% of the bound radioactivity was released in vitro over a 3-day period [103]. To the best of the author’s knowledge, these 188 Re-NPs were not tested in vivo. Similar to 111 In and 67 Ga, liposomes can be labeled with 99m Tc with an after-loading approach by using a lipophilic 99m Tc complex. In the method introduced by Phillips and co-workers, 99m Tc-HMPAO (Fig. 17.3) is used to shuttle the radiometal through the bilayer of the liposome [104]. Preloaded glutathione reduces the complex to a more hydrophilic species that is trapped in the aqueous core and retained in vivo. This labeling approach is highly efficient (85–90% 99m Tc incorporation) and it is not influenced by the size of the liposomes [105], by the lipid composition [106, 107], or by the presence of PEG on the liposome surface [108]. Another lipophilic moiety used for after-loading liposome labeling is the SNS/S complex 99m Tc-BMEDA (Fig. 17.3) [109]. This complex contains tertiary amino functions that become protonated at low pH, and therefore it can be trapped inside the liposome core by using a pH gradient. Bao and co-workers used this approach to label Doxil with good yields (40–80% 99m Tc encapsulation) and proved that this labeling procedure does not alter the in vivo distribution of the NP [110]. However, slow release of the 99m Tc complex from the liposome in vivo caused some radioactivity accumulation in the kidney in rats. Surprisingly, when doxorubicin-free liposomes were labeled with 186 Re-BMEDA with the same procedure, a higher stability was observed both in vitro and in vivo [111]. BFC-based strategies to label liposomes with 99m Tc have also been developed [40]. At first, the use of single-chain fatty acid to anchor the DTPA to the lipid bilayer gave relatively poor in vivo stability, reasonably due to lipid exchange with blood components [112]. The stability of 99m Tc-liposomes was significantly improved by attaching the BFC to a double-chain fatty acid [113]. More recently, Laverman and co-workers developed a method to 99m Tc-label liposomes conjugated to the HYNIC chelator, using tricine as a coligand (Fig. 17.4) [114]. When compared with the after-loading method based on 99m Tc-HMPAO, the HYNIC method showed higher incorporation efficiency, with no need for postlabeling purification, and less in vitro transchelation in the presence of excess DTPA and Cys. Due to the high stability, higher radioactivity uptake in bacterial infection loci and lower radioactivity uptake in kidneys were observed in rats administered 99m Tc-HYNIC liposomes compared to rats injected with a 99m TcHMPAO liposomal preparation. Significant differences were also observed in liver and spleen uptakes, reasonably due to the different surface chemistries of these two NP preparations [114].

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H N

O

N N O

O

N

O

O

M HO HO

N

O

N

OH OH

OH

FIGURE 17.4 Structure of the 99m Tc-HYNIC(tricine)2 complex.

17.3.6 Iodine Radiolabeling Iodine has a variety of radioisotopes that can be used in radioactive assays and nuclear medicine procedures, with a wide range of half-lives and emitted radiations. The long-lived ␥ -emitter 125 I (59.4-day half-life) emits low-energy photons and is mostly used for in vitro assays, ex vivo biodistribution, and autoradiography studies. The ␥ - and ␤− -emitter 131 I (8day half-life) was used for both radiotherapy of thyroid pathologies and diagnostic imaging for over four decades. Nowadays, 123 I (13.2-hour half-life) has replaced this “classic” iodine isotope in SPECT imaging because of the lower radiation dose and because of the 159-keV ␥ emission, ideal for ␥ -cameras. Finally, in recent years the ␤+ -emitter 124 I (4.2-day half-life) has received much consideration for PET imaging [85]. For decades, antibodies and proteins have been radioiodinated using radioiodide in the presence of mild oxidizing agents such as Iodogen and CAT. These reagents oxidize iodide to iodine, which then reacts at the position ortho of Tyr residues in situ. The same strategy is applied to polymers containing Tyr residues or phenol groups. For instance, the Iodogen method [115, 116] and CAT [117] were used to radioiodinate HPMA polymers obtained by radical polymerization in the presence of small amounts (0.3–1 mol %) of methacryloyltyrosinamide. [125 I]Iodide and CAT were also used to label several polyester dendrimers containing phenol rings or tyramine groups in their backbone [118–120]. Interestingly, the radioiodide/CAT approach was applied also to NP lacking phenol groups. For instance, Wang and co-workers used this method to covalently bind 125 I to SWNTols, presumably at defect sites on the wall and at the open ends of the tubes [26]. As an alternative to labeling phenol groups, NPs containing amino functions are reacted with the Bolton–Hunter reagent [121, 122] or with succinimidyl 3-[125 I]iodobenzoate [123, 124]. Cell internalization followed by lysosomal degradation is a major concern when using compounds radioiodinated on Tyr residues. In fact, the resulting lipophilic radiocatabolite (iodotyrosine) diffuses back through cellular membranes and is poorly retained in tissues [125]. Furthermore, if dehalogenation occurs in vivo, the released radioiodide concentrates in organs expressing the sodium–iodide symporter, such as thyroid, stomach, and salivary gland [93]. The degradation processes, however, may happen slowly when the radiolabel is inserted within the core of a macromolecular assembly. In addition, the degradation mechanisms may change depending on the structure and composition of the NP.

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Interestingly, when evaluating the influence of molecular weight on the biodistribution and tumor accumulation of radioiodinated poly(HPMA) in rats, small polymers (20–30 kDa) showed a higher release of radioiodide compared to a ∼60-kDa polymer, as suggested by the different radioactivity uptake in the thyroids in the scintigraphic images [115, 116]. This may reflect the higher chance of the radioiodinated phenols to be situated close to the surface, exposed to enzymatic degradation, in smaller NPs compared to larger ones. Unfortunately, the authors of these studies did not determine the content of radioactivity in the rat stomach and did not analyze the urine to quantify the amount of radioiodide that was released in vivo. On the contrary, no significant release of radioiodide from the Tyr residues in a series of PEGylated polyester dendrimers was detected in vivo [119]. However, 48 hours after NP injection in mice, Guillaudeu and co-workers [119] observed the presence of a high amount of small molecular weight radiometabolites in the urine (65% of the total activity), suggesting dendrimer degradation. The nature of these degradation products was not investigated further. A yet different behavior was observed by Wilbur et al. [124] and Sato et al. [123] after the injection in mice of biotinylated PAMAM dendrimers labeled with succinimidyl 3-[125 I]iodobenzoate. In these studies, very low radioactivity accumulation was observed in thyroid and stomach 4 hours after the NP administration [124] and only intact dendrimer was detected in urine, suggesting in vivo NP stability [123]. High in vivo stability was observed also for the 125 I-SWNTols. In the study by Wang et al. [26], these radioiodinated NPs were administered in mice though different routes (intravenously, subcutaneously, intraperitoneally, and orally). In all cases, the SWNTols exhibited a rapid distribution throughout the entire body and a fast urinary excretion. Noticeably, the amount of radioactivity in the stomach was higher than that observed after administration of a comparable amount of free radioiodide [26] and similar to that reported by Liu et al. [53] for PEGylated SWNT labeled with 64 Cu. Therefore, although the metabolism of 125 I-SWNTols was not investigated, the authors concluded that the presence of radioactivity in the stomach was probably due to accumulation of 125 I-SWNTols in this organ and not to dehalogenation. 17.3.7 Bromine Radiolabeling In recent years, there has been considerable interest in using 76 Br for PET evaluation of long-circulating molecules, especially antibodies [85]. The high abundance ␤+ emission and high production yields in medical cyclotrons [126] combined with the convenient halflife (16.2 hours) make this PET radionuclide a valid alternative to 124 I. As with iodine, NPs containing phenol groups can readily be radiolabeled using electrophilic [76 Br]bromine. This reactive species is obtained in situ from bromide with oxidizing agents (CAT, peracetic acid, etc.) or enzymes (bromoperoxidase/H2 O2 ) [127–129]. One advantage of bromine over iodine is that 76 Br tracers are less prone to dehalogenate in vivo compared to radioiodinated analogs due to the stronger C Br bond. Unfortunately, if in vivo debromination occurs, the elimination of [76 Br]bromide from the system is very slow (10–12 days) [130]. Almutairi et al. [131] used 76 Br and small animal PET imaging to evaluate the in vivo distribution of PEGylated dendrimers targeting the ␣v ␤3 integrin. In this study, the radiohalogen was efficiently bound to the Tyr functionalized core of the dendrimer using CAT. The nanoprobe cleared effectively from the blood (less than 1 % ID/g 48 hours postinjection in rats) and a small amount of a small molecular weight radiometabolite was detected in plasma (approximately 30% of the radioactivity circulating at 48 hours). No

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significant accumulation of radioactivity was detected in brain and stomach [131], target organs for free radiobromide [132], confirming the in vivo stability of this 76 Br-NP. 17.3.8 Fluorine Radiolabeling Fluorine-18 is probably the most important PET radionuclide, as it is used to produce the widely used 18 FDG and many other tracers. Fluorine-18 decays 97% by ␤+ emission and 3% by EC. The maximum energy of the emitted ␤+ (0.633 MeV) is ideal for imaging with PET cameras because of the short mean ␤+ range (0.69 mm) providing the highest resolution of all PET radionuclides. The relatively long half-life of 18 F (109.8 min) allows the distribution of [18 F]fluoride within a few hours drive from commercial sources and a variety of synthetic strategies can be carried out. However, considering the length of the syntheses and purification steps, the use of PET scanners to assess the distribution of 18 F tracers in living systems is limited to a few hours. NP radiolabeling with 18 F can be accomplished by using small reactive 18 F-labeled synthons that bind to chemical functionalities on the NP surface. For instance, Ducong´e and co-workers functionalized the surface of PEGylated phospholipid micelles encapsulating QDs with thiol groups and then used a 18 F-maleimido agent (18 FPyME) for radiolabeling [133]. The complete 18 F-NP radiosynthesis was accomplished within 145 min from the 18 F production with a high specific activity. The stability of the 18 F-QDs was confirmed both in vitro and in vivo. Dynamic PET imaging in mice (90 min) showed a long blood circulation (approximately 2-hour half-life) and delayed MPS uptake for these PEGylated QDs. The QD uptake in liver and spleen macrophages was also monitored at the cellular level by in vivo fiber confocal fluorescence imaging and then confirmed ex vivo by fluorescence microscopy [133]. Willmann et al. [134] used a different 18 F-synthon (succinimidyl [18 F]fluorobenzoate) to label the surface of targeted MBs designed for US imaging of angiogenesis. In this case, 18 F labeling was performed on the targeting ligand (a biotinylated antibody) to retain the integrity of the gas-filled NPs. After radiosynthesis, the 18 F-antibody was bound to the NP surface using biotin–streptavidin interaction. Dynamic PET imaging (90 min) confirmed very short MB circulation in blood (few minutes) and rapid uptake in MPS organs in mice. However, incomplete blood clearance and slow accumulation of radioactivity in the kidney suggested partial degradation of 18 F-MBs in vivo. The direct nucleophilic radiofluorination of NPs was also reported. In a recent study, Matson and co-workers synthesized crosslinked polymeric micelles carrying mesylate groups in the outer shell [135]. These robust NPs were efficiently and safely labeled with [18 F]fluoride in acetonitrile in the presence of Kriptofix-222 (1-hour reaction at 120 ◦ C). The biologic studies on these 18 F-NPs are ongoing. Fluorine-18 is used also for liposome radiolabeling. At first, 18 FDG was encapsulated in the liposome core during manufacturing [136]. With this straightforward labeling approach, Oku and co-workers performed a real-time analysis of liposomal in vivo trafficking with PET and thoroughly evaluated the influence of size, composition, and charge on the early biodistribution of long-circulating stealth liposomes encapsulating 18 FDG [137]. Recently, a more general approach to 18 F-label NPs comprising a lipophilic domain was reported by Marik and co-workers [138]. In this study, the authors synthesized a dipalmitin derivative carrying a tosylate group, which was 18 F-labeled with high specific activity (greater than 110 GBq/mmol) and incorporated in the membrane of long-circulating liposomes during manufacturing (approximately 2 hours for the complete radiosynthesis). As expected, dynamic PET showed a near-constant level of radioactivity in blood and major organs during

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FIGURE 17.5 Representative full body PET projection (90-min continuous bed motion scan) of a normal rat injected with long-circulating PEGylated liposomes radiolabeled with [18 F]fluorodipalmitin. (Adapted from Marik et al. [138].)

the 90-min scan (Fig. 17.5) [138]. The short half-life of 18 F did not allow extending the imaging study beyond this point. The same 18 F-dipalmitin derivative was used also to radiolabel the shell of lipid MBs [139]. Compared to stealth liposomes, 18 F-MBs exhibited a much shorter blood circulation and a higher spleen uptake. Dynamic PET imaging was also capable of quantifying lipid material deposition in the kidney as the result of MB disruption following US pulses.

17.4 ANGIOGENESIS IMAGING WITH RADIOLABELED NANOPLATFORMS Several formulations containing NPs chemically bound to or entrapping a chemotherapeutic agent have reached the market or are in clinical trials for the treatment of cancer [140]. These agents accumulate in cancer tissues mainly through passive targeting (EPR effect) [22]. By attaching specific ligands to the NP surface, active targeting of tumor cells can be accomplished. For instance, NPs conjugated to folic acid [28, 72, 141] and transferrin [48, 142, 143] are known to specifically enter cancer cells through endocytic pathways triggered by the interaction with overexpressed membrane receptors both in vitro and in vivo. However, in solid tumors the rate-determining step for NP accumulation is the extravasation from the blood vessels. For this reason, the use of targeting ligands has

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generally failed in increasing the overall accumulation of NPs in solid tumors and specific accumulation of targeted NPs was seen only in disseminated lymphoma models [68] and ascitic tumor cells [144]. Obviously, this apparent lack of specificity is a drawback when aiming at using NPs for molecular imaging of cancer, especially with nonactivatable labels such as radionuclides. Furthermore, many candidate imaging NPs are too large or leave the bloodstream too rapidly to diffuse out of the blood vessels into the tumor interstitial space. Therefore the evaluation of targeted NPs for molecular imaging of cancer has been mostly limited to vascular markers of disease (e.g., markers of angiogenesis) that are readily accessible and do not require extravasation. Angiogenesis—the formation of new capillaries by cellular outgrowth from existing microvessels—is a fundamental process in tumor development. Small tumors (1–2 mm) rely on oxygen and nutrient diffusion from the surrounding tissues to survive (avascular phase). Tumor progression and metastasis propagation occur only after induction of a tumor vasculature (angiogenic switch) [145]. Besides tumors, angiogenesis takes place also in other tissues as an adaptation to hypoxic stress. For instance, neovascularization occurs as part of the natural healing process after ischemic injury to restore tissue oxygenation [146]. The presence of hypoxia within thickened atherosclerotic lesions causes the formation of neovessels, which may mark plaque vulnerability [147]. Therefore molecular imaging of angiogenesis could be critical for defining the pathophysiology of a variety of diseases including cancer and atherosclerosis, and for assessing the efficacy of anti- and proangiogenic therapies. An altered expression of integrins and growth factor receptors on endothelial cells are known molecular events associated with the angiogenic process. The next sections summarize the progress in molecular imaging of angiogenesis with radiolabeled NPs targeting the ␣v ␤3 integrin and VEGFR-2. 17.4.1 Imaging of ␣ v ␤3 Integrins Integrins are a family of ␣␤ heterodimeric transmembrane glycoproteins that mediate cell–cell and cell–matrix adhesion through interactions with specific ligands [148]. Integrin signaling plays a key role in immune response, hemostasis, wound healing, and tissue remodeling. Integrin binding supports the development of tumor vasculature, tumor growth, and metastasis. In vascular, inflammatory, autoimmune, or hyperproliferative diseases, integrin function can also become dysregulated and contribute to pathogenesis [148]. It is currently accepted that the vitronectin receptor ␣v ␤3 is absent on quiescent endothelial cells in established vessels and that its expression is induced on endothelial cells undergoing the process of active angiogenesis. However, the exact role of ␣v ␤3 as a proangiogenic or antiangiogenic factor is still a matter of debate [149]. A subset of integrins including the ␣v ␤3 recognizes a common motif in their ligands: the RGD sequence. An arsenal of contrast agents containing single or multiple RGD peptides has been developed for angiogenesis imaging and some of them have made the transition from bench to bedside [150]. To date, there has been only one report on SPECT imaging of angiogenesis with radiolabeled NPs. Hu and co-workers [71] tested 111 In-labeled perfluorocarbon particles functionalized with a quinolone-based vitronectin receptor antagonist [151] in a nascent Vx-2 lung carcinoma tumor implanted in rabbit hindlimbs. The presence of neovessels at the capsular interface between the tumor tissue and the thigh muscle 2 weeks after tumor implant was confirmed with a stain for vascular endothelium. In this animal model, the

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tumor-to-muscle ratio obtained with targeted 111 In-NPs was approximately twofold higher than that observed with nontargeted NPs. In addition, the administration of a threefold excess of nonradioactive targeted NPs competitively blocked the uptake in the neovascularized tumor region, suggesting specific interaction with the ␣v ␤3 integrin. In this study, the uptake in the tumor was visible by ␥ imaging as early as 15 min postinjection and persisted for at least 18 hours [71]. This was a significant improvement with respect to a previous imaging experiment performed with a paramagnetic version of the NP [152]. In fact, the MRI contrast in the tumor periphery was markedly lower than that obtained by ␥ imaging and visible only at later time points. In recent years, the group at Stanford University dedicated major efforts to the development of NP-based probes for multimodal imaging of ␣v ␤3 integrins in tumor neovasculature. Copper-64-labeled QDs and IOs carrying RGD moieties were evaluated for PET/NIRF imaging [49] and PET/MRI imaging [52], respectively. The specific targeting capabilities of these NPs were investigated in mice bearing human glioblastoma xenografts (U87MG), a highly malignant tumor hallmarked by extensive invasion and angiogenesis [153, 154]. Following the intravenous administration of 64 Cu-labeled RGD-QDs in mice, small animal PET confirmed the rapid and extensive QD accumulation in MPS organs. Despite the high MPS uptake, the U87MG xenografts were clearly visible as early as 5 hours postinjection and the contrast increased over time (4.0 ± 1.0% ID/g at 18 hours postinjection) [49]. The specific interaction between RGD-QDs and ␣v ␤3 integrins was confirmed both in vitro (competitive blocking experiment) and in vivo (less than 1% ID/g in tumors in mice injected with nontargeted QDs). The confinement of the vast majority of RGD-QDs into the tumor vasculature expressing ␣v ␤3 was assessed ex vivo by immunohistochemistry and fluorescence microscopy. Unfortunately, due to the higher sensitivity and deeper tissue penetration of PET with respect to optical imaging, the small amount of 64 Cu-QDs used in this PET/NIRF study (20 pmol QDs per mouse) did not allow noninvasive fluorescence imaging. However, a good correlation between the PET and NIRF signals was observed ex vivo in isolated organs [49]. Yet, the findings of these experiments agree with the semiquantitative results of a previous NIRF imaging study carried out with nonradiolabeled RGD-QD in U87MG xenografted mice (200 pmol QDs injected per mouse) [155]. Real dual-modality imaging of tumor angiogenesis was achieved with 64 Cu-labeled RGD-IOs and PET/MRI [52]. In this experiment, small amounts of 64 Cu-labeled NPs were diluted with nonradioactive IOs and injected in mice bearing U87MG xenografts (3.7 MBq 64 Cu and 300 ␮g iron per mouse). The tumors were clearly visualized with small animal PET in mice administered IO-RGD (10.1 ± 2.1% ID/g 4 hours postinjection). On the contrary, almost no contrast was observed in mice injected with nontargeted IOs or coinjected with IO-RGD and c(RGDyK) (10 mg/kg mouse body weight). In vivo T2-weighted MRI was also performed on a clinical 3-T scanner 4 hours post-IO injection. In tumors, a decrease in signal intensity was observed after the injection of DOTA-IO-RGD compared to DOTA-IO and DOTA-IO-RGD with blocking agent (qualitative measurements) [52]. In another study by the same group, SWNTs were functionalized with DOTA and c(RGDyK) functionalities appended at the end of phospholipid-PEG chains of two different lengths (2 and 5.4 kDa) [53]. Quantitative analysis of the PET images in mice confirmed a significantly longer blood circulation of 64 Cu-labeled PEG5400 SWNTs (2hour half-life) with respect to the PEG2000 derivatives (30-min half-life). Surprisingly, the SWNTs functionalized with the shorter RGD-PEG chain exhibited only a slightly higher retention in U87MG tumors (∼3–4 % ID/g over a 24-hour period) with respect to the nontargeted analogs (∼2.5% ID/g). On the contrary, the tumor uptake of the RGD-PEG5400

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functionalized SWNTs was significantly higher than that of the nontargeted analogs at each considered time point (∼13% ID/g and ∼4% ID/g, respectively, over a 24-hour period). Furthermore, this tumor uptake was competitively blocked by coinjecting c(RGDyK) (15 mg/kg mouse body weight). The quantitative measurements from PET image analysis were confirmed by Raman spectroscopy [53]. The reason why long-circulating SWNTs accumulated more in tumor compared to the short-circulating analogs in a ␣v ␤3 -dependent fashion was not investigated further by the authors. Previous work carried out with different CNTs revealed that these NPs easily pass through a number of compartments to reach targets in the whole mouse [92]. Therefore, possibly, the long-circulating RGD-SWNTs extravasated from the tumor blood vessels with a mechanism that was initiated or facilitated by ␣v ␤3 -mediated adhesion to the endothelial wall. Alternatively, the high tumor uptake of long-circulating RGD-SWNTs could have been the result of ␣v ␤3 -mediated internalization in endothelial cells followed by integrin recycling to the cell surface. Recently, Almutairi et al. [131] designed a dendritic imaging NP carrying multiple RGD copies appended to PEG chains and labeled with 76 Br for PET imaging of angiogenesis. This nanoprobe was tested in a mouse model of hindlimb ischemia producing slow revascularization of the thigh muscle [156, 157]. In normal rodents, the dendritic NP exhibited sustained blood circulation and did not accumulate significantly in nontarget organs (including the MPS organs). Seven days after the surgical intervention in the mouse hindlimbs, intravenous injection of the ␣v ␤3 -targeted nanoprobe resulted in radioactivity accumulation in the entire ischemic muscle compared to the nonischemic limb (sham surgery), particularly in the area distal to the femoral artery excision (Fig. 17.6). On the contrary, the

(a)

(b)

(c)

FIGURE 17.6 (A) Structure of a 76 Br-labeled dendritic nanoprobe targeting the ␣v ␤3 integrin designed for PET images of angiogenesis. (B) Representative small animal PET/CT coronal and sagittal slices obtained 24 hours after the injection of the RGD nanoprobe (left mouse) and nontargeted nanoprobe (right mouse) in a mouse model of hindlimb ischemia. The images show high RGD nanoprobe uptake in the ischemic limb with respect to the control limb (sham injury) and with respect to the nontargeted nanoprobe. (C) Quantitative analysis of PET images (right-to-left hindlimb ratios) showing an approximate twofold higher uptake of the RGD nanoprobe in the distal and proximal areas of the ischemic thigh respect to the non-targeted nanoprobe. (Adapted from Almutairi et al. [131].)

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administration of nontargeted nanoprobe did not produce significant contrast [131]. This behavior was similar to that observed with a small 99m Tc-RGD derivative in a similar animal model [157], thus suggesting RGD-dendrimer uptake in angiogenic vessels.

17.4.2 Imaging of VEGFR-2 Among the seven structurally related VEGF molecules, VEGF-A (usually referred to as VEGF) is the major mediator of tumor angiogenesis [158]. VEGF is a homodimeric glycoprotein existing in several splice variants differing in the number of amino acids and heparin-binding properties. VEGF121 is a freely diffusible protein. The dominant isoform VEGF165 is also secreted, but a significant fraction is bound to heparin-like moieties on the cell surface and ECM. The longer VEGF189 and VEGF206 are almost completely sequestered by the ECM and can be released in diffusible form by the action of heparin or heparinase [159]. Most types of human cancer cells express VEGF as a response to environmental factors (hypoxia, pH, inflammatory cytokines, sex hormones, etc.) or genetic changes (activation of oncogenes or inactivation of tumor-suppressor genes) and increased expression is often associated with poor prognosis [159]. VEGF signals mainly through the tyrosine kinase VEGFR-2, which is expressed at elevated levels by endothelial cells. VEGF diffusion and binding to VEGFR-2 leads to a cascade of signaling events resulting in proliferation, migration, and survival of endothelial cells and increased vascular permeability, ultimately leading to the formation of tumor neovessels [158, 159]. The pivotal role of VEGF/VEGFR signaling in tumor angiogenesis has prompted significant efforts in the search for antiVEGF treatments to “starve” solid tumors [160]. Besides tumors, VEGF/VEGFR signaling is involved also in several hematologic malignancies, intraocular neovascularization syndromes, inflammatory disorders, and other pathologies and plays an important role in bone healing and in the angiogenesis associated with myocardial infarction and ischemia [146, 159]. The development of reliable VEGF/VEGFR-targeted imaging tracers has also been the subject of extensive research. Several VEGF121 and VEGF165 derivatives, anti-VEGF antibodies, and VEGFR-2-specific fusion proteins were functionalized with radionuclides and optical labels or conjugated to detectable NPs for angiogenesis imaging and therapy monitoring [161]. Chen and co-workers conjugated VEGF and DOTA to QDs to perform dual-modality imaging of tumor angiogenesis and evaluated these nanoconjugates in U87MG xenografts with PET (64 Cu) and NIRF (Fig. 17.7) [61]. Small animal PET showed a significantly higher tumor uptake of 64 Cu-labeled VEGF-QDs (up to 4.5 ± 0.5% ID/g 24 hours postinjection) compared to nontargeted QDs (<1% ID/g). Also in this case, as for 64 Cu-labeled RGD-QDs, NIRF imaging was carried out on a different group of mice, administered a 10-fold higher amount of QDs, and showed good agreement with PET (Fig. 17.7). Immunofluorescence staining on tumor slices treated with an anti-VEGFR-2 antibody showed a precise overlay between the receptor expression and QD fluorescent signal. Interestingly, prominent receptor expression was visible only on small and immature tumor vessels and not on bigger ones [61]. Nonhomogeneous distribution of VEGFR-2 on the U87MG tumor vasculature and a correlation between tumor size, VEGFR-2 expression, and 64 Cu-DOTAVEGF121 uptake were the findings of another study published by Chen and co-workers [162]. These results further stress the importance of imaging VEGFR expression before

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(a)

(b)

(c)

O O N H

N

NH—VEGF

S O

O

ZnS HN O N

HO C N

64Cu N

CO H

N

HO C

FIGURE 17.7 (A) Representative structure of a 64 Cu-labeled DOTA-QD-VEGF conjugate designed for PET/NIRF imaging of angiogenesis. (B) Representative small animal PET coronal slices obtained 1 hour and 24 hours after the injection of 10 pmol nontargeted QDs (top) and VEGF-QDs (bottom) in mice bearing U87MG xenografts. (C) in vivo NIRF images of mice bearing U87MG xenografts 1 hour after the injection of 200 pmol nontargeted QDs (top) and VEGF-QDs (bottom). White arrows indicate the tumors. The images show high radioactivity uptake in liver and spleen. The tumors of the mice administered QD-VEGF are clearly visible with both modalities. (Adapted from Chen et al. [61].)

and during anti-VEGFR cancer treatments, due to the dynamic nature of VEGFR expression through tumor progression and to the high variability among different tumor types. Recently, the feasibility of anti-VEGFR-2 therapy monitoring with targeted MBs and contrast-enhanced ultrasonography was proved in a murine model of human pancreatic adenocarcinoma [163]. Hence Willmann and co-workers used a 18 F-labeled anti-VEGFR-2 to functionalize MBs and performed dynamic small animal PET imaging to assess the whole body distribution and absolute tumor targeting capabilities of the targeted MBs [134]. As expected, the MBs disappeared from the blood circulation within a few minutes from intra-venous injection and accumulated in MPS organs. However, the MB retention in subcutaneous angiosarcomas was significantly higher than that in the adjacent skeletal muscle during the whole imaging session (1.14% ID/g in tumor and 0.84% ID/g in muscle, averaged over 60 min) [134]. The tumor targeting capabilities of these MBs were further enhanced by using a dual-targeting approach [164]. In fact, the administration of MBs targeting both VEGFR-2 and ␣v ␤3 produced significantly higher US contrast in ovarian adenocarcinoma xenografts compared to that obtained with MBs targeting only VEGFR-2 or ␣v ␤3 , injected separately or in combination.

17.5 CONCLUSION Radionuclide-based assays and noninvasive imaging techniques are powerful tools in the hands of scientists for assessing the distribution and targeting capabilities of new molecules. These versatile and highly sensitive techniques allow a more comprehensive understanding of the behavior of candidate drugs at the cell and tissue level and can expedite enormously the path from small to large animals and, finally, to humans. For this reason, the production of new clinical and research PET, SPECT, and hybrid imaging devices has flourished over

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the last decade. Nanomedicine also benefits greatly from the use of these techniques. In fact, more and more material scientists adopt radionuclide-based methods to study the characteristics of new nanomaterials in vitro, in vivo, and ex vivo. It is expected that this trend will continue with the increased availability of nonstandard medium- to long-lived radionuclides such as 64 Cu, 89 Zr, 124 I, and 76 Br and with the distribution of radionuclide generators in medical centers and research institutions worldwide. The thrilling achievements in the field of angiogenesis imaging reported in the previous paragraphs, obtained with the support of PET, SPECT, and other imaging modalities, are only one example of the great potential of NPs for molecular imaging of disease. In the near future, the close partnership among material scientists, chemists, biologists, clinicians, engineers, physicists, and mathematicians will produce more sophisticated nanobullets, better targets, higher sensitivity, and higher resolution imaging devices and better reconstruction algorithms for the needs of the healthcare of tomorrow.

ACKNOWLEDGMENTS Special thanks to Dr. Sander Langereis, Dr. Marc Robillard, and Dr. Holger Gr¨ull for stimulating discussions and assistance in writing this chapter.

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92. Guo, J.; Zhang, X.; Li, Q.; Li, W. Biodistribution of functionalized multiwall carbon nanotubes in mice. Nucl. Med. Biol. 2007, 34(5), 579–583. 93. Zuckier, L. S.; Dohan, O.; Li, Y.; Chang, C. J.; Carrasco, N.; Dadachova, E. Kinetics of perrhenate uptake and comparative biodistribution of perrhenate, pertechnetate, and iodide by NaI symporter-expressing tissues in vivo. J. Nucl. Med. 2004, 45(3), 500–507. 94. Agashe, H. B.; Babbar, A. K.; Jain, S.; et al. Investigations on biodistribution of technetium99m-labeled carbohydrate-coated poly(propylene imine) dendrimers. Nanomedicine 2007, 3(2), 120–127. 95. Park, K. H.; Song, H. C.; Na, K.; et al. Ionic strength-sensitive pullulan acetate nanoparticles (PAN) for intratumoral administration of radioisotope: ionic strength-dependent aggregation behavior and 99 mtechnetium retention property. Colloids Surf. B. Biointerfaces 2007, 59(1), 16–23. 96. Reddy, L. H.; Sharma, R. K.; Chuttani, K.; Mishra, A. K. Etoposide-incorporated tripalmitin nanoparticles with different surface charge: formulation, characterization, radiolabeling and biodistribution studies. AAPS J. 2004, 6(3), 23. 97. Thakkar, H.; Kumar Sharma, R.; Murthy, R. S. Enhanced retention of celecoxib-loaded solid lipid nanoparticles after intra-articular administration. Drugs R&D 2007, 8(5), 275–285. 98. Cahouet, A.; Denizot, B.; Hindr´e, F.; et al. Biodistribution of dual radiolabeled lipidic nanocapsules in the rat using scintigraphy and ? counting. Int. J. Pharm. 2002, 242(1-2), 367–371. 99. Ballot, S.; Noiret, N.; Hindr´e, F.; et al. 99m Tc/188 Re-labelled lipid nanocapsules as promising radiotracers for imaging and therapy: formulation and biodistribution. Eur. J. Nucl. Med. Mol. Imaging 2006, 33(5), 602–607. 100. Allard, E.; Hindre, F.; Passirani, C.; et al. 188 Re-loaded lipid nanocapsules as a promising radiopharmaceutical carrier for internal radiotherapy of malignant gliomas. Eur. J. Nucl. Med. Mol. Imaging 2008, 35(10), 1838–1846. 101. Liang, S.; Wang, Y.; Yu, J.; Zhang, C.; Xia, J.; Yin, D. Surface modified superparamagnetic iron oxide nanoparticles: as a new carrier for bio-magnetically targeted therapy. J. Mater. Sci. Mater. Med. 2007, 18(12), 2297–2302. 102. Liang, S.; Wang, Y.; Zhang, C.; Liu, X. Synthesis of amino-modified magnetite nanoparticles coated with Hepama-1 and radiolabeled with 188 Re for bio-magnetically targeted radiotherapy. J. Radioanal. Nucl. Chem. 2006, 269(1), 3–7. 103. Cao, J.; Wang, Y.; Yu, J.; et al. Preparation and radiolabeling of surface-modified magnetic nanoparticles with rhenium-188 for magnetic targeted radiotherapy. J. Magn. Magn. Mater. 2004, 277(1-2), 165–174. 104. Phillips, W. T.; Rudolph, A. S.; Goins, B.; Timmons, J. H.; Klipper, R.; Blumhardt, R. A simple method for producing a technetium-99m-labeled liposome which is stable in vivo. Int. J. Rad. Appl. Instrum. B 1992, 19(5), 539–547. 105. Awasthi, V. D.; Garcia, D.; Goins, B. A.; Phillips, W. T. Circulation and biodistribution profiles of long-circulating PEG-liposomes of various sizes in rabbits. Int. J. Pharm. 2003, 253(1-2), 121–132. 106. Mirahmadi, N.; Babaei, M. H.; Vali, A. M.; Johari Daha, F.; Kobarfard, F.; Dadashzadeh, S. 99m Tc-HMPAO-labeled liposomes: an investigation into the effects of some formulation factors on labeling efficiency and in vitro stability. Nucl. Med. Biol. 2008, 35(3), 387–392. 107. Carmo, V. A.; Ferrari, C. S.; Reis, E. C.; et al. Biodistribution study and identification of inflammation sites using 99m Tc-labelled stealth pH-sensitive liposomes. Nucl. Med. Commun. 2008, 29(1), 33–38. 108. Lee, C.-M.; Choi, Y.; Huh, E. J.; et al. Polyethylene glycol (PEG) modified 99m Tc-HMPAOliposome for improving blood circulation and biodistribution: the effect of the extent of PEGylation. Cancer Biother. Radiopharm. 2005, 20(6), 620–628.

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127. Lovqvist, A.; Sundin, A.; Ahlstrom, H.; Carlsson, J.; Lundqvist, H. 76 Br-labeled monoclonal anti-CEA antibodies for radioimmuno positron emission tomography. Nucl. Med. Biol. 1995, 22(1), 125–131. 128. Lovqvist, A.; Sundin, A.; Ahlstrom, H.; Carlsson, J.; Lundqvist, H. Pharmacokinetics and experimental PET imaging of a bromine-76-labeled monoclonal anti-CEA antibody. J. Nucl. Med. 1997, 38(3), 395–401. 129. McElvany, K. D.; Welch, M. J. Characterization of bromine-77-labeled proteins prepared by using bromoperoxidase. J. Nucl. Med. 1980, 21(10), 953–960. 130. Gudjonsson, O.; Bergstr¨om, M.; Kristjansson, S.; et al. Analysis of 76 Br-BrdU in DNA of brain tumors after a PET study does not support its use as a proliferation marker. Nucl. Med. Biol. 2001, 28(1), 59–65. 131. Almutairi, A.; Rossin, R.; Shokeen, M.; et al. Biodegradable dendritic positron emitting nanoprobes for the non invasive imaging of angiogenesis. Proc. Natl. Acad. Sci. U.S.A. 2009, 106(3), 685–690. 132. Soremark, R.; Ullberg, S. Distribution of bromide in mice. An autoradiography study with Br-82. Int. J. Appl. Radiat. Isot. 1960, 8, 192–197. 133. Duconge, F.; Pons, T.; Pestourie, C.; et al. Fluorine-18-labeled phospholipid quantum dot micelles for in vivo multimodal imaging from whole body to cellular scales. Bioconjug. Chem. 2008, 19(9), 1921–1926. 134. Willmann, J. K.; Cheng, Z.; Davis, C.; et al. Targeted microbubbles for imaging tumor angiogenesis: assessment of whole-body biodistribution with dynamic micro-PET in mice. Radiology. 2008, 249(1), 212–219. 135. Matson, J. B.; Grubbs, R. H. Synthesis of fluorine-18 functionalized nanoparticles for use as in vivo molecular imaging agents. J. Am. Chem. Soc. 2008, 130(21), 6731–6733. 136. Oku, N.; Tokudome, Y.; Tsukada, H.; Okada, S. Real-time analysis of liposomal trafficking in tumor-bearing mice by use of positron emission tomography. Biochim. Biophys. Acta 1995, 1238(1), 86–90. 137. Oku, N. Delivery of contrast agents for positron emission tomography imaging by liposomes. Adv. Drug Deliv. Rev. 1999, 37(1-3), 53–61. 138. Marik, J.; Tartis, M. S.; Zhang, H.; et al. Long-circulating liposomes radiolabeled with [18 F]fluorodipalmitin ([18 F]FDP). Nucl. Med. Biol. 2007, 34(2), 165–171. 139. Tartis, M. S.; Kruse, D. E.; Zheng, H.; et al. Dynamic microPET imaging of ultrasound contrast agents and lipid delivery. J. Control. Release 2008, 131(3), 160–166. 140. Cho, K.; Wang, X.; Nie, S.; Chen, Z.; Shin, D. M. Therapeutic nanoparticles for drug delivery in cancer. Clin. Cancer Res. 2008, 14(5), 1310–1316. 141. Leamon, C. P.; Reddy, J. A. Folate-targeted chemotherapy. Adv. Drug Deliv. Rev. 2004, 56(8), 1127–1141. 142. Qian, Z. M.; Li, H.; Sun, H.; Ho, K. Targeted drug delivery via the transferrin receptor-mediated endocytosis pathway. Pharmacol. Rev. 2002, 54(4), 561–587. 143. Sahoo, S. K.; Labhasetwar, V. Enhanced antiproliferative activity of transferrin-conjugated paclitaxel-loaded nanoparticles is mediated via sustained intracellular drug retention. Mol. Pharm. 2005, 2(5), 373–383. 144. Gabizon, A.; Horowitz, A. T.; Goren, D.; Tzemach, D.; Shmeeda, H.; Zalipsky, S. In vivo fate of folate-targeted polyethylene-glycol liposomes in tumor-bearing mice. Clin. Cancer Res. 2003, 9(17), 6551–6559. 145. Folkman, J. Angiogenesis in cancer, vascular, rheumatoid and other disease. Nat. Med. 1995, 1, 27–30. 146. Sinusas, A. J. Imaging of angiogenesis. J. Nucl. Cardiol. 2004, 11(5), 617–633.

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147. Doyle, B.; Caplice, N. Plaque neovascularization and antiangiogenic therapy for atherosclerosis. J. Am. Coll. Cardiol. 2007, 49(21), 2073–2080. 148. Staunton, D. E.; Lupher, M. L.; Liddington, R.; Gallatin, W. M.; Frederick, W. A. Targeting integrin structure and function in disease. Adv. Immunol. 2006, 91, 111–157. 149. Hodivala-Dilke, K. ␣v ␤3 Integrin and angiogenesis: a moody integrin in a changing environment. Curr. Opin. Cell. Biol. 2008, 20(5), 514–519. 150. Beer, A.; Schwaiger, M. Imaging of integrin ␣v ␤3 expression. Cancer Metastasis Rev. 2008, 27(4), 631–644. 151. Harris, T. D.; Kalogeropoulos, S.; Nguyen, T.; et al. Design, synthesis, and evaluation of radiolabeled integrin ␣v ␤3 receptor antagonists for tumor imaging and radiotherapy. Cancer Biother. Radiopharm. 2003, 18(4), 627–641. 152. Winter, P. M.; Caruthers, S. D.; Kassner, A.; et al. Molecular imaging of angiogenesis in nascent Vx-2 rabbit tumors using a novel ␣v ␤3 -targeted nanoparticle and 1.5 tesla magnetic resonance imaging. Cancer Res. 2003, 63(18), 5838–5843. 153. Millauer, B.; Shawver, L. K.; Plate, K. H.; Risaui, W.; Ullrich, A. Glioblastoma growth inhibited in vivo by a dominant-negative Flk-1 mutant. Nature 1994, 367(6463), 576–579. 154. Sallinen, S-L; Sallinen, P. K.; Haapasalo, H. K.; et al. Identification of differentially expressed genes in human gliomas by dna microarray and tissue chip techniques. Cancer Res. 2000, 60(23), 6617–6622. 155. Cai, W.; Shin, D.-W.; Chen, K.; et al. Peptide-labeled near-infrared quantum dots for imaging tumor vasculature in living subjects. Nano. Lett. 2006, 6(4), 669–676. 156. Couffinhal, T.; Silver, M.; Zheng, L. P.; Kearney, M.; Witzenbichler, B.; Isner, J. M. Mouse model of angiogenesis. Am. J. Pathol. 1998, 152(6), 1667–1679. 157. Hua, J.; Dobrucki, L. W.; Sadeghi, M. M.; et al. Noninvasive imaging of angiogenesis with a 99m Tc-labeled peptide targeted at ␣v ␤3 integrin after murine hindlimb ischemia. Circulation 2005, 111(24), 3255–3260. 158. Kerbel, R. S. Tumor angiogenesis. N. Engl. J. Med. 2008, 358(19), 2039–2049. 159. Ferrara, N. Vascular endothelial growth factor: basic science and clinical progress. Endocr. Rev. 2004, 25(4), 581–611. 160. Tonra, J. R.; Hicklin, D. J. Targeting the vascular endothelial growth factor pathway in the treatment of human malignancy. Immunol. Invest. 2007, 36(1), 3–23. 161. Cai, W.; Chen, X. Multimodality imaging of vascular endothelial growth factor and vascular endothelial growth factor receptor expression. Front. Biosci. 2007, 12, 4267–4279. 162. Chen, K.; Cai, W.; Li, Z.-B.; Wang, H.; Chen, X. Quantitative PET imaging of VEGF receptor expression. Mol. Imaging Biol. 2009, 11(1), 15–22. 163. Korpanty, G.; Carbon, J. G.; Grayburn, P. A.; Fleming, J. B.; Brekken, R. A. Monitoring response to anticancer therapy by targeting microbubbles to tumor vasculature. Clin. Cancer Res. 2007, 13(1), 323–330. 164. Willmann, J. K.; Lutz, A. M.; Paulmurugan, R.; et al. Dual-targeted contrast agent for US assessment of tumor angiogenesis in vivo. Radiology 2008, 248(3), 936–944.

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PART III

NANOPARTICLE PLATFORMS AS MULTIMODALITY IMAGING AND THERAPY AGENTS

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CHAPTER 18

Lipoprotein-Based Nanoplatforms for Cancer Molecular Imaging IAN R. CORBIN Department of Medical Biophysics, University of Toronto, and Division of Biophysics and Bioimaging, Ontario Cancer Institute, Toronto, Ontario, Canada

KENNETH NG Institute of Biomaterials and Biomedical Engineering, University of Toronto, Toronto, Ontario, Canada

GANG ZHENG Department of Medical Biophysics and Institute of Biomaterials and Biomedical Engineering, University of Toronto, and Division of Biophysics and Bioimaging, Ontario Cancer Institute, Toronto, Ontario, Canada

18.1 LIPOPROTEINS: NATURE’S ENDOGENOUS NANOCARRIER Efficient transport of key lipids, such as triacylglycerols (TAGs) and cholesterol, is an essential process in mammalian systems. These molecules respectively serve as a high-energy fuel source and as a structural building block for all cells. Like other essential nutrients and substrates these molecules must be delivered to cells through the blood. The hydrophobicity of these molecules, however, poses a daunting problem for their transport through the aqueous channels of the vascular system. Nature has uniquely addressed this dilemma through the lipoprotein carrier system. Plasma lipoproteins are macromolecular vehicles that transport neutral lipids (fat and cholesterol) through the circulatory system and extracellular fluid compartments of the body. These spherical emulsion–like complexes display a range of physicochemical properties; however, they have a common structural organization (Fig. 18.1.) consisting of an apolar core of TAG and cholesterol esters surrounded by a monolayer of phospholipids and free cholesterol. Interspersed across the phospholipid monolayer are specific amphipathic proteins (apolipoproteins) whose hydrophilic domains extend toward the aqueous environment. These apolipoproteins act to stabilize and confer structural framework to the lipoproteins, modulate enzyme activity, as well as serve as receptor ligands for the lipoprotein delivery system [1].

Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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FIGURE 18.1 General structure of a lipoprotein.

18.1.1 Classification and Composition of Plasma Lipoproteins Plasma lipoproteins can be designated into various classes based on numerous physical parameters (e.g., electrophoretic mobility, diameter). The most commonly accepted classification is based on the density of the different lipoprotein species (see Table 18.1). According to this classification scheme the major density categories include the following. (1) Chylomicrons (d < 0.95 g/mL) are TAG-rich emulsions particles (80–88% by weight) that are synthesized by the intestine after a fatty meal. Chylomicrons are the largest particles in the lipoprotein family (80 nm to 1 ␮m in diameter) and have the highest lipid-to-protein ratio. (2) Very low density lipoproteins (d = 0.95–1.006 g/mL) are also TAG-rich particles, however, they are synthesized by the liver. They are smaller than chylomicrons (30–80 nm in diameter) and contain relatively less TAG but more cholesterol and protein. (3) Low density lipoproteins (d = 1.020–1.063 g/mL) are particles formed from the intravascular catabolism of VLDLs. The LDL core is predominantly cholesterol ester molecules. Finally, (4) high density lipoproteins (d = 1.063–1.210 g/mL) These carriers are the smallest (6–12 nm in diameter) member of the lipoprotein family. Their core is mainly composed of cholesterol esters and they are composed of a relatively high proportion of protein (35–56% by weight). TABLE 18.1 Density, Size, Electrophoretic Mobility, and Chemical Composition of Lipoproteins [2] Classa

Density (g/mL)

Chylomicrons <0.95 VLDL 0.95–1.006 LDL 1.020–1.063 HDL 1.063–1.210 a b

Diameter Electrophoretic Cholesteryl (nm) Mobility Proteinb Phospholipidb Estersb Cholesterolb TAGa,b 75–100 30–80 18–25 6–12

— Pre-␤ ␤ ␣

1–2 6–10 18–22 35–56

3–6 15–20 18–24 26–32

2–4 6–10 45–50 15–20

1–3 4–8 6–8 3–5

VLDL, very low density lipoprotein; LDL, low density lipoprotein; HDL, high density lipoprotein; TAG, triacylglycerol. Expressed in % dry weight.

80–95 45–65 4–8 2–7

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18.1.2 Plasma Apoproteins Apoproteins play an important role in imparting structure and function to the lipoproteins [1]. To date at least ten different species of apoproteins have been identified in human plasma. These include apo A-I, apo A-II, apo A-IV, apo B-100, apo B-48, apo C-I, apo C-II, apo C-III, apo D, and apo E. Based on structural homologies and solution properties, apolipoproteins can be divided into three groups. Group 1 apoproteins are A-I, A-II, A-IV, C-I, C-II, C-III, and E. These apoproteins, also termed the exchangeable apoproteins, are all water-soluble proteins that transfer between the different classes of lipoproteins within the circulation. With the exception of apo A-II, which is a dimer, all of the group 1 apoproteins consist of a single polypeptide chain. Another unifying feature of this group of proteins is their high degree of amphipathic alpha helix secondary structure [3, 4]. The distribution of the amino acid residues in this structural motif is such that they form a lipid binding domain with polar and nonpolar surfaces. The polar side faces and protrudes into the surrounding aqueous environment, while the nonpolar face penetrates into the hydrophobic region of the phospholipid monolayer that surrounds the neutral lipid core [3, 4]. Apo A-I and A-II are the major protein components on HDL. Both of these apoproteins interact and bind to the scavenger receptor type I (SR-BI) (HDL receptor) [5]. Apo A-I also acts as a cofactor for the enzyme lecithin-cholesterol acyltransferase (LCAT) which catalyzes the conversion of cholesterol and phosphatidylcholine to esterified cholesterol and lysophosphatidylcholine [6]. Furthermore, apo A-I when combined with phospholipids can facilitate the process of reverse cholesterol transport (the transfer of cholesterol from cell membranes back to the liver) [7]. The C apoproteins are present on TAG-rich lipoproteins and HDLs. They work in a coordinated manner to regulate the activity of lipoprotein lipase at the capillary bed. This enzyme hydrolyzes TAGs present in the core of lipoproteins. Apo C-I and C-II activate lipoprotein lipase [8, 9], while apo C-III is postulated to inhibit the activation of this enzyme [10]. Apo E is a component of several lipoprotein species. It serves as an important ligand for LDL receptor-mediated catabolism of lipoproteins [11]. Group 2 apoproteins include the B apoproteins. Apo B proteins are large hydrophobic proteins that contain a large portion of beta sheet and relatively little alpha helix secondary structure [12]. As such, these apoproteins do not exchange between the various lipoproteins. Apo B is present in the plasma in two isoforms: apo B-100 and apo B-48. The apo B gene encodes for a large 549-kDa protein product called apo B-100. This apoprotein is synthesized by the liver and is found on the surface of LDLs and VLDLs, where it acts as a ligand for LDLR. The second isoform of apo B, apo B-48, is found exclusively on chylomicrons. This apolipoprotein is synthesized by the intestine and is a truncated version of apo B-100 consisting of only 48% of the normal gene product [13]. The presence of a translational stop codon in the intestine only translates approximately half of the N-terminal region of apo B-100 [13]. Given that the LDLR binding domain of apo B-100 is located in the carboxy terminal region of the sequence, apo B-48 does not contain any receptor binding properties and thus does not interact with LDLR [14]. Finally, apo D belongs to group 3 apoproteins. These proteins bind and associate with monomeric lipids. Apo D has been found to associate with HDL, cholesterol, progesterone, pregnenolone, bilirubin, and arachidonic acid. The function of apo D has yet to be elucidated [15].

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FIGURE 18.2 Schematic representation of lipoprotein metabolism. See text for details.

18.1.3 Lipoprotein Metabolism Circulating lipoproteins operate in a complex and highly coordinated manner to enable the efficient transport of neutral lipids throughout the body (Fig. 18.2). The lipoprotein carrier system transports up to 200–300 g of TAG and cholesterol between the intestine/liver and peripheral tissues each day [16]. This load of neutral lipids arises from both exogenous (dietary) and endogenous (hepatic synthesis) sources. Dietary lipids are packaged into chylomicrons within the enterocytes of the intestine. These fat-laden particles are released into the circulation, where they are acted upon by extrahepatic capillary bound lipoprotein lipases to liberate fatty acids to peripheral cells. En route within the circulation the chylomicron acquires exchangeable apoproteins (apo A, E, and C) from other lipoproteins. As the chylomicron continues to lose its fatty acid cargo, a resulting smaller chylomicron remnant is eventually recognized (via apo E) and internalized within hepatocytes by LDLR and non-LDLR uptake pathways. Endogenous sources of neutral lipid arise from the liver in the form of VLDL. Like the chylomicron, VLDL circulates to the periphery to release its fatty acid cargo by capillary bound lipoprotein lipases. The resulting TAG depleted VLDL transition through an intermediate density lipoprotein stage, which is partly extracted by the liver via apo E mediated uptake, finally giving rise to the cholesterol ester rich LDL. LDL is metabolized by peripheral and hepatic tissue via the apo B-100 and LDLR pathway. The HDL class of lipoproteins first enters the circulation as lipid-free or lipid-poor apo A-I (secreted by the liver and intestine). In the extracellular medium of peripheral tissues the ATP-binding cassette (ABC)-A1 mediates the efflux of phospholipids and free cholesterol from cells to apo A-I to form nascent HDL particles (initial step of reverse

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cholesterol transport). These newly formed HDL particles are discoidal in shape, consisting only of a phospholipid bilayer (interdigitated with free cholesterol) surrounded by a “belt” of apo A-I protein [17]. Acquisition of core molecules (CE) and its transformation into mature spherical HDL is accomplished through the activity of LCAT, which esterifies the 3-hydroxyl position of cholesterol to form CE. The CE molecule then moves into the center of the nascent HDL to form a hydrophobic core, thereby making the particle spherical. Within the circulation, HDLs are constantly remodeled as LCAT and phospholipid transfer protein (PLTP) introduce CE and phospholipids to HDLs, while cholestryl ester transfer protein (CETP) and hepatic lipases mediate the exchange of CE and TAGs between HDLs and other lipoprotein species. In addition to apo A-I, other exchangeable apoproteins such as apo A-II, C, and E may be incorporated onto the HDL surface from other lipoproteins. HDLs are finally catabolized through interactions with scavenger receptor B1 (SR-BI) mainly in the liver, where core CE are selectively transferred into SRB1 expressing cells without whole particle uptake [18]. The remaining apo A-I on the remnant lipid-free (poor) HDL is removed from the circulation by renal filtration. 18.1.4 Drug Interactions with Plasma Lipoproteins Circulating lipoproteins are highly dynamic macromolecules that are under a constant flux of interchanging lipid and apoprotein components. During their transit time in the vascular system other hydrophobic molecules may also associate with the lipophilic compartments of lipoproteins to assume a more favorable thermodynamic state. This is clearly demonstrated by the fact that several hydrophobic/basic drugs and xenobiotics (e.g., vitamins, cyclosporin A, amiodarone, and amphotericin B) passively associate with plasma lipoproteins [19–21]. While many drugs simply associate with lipoprotein through thermodynamic mechanisms, there has been some evidence to suggest that cholesterol ester transfer protein may facilitate the transfer of drugs between the lipoprotein classes [22, 23]. Regardless of the mechanism of uptake, the compartmentalization of the drug within lipoproteins can occur by absorption either within the apolar core or on the external polar milieu of the particle. The relative preference of the drug to reside in the core or surface of the lipoprotein is likely due to the log P value (n-octanol–water partition coefficient) of the drug. Conceivably, a drug might bind to a site within the apolipoprotein; however, no specific examples of such interactions have been reported in the literature. Once associated with lipoproteins the natural course and disposition of the drug/xenobiotic can become significantly altere [24]. Thus binding or encapsulation of drugs with lipoproteins may facilitate receptor-mediated uptake and sequestration of the drug into tissues expressing high levels of lipoprotein receptors, thereby altering the pharmacokinetic, pharmacodynamic, and toxicological profile of the drug. Additionally, the impact that lipoproteins can have on drug disposition is further exaggerated in conditions of dyslipidemia (disruption in normal distribution of lipoprotein classes within the plasma) [24]. 18.1.5 Cholesterol and Cancer To date, numerous studies have been performed documenting the central role of cholesterol in the development of cardiovascular disease. Substratified analyses of cancer patients within these large clinical studies yielded unexpected results. An association between low plasma cholesterol levels and an increased incidence of cancer was observed [25]. This correlation was found to hold for a number of cancers of different origin [26].

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Hypocholesterolemia, resulting from lower total plasma, LDL, and HDL cholesterol is believed to be a consequence rather than a cause of cancer [27, 28]. This is further supported by the observation that cholesterol levels decrease with disease progression. As well, blood cholesterol levels have been shown to normalize relative to pretreatment levels with successful chemotherapy or remission of the disease [29]. The expression of lipoprotein receptors is markedly elevated on many cancer cells. Early reports of Ho et al. [30] cited that acute myelogenous leukemia has 3–100 times higher LDL receptor activity over normal white blood cells. Similarly, Gal et al. [31] reported that numerous human gynecological cancer cell lines exhibit higher uptake of LDL than corresponding normal tissue. More recently, numerous investigators have documented high levels of HDL uptake in breast cancer cells. It is generally believed that upregulated expression of lipoprotein receptor among cancer cells functions to provide high amounts of cholesterol needed for active membrane turnover [32]. This sequestration of cholesterol to an increasing burden of malignant cells may explain the corresponding low levels of plasma cholesterol found in cancer patients.

18.2 CANCER IMAGING WITH LIPOPROTEIN VEHICLES The previous sections have highlighted (1) the tendency of exogenous lipophilic agents (drugs) to associate with circulating lipoproteins and (2) the strong proclivity that cancer cells have toward acquiring lipoproteins. Given these findings, the next logical step would be to preload or formulate lipoproteins to carry exogenous agents for cancer treatment or detection. In fact, many investigators in the fields of drug delivery and oncology have attempted this dating as far back as 1981 [31]. A number of excellent reviews have been written on the utility of lipoproteins for the delivery of anticancer drugs [33–36]. In this chapter we focus solely on the use of lipoprotein (primarily LDL and HDL) for cancer imaging. 18.2.1 Methods of Probe Incorporation In general, there are three main strategies employed to incorporate probes into lipoproteins. These strategies include covalent modification of the apoprotein component, intercalation of probes within the phospholipid monolayer, and lastly packaging the probe into the lipoprotein core using the “reconstitution” procedure (Fig. 18.3).

Covalent Modification The first strategy in which probes can be incorporated in lipoproteins is through covalent modification of the amino acid residues of the apoproteins. Various residues have been selected as sites for modification; the most common residues include the ε-amino groups of lysine and arginine, the phenolic groups of tyrosine, and the thiol group of cysteine. The number of amino acids that make up the different apoprotein species varies greatly. Apo C-I consist of only 57 amino acids whereas over 4500 amino acids make up apo B-100 [1, 12]. Given the complexity of these multifunctional proteins to conduct lipid binding, receptor recognition, and an array of enzyme catalyzed reactions, these apoproteins should be subject to minimal modification in order to retain normal lipoprotein activity. In the case of apo B-100, the lysine residues that are a common site for chemical modification also play a central role in LDLR recognition. Of the 357 lysine residues within the apo B-100 sequence, 225 are exposed and available for

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FIGURE 18.3 Sites for probe incorporation on or into lipoprotein particles. (1) Covalent modification with amino acid residues in apoprotein component. (2) Surface labeling through intercalation of probes within the phospholipid monolayer. (3) Loading of probe within the lipoprotein core through reconstitution methods.

lipid–protein or protein–protein interactions [37]. However, modification to only 20% of these exposed lysines will abolish apo B-100 capacity to recognize LDLR [37]. Given that apo E is also a ligand for LDLR, it is anticipated that a similar degree of modification to apo E would also result in a similar disruption of receptor recognition. For apo A-I (HDL) its interaction with SR-BI is markedly different from that of the LDL–LDLR system. The apo A-I/SR-BI interaction is complex in that there is not a specific amino acid residue sequence or unique recognition site in apo A-I that mediates binding. Instead, it is the structure motif of the amphipathic ␣-helix present in this apoprotein that is recognized by the SR-BI receptor [38, 39]. As such, several apoproteins (apo A-II, apo-E and apo-C) and multiple sites in the apoprotein may be recognized by SR-BI [38, 40]. Therefore when labeling HDL through covalent modification, the integrity of the amphipathic ␣-helix (clusters of basic amino acid residues located at the polar–nonpolar interfaces of the helix and acidic residues located at the center of the polar face) must be preserved in order to retain normal SR-BI binding. Although covalent modification provides a simple and stable means of labeling lipoproteins, relatively few alterations of the apoprotein are permitted if regular lipoprotein receptor interactions are to be retained. As a result of this stringency, covalent modification methods are limited to probes with high sensitivity, since only a few probe molecules can provide the necessary signal contrast for imaging. Contrast agents such as radiolabels or fluorescent probes lend themselves to this procedure [41, 42]. Numerous reports in the literature have used these methods to study lipoprotein metabolism. Interestingly, excessive modification of apoprotein components has also served as a means to intentionally

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disrupt lipoprotein/receptor interactions (e.g., preparation of acetylated LDL or rerouting of lipoproteins—see Section 18.4) [43, 44].

Intercalation The next strategy in which imaging probes can be incorporated into lipoproteins is through noncovalent association with the phospholipid monolayer. More specifically, the imaging compound nestles or “intercalates” between the phospholipid molecules at the lipoprotein surface. To facilitate this process, the diagnostic agent should be amphiphilic in nature like the phospholipids intended to surround it. Hence for labeling via intercalation, the imaging molecule should contain polar and nonpolar termini. The majority of conventional imaging agents are hydrophilic; thus amphiphilic analogs of these compounds are often prepared by covalently attaching a long hydrocarbon chain or sterol moiety to the parent molecule. Depending on the polarity of the lipophilic conjugate, the analog may partition closer to the hydrophilic surface or the apolar core. This method of labeling is rather attractive as it is relatively simple: usually involving a single incubation step followed by sample purification to separate the free probe from that complexed to the lipoprotein. In addition, very high yields can be achieved with this method. The capacity to ferry intercalated agents correlates with the amount of surface area available on the lipoprotein. Whereas LDL has been shown to carry 200 paramagnetic probes [45] HDL has been cited to only transport 10 [46]. The main caveat to this strategy is that intercalated compounds have a high propensity of non-specific exchange with other membrane surfaces within their immediate vicinity. Such events likely occur in the narrow spacing of distal capillaries, where the lipoproteins are in close apposition with the endothelial lining of the vessel. The mechanism of lipid transfer is a general phenomenon that operates through contact-mediated pathways of intermembrane lipid exchange or mixing that is free of a cellular energy source or complex transport processes. The high free energy stored within the lipoprotein membrane itself (i.e., monolayer bending modulus) likely provides sufficient energy for the thermodynamically favorable exchange of lipids [47]. Lipid exchange may also occur between different lipoprotein species through enzymemediated processes. These reactions are catalyzed by phospholipid transfer protein or cholesterol ester transfer protein [22, 48]. Although direct evidence of enzyme-mediated transfer of exogenous agents has not been documented in the literature, speculative observations do warrant consideration of this process. Regardless of the mechanism of lipid exchange, serum stability tests should be performed to ensure integrity of the surface labeled lipoprotein complex. Core-Loading “Reconstitution” Procedure The previous two strategies for incorporating probes into lipoproteins can be viewed as surface-loading methods as they involve attaching the probes to the protein or phospholipid components of lipoproteins. The third strategy that will be covered is the core-loading strategy. This method of inserting exogenous agents into the interior of lipoproteins, called “reconstitution,” has been most extensively described for LDL particles. This procedure was pioneered by Gustafson and colleague [49] and was later popularized by Krieger in the late 1970s [50, 51]. The method involves the selective extraction of the core lipids using highly nonpolar solvents (e.g., heptane) in the presence of starch. Throughout this procedure the apoprotein is protected and stabilized by the starch; as well, the phospholipid shell remains “intact” in the apolar solvent system. Exogenous compounds solubilized in apolar solvents are then added to the starch–LDL residue, followed by evaporation of the solvent with an inert gas. During

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the course of the solvent evaporation the LDL particle spontaneously reassembles as the exogenous agents are concentrated into the LDL core. This method of labeling lipoproteins is rather attractive as the diagnostic agent is shielded from the external environment until it reaches its designated cellular target, where it is released by LDL catabolism. Another merit of the reconstitution procedure is the carrying capacity of the lipoprotein core. The number of exogenous compounds that can potentially be packaged into the lipoprotein interior can be estimated from the number of endogenous lipid molecules it can normally transport in its core. For example, VLDL can carry an estimated 15,000 molecules of cholesterol esters and triglycerides, LDL 1500 molecules, and HDL2 (10-nm diameter) 109 molecules [52]. In light of these values a study by Li et al. [53] showed that 400 molecules of the fluorophore/photosensitizer derivative tetra-t-butyl silicon phthalocyanine bisoleate could be reconstituted into the core of LDL. The versatility of this method does have some restrictions as reconstituted agents must have sufficient hydrophobicity to be loaded into the LDL core. It should be noted that the reconstitution process is facilitated if the compound contains a long unsaturated hydrocarbon chain (e.g., oleic or linoleic acid) or branched side chains (e.g., polyisoprenes such as phytol) [54]. The sterol ester moiety is not required as modified triglycerides and fatty acids have been loaded into LDL. Furthermore, numerous contrast agent analogs containing either unsaturated fatty acyl chains or sterol ester moieties have also been successfully reconstituted into the LDL core [41, 55, 56]. To date, the core extraction/reconstitution procedure has only been reported for LDL in the literature. The seminal work of Pitman and colleagues demonstrated that synthetic HDL particles can be made from apo A-I and commercial lipids [57]. This process of building HDL from individual constituents has also been cited as a reconstitution method; hence in the literature these synthetic constructs are referred to as “reconstituted HDL particles.” This method of preparing HDL provides the flexibility of introducing exogenous agents into the particle formulation. Similar to LDL, compounds intended for HDL core-loading must be hydrophobic and contain an unsaturated hydrocarbon chain (>10 carbons) or a sterol ester moiety. Several examples of synthetic HDL reconstituted with drug or imaging probes have been described [58–60].

18.2.2 Radionuclide Imaging Investigators in the field of nuclear medicine were the first to exploit the natural cancer targeting capability of lipoproteins for tumor detection and characterization. The small amounts of radiotracers needed for detection and the facile chemistry used for labeling lipoprotein make this modality highly amenable for lipoprotein-based cancer imaging. The two most common radiopharmaceuticals used for labeling lipoproteins have been iodine and technetium. Radioiodination and technetium labeling can be performed by fixing the radionucleotide to the lipoprotein surface. With this method the majority of the radioactive probe will be conjugated to the amino acid residues of the apoprotein component. However, despite this fact, as much as 20% of the radioactivity has been known to become associated with the surface lipids [61, 62]. Alternatively, analogs of lipoprotein core components can be synthesized labeled with a radiopharmaceutical and incorporated into the lipoprotein core through reconstitution or detergent solubilization methods (Fig. 18.4) [55]. Radioiodinated cholesteryl esters and hexaiodinated diglycerides are two such examples [63, 64]. In addition, intercalation of amphiphilic metal chelating complexes, gallium-68

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FIGURE 18.4 Radiolabeled probes for lipoprotein incorporation. (a) Radioiodinated cholesteryl ester. (b) Hexaiodinated diglyceride. (c) Indium-111 DTPA(stearylamide).

or indium-111 DTPA(stearylamide) (see Fig. 18.4), into the phospholipid monolayer can be performed to construct radiolabeled lipoprotein complexes [65]. A great number of studies have been performed using radiolabeled lipoproteins to monitor and characterize lipoprotein metabolism in normal organs, particularly for the liver, adrenal gland, and reproductive tissues [41]. While the number of papers focused on tumor detection has not been as numerous as that of the former, the findings from the tumor studies have been promising. Gal and co-workers were one of the first groups to use radiolabeled lipoproteins to screen a number of different cancer cells of gynecological origin [66]. In these studies they demonstrate that neoplastic cells possess higher binding capacities (15– 30×) and metabolized LDL at an enhanced rate (up to 20×) compared to nonneoplastic cells [66]. In vivo studies by Norata’s group and Hynd’s group went on to support Gal’s initial findings. Norata and co-workers found that murine MS-2 fibrosarcomas took up high levels of iodinated and cyclohexanedione-modified iodinated LDL (via receptormediated and independent pathways) exceeding even that of the host liver (∼2.5-fold)(Fig. 18.5) [67]. Similarly, Hynds and colleagues used scintillation measurements to show that moderate radioactivity was associated with MAC 13 tumors following intravenous (IV) injection of iodinated LDL into tumor bearing mice. The tumor radioactivity was second only to that found in the liver. Furthermore, the investigators showed that they could enhance the tumor targeting selectivity of radiolabeled LDL by pretreating animals with sodium taurocholate and hydrocortisone sodium succinate, which selectively downregulate LDL receptor-mediated uptake in the liver and adrenal glands without affecting the tumor activity [68]. All of the aforementioned studies have relied on ex vivo measurements to demonstrate the assimilation of radiolabeled lipoproteins into tumors. To date, few studies have gone on to assess the utility of these probes for in vivo tumor detection. One such study performed by Ponty and co-workers utilized 99m Tc-LDL to visualize B16-melanoma xenographs in mice [69]. Scintigraphic images demonstrated strong accumulation of the labeled probe in the upper abdomen and in the region of the tumor xenograft (Fig. 18.6). Subsequent

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Relative LDL Uptake (cpm.g-1/plasma cpm)

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FIGURE 18.5 Relative uptake of I-LDL in tissues from BALB/c mice bearing a fibrosarcoma tumor. The radiolabeled LDL was injected 8–10 days after the injection of 106 cells in the hind leg and the mice were killed 24 h later. Data represents means ±SEM for 10 animals. C = Controls; T = tumor bearing mice. (Reproduced with permission from Norata et al. [67].)

biodistribution analysis of excised tissues revealed that the gamma counts in the tumor were 1.4-fold greater than that of host liver. Interestingly, the authors found that the magnitude of LDL uptake by the tumor was associated with the tumor growth rate, where higher levels of LDL uptake occurred during the early period of tumor growth [69]. One caveat associated with this nuclear imaging method is the poor spatial resolution of this modality. Clear demarcation of tumor and host tissue boundaries is often critical during radiographic tumor detection. Alternate modalities may be enlisted to help address this issue.

UA

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FIGURE 18.6 Scintigraphic images of a (a) normal mouse and (b) B16 tumor-bearing mouse taken 18 h after the injection of 99m Tc-LDL[2 2.2 MBq (600 ␮Ci) per mouse]. (Reprinted with permission of Ponty et al. [69].)

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18.2.3 Near-Infrared Fluorescent Optical Imaging Near-infrared (NIR) optical imaging is a sensitive, safe, and inexpensive imaging modality that can be applied for cancer detection [70]. The spectral domain of NIR probes provides a substantial optical window where hemoglobin and water absorption are minimal (650– 900 nm). Differentiation of tumor and normal tissue is based on differences in tissue absorption and fluorescence. Given that tissues are relatively transparent to NIR light, selective delivery of NIR probes to malignant lesions would facilitate the high sensitivity and specificity needed for tumor detection. Packaging these fluorophores into lipoproteins could provide a means whereby NIR probes could be selectively delivered to tumors. Krieger and colleagues were the first to demonstrate that fluorescent probes could be stably formulated into lipoproteins. In these early studies the authors showed that oleoylconjugated fluorophores (3-pyrenemethyl-23, 24-dinor-5-cholen-22-oate-3 beta-yl oleate and dioleyl fluorescein) and a fluorescent derivative of cholesteryl ester (cholesteryl 12-O[methyl 3-O-methyl-5 (6 )-carboxyfluorescein]ricinoleyl carbonate) could be reconstituted into LDL to monitor the LDLR activity in normal and LDLR deficient cells [51, 71]. Several years later, Zheng et al. [72] utilized a similar strategy where they conjugated a tricarbocyanine NIR dye (IRD41; Li-Cor Biotechnology Division, Lincoln, NE) to cholesterol laurate through a thiourea linkage (Fig. 18.7a). NIR carbocyanine dyes are excellent fluorophores due to their long wavelength absorption (750–850 nm), high extinction coefficients (>100,000), and fluorescent quantum yields, making them an ideal probe for cancer detection [73]. These investigators also demonstrated that the visible carbocyanine dye, 1,1 -dioctadecyl-3,3,3 ,3 -tetramethylindocarbocyanine perchlorate (DiI), could also be used to label LDL [74]. Both carbocyanine probes were attached to LDL through surfaceloading methods. They went on to show that these fluorophore-labeled LDL particles could be selectively delivered to various tumors through LDLR-mediated processes [72, 74]. Low-temperature three-dimensional (3D) redox scanning showed that hepatoma, HepG2 xenograft, localized the labeled LDL complexes into the tumor core. Conversely, the distribution of these probes in B16 melanoma xenografts was limited to the tumor periphery due the large central necrosis typically found in these tumors [72, 74]. These findings were also confirmed with a neutral NIR fluorescent bioconjugate, pyropheophorbide cholesterol ester (Pyro-CE) (Fig. 18.7b), which can be loaded within the core of LDL [56]. The tumor distribution of Pyro-CE-LDL appears to be dictated by the perfusion properties of the xenograft, where the B16 melanoma was only well perfused at its periphery while HepG2 tumor was uniformly perfused (Fig. 18.8). Pyropheophorbide is an attractive agent as it can serve both as an NIR fluorescent probe and as a PDT agent [75]. Subsequent studies by this group have shown that several hydrophobic PDT analogs can be loaded in LDL for therapeutic purposes [53, 76]. Complementary work from the lab of Therien showed that NIR fluorescent supermolecular meso-to-meso ethyne-bridged tris[(porphinato)zinc (II)] (PZn3 ) fluorophores can be successfully reconstituted into the core of LDL [77]. The potent fluorescence emission of this agent combined with the minimal cytotoxicity profile of the LDL(PZn3 ) complex make this a promising targeting contrast agent. Indeed, studies from this lab went on to show that LDL(PZn3 ) is actively taken up in B16 melanoma cells and can be imaged at very low doses in the nanomolar range [77]. In vivo fluorescent imaging of tumors using fluorophore-labeled LDL was accomplished with the carbocyanine derivative DiR-BOA (1,1 -dioctadecyl-3,3,3 ,3 tetramethylindotricarbocyanine iodide bis-oleate). Fluorescent LDL was prepared by

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(a)

Me (b) Me

NH N N HN

H Me

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FIGURE 18.7 Structures of (a) tricarbocyanine dye–cholesteryl laurate conjugate and (b) pyropheophorbide cholesterol oleate.

reconstituting DiR-BOA into the lipoprotein core. Figure 18.9 displays serial in vivo optical images of a HepG2 tumor bearing mouse following an IV injection of r-DiRBOA-LDL (molar ratio 90:1). In the immediate postinjection period, low fluorescence can be seen throughout the body. At 1 hour, high-intensity signal can be seen in the abdomen (liver) and by 6 hours high fluorescence can also be seen in the tumor. This pattern persists through to 24 hours postinjection. Similar to the earlier described scintigraphic images, the in vivo fluorescent images show that the majority of the diagnostic signal emanates from the liver and the tumor. 18.2.4 Magnetic Resonance Imaging The high-resolution tomographic capacity that magnetic resonance imaging offers makes it an attractive modality for cancer imaging. The use of amphiphilic chelates has provided a means whereby gadolinium (Gd) contrast agents can be tethered to LDL. The prototype

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r-pyropheophorbide-LDL

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FIGURE 18.8 Redox imaging of the reconstituted pyropheophorbide cholesterol oleate labeled low-density lipoprotein in the HepG2 and B16 tumors.

amphiphilic contrast agent (PTIR267) was the first in this series; it contained both a fluorophore and a Gd(III) moiety along with two palmitate hydrocarbon chains to facilitate intercalation into the LDL phospholipid monolayer (Fig. 18.10) [78]. Approximately 50 molecules of PTIR267 can be attached to each LDL macromolecule. Intravenous administration of this agent into a B16 tumor bearing mouse demonstrated that the labeled LDL accumulated within the liver and tumor (Fig. 18.10) [78]. However, the bulky nature of the PTIR267 complex limits the number of probes that can be loaded onto the LDL surface (upper limit of 50). This becomes a critical issue for MR molecular imaging as high paramagnetic payloads are required to overcome the low inherent sensitivity of MRI. To address this limitation a study by Corbin et al. [45] synthesized a simple and less bulky amphiphilic Gd chelate, DTPA-Bis(stearylamide). The authors cited that it was possible to incorporate much higher payloads (150–500) of the Gd(III) contrast agent into LDL than was previously described. Modest relaxivities were achieved on a per Gd(III) basis, but more compelling relaxation enhancements were identified when viewed with respect to the entire LDL nanoparticle. Coronal MR images of tumor bearing mice display clear enhancement of the tumor xenografts 24 and 36 hours following IV injection of Gd-DTPA-SA-LDL (Fig. 18.11) [45]. These studies represent initial attempts at using LDL as a delivery vehicle for Gd(III). Follow-up investigations by Crich and co-workers further improved upon this system by using novel lipophilic Gd complexes possessing two coordinated water molecule (q = 2)

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FIGURE 18.9 Uptake of core-loaded DiR-LDL in HepG2 tumor-bearing mouse. Intravenous injection consists of 73.9 nmol of DiR and 0.82 nmol of LDL carrier.

coordination sites [79]. The Gd complex, Gd-AAZTAC17, displayed markedly improved relaxivity when inserted into LDL over the initial Gd- DTPA-SA-LDL construct (22 versus 6.5 mM-1 ·s−1 ). In vivo experiments displayed in Figure 18.12 showed that the highest levels of tissue enhancement were achieved in the early postinjection period (first 4 hours) and steadily decreased thereafter over the remaining 48-hour study period.

FIGURE 18.10 (a) Structure of PTIR267, a prototype of dual MRI and fluorescent contrast agent. (b) T1 maps of B16 melanoma in vivo before and 16 h after injection of PTIR267-labeled LDL. The tumors were outlined by dashed contours. The gray level of each pixel represents its T1 value as indicated by the gray scale bar, which presents the T1 value in milliseconds (T1 value greater than 2000 ms is saturated). (Reprinted with permission of Li et al. [74].)

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Pre-contrast

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FIGURE 18.11 T1-weighted coronal spin-echo images of a nude mouse with a subcutaneously implanted HepG2 tumor. Images are from the mouse prior to administration of Gd-DTPA-SA-LDL (precontrast) and at various times following the intravenous administration of MR contrast agent (5, 24, and 36 h). Arrow indicates tumor. (Reprinted with permission from Corbin et al. [45].)

To date, only gadolinium-labeled lipoproteins have been used for tumor targeting. For a number of years there has been considerable interest in incorporating superparamagnetic agents (iron oxide) into lipoproteins for this purpose, but until recently this approach remained elusive. In a recent publication by Cormode et al. [80], oleate-coated iron oxide nanocrystals were successfully formulated into the core of HDL particles. This new technology may now open opportunities for targeted superparamagnetic–lipoprotein-based cancer imaging.

FIGURE 18.12 Structure of Gd-AAZTAC17. Fat-suppressed T1-weighted multislice multiecho MR images of C57BL/6 mice grafted subcutaneously with B16 melanoma cells. Images were obtained (a) before and 8 h after the administration of (b) Gd-AAZTAC17 and (c) Gd-AAZTAC17/LDL. Arrows show the enhanced tumor region (outlined in blank). (d) Course of SI enhancement measured in the ROI corresponding to the whole tumors and muscle from mice injected with Gd-AAZTAC17 (full and empty blue circles, respectively) and Gd-AAZTAC17/LDL (full and empty red squares, respectively). Graph shows the standard error of the means (SEM). (Reprinted with permission from Crich et al. [79].)

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18.2.5 Computed Tomography Similar to MRI, X-ray computed tomography (CT) offers excellent spatial resolution and multiplane imaging capacity. However, the high content of contrast media required to elicit significant Housfield unit enhancement precludes the use of X-ray contrast agents for many nanoparticle/molecular imaging approaches [81]. To date, neither LDL nor HDL particles have been used as a tumor targeting vehicles for X-ray contrast media. While the finite loading capacity of LDL and HDL limits their use for such applications, this may not be the case for the much larger VLDL or chylomicron-like particles. Weichart and colleagues have shown that polyiodinated triglyceride analogs (50% by weight iodine) can be loaded into synthetic chylomicron-like lipid emulsions for liver imaging [82, 83]. These particles proved to be highly hepatoselective, as over 70% of the administered dose of these particles localized in the liver within 30 minutes after injection (Fig. 18.13) [82]. Although chylomicrons do not naturally home to tumors, it may prove interesting to reroute

FIGURE 18.13 (a) Chemical structure of the polyiodinated triglyceride compound, 2oleoylglycerol-1,3-bis[7-(3-amino-2,4,6-triiodophenyl)heptanoate] (DHOG). The aromatic iodine in the compound makes up approximately 50% of the weight of the molecule. (b) Volume-weighted particle size distribution profile for a typical 125 I-DHOG-LE formulation. The mean diameter of the formulation was 219 nm with a standard deviation about the mean of 68 nm (31.2%). (c) Representative CT images after injection of DHOG-LE formulations in normal rats. Images were taken before and 30 min, 2 h, or 24 h after injection of (i) 50 mg I/kg or (ii) 90 mg I/kg. (Reprinted with permission from Bakan et al. [82].)

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this iodinated chylomicron particles to tumors/tumor vessels for targeted CT imaging of cancer (see Section 18.4).

18.2.6 Ultrasound To date, there have been no attempts to use lipoproteins to ferry ultrasound contrast agents for cancer detection.

18.3 HDL FOR TUMOR TARGETING Of all the lipoprotein species, LDL has been most extensively documented as a drug delivery vehicle for targeting cancer cells. More recently, there has been an increasing interest in the association of HDL with cancer. In a large clinical study consisting of 530 cancer patients, serum HDL cholesterol, like LDL cholesterol, was seen to be consistently low compared to noncancer subjects [84]. Within this study group, patients with hematological malignancies displayed the lowest HDL cholesterol while those with breast tumors showed the highest [84]. In addition to supplying cholesterol for active membrane turnover, there is also evidence to suggest that HDL may have a mitogenic effect on various cell types [85]. Early work by Favre et al. [86] showed that HDL or HDL components were able to stimulate proliferation of adenocarcinoma cell lines. Since then, others have also concurred with these findings and have reported on the stimulating effects of HDL on various other cancer cell lines. Furthermore, studies by Cao and co-workers showed that by disrupting the integrity of the HDL receptor (mutation of residues at the carboxy terminus) the proliferation of breast cancer cells can be markedly inhibited [87]. Interestingly, HDL stimulation of cellular proliferation is not the only mechanism for augmenting cell growth, as other studies have suggested that HDL can induce antiapoptotic effects in various cell types [88]. Finally, high levels of HDL receptor have also been reported among cancer cells over surrounding normal cells [89], thus providing malignant cells a selective advantage for acquiring this potent growth stimulant. These collective findings have led several investigators to explore the utility of HDL as a tumor targeting vehicle for drug delivery. Unlike LDL, HDL-like particles (reconstituted HDL) can be prepared from apoproteins (apo A-I with or without apo A-II) and commercial lipids [57]. This affords the opportunity to easily introduce foreign constituents to the HDL particles. Lacko and colleagues have formulated the microtubule toxin paclitaxel into HDL and have shown enhanced efficacy of this complex against breast cancer cells compared to the conventional paclitaxel formulation (Fig. 18.14) [90]. Similar results were also observed by Lou et al. [91], who incorporated aclacinomycin into HDL particles to treat hepatoma cells. To date, there have not been any reports in the literature of labeling HDL particles with diagnostic probes for cancer detection. However, the laboratory of Fayad has published a number of papers utilizing paramagnetic and fluorescent labeled HDL particles for MRI and optical detection of atherosclerotic plaques (Fig. 18.15) [46, 92]. More recently, this group has demonstrated that gold, iron oxide, or quantum dot nanocrystals may also be incorporated into HDL particles for CT, MRI, and fluorescent imaging, respectively (Fig. 18.16) [80]. The convergence of the HDL delivery system with nanocrystal-based probes will undoubtedly open new opportunities for imaging studies in biomedical and cancer research [93].

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FIGURE 18.14 Cytotoxicity of the reconstituted high-density lipoprotein (rHDL)/paclitaxel (PTX) nanoparticles versus free PTX estimated by the MTT assay. (Reproduced with permission from McConathy et al. [90].)

18.4 REROUTING LIPOPROTEINS TO CANCER SPECIFIC TARGETS Although contrast agent-bearing lipoprotein nanoparticles are able to accumulate in cancer cells, this platform suffers from limited tumor selectivity over normal tissues. The liver, adrenal glands, and reproductive organs all express high levels of LDLR [94, 95] and thus would effectively compete and be exposed to high levels of the administered imaging agent. Even though the lipoprotein receptor expression levels in these tissues may be modulated by the concomitant use of steroids, bile acids, or dietary fats [96, 97], the downregulation of lipoprotein receptor is not absolute and the extent of this protective effect is uncertain. Moreover, the range for potential lipoprotein-based applications in cancer medicine is

apo A-1 NDB-DPPE

Phospholipid Cholesterol Ester Triglyceride

Gd-DTPA-DMPE

Pre

Unesterified Cholesterol

24 h

FIGURE 18.15 Gd (III) labeled HDL-like nanoparticle contrast agent displays selective enhancement for atherosclerotic plaques 24 h following intravenous injection of the agent. Schematic of the HDL-like MRI contrast agent is also shown. (Reprinted with permission of Frias et al. [46].)

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FIGURE 18.16 Nanocrystal core high-density lipoprotein. (a) Schematic depiction of the different agents in this study. (b) Negative stain TEM images of FeO-HDL. (c) Photograph of the QD-HDL in (left) normal light and (right) under UV illumination. (d) Phantoms of Au-HDL imaged using CT. (Reprinted with permission of Cormode et al. [80].)

limited by the narrow purview of lipoprotein receptor-positive tumors. Currently, there are a number of “cancer signatures” that bind to surface receptor/epitopes that are much more tumor specific than LDLR/HDLR; these include, for example, Her2/neu [98], EGF [99], somatostatin [100], folate (FA [101], and ␣v ␤3 integrins [102]. Incorporating such targeting ligands into lipoprotein carriers would greatly improve both the targeting specificity and range of cancers reached by this system. In fact, the laboratory of Zheng and co-workers has recently reported that lipoprotein particles can be redirected from their native receptors to alternate surface receptors or epitopes of choice [43]. The redirecting strategy is based on the chemistry of the exposed active lysine residues in the apoprotein component. These residues play a central role in recognition and binding of lipoprotein to lipoprotein receptor; thus by conjugating homing ligands to the ε-amino groups of these lysine residues one effectively abolishes the affinity of the lipoprotein particle for its lipoprotein receptor and simultaneously redirects the particles to alternate receptors. Proof of the redirection strategy was first demonstrated with fluorescent DiI labeled and photosensitizer reconstituted LDL, using folic acid (FA) modification of apo B-100 to redirect the LDL particle to the folate receptor (FR) [43]. Apo B-100 contains 357 Lys residues; of these, 225 are exposed and are available for ligand conjugation [37]. Fifty-three of the exposed Lys ε-amino groups are considered “active” with a low pK a value of 8.9, while the remaining 172 Lys are “normal” with a pK a

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FIGURE 18.17 (a) Schematic of DiR-LDL-FA. (b) Confocal images of (a) KB (FR+) alone, (b) KB + DiR-LDL-FA, (c) KB+DiR-LDL-FA + 200-fold FA, (d) HT1080 (FR-) alone, and (e) HT1080 + DiR-LDL-FA. (c) Fluroescence images of tissues and tumors excised from tumor-bearing mouse 24 h following the intravenous injection of DiR-LDL-FA (5.8 ␮M). (Reprinted with permission of Chen et al. [103].)

value of 10.5 [37]. Thus by adjusting the pH of the conjugation reaction, the majority of the active lysines can be conjugated with FA. FA-conjugated LDL was shown to be selectively taken up in FR-positive cancer cells (KB cells) over FR-deficient cells (HT1080 or CHO) or cells expressing high levels of LDLR (HepG2) (Fig. 18.17) [43]. Similar results were also generated in vivo in nude mice bearing dual tumor xenografts of KB and HT1080. Although these findings demonstrated that FR targeting can be achieved with the LDL nanoparticle, competitive uptake processes (RES, opsonization, and possible absorption of exchangeable apoproteins) by the liver and spleen compromise the tumor uptake of FA-LDL (Fig. 18.17) [103]. The rerouting technology was next investigated with HDL particles due to their small size and intriguing pharmacokinetic properties. HDL was first reconstituted with the lipophilic NIR optical dye (DiR-BOA); this served as an optical tracer for the activity and distribution of the nanoparticle. In a similar manner as with LDL, folic acid moieties were conjugated to the lysine residues of HDL’s apoprotein component (apo A-1) [60]. The Lys residues in apo A-I display various pK a values ranging from 8.3 to 10.5. Thus careful titration of the reaction pH is critical to saturate all the Lys residues in apo A-I. Following successful preparation of this FA-conjugated fluorescent HDL, rerouting from SR-BI to the FR was demonstrated in in vitro and in vivo experiments. The in vivo data showed particularly impressive results as exceptional targeting activity was displayed by the FA-conjugated HDL nanoparticle (Fig. 18.18). High levels of nanoparticle accumulation were detected within the target FR-expressing tumor while minimal uptake was evident in the FR-negative tumor or in tissues of the RES [60]. Ongoing experiments are being conducted to explore other ligands such as EGF and PSMA for studies targeting HDL to lung and prostate cancer, respectively.

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FIGURE 18.18 (a) In vivo fluorescent images of dual tumor (KB/HT-1080) bearing mouse before and at several times following intravenous injection of (DiR-BOA)rHDL-FA (15 ␮M). (b) In vivo inhibition study was performed with 30-fold excess free folic acid.

18.5 FUTURE DIRECTION FOR LIPOPROTEIN-BASED CARRIERS One of the major drawbacks of utilizing lipoproteins as a delivery vehicle for exogenous agents is the need to isolate particles or their respective apoproteins from fresh donor plasma. Relying on donor plasma to acquire lipoprotein is problematic because lipoprotein samples may vary from batch to batch, methods for isolating the various lipoprotein species can be lengthy (up to 72 hours), large quantities of lipoproteins can be difficult to attain, and safety issues surrounding the use of blood products remains a matter of concern. Furthermore, isolated lipoproteins can only be stored for finite periods before aggregation and degradative processes compromise the integrity of the sample. As a result of these limitations, attempts have been made to prepare synthetic lipoprotein-like particles. In the case of LDL, Nikanjam et al. [104] have demonstrated that synthetic LDL-like particles can be prepared from a lipid emulsion and a synthetic bifunctional 29 amino acid peptide, which contains a lipid binding motif and the LDLR binding domain of apo B-100. These LDL-like particles were shown to mimic the behavior of native LDL and target the LDLR on carcinoma cells. The authors also went on to show that a lipid derivative prodrug of

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Nonpolar face (a) Ala 17

Leu 3

Val 10

Leu 14

Tyr 7

Lys 13

Phe 18 Ala 11 Lys 4

Trp 2 Lys + 9

Lys 15 Asp 8 –

Phe 6

+

–

Glu 16

Ala 5

Glu 12

Asp 1

+ +

–

–

(b)

Polar face 18A 40

A. 1

30 20 10 80

120

160

200

240

280

320

FIGURE 18.19 (a) Axial projection of 18A peptide showing the relative location of the amino acid residues with segregation of the hydrophobic and hydrophilic amino acids. (b) Size distribution and transmission electron microscopy of discoidal complexes. (Reprinted with permission from Anantharamaiah et al. [106].)

paclitaxel, paclitaxel oleate, could be loaded in these particles and effectively delivered and activated in glioblastoma multiforme cancer cells via the LDLR-mediated pathway [105]. For a number of years the pharmaceutical industry has invested a considerable amount of effort into investigating apo A-I mimetic peptides as a potential cardiovascular therapeutic. This investigation led to the synthesis of an 18 amino acid peptide known as 18A [106]. The sequence of 18A (D-W-L-K-A-F-Y-D-K-V-A-E-K-L-K-E-A-F) does not have any sequence homology to apo A-I. However, the peptide displays characteristic features of apo A-I, a Class A amphipathic ␣-helix protein. The peptide contains opposing polar and nonpolar faces orientated along the long axis of the helix with positively charged lysine residues residing at the polar–nonpolar interface and clusters of negatively charged residues at the apex of the polar face (Fig. 18.19a) [107]. This unique amphipathic ␣ helix orientation imparts strong lipid binding properties to the peptide. Early work by Anantharamaiah’s group showed that these peptides could not only facilitate reverse cholesterol transport but when mixed with phospholipids they also possessed capacity to spontaneously form discoidal complexes similar to nascent HDL particles (Fig. 18.19b) [106, 107]. A few groups have recently explored the utility of these peptide stabilized disks as a delivery platform for drugs or imaging agents. Tufteland and co-workers showed that peptide–phospholipid disks could be formulated with the antifungal amphotericin B [108]. These complexes

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displayed potent antifungal activity. Thereafter, Cormode and colleagues used a dimer of 18A to make a paramagnetic labeled peptide–phospholipid disk to image atherosclerotic plaques [109]. Investigations utilizing fully synthetic lipoprotein-like particles for drug and contrast agent will undoubtedly continue, and over the next few years it is anticipated that many more studies in this area will be published. As progress continues in this field, issues of peptide immunocompatibility and particle/payload stability of these synthetic lipoproteinlike particles should be addressed. Although several issues remain to be resolved, this technology shows great promise and will certainly advance the drug delivery field. Providing synthetic lipoprotein-like particles that are free of plasma-derived components will help overcome the many obstacles that have hampered the progress of lipoprotein-based delivery systems in both basic and clinical research. 18.6 CONCLUSION Lipoproteins are the main transport system for important hydrophobic molecules such as cholesterol and fatty acids. The compartmentalized organization of these carriers, which allows for the native transport of these molecules, also makes them amenable to facile incorporation of exogenous compounds. From the perspective of cancer molecular imaging, loading of contrast agents into modified lipoproteins confers several advantages, such as payload delivery of imaging probe, improved pharmacokinetics, and the ability to selectively target these probes to cancer cells. In summary, this chapter provides an introduction to the biology of lipoproteins: from the metabolism of lipoproteins to its relationship with malignant tumors. In addition, it has outlined existing techniques for incorporating imaging agents into lipoproteins, as well as introducing novel methods of rerouting lipoprotein to cancer-specific eptitopes. While lipoproteins remain an intriguing candidate for delivery of diagnostic agents to cancer cells, the future clinical application of lipoproteins remains unclear. This is due in large part to issues involving the utilization of human blood products as the starting material. However, recent improvements in protein synthesis and recombinant technology have provided an alternative means for starting materials. Development of recombinant apoproteins and synthetic apoprotein mimetics provides a plausible strategy for future application of this technology and may herald a new age for cancer-targeted diagnostic imaging. ACKNOWLEDGMENT Work from the lab of G.Z. was supported by the Canadian Institutes of Health Research (82847), the Ontario Institute for Cancer Research through funding provided by the Government of Ontario, the Canadian Cancer Society (018510), the National Institute of Health (N01 CO37119 and R21/R33 CA114463), and the Joey and Toby Tanenbaum/Brazilian Ball Chair in Prostate Cancer Research. REFERENCES 1. Mahley, R. W.; Innerarity, T. L.; Rall, S. C., Jr.; Weisgraber, K. H. Plasma lipoproteins: apolipoprotein structure and function. J. Lipid Res. 1984, 25, 1277–1294.

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60. Corbin, I.; et al. Enhanced cancer-targeted delivery using engineered high-density lipoproteinbased nanocarriers. J. Biomed. Nanotechnol. 2007, 3, 367–376. 61. Eisenberg, S.; Windmueller, H. G.; Levy, R. I. Metabolic fate of rat and human lipoprotein apoproteins in the rat. J. Lipid Res. 1973, 14, 446–458. 62. Sigurdsson, G.; Noel, S. P.; Havel, R. J. Quantification of the hepatic contribution to the catabolism of high density lipoproteins in rats. J. Lipid Res. 1979, 20, 316–324. 63. DeForge, L. E.; Ruyan, M. K.; Schwendner, S. W.; Newton, R. S.; Counsell, R. E. Synthesis and evaluation of radioiodinated cholesteryl ethers as lipoprotein probes. Bioconjug. Chem. 1991, 2, 254–260. 64. Xiao, W.; Wang, L.; Ryan, J. M.; Pater, A.; Liu, H. Incorporation of an (125)I-labeled hexaiodinated diglyceride analog into low-density lipoprotein and high specific uptake by cells of cervical carcinoma cell lines. Radiat. Res. 1999, 152, 250–256. 65. Moerlein, S. M.; Daugherty, A.; Sobel, B. E.; Welch, M. J. Metabolic imaging with gallium-68and indium-111-labeled low-density lipoprotein. J. Nucl. Med. 1991, 32, 300–307. 66. Gal, D.; MacDonald, P. C.; Porter, J. C.; Simpson, E. R. Cholesterol metabolism in cancer cells in monolayer culture. III. Low-density lipoprotein metabolism. Int. J. Cancer 1981, 28, 315–319. 67. Norata, G.; et al. In vivo assimilation of low density lipoproteins by a fibrosarcoma tumour line in mice. Cancer Lett. 1984, 25, 203–208. 68. Hynds, S. A.; et al. Low-density lipoprotein metabolism in mice with soft tissue tumours. Biochim. Biophys. Acta 1984, 795, 589–595. 69. Ponty, E.; et al. Biodistribution study of 99m Tc-labeled LDL in B16-melanoma-bearing mice. Visualization of a preferential uptake by the tumor. Int. J. Cancer 1993, 54, 411–417. 70. Sevick-Muraca, E. M.; Lopez, G.; Reynolds, J. S.; Troy, T. L.; Hutchinson, C. L. Fluorescence and absorption contrast mechanisms for biomedical optical imaging using frequency-domain techniques. Photochem. Photobiol. 1997, 66, 55–64. 71. Krieger, M.; Ho, Y. K.; Falck, J. R. Reconstitution of LDL with lipophilic fluorescein derivatives: quantitative analysis of the receptor activity of human lymphocytes. J. Receptor Res. 1983, 3, 361–375. 72. Zheng, G.; et al. Tricarbocyanine cholesteryl laurates labeled LDL: new near infrared fluorescent probes (NIRFs) for monitoring tumors and gene therapy of familial hypercholesterolemia. Bioorg. Med. Chem. Lett. 2002, 12, 1485–1488. 73. Mujumdar, R. B.; Ernst, L. A.; Mujumdar, S. R.; Waggoner, A. S. Cyanine dye labeling reagents containing isothiocyanate groups. Cytometry 1989, 10, 11–19. 74. Li, H.; et al. Carbocyanine labeled LDL for optical imaging of tumors. Acad. Radiol. 2004, 11, 669–677. 75. Blessington, D.; et al. Detection and imaging of the reconstituted pyropheophorbide-cholesterol oleate labeled low-density lipoprotein in the hepG2 tumor. Proc. SPIE 2003, 4955, 497–504. 76. Song, L.; et al. Naphthalocyanine-reconstituted LDL nanoparticles for in vivo cancer imaging and treatment. Int. J. Nanomed. 2007, 2, 767–774. 77. Wu, S. P.; et al. Near-infrared optical imaging of B16 melanoma cells via. lowdensity lipoprotein-mediated uptake and delivery of high emission dipole strength tris[(porphinato)zinc(II)] fluorophores. Bioconjug. Chem. 2005, 16, 542–550. 78. Li, H.; et al. MR and fluorescent imaging of low-density lipoprotein receptors. Acad. Radiol. 2004, 11, 1251–1259. 79. Crich, S. G.; et al. Magnetic resonance imaging detection of tumor cells by targeting low-density lipoprotein receptors with Gd-loaded low-density lipoprotein particles. Neoplasia (N.Y.) 2007, 9, 1046–1056.

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80. Cormode, D. P.; et al. Nanocrystal core high-density lipoproteins: a multimodality contrast agent platform. Nano Lett. 2008, 8, 3715–3723. 81. Speck, U. Contrast agents: X-ray contrast agents and molecular imaging—a contradiction? In Handbook of Experimental Pharmacology, Springer: New York, 2008, pp. 167–175. 82. Bakan, D. A.; Weichert, J. P.; Longino, M. A.; Counsell, R. E. Polyiodinated triglyceride lipid emulsions for use as hepatoselective contrast agents in CT: effects of physicochemical properties on biodistribution and imaging profiles. Invest. Radiol. 2000, 35, 158–169. 83. Weichert, J. P.; et al. Polyiodinated triglyceride analogs as potential computed tomography imaging agents for the liver. J. Med. Chem. 1995, 38, 636–646. 84. Fiorenza, A. M.; Branchi, A.; Sommariva, D. Serum lipoprotein profile in patients with cancer. A comparison with non-cancer subjects. Int. J. Clin. Lab. Res. 2000, 30, 141–145. 85. Pussinen, P. J.; et al. The human breast carcinoma cell line HBL-100 acquires exogenous cholesterol from high-density lipoprotein via CLA-1 (CD-36 and LIMPII analogous 1)-mediated selective cholesteryl ester uptake. Biochem. J. 2000, 349, 559–566. 86. Favre, G.; et al. High density lipoprotein3 binding sites are related to DNA biosynthesis in the adenocarcinoma cell line A549. J. Lipid Res. 1993, 34, 1093–1106. 87. Cao, W. M.; et al. A mutant high-density lipoprotein receptor inhibits proliferation of human breast cancer cells. Cancer Res. 2004, 64, 1515–1521. 88. Suc, I.; Escargueil-Blanc, I.; Troly, M.; Salvayre, R.; Negre-Salvayre, A. HDL and Apo A prevent cell death of endothelial cells induced by oxidized LDL. Arterioscl. Thromb. Vascular Biol. 1997, 17, 2158–2166. 89. Lacko, A. G.; Nair, M.; Paranjape, S.; Johnso, S.; McConathy, W. J. High density lipoprotein complexes as delivery vehicles for anticancer drugs. Anticancer Res. 2002, 22, 2045–2049. 90. McConathy, W. J.; Nair, M. P.; Paranjape, S.; Mooberry, L.; Lacko, A. G. Evaluation of synthetic/reconstituted high-density lipoproteins as delivery vehicles for paclitaxel. Anti-cancer Drugs 2008, 19, 183–188. 91. Lou, B.; et al. High-density lipoprotein as a potential carrier for delivery of a lipophilic antitumoral drug into hepatoma cells. World J. Gastroenterol. 2005, 11, 954–959. 92. Frias, J. C.; Ma, Y.; Williams, K. J.; Fayad, Z. A.; Fisher, E. A. Properties of a versatile nanoparticle platform contrast agent to image and characterize atherosclerotic plaques by magnetic resonance imaging. Nano Lett. 2006, 6, 2220–2224. 93. Corbin, I.; Scott, B.; Zheng, G. Convergence of nanocrystal imaging and lipoprotein delivery. Nanomedicine 2009, 4, 7–10. 94. Carew, T. E.; Pittman, R. C.; Steinberg, D. Tissue sites of degradation of native and reductively methylated [14 C]sucrose-labeled low density lipoprotein in rats. Contribution of receptordependent and receptor-independent pathways. J. Biol. Chem. 1982, 257, 8001–8008. 95. Pittman, R. C.; Steinberg, D. Sites and mechanisms of uptake and degradation of high density and low density lipoproteins. J. Lipid Res. 1984, 25, 1577–1585. 96. de Smidt, P. C.; van Berkel, T. J. LDL-mediated drug targeting. Crit. Rev. Ther. Drug Carrier Syst. 1990, 7, 99–120. 97. Firestone, R. A. Low-density lipoprotein as a vehicle for targeting antitumor compounds to cancer cells. Bioconjug. Chem. 1994, 5, 105–113. 98. Slamon, D. J.; et al. Studies of the HER-2/neu proto-oncogene in human breast and ovarian cancer. Science (N.Y.) 1989, 244, 707–712. 99. Wong, A. J.; et al. Structural alterations of the epidermal growth factor receptor gene in human gliomas. Proc. Natl. Acad. Sci. U.S.A. 1992, 89, 2965–2969. 100. Becker, A.; et al. Receptor-targeted optical imaging of tumors with near-infrared fluorescent ligands. Nat. Biotechnol. 2001, 19, 327–331.

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101. Leamon, C. P.; Low, P. S. Delivery of macromolecules into living cells: a method that exploits folate receptor endocytosis. Proc. Natl. Acad. Sci. U.S.A. 1991, 88, 5572–5576. 102. Ruoslahti, E. Specialization of tumour vasculature. Nat. Rev. Cancer 2002, 2, 83–90. 103. Chen, J.; et al. Ligand conjugated low-density lipoprotein nanoparticles for enhanced optical cancer imaging in vivo. J. Am. Chem. Soc. 2007, 129, 5798–5799. 104. Nikanjam, M.; et al. Synthetic nano-low density lipoprotein as targeted drug delivery vehicle for glioblastoma multiforme. Int. J. Pharm. 2007, 328, 86–94. 105. Nikanjam, M.; Gibbs, A. R.; Hunt, C. A.; Budinger, T. F.; Forte, T. M. Synthetic nano-LDL with paclitaxel oleate as a targeted drug delivery vehicle for glioblastoma multiforme. J. Control. Release 2007, 124, 163–171. 106. Anantharamaiah, G. M.; et al. Studies of synthetic peptide analogs of the amphipathic helix. Structure of complexes with dimyristoyl phosphatidylcholine. J. Biol. Chem. 1985, 260, 10248–10255. 107. Anantharamaiah, G. M. Synthetic peptide analogs of apolipoproteins. Methods Enzymol. 1986, 128, 627–647. 108. Tufteland, M.; Pesavento, J. B.; Bermingham, R. L.; Hoeprich, P. D., Jr.; Ryan, R. O. Peptide stabilized amphotericin B nanodisks. Peptides 2007, 28, 741–746. 109. Cormode, D. P.; et al. An Apo A-I mimetic peptide high-density-lipoprotein-based MRI contrast agent for atherosclerotic plaque composition detection. Small (Weinheim, Germany) 2008, 4, 1437–1444.

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CHAPTER 19

Protein Cages as Multimode Imaging Agents MASAKI UCHIDA, LARS LIEPOLD, MARK YOUNG, and TREVOR DOUGLAS Department of Chemistry and Biochemistry and Department of Plant Sciences, Center for Bio-Inspired Nanomaterials, Montana State University, Bozeman, Montana, USA

19.1 INTRODUCTION Protein cage architectures such as those of small virus capsids and ferritins are selfassembled from a limited and defined number of subunit building blocks [1, 2]. These protein cage architectures typically range in size from tens to hundreds of nanometers and are spherical or rod shaped. By combining both chemical and genetic modifications of the subunits, one can impart novel functions to these protein cage architectures that are quite different from their native function in biology. There are three distinct surfaces of the assembled protein cage architecture that can be manipulated to impart function by design [1, 2]: these are the external surface, the interior surface, and the surface that forms the interface between the subunits. Manipulation of the interactions at the subunit interface allows control over disassembly and reassembly [3, 4]. The interior interface can be used for chemical attachment of small molecules and for encapsulation of nanoparticles, sequestered inside the cage, useful for imaging and therapeutics. The exterior surface can also be modified by small molecule attachment or by presentation of cell-specific targeting ligands (peptides, antibodies, carbohydrates) [5–10]. Considerable effort has gone into the development of protein cage architectures as robust systems that combine advanced imaging capacity with highly selective cell targeting within the same protein cage platform [2, 11]. A library of protein cages, which support a range of synthetic chemistries and genetic manipulations useful for creating high performance multimode imaging agents, have been employed (Fig. 19.1) [12, 13]. These include the 28-nm capsid of Cowpea chlorotic mottle virus (CCMV) [14], the 30-nm capsid of Cowpea mosaic virus (CPMV) [15], the bacteriophages MS2 (27 nm) [16] and Q␤ (30 nm) [17], the 12-nm small heat shock protein from Methanococcus jannaschii [18], and the 12-nm human ferritin cage [19]. Many closed shell protein cage architectures define an interior space that can be used to selectively direct synthesis or entrapment of a guest cargo. A significant advantage to the use of protein cages is the ability to produce monodisperse nanoparticles with a precisely Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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FIGURE 19.1 Protein cages including the viral capsids CCMV (28-nm diameter), CPMV (30-nm), MS2 (27-nm), and Q␤ (30-nm), the heat shock protein Hsp (12-nm), and a ferritin Fn (12-nm). The images of CCMV, CPMV, MS2, and Q␤ were reproduced from VIPER (http://viperdb.scripps.edu/) at the Scripps Research Institute. Those of Hsp and Fn were reproduced using the UCSF Chimera package (http://www.cgl.ucsf.edu/chimera/) from the Resource for Biocomputing, Visualization, and Informatics at the University of California (supported by NIH P41 RR-01081).

defined size and shape and since the properties of nanophase materials are intimately related to their dimensions, any heterogeneity in size is reflected as heterogeneity in their physical properties. The controlled formation of nanomaterials within protein cage architectures thus has broad applications in a variety of emerging technologies including drug delivery and medical diagnostics. Using approaches that mimic the structure–function relationships in virus particles, that is, selective encapsulation, homogeneous size, and targeted cell delivery, creates the potential for protein cage nanoparticles for imaging and therapeutic applications [20]. Key to the success of this approach is an ability to produce and to modify protein cages to impart novel functionality. The highly symmetrical, spatially controlled presentation of functional groups has allowed the attachment of organic species, such as targeting peptides [6, 8], fluorescent labels [11, 21, 22], metal chelating groups [23, 24], and inorganic species such as magnetic nanomaterials [25, 26] and gadolinium ions [27–31], specifically to either the interior or exterior of the protein cages [32]. These agents represent a wide range of potential imaging modalities and, when coupled with targeting capability on the outside of the cages, provide a powerful platform for medial imaging.

19.2 PROTEIN CAGE BASED MAGNETIC MATERIALS AS MRI CONTRAST AGENTS The high resolution of magnetic resonance imaging (MRI) affords excellent anatomical detail; however, the ability to detect low copy number tissues is limited by the technique’s low sensitivity. Targeted contrast agents increase the sensitivity of MR to tissues of interest

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by accelerating the NMR relaxation rate of water protons in the vicinity of these tissues. Two relaxation processes, either T1 relaxation (spin–lattice or longitudinal) or T2* relaxation (spin–spin or transverse), are probed to produce MR images. The bulk of the image is produced from the nuclear spins of protons that are in water or fatty molecules. Contrast agents are materials that catalyze this realignment of nuclear spins and in doing so they increase the relaxation rate and therefore increase the contrast of and image. This effect is called paramagnetic relaxation enhancement (PRE) and a common class of contrast agents contain gadolinium (Gd) ions or Fe3 O4 nanoparticles as catalysts for this process.

19.3 PROTEIN CAGES AS T1 AGENTS Gd agents typically are used to enhance the T1 relaxation process and “T1” specifically refers to a time in units of seconds equal to a realignment decay constant for the proton’s spin in a given sample. The T1 of water, for example, is approximately 3 seconds. The PRE of a contrast agent is given in units of mM-1 ·seconds-1 (mM-1 ·s-1 ) and is referred to as relaxivity. The Gd concentration (mM) is used for the calculation of ionic relaxivity while the concentration of a particle (mM) containing one or more Gd ions is used to calculate the particle relaxivity. For example, a 1 mM aqueous solution of Magnevist, a clinically used contrast agent, has a relaxivity of r1 ≈ 5 mM-1 ·s-1 with a T1 of 0.19 second based on Eq. (19.1) [33].
T1observed =

1 T1sample

+ r1 CA ConcentrationCA

−1 (19.1)

There are a few parameters that must be optimized with regard to development of Gdbased contrast agents. The first consideration is that Gd is toxic and must therefore be used in the form of a tightly associated metal ion chelate. Since the Gd-based relaxivity is a spin–lattice effect, direct water interaction with the Gd ion is necessary through an inner sphere mechanism. The number of metal-bound water molecules that can occupy a coordination sphere of Gd is referred to as “q” and, in general, as q increases the stability of the chelator–metal interaction decreases. Therefore a balance must be struck between increasing the number of water coordination sites and the overall stability of the metal–chelator interaction. Another parameter that must be optimized is the residence time of the Gd-bound water, referred to as ␶ M . It is important for the water to remain bound to Gd long enough for the proton spins to relax. However, when ␶ M is too high, and the water molecule remains bound to Gd long after it has been relaxed, the result is an effective decrease in the “duty cycle” of the Gd ion resulting in a decrease of the overall relaxivity. The ␶ M value can be modulated by varying the chemical nature of the metal chelator. The final parameter that is important for efficient PRE is the rotational correlation time of the Gd ion (␶ R ). Larger Gd complexes result in greater ␶ R values, which yields more efficient PRE. Figure 19.2 shows the parameters that effect PRE. Protein cages are well suited as vehicles for Gd contrast agents since their large size provides a large exterior and interior surface area as well as an interior volume, all of which can be functionalized to carry high densities of Gd. Also, the large size of these particles

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q

τM τR

FIGURE 19.2 A schematic of DTPA-Gd and the important physical parameters that affect the relaxivity, including the number of water molecules that directly interact with the Gd ion (q), the exchange lifetime of the metal-bound water molecule (␶ M ), and the rotational correlation time of the Gd chelate (␶ R ).

results in ideal rotational properties (␶ R ) and therefore theoretical maximum relaxivity values can be realized in these systems. To date, several groups have reported the utilization of viral capsids or other protein cage structures as MRI contrast agents. In the first report of this kind, Allen and co-workers utilized the endogenous metal binding site in the plant virus CCMV and bound Gd3+ ions to endogenous Ca-binding sites in the capsid [27]. Since the metal ion in this capsid is situated at the interface of three subunits, it is likely that the ␶ R for Gd is similar to ␶ R of the entire cage, therefore this Gd–protein cage construct should exhibit near ideal ␶ R values. This construct has a relatively weak affinity for the Gd ion (K d = 31 ␮M), which along with analysis of the crystal structure of the metal binding pocket both suggest that q is likely to be greater than one. Recently performed fitting of the CCMV construct’s r1 NMRD profile to the Solomon–Bloembergen–Morgan (SBM) analytical model for relaxivity returned values of q = 4, ␶ M = 3 × 10-9 second, and ␶ R > 10-7 second, all of which are highly desirable. Not surprisingly, the optimal relaxivity properties of this construct resulted in extremely high ionic relaxivities (r1 = 202 mM-1 ·s-1 @ 1.5 T) [27]. Unfortunately, the low Gd affinity of this construct makes the clinical applicability impractical. Several groups have developed a range of next generation constructs, which can now be divided into three categories: (1) endogenous metal binding sites, [27], (2) genetic insertion of a metal binding peptide [30], and (3) chemical attachment of small molecule chelates [28–32]. Of these three categories the first two are less significant from a clinical standpoint due to their low Gd binding affinities. The third category is the most common approach with three groups reporting similar results. Figure 19.3 outlines the single step coupling of chelated Gd ions to the CCMV protein cage. Anderson et al. [28] attached DTPA-Gd to lysine residues on the surface of the MS2 capsid (MS2 DTPA ITC). Despite the relatively low ionic relaxivity of this construct, the high density of Gd labeling resulted in the highest particle relaxivity for constructs with sufficient Gd binding stability and is approximately three orders of magnitude higher than the small molecule contrast agent. Finn and co-workers produced a similar construct by using click chemistry to attach DOTA-Gd to CPMV and Q-beta and this scheme also produced a construct with a relatively low ionic relaxivity (r1 = 16 mM-1 ·s-1 @1.5 T); however, high levels of Gd ions were attached [31]. A CCMV-DOTAGd construct and a MS2 Bis (HOPO) TAM construct both have higher ionic relaxivities

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O N O

NH2

O O

O

N

N

N

N

+

N

O-

O

-O

O

O

O

O

N

N

N

N

N

O N

O

-O

+ GdCI3

O

O-

-O

O

O-

O-

O

N

Gd3+ N

N O

O-

FIGURE 19.3 The reaction scheme to attach DOTA-Gd to the CCMV viral capsid. Endogenous lysines on the viral capsid are reacted to a DOTA/NHS conjugation. Next, GdCl3 is added to produce a viral capsid conjugated with Gd3+ ions.

(r1 = 46 mM-1 ·s-1 @1.5 T; r1 = 31 mM-1 ·s-1 @1.5 T, respectively) but lower density of Gd labeling resulted in lower particle relaxivities [29]. To optimize the relaxivity properties of Gd-based contrast agents, it is helpful to have quantitative information of the parameters important to PRE. Since ␶ R , ␶ M , and q have different dependencies on the external magnetic field, it is possible to extract information about these values from a plot of relaxivities obtained for one contrast agent taken over a range of field strengths. These plots are referred to as nuclear magnetic resonance dispersion (NMRD) plots and they can be fit to the SBM analytical model for PRE and values for ␶ R , ␶ M , and q can be obtained [34]. Francis and co-workers labeled the MS2 viral capsids on the interior surface and on the exterior surface with a novel Gd chelator (Fig. 19.4). A fit of their NMRD data with a modified SBM model suggests that the ␶ M for this construct is near optimal and q = 2 [32]. However, local mobility in the chelator attachment scheme partially negated the ideal rotational properties of this large construct. Furthermore, these authors reported a relaxivity difference between the interior and exterior labeled versions of the MS2 construct that was attributed to a difference in the rigidity of the these two anchoring schemes. The ␶ R values

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FIGURE 19.4 The MS2 viral capsid labeled externally (4) and internally (5) with a Gd chelate. The crystal structure of the MS2 virus capsid is shown, highlighting the amino groups on the exterior: red, K106; yellow, K113; and orange, the N terminus. Y85 is highlighted in green on the interior surface. (Copyright © 2008 American Chemical Society.)

were higher for the interior labeled construct, which has higher relaxivity values, suggesting that local mobility can dampen the ideal rotational effects of protein cages. Recently, a strategy has been developed for synthesizing a branched structure, via azide alkyne cycloaddition reactions, that projects into the interior cavity of the heat shock protein (Hsp) cage from Methanococcus jannaschii (Fig. 19.5) [35]. With this synthetic strategy a high Gd payload was incorporated in the interior cavity of a protein cage by anchoring a Gd chelator to this interiorly grown branched polymer (Fig. 19.5). In doing so, very high Gd densities in the protein cage and reasonably high ionic relaxivity values were achieved. Therefore the resulting particle has exceptional particle relaxivity properties. The most promising aspect of coupling chelated Gd to the polymer is the development of a method to take advantage of the interior volume of the protein cage. To date, only the

FIGURE 19.5 Shown here is the synthesis of the branch polymer in the interior of the Hsp protein cage via the alkyne–azide click coupling reaction. Amines on the polymer are functionalized with DTPA-Gd after polymer growth.

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surface of protein cages including viral capsids have been modified to carry Gd chelators by single step chemical labeling. Here we expand the chemical modification to fill the interior volume with a DTPA-Gd labeled polymer network. This results in a particle that has a per particle relaxivity of 4200 mM-1 ·s-1 , which, while lower than the MS2-DTPA-Gd value (7200 mM-1 ·s-1 ), exhibits relaxivity properties that are clinically significant. The MS2 construct is significantly larger than the Hsp, occupying 12 times the volume. When the relaxivity behaviors of these materials are compared on a volume basis, the Hsp construct exhibits a relaxivity that is seven times that of the MS2 construct. Until now the relaxivity properties of wtCCMV-Gd seemed unattainable in systems with high-affinity chelators. This has been achieved by filling the interior cavity of the Hsp with DTPA-Gd and the per volume relaxivity or the “relaxivity density” values are equal to that of the wtCCMV-Gd construct. The advantages of a particle with high relaxivity densities become apparent when one considers the relaxivity properties required to image a cell. A previously reported equation returns the amount of contrast agent required to image a mammalian cell as a function of the contrast agent’s particle relaxivity [36]. We used this equation along with values for the surface area of the contrast agent and the surface area of a mammalian cell to calculate the percentage of the cell that would be covered with the contrast agent for successful imaging, assuming that the contrast agent would bind irreversibly and would not be taken into the cell. These calculations resulted in 11% of the mammalian cell surface covered with Hsp-BP-(DTPA-Gd) for successful imaging, while the MS2 construct would cover 29% of the surface and Magnevist would cover 72%. While the details of cellular imaging are likely to be more complex than this simple calculation assumes, the importance of a contrast agent with high relaxivity densities is apparent. There are significant benefits—namely, increased stability and efficient particle relaxivity properties—that result from a synthetic approach that takes full advantage of interior volume of protein cages. This strategy can be extended to larger protein cages and it will be interesting to see if the ionic relaxivity remains constant in larger systems. Also, developing similar protein cage–polymer hybrid constructs with Gd chelators that have more desirable q and ␶ M values should result in particles with very attractive relaxivity properties.

19.4 PROTEIN CAGES AS T2 MRI CONTRAST AGENTS The T2 relaxation is a time constant in units of seconds that describes the length of time for the precession of water proton to lose coherence and become randomized after an excitation pulse. This is dependent on the local inhomogeneities in the magnetic field, which can be controlled by the presence of magnetic nanoparticles. By using a T2-weighted pulse sequence, the MR scanner is biased to detect the T2 relaxation process, an extremely useful clinical option for some diagnostic applications. Thus T2 contrast agents are used to enhance contrast in diagnostic imaging and have been manipulated to yield maximal efficiency of T2 relaxivity. Among the variety of protein cage architectures, ferritin has been the most intensively investigated for its application as a T2 MRI contrast agent. Ferritin’s inherent biological function is the sequestration and storage of iron as an amorphous iron oxide nanoparticle of ferrihydrite. Ferritins are assembled from 24 subunits and exhibit cage-like structure 12 nm in diameter with an interior cavity of 8 nm where ferrihydrite is encapsulated [19]. This weakly superparamagnetic ferrihydrite [37] functions as an endogenous contrast agent that

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significantly shortens the T2 of surrounding water protons. Since iron is well known to play an important role in a number of diseases, endogenous ferritin in humans has been considered as a useful marker for MRI examination to assess the amount of iron in liver, spleen, and brain [38–40]. In addition, ferritin has recently been proposed as a reporter protein to monitor transgene expression using MRI [41, 42]. These studies suggest ferritin could be used like green fluorescence protein and luciferase for in vivo optical imaging of gene expression. However, here we will focus on development of synthetic ferritin–iron oxide nanocomposites for MRI contrast agents rather than exploitation of endogenous ferritin. One disadvantage of endogenous ferritin as an MRI contrast agent is the fact that it exhibits 10–100-fold lower relaxivity per iron than commercially available synthetic iron oxide MRI contrast agents. In order to overcome this drawback, demineralized ferritin (apoferritin) was developed as a size-constrained reaction environment for the synthesis of superparamagnetic iron oxide nanoparticles (magnetoferritin) which exhibit much higher R2 compared to endogenous ferritin [43, 44]. Recently, recombinant ferritin cages have been developed, which have demonstrated advantages for the encapsulation of magnetite or maghemite (Fig. 19.6) [45, 46]. Native mammalian ferritins are composed of a mixture of two different types of subunit, light (L) chain and heavy (H) chain [37]. A catalytic ferroxidase site, which catalyzes the oxidation of Fe(II) to Fe(III) exists in the H chain, but is absent in the L chain. Oxidation of Fe(II) to Fe(III) might be a critical step in the formation of iron oxide because of the low solubility of Fe(III) [19], even in synthetic systems. Using recombinant human H chain ferritin cages, it has been shown that Fe3 O4 cores formed close to the theoretical core diameter, calculated from input Fe loading with narrow size distribution. The average size of the mineralized Fe3 O4 particles in HFn increases from 3.6 to 5.9 nm with increasing Fe loading from 1000 Fe to 5000 Fe per cage. This means that both the Fe3 O4 core of the mineralized HFn and the exterior diameter are comparable to ultrasmall

FIGURE 19.6 (a) Illustration of synthesis of magnetite (Fe3 O4 ) nanoparticle in HFn. (b) TEM image of mineralized particles within HFn with 5000 Fe/cage (right) and size distribution of the particles analyzed from the TEM image (left).

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TABLE 19.1 Comparison of Size and r1 and r2 Relaxivities of the Mineralized HFn, SPIO, and USPIO HFn1000Fe

HFn3000Fe

HFn5000Fe

Ferumoxides

Ferumoxtran-10

Exterior (nm) Interior particle (nm)

12 3.6 ± 0.7

12 5.1 ± 0.9

12 5.8 ± 0.9

58.5 ± 185.8 15.2 ± 8.9

29.5 ± 23.1 4.9 ± 1.5

r1 (mM·s)−1 r2 (mM·s)−1

2.3 11

4.2 31

8.4 93

7.8 119

12 101

superparamagnetic iron oxide (USPIO) contrast agents such as Ferumoxtran-10 rather than small superparamagnetic iron oxide (SPIO) such as Ferumoxides [47] (Table 19.1). As expected, the r1 and r2 relaxivities of the mineralized HFn increase with increasing Fe per cage (i.e., the particle size) and the r1 and r2 of the HFn loaded with 5000 Fe were comparable to those of commercially available SPIO and USPIO materials [46] (see Table 19.1).

19.5 PROTEIN CAGES AS CONTRAST AGENTS FOR IMAGING ATHEROSCLEROTIC PLAQUES One of the unique features of the mineralized HFn with regard to MRI contrast is that it is readily taken up by macrophage cells and therefore provides a strong T2* effect [46]. Since macrophage cells play a significant role in the progression of inflammatory responses, imaging of macrophages involved in inflammatory diseases such as atherosclerosis is useful to assess and diagnose these diseases [48, 49]. SPIO particles are taken up by macrophages more readily than the mineralized HFn in vitro, likely due to its larger particle size. However, in vivo the SPIO particles are trapped by Kupffer cells in liver tissue during the first pass after injection in vivo [50] and therefore are less useful for imaging macrophages outside the liver. In contrast, USPIO particles have a longer blood circulation time and hence are able to label atherosclerotic plaques in vivo [51, 52], but their uptake efficiency into macrophages is low. Recently, it has been shown that mineralized HFn particles are taken up by macrophages much more efficiently than USPIO in vitro, even though the overall size of the mineralized HFn is similar to that of the USPIO (Fig. 19.7a). These findings hold the promise that mineralized ferritin can be developed as an MRI contrast agent to monitor inflammatory events induced by macrophages such as atherosclerotic plaque progression (Fig. 19.7b). Indeed, fluorescence imaging studies have revealed that fluorescently labeled HFn accumulates significantly more in atherosclerotic plaque lesions formed in the carotid arteries of mice than in normal carotid arteries and colocalizes with macrophage cells in these lesions [53] (Fig. 19.8). Furthermore, in vivo MRI assessment of the atherosclerotic mouse model confirmed reduction of carotid artery lumen caliber size on a T2* enhanced image taken 24 h and 48 h after injection of the mineralized HFn only when the atherosclerotic plaque was induced within the artery [53]. This reduction is likely due to a T2* enhanced effect caused by the mineralized HFn accumulated in the plaque lesion. These results indicate that the mineralized HFn is a promising MRI contrast agent to monitor inflammation in atherosclerosis.

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FIGURE 19.7 (a) Fe content taken up by macrophage cells after incubation with 165 ␮g/mL of various contrast agents. (b) In vitro MR images of macrophage cells incubated with 165 ␮g/mL of various contrast agents for 24 h. (Copyright © 2008 Wiley-Liss, Inc.)

Another unique feature of protein cage architectures is that cell- and/or tissue-specific targeting moieties can be functionalized on their exterior surface using either chemical or genetic approaches. For example, the amino acid sequence RGD-4C (CDCRGDCFC), known to bind integrins ␣v ␤3 and ␣v ␤5 [54], has been genetically introduced as an N-terminal fusion protein into HFn [45]. The HFn mutant (RGD4C-Fn) exhibited a cagelike structure indistinguishable from the wild-type HFn by either size or morphology. In addition, there was no significant difference in the average diameter and size distribution of the mineralized Fe3 O4 formed within the RGD4C-Fn and HFn under the same Fe loading conditions. These results indicate that the introduction of the RGD-4C targeting

FIGURE 19.8 In situ and ex vivo fluorescence images of HFn-Cy5.5 nanoparticles in mouse carotid arteries. The fluorescently labeled HFn is accumulated in atherosclerotic plaques formed in a ligated carotid artery.

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FIGURE 19.9 FACS analysis of THP-1 cells incubated with fluorescence-labeled protein cages. The data are plotted as histograms with their corresponding geometric (geo) mean fluorescence values. The increased level of fluorescence intensity of the cells incubated with RGD4C-Fn indicates specific binding of the cages to THP-1 cells. (Copyright © 2009 Wiley-VCH Verlag GmbH & Co.)

peptide has little effect on either the protein cage assembly or the mineralization behavior of this protein cage mutant. Analysis using fluorescence-activated cell sorting demonstrated that the RGD4C-Fn clearly exhibited enhanced targeting interaction with C32 amelanotic melanoma cells and THP-1 monocyte cells, which are known to overexpress integrins ␣v ␤3 [55] (Fig. 19.9). Furthermore, TEM observation of macrophage cells incubated with the Fe3 O4 mineralized HFn or RDG4C-Fn revealed that the RDG4C-Fn particles both bind to the cell surface and were also internalized into the cells more efficiently than the nontargeted HFn controls [55]. These results suggest that introduction of cell and/or tissue targeting moieties onto the exterior of a protein cage architecture shows promise for enhancing and directing the MR imaging capability.

19.6 PROTEIN CAGES AS VEHICLES FOR TARGETED DELIVERY OF IMAGING AGENTS TO BIOFILM INFECTIONS Another promising application of protein cage-based imaging agents is the diagnosis and imaging of bacterial infection in biofilms. Chronic infections associated with device implantation, also known as biomaterial centered infections, are rare but the consequences are dire for the patient. It is now well accepted that biofilms, which are structured communities of microorganisms that develop on tissue or material surfaces, are the etiological agents of biomaterial centered infections. In terms of our current understanding, the most cost-effective approach for treatment of biofilm infections would be early detection followed by early treatment with an appropriate antimicrobial. In principle, there are currently a variety of noninvasive imaging techniques that could be used to diagnose biomaterial centered infections. Among various types of imaging techniques, MRI has some advantages for imaging biofilms. The quality of MR images of soft tissues is superior to those produced by any other technique. In addition, it does not require exposure to potentially harmful radiation or administration of radioisotopes. However, MRI is a relatively insensitive technique compared to radionuclide imaging. In addition, the expected size of biofilms

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(a)

(b)

(c)

(d)

FIGURE 19.10 Targeting of CCMV to a pathogenic bacterium: (a) targeting scheme—SpA, protein A; Ab, monoclonal antibody against protein A; B, biotin; StAv, streptavidin; IA, imaging agent; (b,c) SEM micrographs showing the high density of targeting; and (d) SEM micrograph of cell exposed to CCMV without Ab targeting.

colonizing implants is small (less than 1 mm in dimension). Therefore developing the capability to image biofilms in vivo by MRI is nontrivial and it will be necessary to develop new methods for targeted delivery of imaging agents to achieve early detection of biofilms by MRI. Protein cage architectures are an ideal platform to accomplish loading of designed targeting moieties, imaging agents, and possibly therapeutic agents simultaneously. Cowpea chlorotic mottle virus (CCMV) has been used as a delivery vehicle and Staphylococcus aureus (S. aureus) as a model organism for some initial studies. Staphylococcus aureus is a primary culprit in prosthetic joint and graft infections [56, 57] and also tissue infections such as endocarditis, osteomyelitis, and native heart valve infections that conform to the biofilm model. In general, bacterial pathogens express cell surface proteins that are distinct from those of the host tissues and can be used for targeting. For example, S. aureus expresses at least 22 different cell wall anchored proteins, many of which are involved in conferring virulence and are likely to be expressed at a high level in infections [58, 59]. Among these is protein A (SpA), which binds to the Fc region of IgG, and thereby may serve to conceal the cells from immune detection. We designed the CCMV capsid to target protein A displayed on S. aureus using streptavidin (StAv) to mediate the interaction between biotinylated monoclonal antibody (mAb) against protein A and CCMV (Fig. 19.10). This provided a clear experimental advantage since mAb against protein A is commercially available. By using the indirect targeting scheme outlined in Figure 19.10a, a relatively high density of CCMV binding to the cell surface of S. aureus could be achieved (Fig. 19.10b,c)

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[60]. Negative controls (e.g., Fig. 19.10d) indicated binding was specifically mediated by the mAb–SpA interaction. The specificity of the binding was corroborated by the results of epifluorescence microscopy, flow cytometry, and TEM [60]. An additional issue to be considered is the penetration of the protein cage into the biofilm since biofilm cells are enclosed in a matrix of extracellular polymeric substances (EPSs) that can impede transport of nanoplatforms. In biofilms of S. aureus treated with CCMV that was dual labeled with biotin and a fluorescent tag, it was shown that the functionalized CCMV penetrated into the biofilm microcolonies to a depth of about 20 ␮m during a 90-min exposure period [60]. Thus cells in relatively thin biofilms (<10 ␮m) such as those expected to be involved in biomaterial centered infections [61–63] can be fully penetrated during an exposure period that is feasible to achieve clinically. More recently, we have successfully prepared a “Janus-like” protein cage nanoparticle using the DNA-binding protein from Listeria innocua (LiDps) [64]. The cage is dual labeled with biotin and a fluorescent tag similar to CCMV, but possesses two spatially distinct faces, one labeled only with biotin and the other with a fluorescent tag. These Janus-like LiDps cages can be used as a “plug and play” nanoplatform for antibody directed cell targeting with the potential for control over the orientation of the targeted nanoplatform [65]. In addition, the protein cages, such as CCMV, can be conjugated with DOTA, a clinically relevant Gd chelate, to load CCMV with the Gd-based MRI contrast agent [60]. Using this approach, the mean number of Gd ions bound to this biotinylated construct was estimated at 166 per cage. Because of the selected targeting of the CCMV cage to S. aureus discussed above, roughly 1.8 × 105 Gd atoms per cell were delivered. A typical S. aureus biofilm is quite dense (approximately 3 × 1010 cells/cm3 ) with a distance between cells of less than 3 ␮m [66]. At this cell density the level of Gd loading onto S. aureus cells would result in about 10 mM Gd chelate. It has been estimated that about 50 mM Gd chelate is required for discrimination of tissues targeted with MRI contrast agent [67]. This estimate is based on a Gd chelate with an ionic relaxivity of about 5 mM-1 ·s-1 . The concentration of Gd chelate required for discrimination of tissues targeted with contrast agent is inversely proportional to R1 . As discussed above, a large effort to increase relaxivity of targeted contrast agents is ongoing and we, and others, have demonstrated that ionic relaxivities can be increased by more than an order of magnitude by coupling Gd to a protein cage [27]. In addition, previous levels of labeling with fluorescence imaging agents suggests that Gd loading onto CCMV can be increased substantially [68], effectively producing large particle relaxivities and increasing the localized Gd concentration above 50 mM. Even more promising in this respect are protein cages in which the Gd is incorporated into a crosslinked polymer that fills the protein cage interior (Fig. 19.5). The possibility of high density targeting combined with the high particle relaxivities of Gd coupled to protein cages suggests that protein cages have great potential as vehicles for targeted delivery of Gd to pathogenic microbes for MR imaging of biofilm infections.

19.7 CONCLUSION In this chapter, we outlined how protein cages can be used as platforms for MRI contrast agents and fluorescent imaging agents. Protein cages have a number of advantages as templates for development of nanoplatforms for targeted delivery of both T1 and T2 contrast agents, fluorescent agents, and potentially other imaging modalities as well. One salient

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feature in this respect is their unique versatility, which allows many different design strategies for incorporation of both cell-specific targeting ligands and imaging agents (or therapeutic agents) to be implemented. In addition, monodisperse preparations of nanoparticles can be produced with relative ease. We have introduced examples of targeted delivery of protein cages to both atherosclerotic plaques and microbial biofilms. Coupling of T1 contrast agents to nanoplatforms can increase their relaxivity substantially and protein cages can be loaded with contrast agents both on the exterior and interior surfaces and in the interior cavity. T2 contrast agents such as iron oxides can also be loaded into the interior cavity with control over both the size and phase. The synthetic manipulation of the cages to allow incorporation of metal chelating agents, and other small organics, opens up the promise for using these multivalent platforms with other imaging modalities such as PET and CT. The development of protein cages as a biomaterial has progressed only for a decade. There are a wide variety of proteins that assemble into cage-like structures. We anticipate that the chemical and biophysical diversity of the family of protein cages will lead to an expanding range of biomedical applications.

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65. Suci, P. A.; Kang, S.; Young, M.; Douglas, T. A streptavidin–protein cage janus particle for polarized targeting and modular functionalization. J. Am. Chem. Soc. 2009, 131(26), 9164. 66. Leid, J. G.; Shirtliff, M. E.; Costerton, J. W.; Stoodley, P. Human leukocytes adhere to, penetrate, and respond to Staphylococcus aureus biofilms. Infect. Immun. 2002, 70(11), 6339–6345. 67. Ahrens, E. T.; Rothbacher, U.; Jacobs, R. E.; Fraser, S. E. A model for MRI contrast enhancement using T-1 agents. Proc. Nat. Acad. Sci. U.S.A 1998, 95(15), 8443–8448. 68. Gillitzer, E.; Willits, D.; Young, M.; Douglas, T. Chemical modification of a viral cage for multivalent presentation. Chem. Commun. 2002, 20, 2390–2391.

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CHAPTER 20

Biomedical Applications of Single-Walled Carbon Nanotubes WEIBO CAI Departments of Radiology and Medical Physics, School of Medicine and Public Health, University of Wisconsin–Madison, and University of Wisconsin Carbone Cancer Center, Madison, Wisconsin, USA

TING GAO Tyco Electronics Corporation, Menlo Park, California, USA

HAO HONG Departments of Radiology and Medical Physics, School of Medicine and Public Health, University of Wisconsin–Madison, Madison, Wisconsin, USA

20.1 INTRODUCTION Nanotechnology, an interdisciplinary research field involving chemistry, engineering, biology, material science, and medicine, has enormous potential for early detection, accurate diagnosis, and personalized treatment of diseases. With the size of many orders of magnitude smaller than human cells, nanoparticles can offer unprecedented interactions with biomolecules both on the surface of and inside the cells, which may revolutionize disease diagnosis and treatment. Over the last decade, there have been numerous nanotechnology centers established worldwide and it is expected that nanotechnology will mature into a clinically useful field in the near future [1, 2]. To date, the most well-studied nanoparticles include quantum dots [3, 4], paramagnetic nanoparticles [5], nanowires [6], carbon nanotubes (CNTs) [7], gold nanoparticles [8], nanoshells [9], and many others [10, 11]. No other nanoparticles have been investigated for such a wide variety of biomedical applications like the single-walled carbon nanotubes (SWNTs). Over the last decade, SWNTs have been explored for sensing, drug/gene/vaccine delivery, thermal therapy, molecular imaging, tissue engineering, biomedical instrumentation, and other applications. With the capacity to provide enormous sensitivity, throughput, and flexibility, SWNTs have the potential to profoundly impact disease diagnosis and patient management in the near future. In this chapter, we summarize the current state-of-the-art of SWNTs for biomedical Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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applications. Big strides have been made and many proof-of-principle studies have been successfully performed. The future of SWNTs looks brighter than ever yet many hurdles remain to be conquered. Discovery of the multiwalled carbon nanotube (MWNT) and synthesis of SWNTs in the early 1990s opened up a new arena for research on the nanoscale [12–14]. A search in PubMed for “carbon nanotube” returned more than 5000 publications. CNTs are generally produced by three major techniques: electric arc discharge, laser ablation, and thermal or plasma-enhanced chemical vapor deposition (CVD) [15]. A number of purification techniques have been developed for CNTs such as physical separation and gas/liquid phase oxidation in combination with chemical treatment [16]. Structural characterization of CNTs generally includes scanning electron microscopy (SEM), transmission electron microscopy (TEM), atomic force microscopy (AFM), scanning tunneling microscopy (STM), and many other techniques [17]. Among these, TEM and SEM are the most popular choices for imaging CNTs [18–20]. Intriguingly, the confined space inside SWNTs has been utilized for high-resolution TEM imaging of the dynamic behavior of individual small molecules [21, 22]. These findings opened up the possibility of investigating the biological activities of these molecules on an individual basis, which may potentially be applied to a wide range of systems. A number of methods have been explored for the chemical functionalization of CNTs, a prerequisite for their potential biomedical applications [23]. In addition, the unique properties of CNTs, including but not limited to electrical conductance, high mechanical stiffness, light weight, transistor behavior, piezoresistance, thermal conductivity, luminescence, and electrochemical bond expansion, have made them superb materials for a broad spectrum of applications ranging from energy storage (e.g., H2 , Li) to sensing to imaging and therapy.

20.2 SWNT-BASED SENSORS Besides the superb physical properties such as high strength and excellent electrical/thermal conductivity, SWNT also has a large surface area (estimated to be as high as 1600 m2 /g) which provides numerous sites at which chemicals can react [24]. Furthermore, SWNTs can be chemically modified or permeated with other materials to modify their physical, chemical, and electronic properties, which make them very versatile materials for many applications such as sensing. Compared to MWNTs, SWNTs are more durable, have less structural defects, are much easier to purify, and therefore are more suitable for sensing applications. Since the diameters of SWNTs are similar to or smaller than those of proteins, they can serve as high-performance electrical conduits for interfacing with biological systems, which is also highly desirable for sensing applications [25]. Several review articles focusing on CNT-based chemical and biochemical sensors have been published previously [26–28]. Chemical and biochemical sensors based on SWNTs vary by their transduction mechanisms, which can be electrical, electrochemical, optical, or mass resonant in nature [29, 30]. SWNT-based pressure and flow sensors, which can be used in a wide variety of scenarios in the clinic, are also briefly discussed. 20.2.1 Field-Effect Transistor (FET) Sensors SWNTs can be either metallic or semiconducting depending on the tube diameter and the chirality (i.e., the direction in which the graphite sheet is rolled to form a tube) [31, 32].

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FIGURE 20.1 SWNT-based FET sensor: (a) a bottom gate FET, (b) a liquid gate FET, and (c) typical SWNT FET transfer characteristics in air (solid line), using the bottom gate, and in water (dotted line), using the liquid gate. S, source; D, drain. (Adapted from Bradley et al. [47].)

While metallic SWNTs can be used as coulomb islands in single-electron transistors [33, 34], semiconducting SWNTs can be used to build FETs [35, 36]. The electrical properties of SWNTs can be significantly affected by the interaction between the SWNTs and foreign molecular species, which makes them outstanding candidates for chemical nanosensors [37, 38]. To minimize the nonspecific binding (NSB) of SWNTs to chemicals or biomolecules, SWNTs can be functionalized to be biocompatible for the recognition of specific chemical species [26, 39–41]. SWNT FETs are typically fabricated on an SiO2 /Si chip, composed of multiple or single SWNT which connects two closely spaced metal electrodes that serve as the source and drain for transistors [42]. The conductance of the semiconductor between the source and drain is switched on and off by a third gate electrode that is capacitively coupled through a thin dielectric layer. The gate voltage is applied either on the Si chip (bottom gate) or on the liquid through a Pt electrode in which the entire device is immersed in (liquid gate) (Fig. 20.1). The bottom gate and liquid gate configurations are similar in direct current characterization yet they differ in scale because the capacitance between the SWNT and the two gates is different, with one being the oxide insulating layer (bottom gate) and the other being a hydration layer (liquid gate) (Fig. 20.1C). In an SWNT FET, the sensing signal could be monitored as the conductance values determined by the gate configurations (the current–voltage characteristics at various gate voltages) or as the current changes at a fixed gate voltage. The observed changes in conductance at a given gate voltage can be attributed to two phenomena, the Schottky barrier

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modulation at the SWNT/metal contact or the chemical gating at the SWNT channel [43, 44]. Several studies claimed that the majority of the source–drain current and conductance changes originated from the former [45, 46] while a few other reports suggested the latter [47, 48]. Certain small molecule (e.g., NO2 ) adsorption on SWNT can lead to p-type hole doping, which can increase positive hole carriers, thereby increasing the conductance of the SWNT [37]. On the other hand, n-type electron doping of the SWNT, which is caused by molecules such as NH3 , can result in a decrease in conductance. Many groups have investigated the use of SWNT FETs for sensing a wide variety of biolmolecules, such as streptavidin [39, 49], human immunoglobulin G (IgG) [40], carcinoembryonic antigen [50], hemagglutinin [51], Chromogranin A [52, 53], single-stranded DNA (ssDNA) [41], and ethanol vapor [54]. In one report, the enzymatic degradation of starch was monitored electronically using an SWNT-based FET sensor [55]. Incubation of the device in aqueous buffer solutions of amyloglucosidase resulted in the removal of starch from both the silicon surfaces and the side walls of the SWNTs in the FETs, as evidenced by both direct imaging and electronic measurements. In many of these studies, the removal of NSB to SWNTs was found to be critical and could dramatically affect the experimental results. One report suggested that reversible adsorption sites can be created on the SWNT via noncovalent functionalization with amine-terminated molecules and the analyte adsorption is largely affected by the basicity of the surface groups on the SWNT [56]. 20.2.2 Resistance Chemical Sensors Due to the high resistance of SWNT at the gate voltage of 0 V, direct detection of the change in SWNT resistance, when exposed to chemicals and biomolecules, is rarely used. This method has mostly been employed for gas detection with surface functionalized SWNTs. For example, palladium functionalized SWNTs showed good sensitivity and selectivity to H2 at room temperature [57], while palladium-doped SWNTs on an interdigitated electrode resistance sensor exhibited unique adsorption and electronic properties to methane [58]. Polyaniline functionalized SWNTs were demonstrated to have excellent sensitivity to NH3 , more than 60 times higher than that of nonfunctionalized SWNT sensors, as well as good reproducibility upon repeated exposure to high concentrations of NH3 [59]. In one report, network films of SWNT bundles on polyethylene terephthalate were used to detect nerve agent simulants such as diisopropyl methylphosphonate [60]. 20.2.3 Electrochemical Sensors Electrochemical sensors can facilitate direct electron exchange between the electrodes and a biomolecule [61, 62]. SWNTs can be incorporated into the electrode in many ways: random deposition; vertically “standing” on the electrode surface with one end of the SWNT exposed to the electrolyte solution and the other end in contact with the underlying electrode (termed “SWNT forest”); or integration into the electrode material [26, 27, 63–65]. Among these, SWNT forest-assembled electrodes have attracted the most attention, especially after they can be made to be stable for weeks by either electrostatic interactions [66, 67] or carbodiimide-assisted covalent coupling of amine-functionalized substrates with carboxylterminated SWNTs [68–70]. Cyclic amperometry/voltammetry and/or electrical impedance measurements are typically used for the monitoring of these SWNT-based electrochemical sensors.

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SWNT modified electrodes have been used directly for sensing applications. For example, a carbon fiber electrode with SWNT deposited on the surface was used to detect nitric oxide release from single isolated human umbilical vein endothelial cells [71]. SWNT modification of carbon fiber microelectrodes was reported to increase the sensitivity and reduce fouling in the detection of dopamine and serotonin in vivo [72, 73]. SWNT modification of glass carbon electrode also proved useful for continuous and sensitive monitoring of the ascorbate level and its depletion in the rat striatum when induced by global ischemia [74]. In another study, SWNT forest self-assembled on graphite electrodes was demonstrated to be highly sensitive to mediated immunosensing [75]. In one interesting report, a freestanding SWNT film electrode has been studied as a nonenzymatic glucose sensor [76]. Lastly, a number of different polymer–SWNT composites have also been investigated for electrode modification [77–82]. The sensitivity and dynamic range of these sensors vary to a great extent depending on a number of experimental parameters; therefore they cannot be directly compared with each other. In many studies, the SWNTs that were used to modify the electrodes were either functionalized or immobilized with enzymes such as bilirubin oxidase [83], alcohol dehydrogenase [83], glucose dehydrogenase [83], glucose oxidase [84–87], and horseradish peroxidase [88]. These enzyme-containing electrodes exhibited good catalytic activities and some were suggested to also have a promising future in areas beyond sensing, such as biofuel cell applications. ssDNA is another popular choice for SWNT modification and a few examples have been given in the previous discussion regarding SWNT FET sensors [41]. In electrochemical sensing, electrodes coated with ssDNA-modified SWNTs have also been investigated [89, 90]. In one report, hydrogen peroxide reduction was detected with amplified electrochemical response using DNA immobilized on a ferrocene-filled SWNT electrode [91]. Several other studies have also focused on hydrogen peroxide detection. SWNT-protein dopant modified electrode [92] and a ferrocene-filled SWNT-based sensor [93] both exhibited enhanced sensor performance to hydrogen peroxide with good stability and reproducibility. In one intriguing report, the aligned reconstitution of glucose oxidase on the edge of SWNTs linked to an electrode surface was demonstrated for the first time [94]. The SWNT acted as a connector that electrically connected the active site of the enzyme and the electrode (Fig. 20.2). The electrons were transported along distances of greater than 150 nm and the rate of electron transport could be controlled by the length of the SWNTs [94]. These findings clearly demonstrated the compatibility of SWNTs with novel biomaterial hybrid systems that may have fascinating properties.

20.2.4 Capacitance Chemical Sensors Unlike chemiresistor sensors, which typically suffer from slow recovery after chemical exposure [95], because they detect charge transfer from the analytes, SWNT-based chemicapacitors detect the polarization of surface adsorbates. Therefore they are much more sensitive, recover faster, and are completely reversible. In the presence of a dilute chemical vapor, molecular adsorbates are polarized by the fringing electric fields radiating from the surface of an SWNT electrode, which causes an increase in its capacitance. By coating thin chemoselective materials on the SWNTs, the sensor can provide a large gain to the capacitance response, hence making it capable of detecting both volatile organic compounds (VOCs) and low vapor pressure explosives [96].

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I / μA

40 30 20 10 0 –10 –0.2 –0.1 0 0.1 0.2 0.3 0.4 0.5 E / V (vs SCE)

FIGURE 20.2 Long-range electrical contact of enzymes by SWNT connectors. (a) Assembly of the electrode. FAD: flavin adenine dinucleotide, cofactor of the enzyme. (b) An AFM image of the glucose oxidase reconstituted on the FAD-functionalized SWNT monolayer associated with the Au surface. (c) Cyclic voltammograms corresponding to the catalyzed oxidation of different concentrations of glucose by the electrode. (Adapted from Patolsky et al. [94].)

The electronic response of SWNTs to trace levels of chemical vapors was explored [97]. It was found that adsorption at the defect sites produced a large electronic response that dominated the SWNT capacitance and conductance sensitivity. Furthermore, it was demonstrated that controlled introduction of oxidation defects can be used to enhance the sensitivity of an SWNT network sensor to a variety of chemical vapors. These initial results indicated that optimization of the chemoselective monolayers (to better cover the SWNTs), in combination with improved sensor design, may result in a new class of sorption-based sensors with high sensitivity and fast response time. To date, this sensing mechanism has not been widely used for biomolecule sensing yet.

20.2.5 Resonant Chemical Sensors Quartz crystal microbalance (QCM), which can measure a mass change by detecting the change in frequency of a quartz crystal resonator, has been employed for sensing applications [98]. In one study, SWNTs functionalized with a rheumatoid arthritis specific peptide were deposited on a quartz crystal [99]. Specific antibody binding was then detected in the serum from rheumatoid arthritis patients with sensitivity (in the femtomole range) significantly higher than that of established enzyme-linked immunosorbent assay (ELISA) and microarray systems. SWNT-based sensors for streptavidin and IgG detection were also reported to have sensitive changes in the electronic conductance and mass uptake [100]. By depositing SWNTs onto SiO2 film coated, ST-X quartz substrate as surface acoustic wave (SAW) sensors, VOCs could be detected [101, 102]. In subsequent studies, SWNTbased QCM, SAW, and silica optical fiber sensors were compared in the detection of

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VOCs [103–105]. It was found that the frequency changes in SWNT-based SAW sensors, due to the conductivity and permittivity of the composite changes arising from chemical adsorption, could be tracked with a loop antenna, which makes them more advantageous for remote monitoring. 20.2.6 Pressure Sensors and Flow Sensors The electronic properties of SWNTs, strongly depending on their chirality and diameter, are affected by distortion (e.g., bending and twisting), which makes them excellent candidates for piezoelectric pressure sensors [28]. SWNT-based pressure sensors can be used in many diagnostic and therapeutic procedures/devices, such as in eye surgery, hospital beds, respiratory devices, patient monitors, inhalers, and kidney dialysis instruments. SWNT films on suspended polysilicon membrane were found to exhibit a change in resistance under uniform air pressure, which can be restored when the pressure is released [106]. Furthermore, randomly oriented SWNT bundle films showed a linear response in the measured change in voltage when subjected to tensile and compressive stress tests. Using first principle quantum transport calculations, molecular dynamics simulation, and continuum mechanics analysis, it was demonstrated that hydrostatic pressure can induce radial deformation, which can lead to electrical transition of SWNTs [107]. In this study, a pressure induced metal-to-semiconductor transition in armchair SWNTs was observed, which provided a basis for the design of novel nanoscale tunable pressure sensors. The fabrication and characterization of pressure sensors based on individual SWNTs as the active electromechanical transducer elements was reported [108]. Electromechanical measurements on strained metallic SWNTs adhering to a membrane revealed a piezoresistive gauge factor of approximately 210 for metallic SWNTs. SWNT bundles, which can produce an electrical signal in response to fluid flow when packed between two metal electrodes, have also been investigated [109]. 20.2.7 Sensors Based on Optical Techniques Several reports have focused on optical techniques for SWNT-based sensing. DNA–SWNT hybrids in aqueous solution, in the presence of glucose oxidase, were optically sensitive to hydrogen peroxide and glucose [110]. The DNA–SWNT hybrid surface could be regenerated by dialytically removing the hydrogen peroxide, which makes it potentially useful in immunoassays and glucose sensing. The dynamics of an SWNT-based optical sensor were compared to a flux-measuring electrochemical sensor for in vivo fluorescence sensing of glucose [111]. Both sensors demonstrated an approximately linear response to blood glucose levels. However, the SWNT-based optical sensor, which transduces glucose concentration, not flux, directly was found to be significantly more stable to membrane biofouling. In another report, hollow-core optical fibers integrated with SWNTs were used for VOC detection [112]. Good sensitivity and fast response time were achieved by characterizing the reflectance of the sensing probe. A different approach for the immobilization of glucose oxidase on SWNTs was also explored [113]. In this report, the near-infrared (NIR; 700–1200 nm) luminescence properties of the aqueous solution of DNA wrapped SWNTs did not change after addition of glucose oxidase. However, with surface bound potassium ferricyanide (PFC) as an electron transfer mediator, quenched emission of the SWNTs (after addition of PFC) was easily restored by glucose injection, as a result of the reaction between PFC and hydrogen peroxide, which

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FIGURE 20.3 SWNT-based protein microarray chip. (a) A schematic view of protein microarray formation on an SWNT substrate. (b) An AFM image of the SWNT film grown by the CVD method. (c) Fluorescence images after reaction of the substrates (containing biotin-BSA and SpA) with Cy5SA, Cy3-IgG, and a mixture of both. SA, streptavidin; SpA, a specific antigen recognized by the IgG. (Adapted from Byon et al. [114].)

was generated during glucose oxidation. The increased SWNT luminescence with rising glucose concentration can therefore be detected using this method. High-yield CVD-grown SWNT films were reported to be efficient substrates for microarray protein chip applications (Fig. 20.3) [114]. A simple one-step treatment of the SWNT films allows both the immobilization of probe proteins and efficient suppression of NSB without the need for additional bovine serum albumin (BSA). Since BSA treatment frequently leads to ambiguous results in various assays, mostly arising from the size and intrinsic NSB characteristics of BSA itself, it was suggested that such a BSA-free system should provide new opportunities for extensive and detailed studies of diverse biomolecular interactions. 20.2.8 Challenges and Future Directions for SWNT-Based Sensors SWNT-based sensors can have enormous impact on both disease diagnosis and treatment monitoring. However, clinical use of SWNT-based sensors is not yet mature. Separation of

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metallic and semiconducting SWNTs to fabricate homogeneous sensing arrays, as well as the purification of SWNTs, is a major obstacle, unlike nanowires made from semiconducting materials [6]. Thus whether SWNT is the ideal material for nanosensing applications is still debatable. More importantly, whether SWNT-based sensing can compete well with DNA microarrays and mass spectroscopy-based proteomics remains unclear. As nanoscale sensing evolves, incorporation of SWNT-based sensors for multiplexed sensing will need to be applied to more complicated proteins and sample systems. Site-specific modification to improve selectivity/reproducibility, protective coatings to minimize NSB of biomolecules, and arraying sensor elements are current approaches to achieving reliable and consistent performance. SWNT arrays for conductivity-based sensing are probably more ideal than other sensing mechanisms in that they do not require labeling, external probes, or optical excitation. In many cases, literature reports have mixed findings. It will be ideal if different sensors (SWNT-based as well as those based on other nanoparticles) can be compared side-by-side using the same model system, which may significantly help in deciding which sensors are the best candidates for potential clinical testing. The National Cancer Institute (NCI) has required each of its funded nanotechnology centers to test their newly developed nanosensors using the same standard samples, which is expected to readily identify which new sensors truly stand out from the large pool of candidates. Extending a similar standard to a much broader range of research laboratories across the country would be highly beneficial to patients. Choosing the right candidate(s) at an early stage not only saves precious research time, but also significantly reduces the cost for new nanosensor development. Breakthrough in several research areas may lead to significant advances in the development of SWNT-based sensors. One area is the development of responsive biosensors that can change their configuration (e.g., size, gap, alignment, or stiffness) to improve sensitivity, selectivity, and/or performance based on the measurement of a specific analyte. The idea of this biosensor design is to use the intrinsic actuation properties of SWNTs to create a device that can be adjusted to improve its sensing capabilities or to add other functionalities such as delivering a drug or controlling a process that is related to an in-body device. Another promising research direction is hybrid biosensors that combine engineering and biological materials to form a more sensitive device, where the key is to find proteins that gate the process based on specific biomolecules. Lastly, the development of nanosensors that could be injected intravenously or implanted into the body, in order to monitor and record key physiological variables continuously, is another vibrant research area. Three major problems must be overcome to develop an in vivo biosensor: communication between the sensor and the outside world, elimination of biofouling, and potential toxicity of the sensor. Synthesis, processing, and device fabrication techniques for SWNTs have greatly improved after intensive research over the last decade. The future of SWNT-based sensors looks brighter than ever, yet many hurdles remain to be overcome.

20.3 GENE DELIVERY WITH SWNTs Gene therapy holds great potential for the treatment of a wide variety of diseases. However, the safety of gene therapy has always been a major concern for scientists, clinicians, and the general public [115]. Currently, gene therapy is generally performed through local administration. In addition, various strategies have been investigated to improve the gene

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expression efficiency and the safety profile of gene therapy for clinical applications [116]. SWNTs have been explored as a novel class of gene delivery vehicles because of their unique chemical/physical properties. In the first example of an SWNT-based gene delivery system, SWNTs were functionalized with ammonium to complex plasmid DNA through ionic interactions [117]. These complexes were proposed to enter cells by an endosome-independent mechanism. It was shown that these complexes gave higher DNA uptake and gene expression in vitro than DNA alone. However, such comparison is of no practical significance. Comparison with other DNA delivery agents, such as viral vectors and cationic polymers [118, 119], should have been carried out. After this proof-of-concept study, which demonstrated the potential of SWNTs for gene delivery applications, the interactions of different types of functionalized SWNTs with plasmid DNA were analyzed by various methods [120]. It was concluded that both the SWNT surface area and charge density were critical for the interaction between the SWNTs and the DNA. Systematic investigations of the cellular uptake mechanism for SWNTs were carried out [121]. Unlike the above-mentioned reports, it was found that intracellular transportation of proteins and DNA by SWNTs is through clathrin-dependent endocytosis, rather than caveolae or lipid rafts. The use of SWNTs as molecular transporters was demonstrated for various cargoes such as small molecules and proteins [122]. Recently, SWNTs have been employed in a number of studies for the delivery of genes such as antisense nucleic acids and small interfering RNAs (siRNAs). The major limitations of antisense therapy in the clinical setting are the rapid degradation of antisense nucleic acids and poor diffusion across the cell membrane [123]. To explore the potential of SWNTs in antisense therapy, SWNTs were used to deliver antisense myc (a oncogene [124]) into cells [125]. When compared with either antisense myc or SWNTs, the SWNT–antisense myc conjugate performed much better in the inhibition of cell proliferation, induction of apoptosis, and downregulation of myc at both the mRNA and protein levels. Recently, it was found that SWNTs can cause the polyA tail of single-stranded mRNA to form a duplex structure [126]. Since all mRNAs in eukaryotic cells have a polyA  tail at the 3 end, this finding may potentially enable the application of SWNTs in targeting specific gene sequences. Gene silencing with siRNA, an attractive approach to probe gene function in mammalian cells, is also expected to be highly efficacious for therapeutic applications in the clinic [127–129]. The major limitation of siRNA-based therapy is the efficiency of siRNA delivery, where cytoplasmic delivery is a prerequisite for silencing mRNA expression through RNA interference (RNAi). Because of their ability to cross cell membranes, SWNTs were used to deliver siRNA and DNA through a cleavable disulfide bond linkage (Fig. 20.4) [130]. Subsequently, SWNTs were explored for siRNA delivery into human T cells and primary cells (Fig. 20.4) [131]. It was shown that the delivery capability and RNAi efficiency of SWNTs exceeded that of nonviral transfection agents such as liposomes. Further, cellular uptake was dependent on the functionalization and the polyethylene glycol (PEG) chain length on the SWNTs. It was proposed that hydrophobic interaction is one of the major driving forces for cellular delivery of siRNA. In another study, delivery of telomerase reverse transcriptase siRNA with positively charged SWNTs was found to suppresses tumor growth both in vitro and in vivo [132]. SWNT-based siRNA delivery, targeting various genes, has also been achieved in primary cardiomyocytes and chronic myelogenous leukemia cells [133, 134].

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FIGURE 20.4 SWNT-based gene (DNA and siRNA) delivery into cells. (a) Functionalized SWNTs were conjugated to the cargo (DNA or siRNA) through cleavable disulfide linkages. (b) Three confocal images recorded with different focal planes (top, middle, and bottom of cells) along the viewing direction for a cell after incubation with SWNT–DNA conjugate. The red color surrounding the nucleus corresponds to Cy3-labeled DNA. (c) Confocal images of untreated human T cells (left), cells treated with siRNA in complex with a commercial transfection agent (middle), and cells treated with SWNT–siRNA complexes (right), respectively, after staining for the cell surface marker targeted by the siRNA. Scale bar: 40 ␮m. (Adapted from Kam et al. [130] and Liu et al. [131].)

These reports clearly demonstrated the potential of SWNTs as gene delivery systems. However, much future research and optimization will be needed before any clinical investigation since systemic delivery of genes (either DNA or siRNA) with SWNTs has not been achieved in living subjects. The barriers for gene delivery in vivo far supersede the barriers posed by the cell membrane in culture. Whether these SWNT conjugates can efficiently extravasate and deliver sufficient quantity of genes into the targeted cells in live animals remains unclear. In particular, delivering foreign DNAs into the cell nucleus is very challenging. The large surface area of SWNTs, which offers ample sites for the incorporation of nuclear localization sequences [135, 136], may be advantageous for DNA delivery and should be explored in future studies.

20.4 DRUG DELIVERY WITH SWNTs The same properties that make SWNTs attractive as gene delivery vehicles also make them desirable for drug delivery applications. In one report, it was found that various proteins adsorb spontaneously on the side walls of SWNTs through NSB [137]. More importantly, the adsorbed proteins could be delivered into the cytoplasm of various mammalian cells, presumably via the endocytosis pathway, which can lead to various biological sequences depending on the proteins used. A recent study showed that dispersion of SWNTs by ultrasonication with phospholipid-PEG can fragment the polymer, thus interfering with its ability to block nonspecific uptake by cells [138]. Since ultrasonication is a technique commonly used to disperse SWNTs [139], this may be a concern for drug delivery applications. One report showed that large surface areas exist for supramolecular chemistry on functionalized SWNTs, either noncovalently or covalently [140]. Water-soluble SWNTs with PEG functionalization can allow for very high degrees of ␲-stacking of aromatic molecules such as anticancer drugs and/or certain fluorescent molecules, which can be

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subsequently released in a controllable fashion by varying the pH value. The dependence of the ␲-stacking strength on the diameter of SWNTs could also be exploited for controlled release of the loaded molecules. The fact that multiple species such as PEG, drugs, and fluorescent tags can be attached onto the surface of the same SWNT makes it highly useful for a wide variety of future applications. Covalent linkages have been explored for SWNT-based drug delivery. One of the major disadvantages of low molecular weight (MW) anticancer drugs is their short circulation halflives, which typically lead to low tumor uptake and poor therapeutic efficacy [141]. Aminefunctionalized soluble SWNTs have been derivatized with cisplatin prodrug conjugates, where multiple prodrug centers were incorporated onto each SWNT [142]. Through pHmediated release following endocytosis in cancer cells, the cytotoxicity of the SWNT– prodrug conjugate was >100-fold more potent than the prodrug itself. In a follow-up study, targeted SWNT-mediated cisplatin prodrug delivery was reported using folate as the targeting ligand to cells that overexpress the folate receptor [143]. A recent report from another group also showed that an SWNT-based, tumor-targeted prodrug delivery system exhibited high potency toward cancer cells [144]. Recently, SWNT-based drug delivery was shown to suppress tumor growth in living mice [145]. Paclitaxel (PTX, Taxol), a widely used anticancer drug [146], was attached to PEG chains on the surface of SWNTs via a cleavable ester bond. The SWNT–PTX conjugate afforded higher efficacy in suppressing tumor growth than PTX in a murine breast cancer model, primarily due to prolonged blood circulation, which led to higher PTX uptake in the tumor through the enhanced permeability and retention (EPR) effect (Fig. 20.5). Such EPR effect is due to the nonspecific accumulation of SWNT–PTX in the tumor, since the tumor vasculature is more leaky than normal blood vessels and there is no lymphatic drainage in the tumor [147]. In future studies, tumor-specific targeting may further enhance the therapeutic index of anticancer drugs delivered by SWNTs. Various antibodies have been attached to SWNTs, either noncovalently or covalently. Due to the large surface area of SWNTs, it was shown that antibodies adsorbed onto SWNT bundles could stimulate T cells more effectively than equivalent concentrations of the antibody itself, potentially useful for applications such as immunotherapy [148]. In another study, a multivalent system based on SWNTs was constructed to target T cells both in vitro and in living animals [149]. Using a biotin–neutravidin linkage, the functionalized SWNTs were linked with antibodies for T cell receptor postsignaling endocytosis, as well as a synthetic fusogenic polymer for lysosome disruption. Although cytoplasmic delivery of SWNTs was accomplished in live animals, whether this system can enable therapeutically efficacious drug/gene delivery in vivo remains to be demonstrated in the future. In summary, SWNTs have been explored for drug delivery applications in cancer cells and animals through NSB, targeted and nontargeted approaches, as well as controlled release. Several studies in animals have been reported, which makes SWNT-based drug delivery one step closer to potential clinical applications than SWNT-based gene delivery. More importantly, the biological barriers faced by drug delivery are much less challenging than those confronted by gene delivery. In many cases, EPR-based drug delivery is already therapeutically efficacious and many anticancer drug formulations have been approved for clinical use without any targeting moiety [150, 151]. Once the potential toxicity of SWNTs is fully addressed, they may find widespread use in the clinic because of their versatile chemistry.

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FIGURE 20.5 Drug delivery with SWNTs for anticancer therapy in live animals. (a) A schematic illustration of PTX conjugation to phospholipid-PEG functionalized SWNT, which allows for PTX release in vivo by cleavage of the ester bond. (b) A photo of representative tumors harvested from untreated mice, Taxol-treated mice, and SWNT–PTX treated mice at the end of the study. (c) TUNEL (indicating apoptosis status) and Ki67 (indicating cell proliferation) staining of tumor slices from mice after different treatment, which confirmed the therapeutic efficacy of SWNT–PTX. The fluorescence signal is shown in red and blue staining indicates the nuclei. Scale bar: 100 ␮m. (Adapted from Liu et al. [145].)

20.5 VACCINE DELIVERY WITH SWNTs A few studies have demonstrated that SWNTs can also be useful for vaccine delivery. For example, a peptide from the foot-and-mouth disease virus, when linked to SWNTs covalently, was able to retain the structural integrity and immunogenicity hence could be recognized by monoclonal and polyclonal antibodies [152]. In another report, a neutralizing and protective B cell epitope was linked to SWNTs via selective chemical ligation [153]. The

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SWNT displayed B cell epitope retained its conformational characteristics because it could be recognized by specific antibodies. Subsequently, it was reported that functionalized SWNTs could also be used for delivering oligodeoxynucleotides into target cells [154]. Further investigation in animal models is required to evaluate the immunogenicity of these SWNT-based complexes in vivo and elucidate their potential therapeutic or adjuvant properties.

20.6 THERMAL THERAPY WITH SWNTs Aside from the use of SWNTs as delivery vehicles, the physical properties of the SWNT itself can also be exploited for therapeutic applications. In a pioneering report, the intrinsic strong optical absorbance of SWNTs was used for selective cancer cell killing [155]. It was reported that continuous NIR radiation of SWNT-loaded cells can cause cell death because of excessive local heating of the SWNTs in vitro. Taking advantage of this phenomenon, selective cancer cell destruction was achieved by functionalization of SWNTs with folate, which can enable internalization of SWNTs into cells overexpressing the folate receptor on the surface (Fig. 20.6). Together, the versatile delivery capabilities of SWNTs combined with their intrinsic optical properties can lead to a new class of novel materials for cancer therapy. Subsequently, SWNTs have been used for photothermal antimicrobial therapy in Escherichia coli [156]. Upon exposure to multipulse lasers at 532 and 1064 nm with nanosecond pulse duration, notable changes in bacteria viability were observed. Thermal ablation of tumor cells with antibody-functionalized SWNTs was also reported recently by another group [157].

FIGURE 20.6 Photothermal cancer cell killing with SWNTs. (Left) Selective internalization of functionalized SWNTs into folate receptor overexpressing (FR+) cells, followed by NIR laser radiation, caused cell death. (Right) Cells that do not express the folate receptor remain viable after the same treatment. Inset: high magnification image of the cells. (Adapted from Kam et al. [155].)

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Heat release from SWNTs in a radiofrequency (rf) field was found to be cytotoxic in cell culture [158]. When treated in the rf field, human cancer cell lines incubated with functionalized SWNTs showed selective, SWNT concentration-dependent thermal destruction. In a rabbit hepatic tumor model, SWNT injection followed by rf field treatment resulted in complete tumor necrosis at 48 hours while control tumors treated with rf without SWNT injection remained completely viable. Conjugation of certain targeting ligands to enhance the SWNT uptake by cancer cells may further improve the therapeutic effect in future studies. Various other nanoparticles have been used for tumor ablation based on their intrinsic physical properties. For example, gold nanoparticles have been tested for photothermal destruction of tumor cells/tissues in vitro and in vivo [8, 159]. Inductively heating a radiolabeled, antibody-linked iron oxide nanoparticle by externally applied alternating magnetic field was shown to cause tumor necrosis in mice bearing human breast cancer tumors [160, 161]. These nanoparticles with intrinsic therapeutic potential, including SWNTs, may have far reaching impact in the treatment of cancer and other diseases, after they are functionalized for targeted drug/gene delivery and/or noninvasive imaging.

20.7 NONRADIONUCLIDE-BASED MOLECULAR IMAGING WITH SWNTs The field of molecular imaging, “the visualization, characterization and measurement of biological processes at the molecular and cellular levels in humans and other living systems” [162], has expanded tremendously over the last decade. In general, molecular imaging modalities include molecular magnetic resonance imaging (MRI), magnetic resonance spectroscopy (MRS), optical bioluminescence, optical fluorescence, targeted ultrasound, single photon emission computed tomography (SPECT), and positron emission tomography (PET) [163]. Many hybrid systems that combine two or more modalities are also commercially available and certain others are under active development [164–166]. Molecular imaging takes advantage of traditional diagnostic imaging techniques and introduces molecular imaging agents (probes) to determine the expression of indicative molecular markers at different stages of diseases. Noninvasive detection of the molecular markers can allow for much earlier diagnosis, earlier treatment, and better prognosis that will eventually lead to personalized medicine [167, 168]. Continued development and wider availability of scanners dedicated to small animal imaging studies, which can provide a similar in vivo imaging capability in mice, primates, and humans, can enable smooth transfer of knowledge and molecular measurements between species, thereby facilitating clinical translation. The physical properties and relative ease for chemical modification have made SWNTs the most versatile nanoparticle for molecule imaging applications. 20.7.1 MRI with SWNTs MRI detects the interaction of protons (or certain other nuclei) with each other and with the surrounding molecules in a tissue of interest [169]. The endogenous tissue contrast can be enhanced by administration of exogenous contrast agents [170, 171]. Traditionally, Gd3+ chelates have been used to enhance the T1 contrast [172] and iron oxide nanoparticles have been used to increase the T2 contrast [5]. Gd3+ -functionalized SWNTs have been studied as MRI contrast agents [173]. The nanoscale loading and confinement of aquated Gd3+ ion clusters within ultrashort (US;

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FIGURE 20.7 Molecular imaging with SWNTs. (a) MRI of Gd3+ filled SWNTs at different pH value. (b) NIRF images of antigen-positive and control cells treated with an SWNT–antibody conjugate. (c) (Left) Photograph of an integrin ␣v ␤3 -positive tumor-bearing mouse depicting the area scanned with Raman spectroscopy (black box). (Right) Raman tumor maps of mice receiving SWNT–RGD conjugate or unconjugated SWNT at 72 h postinjection. (d) Photographs of the tumors in mice and the corresponding photoacoustic subtraction images (green). (e) Two-dimensional projection of the PET images of integrin ␣v ␤3 -positive tumor-bearing mice at 8 h postinjection of SWNT–RGD without and with (denoted as “Block”) coinjection of an excess amount of the RGD peptide. Arrowheads indicate the tumors. (Adapted from Hartman et al. [174], Welsher et al. [186], Zavaleta et al. [204], De la Zerda et al. [215], and Liu et al. [224].)

20–80 nm) SWNTs was found to be linear superparamagnetic molecular magnets with a MRI efficacy of 40 times greater than that of the Gd3+ -based contrast agents in current clinical use [173]. In a follow-up study, it was found that these Gd3+ -containing US SWNTs are also sensitive to the pH value of the environment (Fig. 20.7a) [174]. These findings suggested that Gd3+ -containing SWNTs might be useful in the clinic for early detection of cancer, where the extracellular pH of tumors can drop to pH 7.0 or below.

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20.7.2 Fluorescence Imaging with SWNTs Due to their strong light quenching properties, direct fluorescent labeling of SWNTs has not been very successful for optical imaging applications. However, studies have shown that optical properties of SWNTs can also be manipulated without covalent modification [175, 176]. The major drawback of optical imaging in living subjects is the poor tissue penetration and intense scattering of light [177]. In the NIR region, the absorbance of all biomolecules reaches a minimum, which provides a relatively clear window for optical studies/applications [178, 179]. Therefore NIR fluorescence (NIRF) imaging can provide opportunities for rapid and cost-effective preclinical evaluation in small animal models. Many reports have investigated the NIRF imaging of SWNT and/or SWNT conjugates ex vivo [180–183], in cells (Fig. 20.7b) [184–187], and in living subjects [188]. Among all molecular imaging modalities, no single modality is perfect and sufficient to obtain all the necessary information for a particular question [163]. A combination of multiple imaging modalities can offer synergistic advantages over any single modality alone. In one report, heterostructured complexes formed from iron oxide nanoparticles and SWNTs were investigated as multimodal imaging agents [189]. Macrophage cells that engulf the DNA-wrapped complexes were imaged with both MRI and NIRF mapping. Compared with fluorescent dyes or quantum dots, the more popular fluorescent agents, the quantum efficiency of SWNTs is much lower. Therefore NIRF imaging of SWNTs is quite challenging in live animals (e.g., mice and rats). Another limitation of NIRF microscopy of SWNTs is the diffraction limit to resolution, which is more relevant for high-resolution imaging of cells. Recently, a nonperturbing, far-field optical measurement was reported, which can allow for sub-wavelength mapping of single molecule chemical reaction sites on semiconducting SWNTs [190]. X-ray fluorescence has also been applied for the first time to study macrophages exposed to crude and purified SWNTs [191].

20.7.3 Raman Spectroscopy with SWNTs Raman spectroscopy, which can differentiate the spectral fingerprint of various molecules, can have very high multiplexing capabilities [192, 193]. However, the inherently weak magnitude of the Raman effect limits its sensitivity, which severely hampered the biomedical applications of Raman spectroscopy. SWNT has an intense Raman peak produced by the strong electron–phonon coupling, which causes efficient excitation of tangential vibration in the SWNT upon light exposure [194]. A number of literature reports have focused on Raman imaging and examination of the surface-enhanced Raman scattering (SERS) of SWNTs [195–202]. In a pioneering study, SWNTs and other SERS nanoparticles were used to demonstrate whole-body Raman imaging, multiplexing, as well as in vivo tumor targeting with a small-animal Raman imaging system [203]. Subsequently, an optimized Raman microscope was used to further evaluate tumor targeting and localization of SWNTs in mice (Fig. 20.7c) [204]. The molecular target and the targeting ligand used in this study was integrin ␣v ␤3 and an arginine–glycine–aspartic acid (RGD, potent integrin ␣v ␤3 antagonist [205]) containing peptide, respectively, one of the most extensively studied and validated receptor–ligand pair over the last decade [206, 207]. One major advantage of Raman spectroscopy is its superb potential for multiplexed imaging. In a recent study, multiplexed Raman imaging of live cells with isotopically modified SWNTs was reported [208]. SWNTs with different isotope (12 C and 13 C)

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compositions, which in turn exhibit distinct Raman G-band peaks, were used for multiplexed Raman imaging in different cancer cells overexpressing specific cell surface receptors. Functionalized, macromolecular SWNTs have recently been used as multicolor Raman labels for highly sensitive, multiplexed protein detection in an arrayed format [209]. In future studies, SWNTs with controlled diameters (which also exhibit distinctive Raman peaks in their radial-breathing modes depending on the diameters of the SWNTs [210]) may also be investigated for multiplexed Raman imaging. 20.7.4 Photoacoustic Tomography (PAT) with SWNTs PAT is a cross-sectional imaging technique based on the photoacoustic effect [211]. In PAT, the tissue is usually irradiated by a short-pulsed laser beam to produce thermal and acoustic impulse responses. Locally absorbed light is converted into heat, which is further converted to a pressure rise via thermoelastic expansion of the tissue. The initial pressure rise propagates in the tissue as an ultrasonic wave, referred to as a photoacoustic wave, which is detected by ultrasonic transducers placed outside the tissue to produce electric signals. The electric signals are then amplified, digitized, and transferred to a computer to form a PAT image. A number of contrast agents for photoacoustic imaging have been suggested [212–214], yet most were not used for targeted imaging. In a recent study, SWNTs conjugated with cyclic RGD peptides were used to enhance the contrast of photoacoustic imaging of tumors in living mice [215]. Intravenous administration of the SWNT–RGD conjugate to tumorbearing mice showed eight times greater photoacoustic signal in the tumor than mice injected with SWNT alone (Fig. 20.7d). Taking advantage of the intrinsic Raman signal of the SWNTs, the in vivo PAT imaging results were further validated ex vivo with Raman microscopy. 20.7.5 A Brief Summary The major disadvantage of MRI is its inherent low sensitivity, which can only be partially compensated by working at higher magnetic fields (4.7–14 T), acquiring data for much longer time periods, and using exogenous contrast agents. Although proof-of-principle studies have been reported for molecular MRI of several targets [216], whether molecular MRI can significantly improve patient management remains to be elucidated. Whether Gd3+ -containing SWNTs can be molecularly targeted and perform well in animal studies needs to be investigated before any potential clinical applications can be in place. For clinical applications, optical imaging techniques will only be possible in limited sites such as the tissues and lesions close to the surface of the skin, tissues accessible by endoscopy, and intraoperative visualization [3]. Due to the poor fluorescence properties of SWNTs, they are not suitable for NIRF imaging in potential human studies. However, Raman spectroscopy and PAT may have potential clinical relevance/applications in the future, although the primary limitation is still light penetration beyond a few centimeters of tissue. The key advantages of Raman imaging over fluorescence imaging is the very high multiplexing capability and lack of confounding background signal from autofluorescence. In addition, the use of SWNTs as Raman tags is more advantageous over other Raman nanoparticles because of the inherent Raman signature of SWNTs, which requires no further labeling or encapsulation to produce a Raman signal. PAT compares very favorably

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to other imaging modalities with its precise depth information, submillimeter resolution, and nanomolar sensitivity [215]. With further improvement in background reduction, hardware and image reconstruction software, and the use of lasers with high repetition rates, it is very likely that PAT will find wide use in the future in both basic research and clinical care. The fact that SWNTs can be used as contrast agents in a number of different imaging techniques, which can significantly facilitate the cross-validation of the imaging result, is also another highly desirable trait over other contrast agents.

20.8 RADIONUCLIDE-BASED MOLECULAR IMAGING WITH SWNTs Radionuclide-based imaging techniques (PET and SPECT) have much more clinical relevance and hence are much more widely used in the clinic than molecular MRI and optical imaging. Not only is there no tissue penetration limit for these techniques, but PET and SPECT are also highly sensitive (down to the picomolar level) and quantitative [217]. In one early study, water-soluble hydroxylated SWNTs were labeled with 125 I to study the distribution in mice [218]. It was concluded that these SWNTs moved easily among the compartments and tissues of the body and behaved like small molecules, despite the fact that their apparent molecular weight was tremendously large. In a later report, 111 In-labled water-soluble SWNTs were tested for SPECT imaging [219]. Strikingly, intravenous administration of these functionalized SWNTs indicated that they were not retained in any of the reticuloendothelial system (RES) organs (e.g., liver or spleen) and were rapidly cleared from the circulation system through the renal excretion route. Because of the use of lead collimators to define the angle of incidence, SPECT imaging has a very low detection efficiency (< 10−4 times the emitted number of gamma rays) [220]. PET, on the other hand, has much higher sensitivity (greater than 10%) [221]. It was first developed in the mid-1970s [222] and dedicated PET scanners for small-animal studies were first reported in the late 1990s [223]. The biodistribution of 64 Cu-labeled SWNTs in mice has been investigated by PET, ex vivo biodistribution, and Raman spectroscopy [224]. It was found that these SWNTs are highly stable in vivo and the surface PEG chain length can significantly affect its blood concentration and biodistribution. Most importantly, efficient targeting of integrin ␣v ␤3 -positive tumor in mice was achieved with RGD peptide conjugated SWNTs (Fig. 20.7e). The intrinsic Raman signatures of SWNTs were used to directly probe the presence of SWNTs in mouse tissues and confirm the PET imaging results. All in vivo and ex vivo measurements confirmed that there was minimal renal uptake of the SWNT–RGD conjugate, and the majority of the conjugate was taken up by the tumor and the RES [224]. Subsequently, a few other reports also appeared on radiolabeled SWNTs (both targeted and nontargeted) [225–227]. Radiolabeled nanoparticles represent a new class of probes that have enormous potential for clinical applications. Different from other molecular imaging modalities, where typically the nanoparticle itself is detected, such as quantum dots and iron oxide nanoparticles for optical imaging and MRI, respectively, radionuclide-based imaging detects the radiolabel rather than the nanoparticle itself. The nanoparticle distribution is measured indirectly by assessing the localization of the radionuclide, which can provide quantitative measurement of the tumor targeting efficacy and pharmacokinetics only if the radiolabel on the nanoparticle is stable enough under physiological conditions. However, dissociation of the radionuclide (typically metal) from the chelator, and/or the radionuclide-containing

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TABLE 20.1 The Unique Physical and Chemical Properties of SWNTs Have Made Them Versatile Contrast Agents for a Number of Imaging Modalities Modality MRI NIRF Raman PAT PET/SPECT

Contrast Agent 3+

Gd SWNT SWNT SWNT Radioisotopes

Resolution

Sensitivity

Penetration

References

+++ ++ + +++ ++

+ ++ ++ ++ +++

+++ + + + +++

173, 174, 189 180–182, 185–191 195–204, 208, 209 215 219, 224–227

polymer coating from the nanoparticle, may occur, which can cause significant difference between the nanoparticle distribution and the radionuclide distribution. In several studies, SWNTs were reported to undergo either complete or partial renal clearance in mice, with little uptake by the liver or other organs of the RES [219, 225]. These findings defy the general trend of high RES uptake for nanoparticles [228, 229] and deserve further investigation/validation. Studies have shown that typically only molecules less than 70 kDa (a few nanometers in diameter) undergo renal clearance [230–232]. The SWNTs used in these reports are typically more than 200 nm in length, even up to a few micrometers [219, 224, 225]. It is very unlikely that SWNTs can be cleared from the kidney (dissociated radiolabel and polymer coating can, however) unless severe kidney damage has occurred. Nanoparticle-based imaging has touched upon every single modality of the molecular imaging arena [11]. The unique physical and chemical properties of SWNTs have made them versatile contrast agents for a wide variety of imaging modalities, including MRI, optical, SPECT, and PET imaging. In addition, SWNTs have also been investigated with several other imaging techniques such as Raman spectroscopy and PAT. A summary and comparison of these imaging techniques, with SWNTs or SWNT conjugates as the contrast agents, is shown in Table 20.1. It is expected that SWNT-based molecular imaging will continue to flourish for years to come.

20.9 TISSUE ENGINEERING WITH SWNTs The goal of tissue engineering is to replace diseased or damaged tissue with biologic substitutes that can restore and maintain normal function. Major advances in the areas of cell and organ transplantation [233], as well as advances in materials science and engineering, have aided in the continuing development of tissue engineering and regenerative medicine [234]. As the field of tissue engineering advances, there is a growing need for better monitoring and evaluation tools of engineered tissue along with new biomaterials to facilitate tissue growth. One class of nanomaterials that has the potential for multiple applications in tissue engineering is the SWNTs [235]. 20.9.1 Cell-Based Studies A number of reports have investigated, in cell culture, the potential of SWNTs in tissue engineering applications. Chemically functionalized water-soluble SWNT graft copolymers, when added to the culturing medium, were shown to increase the length of various neuronal

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processes [236]. These SWNT-based conjugates not only exhibited excellent biocompatibility with osteoblast cells but also appeared to modulate the cell phenotype. SWNTs grafted with polyethyleneimine, which could promote neurite outgrowth and branching, have been used as the substrate for neuron culture [237]. Subsequently, similar SWNTs were also shown to be suitable scaffold materials for osteoblast proliferation and bone formation, which represents a potential technological advance in the field of bone engineering [238]. SWNTs have been shown to support neuronal attachment and growth [239]. Simple chemical modification of the SWNTs was employed to control cell growth without interfering with the ongoing neuronal function. More importantly, electrical coupling of SWNTs and neurons was demonstrated, suggesting that SWNTs are flexible materials for tissue engineering applications that involve electrically excitable tissues such as muscles and nerves. Other groups have also demonstrated that SWNTs can be used as scaffolds for cell culture [240–242]. The capability of SWNTs in enhancing long-term cell proliferation can be of significance to in vitro cell amplification on a large scale, tissue regeneration or guided repair, as well as biomedical device applications. Composite polymeric scaffolds from alginate and SWNTs have been shown to improve the adhesion and proliferation of rat heart endothelial cells [243]. Recently, biodegradable SWNT–polymer composites were investigated for their effects on the viability of fibroblast cells [244]. Nearly 100% cell viability was observed on all crosslinked nanocomposites and cell attachment on their surfaces was comparable with that on tissue culture polystyrene, clearly indicating favorable cytocompatibility of these nanocomposites for potential use as scaffolds in tissue engineering applications. Neural stem cells (NSCs) are highly plastic neural precursors capable of adapting to environmental conditions and recreating signal transduction pathways [245]. Successful differentiation of mouse NSCs on layer-by-layer assembled SWNT composites was reported [246]. Mouse embryonic NSCs from the cortex successfully differentiated into neurons, astrocytes, and oligodendrocytes with clear formation of neurites (Fig. 20.8a). A follow-up study demonstrated that the fabrication of layer-by-layer assembled composites from SWNTs and laminin, an essential part of the human extracellular matrix, can be conducive to NSC differentiation and suitable for their successful excitation [247]. Extensive formation of functional neural networks was observed as demonstrated by the presence of synaptic connections, indicating that the protein–SWNT composite can serve as the material foundation of neural electrodes with chemical structure better adapted for long-term integration with the neural tissue. Additional studies to further understand the effect of electrical stimulus on the NSCs during the differentiation cycle and over a longer time frame are needed in the future.

20.9.2 Matrix Enhancement with SWNTs In tissue engineering, scaffolds play a pivotal role in providing temporary structural support, guiding cell growth, assisting nutrient/waste transport, and facilitating functional tissue/organ formation [248]. It is responsible for defining the space the engineered tissue occupies and aiding the process of tissue development. While popular synthetic polymers such as poly lactic acid (PLA) and co-poly lactic acid/glycolic acid (PLGA) have been widely used for tissue engineering, they lack the necessary mechanical strength [249, 250]. In addition, these polymers cannot be functionalized easily while SWNTs can.

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FIGURE 20.8 Applications of SWNTs in tissue engineering. (a) Confocal microscopy images of differentiated neurospheres grown on an SWNT composite. Neurospheres were stained for markers of NSCs (nestin), neurons (MAP2), astrocytes (GFAP), and oligodendrocytes (O4). The markers are shown in red and the cell nuclei are shown in blue. Scale bar: 20 ␮m. (b) Representative microCT images of the rabbit femoral condyle 3 months after implantation of either a poly(propylene fumarate) (PPF) or US–SWNT scaffold. (Adapted from Jan and Kotov [246] and Sitharaman et al. [253].)

Live smooth muscle cell seeded collagen–SWNT composite matrices were investigated [251]. Cell viability in all constructs was consistently above 85% after 1 week, suggesting that such collagen–SWNT composite matrices can be used as scaffolds in tissue engineering or other medical devices. Highly porous scaffolds made of three different materials including SWNTs were fabricated to evaluate the effects of material composition and porosity on the scaffold pore structure, mechanical properties, and marrow stromal cell culture [252]. Based on the in vitro results, it was suggested that functionalized US SWNT nanocomposite scaffolds with tunable porosity and mechanical properties hold great promise for bone tissue engineering applications. Subsequently, the in vivo biocompatibility of US SWNT reinforced porous biodegradable scaffolds was tested in a rabbit model [253]. The US SWNT scaffolds, when implanted in rabbit femoral condyles or subcutaneous pockets, exhibited favorable hard and soft tissue responses at 1 and 3 months after implantation based on micro-computed tomography (microCT), histology, as well as histomorphometry (Fig. 20.8b). At 3 months, a threefold greater bone tissue ingrowth was seen in defects containing US SWNT nanocomposite scaffolds than the control scaffold, suggesting that these US SWNT scaffolds are very useful for assisting osteogenesis. The investigation of SWNTs for tissue engineering applications is still at the early development stage. As discussed in Section 20.2.3, studies have shown that enzymes can be attached to SWNTs and still retain enzymatic activity, which may allow a mixture of differently functionalized SWNTs to be added to a scaffold to create increasingly sophisticated structures. Potentially, enzymatic or protein functionalization, coupled with the

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electrochemical properties mentioned previously, may give rise to novel SWNT-based matrices that can not only provide the mechanical integrity for cell growth but can also monitor simultaneously the growth progress. 20.10 APPLICATIONS OF SWNTs IN BIOMEDICAL INSTRUMENTATION AFM is a promising technique for imaging nanoscale structures (e.g., proteins, nucleic acids, nanoparticles) with high resolution [254]. However, conventional silicon and silicon nitride AFM tips are not ideal for such applications. SWNTs, possessing small diameter, high aspect ratio, mechanical robustness, and unique chemical properties, have been investigated as AFM tips using a variety of assembly techniques such as CVD on the commercial silicon cantilever tip [255]. In an early study, SWNT-based AFM tips were used for multiplexed detection of polymorphic sites and direct determination of haplotypes in 10-kilobase sized DNA fragments [256], based on visual inspection of the AFM images of individual DNA molecules. This technology represents an important first step toward the use of nanotechnology in scanning probe microscopy in the biomedical arena. Attaching a smaller SWNT with subnanometer diameter may enable direct reading of the DNA sequence of these fragments. The high-resolution imaging capabilities of SWNT AFM probes have been demonstrated on a number of nanoparticles and biomolecules (Fig. 20.9) [254].

FIGURE 20.9 AFM with SWNTs as the tip. (a) A field emission/SEM image of SWNT bundles grown from an Si cantilever/tip assembly. (b) Detection of labeled DNA sites with SWNT tips. The arrow points to the streptavidin tag on the DNA. Scale bar: 100 nm. (c) The AFM image of a large scan area, a higher resolution image, and the crystal structure of the protein being imaged. (Adapted from Cheung et al. [255] and Woolley et al. [256].)

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Recently, a simple method to fabricate SWNT probes for AFM using the Langmuir–Blodgett (LB) technique was reported [257]. A two-step transfer process, dipping conventional AFM probes into the Langmuir monolayer followed by lifting the probes from the water surface, was able to assemble well-oriented and robust SWNTs that maintained their shape and direction even after successive scans. More importantly, it was also demonstrated that the LB method is a scalable process capable of simultaneously fabricating a large number of SWNT-modified nanoprobes. Carboxyl-terminated SWNTs have been immobilized onto the apexes of gold tips for STM [258], which gave high-resolution STM images of a diether monolayer formed on the graphite surface. In another study, it was shown that SWNTs at the apex of AFM probes can act as the template for the formation of nanoelectrodes through sputter coating with metal [259]. The resulting disk-shaped nanoelectrode at the AFM tip apex was capable of high-resolution electrochemical and topographical imaging. In addition to measuring topography, chemically functionalized SWNT-based probes can also measure the spatial arrangement of functional groups in a sample. Therefore functionalization of SWNT AFM tips offers great potential for high-resolution, chemically selective imaging of biological structures. Image-guided radiation therapy, with the goal of eradicating tumor cells while sparing healthy tissue, has been widely adopted in the clinic [260]. Image guidance at the time of treatment has been implemented using a variety of systems such as portal images, ultrasound devices, and CT scanners [261, 262]. With the exploding interest in small-animal-based research, there is an increasing need for such a system dedicated to small animals. Recently, CNT field emission technology was used to develop a micro-radiotherapy system for image-guided high precision irradiation of mice [263]. Through the field emission control of its individually addressable X-ray pixel beams, the micro-radiotherapy system can electronically shape the radiation field and form an intensity modulation pattern with temporal and spatial resolutions of approximately milliseconds and <2 mm, respectively.

20.11 OTHER APPLICATIONS OF SWNTs SWNTs have remained at the forefront of nanomaterial research since their discovery in the early 1990s, largely due to their exceptional and unusual mechanical, electrical, optical, and chemical properties. No other nanoparticle has been explored for such a diverse array of biomedical applications (Fig. 20.10). Besides the areas discussed above, SWNTs have also been investigated for certain other uses. One study showed that SWNTs can inhibit DNA duplex association and selectively induce human telomeric i-motif DNA formation by binding to the 5 -end major groove under physiological conditions [264]. This report suggested that SWNTs might have the intriguing potential to modulate human telomeric DNA structures in vivo, which may be useful for future drug design and anticancer therapy. The investigation of surface structure and characteristics of anthrax (Bacillus anthracis) spores as related to their binding by molecular species is very important to the development of countermeasure technologies for the detection and decontamination of anthrax spores [265]. SWNTs have been used as a scaffold for displaying multivalent monosaccharide ligands that bind effectively to anthrax spores with divalent cation mediation to cause significant spore aggregation [266]. Although optimization of such binding and the elucidation

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FIGURE 20.10 The broad applications of SWNTs in the biomedical arena.

of related mechanistic details remain to be investigated, the aggregation of anthrax spores may be exploited for antibioterrorism applications.

20.12 TOXICITY OF SWNTs IN VITRO Like most new technologies, there are concerns about the possible side effects caused by the use of SWNTs. A large number of publications have appeared on this topic over the last several years. The potential toxicity of nanoparticles mainly comes from two aspects. First, nanoparticles can enter the body through the skin, lungs, or intestinal tract, depositing in several organs, and may cause adverse biological reactions [267]. In addition, the toxicity of nanoparticles also depends on whether they get cleared from the body and whether the host can raise an effective response to sequester or dispose of the particles [268]. Second, the toxicity can come from the material itself [269]. Although a few studies indicated that SWNTs exhibit little to no toxicity to cells [270, 271], many more reports have suggested that SWNTs do have adverse and/or toxic effects. However, the proposed mechanism for such effects varies dramatically, including increased oxidative stress and inhibition of cell proliferation due to NF-␬B activation [272], induction of cell apoptosis and decrease in cellular adhesion ability [273], inhibition of cell proliferation [274, 275], indirect cytotoxicity by alteration of the cell culture medium [276], induced oxidative stress and loss of antioxidants [277], expression of inducible stress responsive genes [278], as well as loss of cell viability [279]. A number of parameters were found to have considerable impact on the reactivity/toxicity of SWNTs. One study showed a dose- and time-dependent increase of intracellular reactive oxygen species and a decrease of the mitochondrial membrane potential when cells were treated with commercial SWNTs; however, incubation with purified SWNTs had no effect on the cells [280]. This report, along with a few other publications [281–283], suggested that the metal traces associated with commercial SWNTs are likely responsible for the adverse biological effects. Various studies have also suggested that the degree of SWNT functionalization [284], the diameter of the SWNTs [285, 286], the amount of adsorbed

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proteins on the SWNTs [287], as well as a combination of multiple factors [288] can affect the SWNT–cell interactions. To decipher the apparent contradictory findings of various literature reports, one study demonstrated that SWNTs interact with some tetrazolium salts (commonly used in cellbased assays) but not with the others [289]. This study brought the experimental data of many of the abovementioned reports to question. It was suggested that the cytotoxicity data of SWNTs should be verified with at least two or more independent test systems. Since SWNTs are known to also interact with many colorimetric indicator dyes frequently used in cytotoxicity assays, it was suggested that a clonogenic assay would be more accurate since no dye is used [290].

20.13 TOXICITY OF SWNTs IN RODENTS The toxicological effects of SWNTs have been studied in rodents. In an early study, pharyngeal aspiration of SWNTs elicited pulmonary effects in mice that combined acute inflammation with early onset yet progressive fibrosis and granulomas [291]. Functional respiratory deficiencies and decreased bacterial clearance were also observed in mice treated with SWNTs. A follow-up study showed that such enhanced acute inflammation and pulmonary injury with delayed bacterial clearance after SWNT exposure may lead to increased susceptibility to lung infection in mice [292]. A role for NADPH oxidase-derived reactive oxygen species in determining the course of the pulmonary response to SWNTs was then suggested [293]. Subsequently, the critical role of vitamin E (the major lipidsoluble antioxidant of the antioxidant protective system [294]) in SWNT-induced reactions was also proposed, because lowered levels of antioxidants in vitamin E-deficient mice were associated with higher sensitivity to SWNT-induced acute inflammation and enhanced profibrotic responses [295]. Since the pulmonary levels of vitamin E can be manipulated through diet, its effects on SWNT-induced inflammation may be of practical importance to the optimization of protective strategies. SWNT inhalation was found to be more effective than aspiration in causing inflammatory response, oxidative stress, collagen deposition, and fibrosis, as well as mutations of the K-ras gene locus in the mouse lungs [296]. Intratracheal instillation of SWNTs into mice caused induced alveolar macrophage activation, various chronic inflammatory responses, and severe pulmonary granuloma formation [297]. A series of signal transduction and biological events were suggested to be responsible for these effects. Interestingly, a subsequent report showed that exposure to more dispersed SWNT structures could alter the pulmonary distribution and response [298]. Pharyngeal aspiration of a dispersed preparation of SWNTs in mice caused no granulomatous lesions or epithelioid macrophages, confirming that granulomatous lesions encased by epithelioid macrophages were produced by large agglomerates in the previous intratracheal instillation or aspiration studies of SWNTs. A few other studies also reported that oral exposure or oropharyngeal aspiration of SWNTs did not exhibit significant toxic effect in mice [299, 300]. Most of the abovementioned studies used pristine SWNTs, which are irrelevant to the biomedical applications of SWNTs, in which the SWNTs would have to be functionalized for any potential use. Furthermore, SWNT-based agents will typically be intravenously injected for biomedical applications, rather than pharyngeal aspiration or inhalation. Certainly, SWNT itself is toxic to some extent just like any other nanoparticle. The most important question is not how toxic these “naked” nanoparticles are, but how to modify

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FIGURE 20.11 The distribution and clearance of SWNTs in mice. (a) Biodistribution of SWNTs functionalized with different sized PEG chains (SWNT-l-2kPEG, SWNT-l-5kPEG, and SWNT-br7kPEG) at 1 day postinjection as measured by Raman spectroscopy. (b) The concentrations of SWNTs retained in the liver and spleen of mice over a period of 3 months. Longer PEG chains on the SWNTs gave lower RES uptake. (c) No obvious toxic effect was seen in the liver and spleen of mice injected with SWNTs at 3 months postinjection. Scale bar: 50 ␮m. (Adapted from Liu et al. [301].)

and functionalize them so that they do not exhibit any toxicity and thereby can be useful in biological systems. In a recent study, the long-term fate of functionalized SWNTs intravenously injected into animals was investigated [301]. Using the intrinsic Raman signatures of SWNTs, the blood circulation of intravenously injected SWNTs and their accumulation in various organs and tissues of mice over a period of 3 months was measured (Fig. 20.11). It was found that SWNTs functionalized by branched PEG chains enabled very long blood circulation up to 1 day, relatively low uptake in the RES, and virtually complete clearance from the main

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organs in about 2 months. Raman spectroscopy also detected SWNTs in the feces and other clearance organs, confirming the excretion and clearance of SWNTs from mice via the biliary and renal pathways. Most importantly, no toxic side effect of SWNTs in mice was observed in necropsy, histology, and blood chemistry measurements. Such a favorable in vivo toxicity profile of systematically administered, functionalized SWNTs was also confirmed by a number of other studies [302, 303].

20.14 TOXICITY OF SWNTs IN OTHER SPECIES Interestingly, the potential toxicity of SWNTs has been studied in a variety of other species since these SWNTs will eventually reach the ecosystem. In one study, acute and chronic toxicities of SWNTs were evaluated with the estuarine copepod Amphiascus tenuiremis [304]. It was found that copepods ingesting purified SWNTs exhibited no significant effect on mortality, development, and reproduction across the range of different exposure levels. However, exposure to the more complex commercial SWNTs (not purified) caused a significant increase in mortality, as well as a decrease in the fertilization rates and molting success. This study suggested that the small synthetic by-product fractions caused the increased mortality and delayed development rather than the SWNTs themselves, which corroborated with many of the previously discussed cell-based investigations [280–283]. Another study using zebrafish embryos also suggested that materials associated with raw SWNTs (likely the metal contaminants) have the potential to affect aquatic life when released into the aquatic environment while the SWNTs themselves had minimal toxicity [305]. SWNTs were found to be a respiratory toxicant in rainbow trout [306]. Although the fish were able to manage oxidative stress and osmoregulatory disturbances, pathological examination raised concerns about the cell cycle defects, neurotoxicity, and other unidentified blood-borne factors. A recent study using ciliated protozoan, widely used by ecotoxicologists because of its role in the regulation of microbial populations through the ingestion and digestion of bacteria, showed that SWNTs can be internalized, which may allow them to move up the food chain [307]. Ecotoxicity of SWNTs, along with many other nanoparticles, to aquatic organisms was studied, and the toxicity data generated on the various nanoparticles reflected a wide spectrum of sensitivity that was biological level, test, and endpoint specific [308]. Most importantly, SWNTs were found to be much less toxic than many other nanoparticles. A study on the effects of functionalized and nonfunctionalized SWNTs on the root elongation of six crop species was also carried out [309]. Not surprisingly, it was found that nonfunctionalized SWNTs affected the root length more than functionalized SWNTs.

20.15 FUTURE DIRECTIONS AND CHALLENGES Functionalized SWNTs have made far-reaching impact in the biomedical arena and it is expected that SWNTs will remain a major player in biomedical research for years to come (Fig. 20.12). With the capacity to provide enormous sensitivity, throughput, and flexibility, SWNTs have the potential to profoundly impact disease diagnosis and patient management in the near future. Big strides have been made and many proof-of-principle studies have been successfully performed, the future of SWNTs looks brighter than ever yet many

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FIGURE 20.12 A wide variety of functional moieties have been attached to SWNTs for a broad array of biomedical applications.

hurdles remain to be conquered. Much research effort in the following areas will be needed before clinical applications of SWNT-based agents are possible. 20.15.1 Multimodality Imaging and Early Detection of Diseases A combination of certain imaging modalities can offer synergistic advantages over any single modality alone [11, 207]. Dual-modality agents that combine PET, which is very sensitive and highly quantitative [310], and optical imaging, which can significantly facilitate ex vivo validation of the in vivo data, should be of particular interest for future SWNT-based biomedical research. The relatively large size of SWNTs may potentially allow for simultaneous multiple receptor binding of the targeting ligands on the same particle. Therefore targeting multiple, closely related receptors simultaneously may be more efficacious for molecular imaging and therapy, with both improved sensitivity and better specificity. We envision that peptides or small molecules are better targeting ligands than antibodies for SWNT conjugation, not only to keep the overall size small but also to take fully advantage of the polyvalency effect [311] since significantly more small molecules (MW < 1000) can be attached to an SWNT than the macromolecular antibodies (MW ∼150,000). Nanotechnology can have an early, paradigm-changing impact on how clinicians will detect diseases (e.g., cancer and cardiovascular diseases) in the earliest stage. One of the most pressing needs in clinical oncology is for imaging agents to detect tumors that are far smaller than what is possible with today’s technology, at a scale of 100,000 cells rather than 1,000,000,000 cells. Achieving this level of sensitivity requires better targeting efficacy of the imaging agents and signal amplification, both of which may be achieved with SWNTs. In terms of targeted imaging with SWNTs, vasculature targeting is more likely to succeed than targeting the tumor cells since SWNT-based agents may not extravasate well due to the relatively large size. Another area with near-term potential is the detection of mutations and genome instability in situ. Further work could result in an SWNT-based system capable of differentiating different types of tumors accurately and quickly, information that would be invaluable to clinicians and researchers alike. Implantable SWNT-based sensors may enable clinicians

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to more closely monitor the disease-free status of patients who have undergone treatment or individuals susceptible to the diseases because of various risk factors. SWNT-based nanosensors [17, 28] and in vivo imaging are both critical for future optimization of patient management. Ex vivo diagnostics in combination with in vivo diagnostics can provide a synergistic approach that neither strategy alone can offer. While imaging can give a whole-body perspective of the disease state in patients, detection of blood and urine markers of the disease can provide invaluable and complementary information of the biological responses to therapeutic intervention. Before treatment, after treatment, and more importantly during treatment, patient response can be evaluated by blood analysis and molecular imaging to ensure the accurate differentiation of responders from nonresponders and monitor their response to personalized therapy.

20.15.2 SWNT-Based Multifunctional Nanoplatform Because of their multifunctional capabilities, SWNTs can contain both targeting agents and therapeutic payloads. More importantly, they can be engineered to bypass a variety of biological barriers to enhance the targeting efficacy. The intrinsic physical properties of SWNTs can also offer the opportunity to utilize new approaches to therapy, such as localized heating. The future of nanomedicine lies in multifunctional nanoplatforms, which combine both therapeutic components and multimodality imaging. The ultimate goal is that nanoplatform-based agents can allow for efficient, specific in vivo delivery of therapeutic agents (drugs, genes, etc.) without systemic toxicity, and the dose delivered as well as the therapeutic efficacy can be accurately measured noninvasively over time. The versatile chemistry, as well as the unique physical properties, makes SWNT the most promising candidate for such a multifunctional nanoplatform for both multimodality molecular imaging and combination therapy (“mix-and-match” with suitably selected components for each individual application; Fig. 20.13). The most feasible applications of such nanoplatform-based agents will be in cardiovascular medicine, where there is much less biological barrier for the efficient delivery of nanoparticles, and in oncology, where the leaky tumor vasculature can allow for better tissue penetration than in normal organs/tissues.

20.15.3 Challenges Facing SWNT-Based Agents For in vitro and ex vivo applications, the advantages of state-of-the-art nanodevices (e.g., nanochips and nanosensors, including those based on SWNTs) over traditional assay methods are obvious [312, 313]. However, several barriers exist for in vivo applications of SWNT-based agents, among which are the biocompatibility, acute and chronic toxicity, pharmacokinetics and targeting efficacy, ability to escape the RES, and cost effectiveness. Recently, nanotoxicology has emerged as a new branch of toxicology for studying the undesirable effects of nanoparticles [314, 315]. Development of SWNT-based nanoplatforms for biomedical applications must proceed in tandem with the assessment of any toxicological side effects. The literature reports on SWNT-based toxicity studies clearly indicate that crude, nonfunctionalized SWNTs are toxic while pure, stably functionalized SWNTs are not. Thus, quality control of the SWNTs and robust chemistry on the SWNT surface are the two most important prerequisites for future biomedical and clinical applications of SWNTs. Further studies on well-characterized SWNTs are necessary to determine their safety profile as well as the environmental impact.

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FIGURE 20.13 A multifunctional SWNT-based nanoplatform incorporating multiple receptor targeting, multimodality imaging, and multiple therapeutic entities. Not all functional moieties will be necessary and only suitably selected components are needed for each individual application.

There are many commercial and regulatory challenges to be tackled with the emerging generation of more complex nanoparticles, in part owing to their multicomponent nature. SWNT-based multifunctional nanoplatforms will be difficult and expensive to manufacture on a large scale with appropriate quality, once they reach the stage of clinical investigation. However, some highly complex nanoparticles have reached the clinic for Phase I trials, indicating that complex nanoparticles can be manufactured with current good manufacturing practice (cGMP) and satisfy regulatory requirements, which is very encouraging [151]. On a nonscientific note, intellectual property-related issues may also make the eventually approved SWNT-based products quite expensive. The NCI Alliance for Nanotechnology in Cancer (http://nano.cancer.gov), a comprehensive, systematized initiative encompassing the public and private sectors, was formed in 2005 to accelerate the application of the best capabilities of nanotechnology to cancer. After establishing an interdisciplinary nanotechnology workforce, it is expected that nanotechnology will mature into a clinically useful field in the near future. Although this alliance is focused on cancer, the forefront of the battlefield against human diseases, the same techniques and agents can also be applied to a variety of other pathological disorders. After overcoming the abovementioned challenges, SWNTs have the potential to profoundly impact disease diagnosis and patient management in the future.

ACKNOWLEDGMENTS Funding was provided, in part, by the UW School of Medicine and Public Health’s Medical Education and Research Committee through the Wisconsin Partnership Program, the UW Carbone Cancer Center, and Susan G. Komen for the Cure. Ting Gao acknowledges support from Tyco Electronics.

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273. Cui, D.; Tian, F.; Ozkan, C. S.; Wang, M.; Gao, H. Effect of single wall carbon nanotubes on human HEK293 cells. Toxicol. Lett. 2005, 155, 73–85. 274. Jacobsen, N. R.; Pojana, G.; White, P.; Moller, P.; Cohn, C. A.; Korsholm, K. S.; Vogel, U.; Marcomini, A.; Loft, S.; Wallin, H. Genotoxicity, cytotoxicity, and reactive oxygen species induced by single-walled carbon nanotubes and C(60) fullerenes in the FE1-Mutatrade markMouse lung epithelial cells. Environ. Mol. Mutagen. 2008, 49, 476–487. 275. Raja, P. M.; Connolley, J.; Ganesan, G. P.; Ci, L.; Ajayan, P. M.; Nalamasu, O.; Thompson, D. M. Impact of carbon nanotube exposure, dosage and aggregation on smooth muscle cells. Toxicol. Lett. 2007, 169, 51–63. 276. Casey, A.; Herzog, E.; Lyng, F. M.; Byrne, H. J.; Chambers, G.; Davoren, M. Single walled carbon nanotubes induce indirect cytotoxicity by medium depletion in A549 lung cells. Toxicol. Lett. 2008, 179, 78–84. 277. Sharma, C. S.; Sarkar, S.; Periyakaruppan, A.; Barr, J.; Wise, K.; Thomas, R.; Wilson, B. L.; Ramesh, G. T. Single-walled carbon nanotubes induces oxidative stress in rat lung epithelial cells. J. Nanosci. Nanotechnol. 2007, 7, 2466–2472. 278. Sarkar, S.; Sharma, C.; Yog, R.; Periakaruppan, A.; Jejelowo, O.; Thomas, R.; Barrera, E. V.; Rice-Ficht, A. C.; Wilson, B. L.; Ramesh, G. T. Analysis of stress responsive genes induced by single-walled carbon nanotubes in BJ Foreskin cells. J. Nanosci. Nanotechnol. 2007, 7, 584–592. 279. Kisin, E. R.; Murray, A. R.; Keane, M. J.; Shi, X. C.; Schwegler-Berry, D.; Gorelik, O.; Arepalli, S.; Castranova, V.; Wallace, W. E.; Kagan, V. E.; Shvedova, A. A. Single-walled carbon nanotubes: geno- and cytotoxic effects in lung fibroblast V79 cells. J. Toxicol. Environ. Health A. 2007, 70, 2071–2079. 280. Pulskamp, K.; Diabate, S.; Krug, H. F. Carbon nanotubes show no sign of acute toxicity but induce intracellular reactive oxygen species in dependence on contaminants. Toxicol. Lett. 2007, 168, 58–74. 281. Rotoli, B. M.; Bussolati, O.; Bianchi, M. G.; Barilli, A.; Balasubramanian, C.; Bellucci, S.; Bergamaschi, E. Non-functionalized multi-walled carbon nanotubes alter the paracellular permeability of human airway epithelial cells. Toxicol. Lett. 2008, 178, 95–102. 282. Kagan, V. E.; Tyurina, Y. Y.; Tyurin, V. A.; Konduru, N. V.; Potapovich, A. I.; Osipov, A. N.; Kisin, E. R.; Schwegler-Berry, D.; Mercer, R.; Castranova, V.; Shvedova, A. A. Direct and indirect effects of single walled carbon nanotubes on RAW 264.7 macrophages: role of iron. Toxicol. Lett. 2006, 165, 88–100. 283. Kaiser, J. P.; Wick, P.; Manser, P.; Spohn, P.; Bruinink, A. Single walled carbon nanotubes (SWCNT) affect cell physiology and cell architecture. J. Mater. Sci. Mater. Med. 2008, 19, 1523–1527. 284. Sayes, C. M.; Liang, F.; Hudson, J. L.; Mendez, J.; Guo, W.; Beach, J. M.; Moore, V. C.; Doyle, C. D.; West, J. L.; Billups, W. E.; Ausman, K. D.; Colvin, V. L. Functionalization density dependence of single-walled carbon nanotubes cytotoxicity in vitro. Toxicol. Lett. 2006, 161, 135–142. 285. Kang, S.; Herzberg, M.; Rodrigues, D. F.; Elimelech, M. Antibacterial effects of carbon nanotubes: size does matter! Langmuir. 2008, 24, 6409–6413. 286. Kang, S.; Pinault, M.; Pfefferle, L. D.; Elimelech, M. Single-walled carbon nanotubes exhibit strong antimicrobial activity. Langmuir 2007, 23, 8670–8673. 287. Dutta, D.; Sundaram, S. K.; Teeguarden, J. G.; Riley, B. J.; Fifield, L. S.; Jacobs, J. M.; Addleman, S. R.; Kaysen, G. A.; Moudgil, B. M.; Weber, T. J. Adsorbed proteins influence the biological activity and molecular targeting of nanomaterials. Toxicol. Sci. 2007, 100, 303–315. 288. Nimmagadda, A.; Thurston, K.; Nollert, M. U.; McFetridge, P. S. Chemical modification of SWNT alters in vitro cell-SWNT interactions. J. Biomed. Mater. Res. A 2006, 76, 614–625.

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303. Hamad, I.; Christy Hunter, A.; Rutt, K. J.; Liu, Z.; Dai, H.; Moein, Moghimi S. Complement activation by PEGylated single-walled carbon nanotubes is independent of C1q and alternative pathway turnover. Mol. Immunol. 2008, 45, 3797–3803. 304. Templeton, R. C.; Ferguson, P. L.; Washburn, K. M.; Scrivens, W. A.; Chandler, G. T. Lifecycle effects of single-walled carbon nanotubes (SWNTs) on an estuarine meiobenthic copepod. Environ. Sci. Technol. 2006, 40, 7387–7393. 305. Cheng, J.; Flahaut, E.; Cheng, S. H. Effect of carbon nanotubes on developing zebrafish (Danio rerio) embryos. Environ. Toxicol. Chem. 2007, 26, 708–716. 306. Smith, C. J.; Shaw, B. J.; Handy, R. D. Toxicity of single walled carbon nanotubes to rainbow trout, (Oncorhynchus mykiss): respiratory toxicity, organ pathologies, and other physiological effects. Aquat. Toxicol. 2007, 82, 94–109. 307. Ghafari, P.; St.-Denis, C. H.; Power, M. E.; Jin, X.; Tsou, V.; Mandal, H. S.; Bols, N. C.; Tang, X. S. Impact of carbon nanotubes on the ingestion and digestion of bacteria by ciliated protozoa. Nat. Nanotechnol. 2008, 3, 347–351. 308. Blaise, C.; Gagne, F.; Ferard, J. F.; Eullaffroy, P. Ecotoxicity of selected nano-materials to aquatic organisms. Environ. Toxicol. 2008, 23, 591–598. 309. Canas, J. E.; Long, M.; Nations, S.; Vadan, R.; Dai, L.; Luo, M.; Ambikapathi, R.; Lee, E. H.; Olszyk, D. Effects of functionalized and non-functionalized single-walled carbon nanotubes on root elongation of select crop species. Environ. Toxicol. Chem. 2008, 27, 1922–1931. 310. Phelps, M. E. PET: the merging of biology and imaging into molecular imaging. J. Nucl. Med. 2000, 41, 661–681. 311. Mammen, M.; Chio, S.; Whitesides, G. M. Polyvalent interactions in biological systems: implications for design and use of multivalent ligands and inhibitors. Angew. Chem. Int. Ed. Engl. 1998, 37, 2755–2794. 312. Grodzinski, P.; Silver, M.; Molnar, L. K. Nanotechnology for cancer diagnostics: promises and challenges. Expert. Rev. Mol. Diagn. 2006, 6, 307–318. 313. Sahoo, S. K.; Parveen, S.; Panda, J. J. The present and future of nanotechnology in human health care. Nanomedicine 2007, 3, 20–31. 314. Oberdorster, G.; Oberdorster, E.; Oberdorster, J. Nanotoxicology: an emerging discipline evolving from studies of ultrafine particles. Environ. Health Perspect. 2005, 113, 823–839. 315. Kagan, V. E.; Bayir, H.; Shvedova, A. A. Nanomedicine and nanotoxicology: two sides of the same coin. Nanomedicine 2005, 1, 313–316.

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CHAPTER 21

Multifunctional Nanoparticles for Multimodal Molecular Imaging YANGLONG HOU and RUI HAO Department of Advanced Materials and Nanotechnology, College of Engineering, Peking University, Beijing, China

21.1 INTRODUCTION Inorganic nanoparticles (NPs) are coming forth as potential probes in biomedical imaging and therapy [1, 2]; their comparable size to biofunctional molecules has allowed for ultrasensitive detection of biomolecular targets [3]. Of the various inorganic NPs, magnetic NPs show a unique magnetic resonance (MR) contrast enhancement effect, enabling noninvasive MR imaging of cancer, cardiovascular disease, and cell trafficking [4]. On the other hand, quantum dots (QDs), noble metal, and fluorescent dye hybrid NPs are representative examples of optical nanoprobes [5]. However, to retrieve accurate biological information on a subcellular level is difficult by individually using any of the imaging modalities described above due to their low sensitivity and/or low resolution. Therefore the technology to obtain the detailed information of biological targets becomes an increasingly important need for understanding basic biological phenomena and for early stage diagnosis of various diseases. Current single-model imaging tends to be inadequate. Note that multifunctional NPs can be engineered as platforms for effective multimodal imaging, overcoming the shortcomings that are present for single-model imaging methods. Multifunctional nanoparticle-based probes have great potential for early detection, accurate diagnosis, and personalized treatment of diseases. For example, it has been demonstrated that the simultaneous use of positron emission tomography (PET) and computed tomography (CT) allowed for early detection of cancer in dual modalities [6]. Except for multimodal imaging, the function of therapy can also be integrated into one particle [7]. The multifunctional NPs are emerging as versatile probes for molecular imaging and therapy. So far, as shown in Figure 21.1, there are several types of multifunctional NPs for multimodal probes: (a) functionalized nanoparticle, (b) core–shell nanoparticle, (c) dimer, (d) complex core–shell nanostructure, (e) core–satellite nanostructure, and (f) hybrid nanostructure. In this chapter, we review the current state-of-the-art multifunctional nanoplatforms for multimodal molecular imaging. Single-model molecular imaging is not emphasized here. Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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(a)

(b)

(c)

(d)

(e)

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FIGURE 21.1 Several types of nanostructures for multimodal imaging: (a) functionalized nanoparticle, (b) core–shell nanoparticle, (c) dimer, (d) complex core–shell nanostructure, (e) core–satellite nanostructure, and (f) hybrid nanostructure.

21.2 NANOPARTICLE-BASED PROBES 21.2.1 Magnetic NPs Magnetic NPs exhibit a variety of unique magnetic behavior that is drastically different from that of their bulk counterparts due to the nanoscale effect. Magnetic NPs exhibit a significant potential in biomedical applications such as molecular imaging, drug delivery, thermal therapy, biosensing, and bioseparation. In magnetic resonance imaging (MRI), the larger magnetic moment of NPs could create a larger magnetic field heterogeneity, which affects the surrounding water in living subjects. It is used to produce predominantly T2 relaxation effects, resulting in a negative signal on a T2-weighted image. Superparamagnetic iron oxide (SPIO) NPs were employed as an MRI contrast agent after surface modification of the NPs. Coating materials here are usually biocompatible polymers, such as dextran and polyethylene glycol (PEG). At the present time, there are several commercial SPIO NPs, for example, AMI-25 (Feridex, Endorem) [8], SHU 555A (Resovist) [9], and AMI-227 (Combidex, Sinerem) [10]. Some other magnetic NPs (e.g., FeCo [11], MnFe2 O4 [12]), which offering high magnetic moments and r1/r2 relaxivities, have demonstrated longlasting positive-contrast enhancement for vascular MRI. After conjugation with targeting molecules, functionalized NPs can also be used as probes for targeting specific species on the surface of cells not only as an MRI contrast agent. Montet et al. [13] prepared cross linked iron oxide (CLIO) NPs, which were conjugated with Cy5.5 (an infrared optical fluorescence dye) and bound to cyclic arginine–glycine–aspartic acid (RGD); the resulting NPs were used to target ␣v ␤3 integrins on BT-20 tumor. In this study, the BT-20 tumor cell was successfully observed by fluorescence-based imaging and MRI. A similar route was explored by employing Cy5.5 for fluorescence imaging and CLIO as an MRI contrast

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agent [14]. The difference between the two examples is that the latter chose peptides (EPPT) which specifically recognize uMUC-1 as a targeting group. This multimodal probe could not only detect orthotopically implanted preclinical models of adenocarcinomas but could also track in situ tumor response to chemotherapy in vivo. By using Cy5.5 as the fluorescence probe, thermally crosslinked superparamagnetic iron oxide (TCL-SPIO) NPs were also developed for dual-modality in vivo cancer imaging of Lewis lung carcinoma tumor allograft in mice. Although the NPs here weren’t conjugated with a targeting group, the tumor was unambiguously detected after intravenous injection in T2-weighted MR images as a 68% signal drop as well as in optical fluorescence images. Cancer immunotherapy approach has been developed to assist the body’s immune system to selectively recognize and kill malignant tumor cells. A dual-modality nanoparticle system is capable of selectively labeling and imaging CTL cells expressing T-cell receptors (TCRs) that recognize cognate MHC–peptide complexes on the surface of antigen-presenting tumor cells [15], as shown in Figure 21.2. In this work, it has been demonstrated that the ex vivo T-cell selectivity and reporter functionality allowed development of NPs for both MRI and fluorescence imaging. PET technology is an imaging modality that uses the signal emitted by positron-emitting radiotracers to construct images of tracer distribution in vivo [16]. (By combining PET with MRI, more information on biological structures could be obtained. Jarrett et al. [17] reported 64 Cu radiolabeling of dextran sulfate coated SPIO NPs, which were prepared by coordinating the 64 Cu to the chelating bifunctional ligand, p-SCNBz-DOTA, followed by integrating the NPs. The multifunctional NPs are targeted toward inflammatory events such as atherosclerotic plaques. Polyaspartic acid (PASP) coated iron oxide NPs [18] were also synthesized for integrin ␣v ␤3 recognition and positron-emitting radionuclide 64 Cu labeling by conjugating with cyclic arginine–glycine–aspartic (RGD) peptides and the macrocyclic chelating agent DOTA. In this study, multimodal probes were successfully obtained for dual PET/MRI of tumor integrin ␣v ␤3 expression in vivo using a small-animal model. As is well known, the lymphatic system is an important first line in defense against infection, and it is also used for the metastasis of malignant cancer; hence accurate localization of sentinel lymph nodes (SLNs) is critical for cancer staging [19]. A PET/MRI probe was explored as an in vivo dual-modality imaging agent in an SLN model biological subject [20]. Herein, serum albumin (SA) coated MnFe2 O4 was applied as the MRI probe. In this procedure, the carboxylic and amine groups of SA were crosslinked by using N-(3-dimethylaminopropyl)-N-ethylcarbodiimide hydrochloride and N-hydroxysulfosuccinimide, and then 124 I, which is a PET radionuclide, was conjugated to the tyrosine residue in SA. In the in vivo experiment, two different types of LNs could clearly be identified and accurately localized in a PET/MR fusion image. 21.2.2 Quantum Dots QDs are usually II–VI or III–V semiconductor NPs with a dimensional size of about 10 nm. Compared with small molecule fluorophores, QDs show a narrow, tunable, emission spectrum, broader excitation energy, and better environmental stability. QDs are widely used as biomarkers in molecular imaging and biosensing. Fluorescence imaging/MRI probes were prepared by simple modification of quantum dots, followed by the conjugation of molecular ions such as Gd3+ . Programmed cell death (PCD) is a process of organized cell suicide and occurs in various forms; PCD plays an important role in the physiology as well as the pathology of multicellular organisms. One of the most explored molecular imaging

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(a) NH2 NH

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NH2 NH2 NH2

Alexa Fluor 647

Polyethylene Glycol

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Biotin MHC Complex

Iron Oxide Nanoparticle

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ra ut

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FIGURE 21.2 (a) Schematic illustration of synthesis of NP-PEG-MHC-AF647. Iron oxide nanoparticles were coated with functionalized PEG to which neutravidin was covalently bound via a thioether linkage. Biotinylated peptide–MHC was attached to the PEG termini, lending the particle targeting specificity for CTLs. Neutravidin was prelabeled with the fluorophore Alexa Fluor 647. (b) Surface modification of nanoparticles with PEG and MHC–peptide verified by FTIR. (c) Hydrodynamic size and zeta-potential study of nanoparticle constructs at physiologic pH. (Reproduced with permission from Gunn et al. [15].)

targets of PCD is phosphatidylserine. Prinzen et al. [21] prepared AnxA5-functionalized and Gd3+ conjugated QDs for fluorescence imaging and MRI of phosphatidylserine expressed cell. Jin et al. [22] reported the preparation of Gd3+ -functionalized near-infrared (NIR)emitting QDs as in vivo NIR fluorescence imaging and MRI dual functional probe. The as-prepared probe was prepared based on hydrophobic glutathione coated CdSeTe/CdS QDs bound with Gd3+ –DOTA complex. In an in vivo experiment, a phantom containing 5 mL of Gd3+ –DOTA QDs (10 mM) was embedded into a mouse abdomen to test the probe.

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21.2.3 Noble Metal NPs Noble metal NPs such as gold and silver particles have attracted significant attention owing to their potential applications in molecular imaging, in biosensing, and in plasmons. The strong optical absorption and scattering of noble metal nanoparticles are attributed to the localized surface plasmon resonance [23, 24]. For example, the resonant extinction, which can be tuned from the visible to the near-infrared region by controlling the shape and structure, allows the nanoparticles to serve as molecular contrast agents. The localized heating of resonant absorption drives new thermal therapies and drug delivery mechanisms. The sensitivity of these resonances to their environment leads to simple affinity sensors for the detection of low-level molecular analytes. Coupled with their general lack of toxicity, noble metal nanoparticles are a highly promising class of nanomaterials for new biomedical applications. For example, DTDTPA (dithiolated diethylenetriaminepentaacetic acid) on the surface of gold NPs facilitates conjugation of gadolinium complex; the resulting NPs generate a pronounced enhancement in contrast [25]. The multimodality of Au-Gd-DTDTPA suggests the possible use of these systems in living organisms for diagnostic (MRI) and/or therapeutic applications. In the follow-up work [26], gold@gadolinium chelate complex can freely circulate in the blood pool without undesirable accumulation in the lung, liver, and spleen, suggesting that it is a promising multimodal probe for in vivo MRI and X-ray imaging synchronously. Silver nanoparticles were also applied to the areas of microbial resistance, wound healing, surface enhanced Raman scattering, and metal enhanced fluorescence [27]; however, the multimodal probe based on silver NPs has rarely been reported.

21.2.4 Silica and Other NPs Owing to their good chemical stability and biocompatibility, silica NPs have been widely used in biomedicine. The chemical preparation of silica NPs is easy, low-cost, and convenient; moreover, silica NPs can easily be modified by a silylation agent and encapsulated functional molecular materials. Recently, multifunctional magnetic SiO2 @Ag sandwichlike nanostructures were reported; they are is composed of a magnetic silica core and a continuous Ag shell. The magnetic Ag shells integrated the broad NIR absorption property and superparamagnetic property into one particle, providing promising multimodal probes and magnetic-field targeted photothermal therapy agents [28], as shown in Figure 21.3. MCM-41 type materials, which possess a hexagonal array of one-dimensional channels with diameters of 2–10 nm [29], have been introduced into biomedicine in recent years. For example, Gd-DTPA (DTPA=diethylenetriamine pentaacetic acid) and rhodamine B grafted MCM-4 nanospheres were prepared [30]. In this study, the resulting NPs are a highly efficient T1 contrast agent for intravascular MR imaging and an excellent T2 contrast agent for MR imaging of soft tissues when the NPs were applied at a higher dosage. Polymer or other kinds of organic nanomaterials were also explored as multimodal probes. The complex NPs constructed from Gd and rhodamine B [31], which were linked lipid monomers with diacetylene bonds, were labeled to tumor cells. In this study, labeled cells were inoculated subcutaneouosly into the flanks of C3H mice, and used for imaging the labeled tumor cells via MRI and optical imaging. Cartier et al. [32] reported the synthesis of polyglycidylmethacrylate (EPMA) NPs for multimodality in vivo imaging and detection of the cationic tracers. 111 In, 68 Ga, and Gd3+ were directly bound without the necessity of a chelating conjugate; 68 Ga can be used for PET, while 111 In can be used for quantitative

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FIGURE 21.3 Schematic Illustration of the fabrication procedure of magnetic Ag nanoshells. (Reproduced with permission from Chen et al. [28].)

gamma scintigraphy and Gd3+ for MR imaging. Rhodamine B was also incorporated to examine specific interactions with blood cells. Metal organic framework (MOF) is a new class of hybrid material built from metal ions with well-defined coordination geometry and organic bridging ligands. MOF nanorods containing Gd3+ , Eu3+ , or Td3+ were applied to improve the water signals in T1-weighted images efficiently [33, 34]; the results showed high luminescence upon UV excitation with characteristic red and green luminescence from Eu3+ and Tb3+ .

21.3 COMPLEX NANOSTRUCTURE-BASED PROBES 21.3.1 Core–Shell NPs Core–shell NPs, which enrich different components and chemical structures, showed layered spherical structures. Different layers of core–shell NPs can be designed for various needs, possessing multifunctional properties in one unit. Hence it is easy to manufacture multimodal probes by using core–shell nanostructures. Of the core–shell nanostructures, silica-based core–shell NPs have been widely used because of the flexible synthetic route and surface modification. For time-resolved detection of fluorescence cell imaging, Wu et al. [35] prepared multifunctional iron oxide@silica NPs through the copolymerization of Eu3+ complex, 4,4 -bis(1 ,1 ,1 -trifluoro-2 ,4 -butanedion-4 -yl) chlrorosulfoO-terphenyl-Eu3+ (APS-BTBCT-Eu3+ ), free (3-aminopropyl) triethoxysilane (APS), and tetraethyl orthosilicate (TEOS) in the presence of poly(vinylpyrrolidone) (PVP) stabilized magnetic Fe3 O4 NPs. By using as analogous route, fluorescein-isothiocyanate (FITC) functionalization of iron oxide@silica NPs was achieved to improve the specificity of SPIO for targeting glioma cells [36]. Hsiao et al. [37] synthesized small spherical bifunctional Gd-Dye@mesoporous silica NPs (MSN), which were conjugated with FITC and diethylenetriaminepentaacetic acid (DTPA). The functionalized NPs are paramagnetic and emit green fluorescence. MSN was evaluated for its capability as an effective bifunctional fluorescent and T1-enhanced tracker in human mesenchymal stem cells. QDs are good fluorescent agents. The major drawbacks for clinical translation of QDs are inefficient delivery, potential toxicity, and lack of quantification [1]. However, with further improvement of the conjugation strategy, it is expected that QDs may achieve optimal tumor targeting efficacy with acceptable toxicity profile for further application.

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FIGURE 21.4 Iron oxide@Au nanoparicle and its blue and red shift achieved by adjusting the coating layer. (Reproduced with permission from Xu et al. [43].)

Yang et al. [38] developed QDs with silica coating for multimodal imaging. Paramagnetic Gd(III) functionalization via a metal chelating silane coupling agent (TSPETE) to the yellow fluorescent, silica-coated CdS:Mn@ZnS core–shell quantum dots resulted in multimodal NPs that can be imaged optically and by MRI. After silica-coated NPs were covalently linked to Gd3+ ion chelator [39], tetra-azacyclo-dodecane-tetra-acetic acid (DOTA), QDs@silica and Au@silica nanoparticles for dual-modality imaging were obtained. Similar core–shell NPs exhibited high photoluminescence quantum yield, good positive MRI contrast, low cytotoxicity, and intracellular delivery in viable cells [40]. Mn2+ can be used as a T1-weighted MRI probe, while CdSe quantum dots are extremely high luminescence material; however, it is difficult to directly dope Mn2+ into CdSe. Wang et al. [41] maintained optical/MRI dual-modality core–shell structural probes with luminescent CdSe nanoparticles as the core and a paramagnetic Mn-doped ZnS surface. If Gd2 O3 nanoparticles are used for T1-weighted imaging, a dual functional probe can be prepared by conjugating with a fluorescent group. Gadolinium oxide@polysiloxane NPs were functionalized by the organic dye PEG for in vivo dual-modality magnetic resonance and fluorescence imaging [42]. Gold nanostructures, which possess good biocompatibility and surface plasmon resonance, can be used as optical probes. By combining iron oxide with gold, magic iron oxide@Au nanostructures [43] exhibit great potential for biomedical applications due to their integration of magnetic and optical properties (Fig. 21.4). Iron oxide@Au NPs were applied to molecular specific bimodal contrast agents for both magnetic and optical imaging [44]. Kim et al. [45] designed a silica@Au core–shell nanostructure that embeded iron oxide nanoparticle in the gold layer. Anti-HER2/neu was linked to the core–shell structure for targeted MRI and NIR photothermal therapy of cancer cells. A novel multifunctional magnetic particle@silica@Au core–shell structure was reported for targeted cancer detection by MRI as well as for a synchronous cancer therapy via therapeutic antibody and hyperthermia effects [46]. In this study, MnFe2 O4 NP clusters were embedded in silica and further coated with a gold nanolayer. 21.3.2 Heterostructures Similar to core–shell structures, heterostructural NPs are integral domains of different functional components and show multiple properties within one particle. Moreover, heterostructures introduce anisotropy into NPs. For example, they offer more functional surfaces than core–shell NPs, so it is easy to attach different kinds of molecules for specific purposes. Primary research is focused on dimer structures in biomedicine. Choi et al. [47] prepared FePt-Au dimer NPs. In this procedure, FePt NPs were first prepared and Au NPs were added on to the FePt NPs by heteroepitaxial growth. By Au-S linkage modification, the

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dimer structures showed high water solubility and biocompatibility. While Au was used for chip-based biosensing, the FePt part was used for MRI because of its superparamagnetism. Sun and co-workers demonstrated that the biocompatible dimer Au–Fe3 O4 NPs can be suitable for linking different functional molecules to each end of the structure through proper surface functionalization [48]. Similar dimer NPs composed of various combinations of metal (Au, Ag, Pt, or Ni) and oxide (Fe3 O4 or MnO) can readily be synthesized from thermal decomposition of mixtures of metal–oleate and metal–oleylamine complexes [49]. Au-MnO and Au-Fe3 O4 dimer NPs have been tested as multimodal imaging probes. Another kind of dimer structure, Fe3 O4 @SiO2 NPs, were prepared first [50]; then Fe3 O4 @SiO2 NPs attached with FITC were produced by using the as-prepared core–shell structures and FITC as precursor. These NPs are fluorescent, magnetic, and porous, exhibiting potential applications in bimodal imaging probes and drug reservoirs.

21.4 NANOHYBRID-BASED PROBES Nanohybrids are a kind of composite material prepared by simple linkage of chemical bonds and/or molecular interactions. Recently, Cheon and co-workers fabricated new “core–satellite” structured dual functional NPs by employing dye doped silica as core and multiple magnetic NPs as satellites [51]. The integrated structures can be used for simultaneous optical and MR imaging of neuroblastoma cells expressing polysialic acids (PSAs) (Fig. 21.5). Yang et al. [52] synthesized antibody conjugated anticancer drug doxorubicin (DOX)– magnetic PLGA NPs (HER-DMPNP) for the detection and treatment of cancer. The welltailored DMPNP was prepared using a surfactant through a nanoemulsion method, in which magnetic NPs were embedded in polymeric NPs through hydrophobic interaction. DOX acts as both a drug and fluorescent group here; the NPs showed excellent sensitivity as MR probes for detection of cancer cells. Cetuximab conjugated fluorescent magnetic nanohybrids were also developed to detect cancer using MR and optical imaging, as shown in Figure 21.6 [53]. Another fluorescent magnetic nanohybrid (FMNH), which was stabilized by pyrenelabeled PCL-b-PMAA, was explored for simultaneous cancer-targeted magnetic resonance (MR) or optical imaging and magnetically-guided drug delivery [54]. This platform was composed of inorganic nanocrystals of Fe3 O4 or CdSe/ZnS, in addition to DOX embedded in biodegradable PLGA polymer NPs. Pegylated folate was coated on the surface of the polymer NPs for active targeting of cancer cells. Luminescent [Ru(bpy)3 ]Cl2 (bpy = 2,2 -bypyridine) was used to build the multifunctional probes. Recently, Ru(bpy):GdIII @SiO2 complex NPs have been demonstrated to be photostable, radio-opaque, and paramagnetic, emerging as multimodal imaging probes, visualized using CT, MRI, and diffuse optical tomography. The surface reactive groups of these probes can be modified to contain both ligands and antibodies, allowing for the detection of cellular events in vivo [55]. Robust luminescent and paramagnetic hybrid silica NPs showed high payloads of magnetic centers for in vitro optical and T1- and T2-weighted MR imaging; in similar work [56], hybrid silica NPs containing [Ru(bpy)3 ]Cl2 were employed for the fluorescent part, and a monolayer coating of silylated Gd complexes were used for MRI. By using layer-by-layer (LbL) self-assembly [57], the developed NPs were further deposited with cationic Gd(III)-DOTA or FITC oligomer and polystyrenesulfonate as an anion layer via electrostatic interactions. By using in vitro MR imaging, it was demonstrated that LbL particles efficiently targeted cancer cells with attached K7 RGD peptide

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FIGURE 21.5 Core–satellite heterostructures: (a) schematic diagram for the synthesis of core–satellite DySiO2 –(Fe3 O4 )n NPs.(b–d) TEM images of (b) rhodamine-doped silica (DySiO2 ), (c) Fe3 O4 , and (d) core–satellite DySiO2 –(Fe3 O4 )n NPs. (Reproduced with permission from Lee et al. [51].)

FIGURE 21.6 Schematic illustration for preparation of fluorescence magnetic nanohybrids as multimodal imaging agents for cancer detection. (Reproduced with permission from Yang et al. [53].)

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noncovalently. This strategy not only provides NPs with transnormally high MR relaxivities but also provides an efficient route for noncovalent functionalization with affinity molecules. 21.5 CONCLUSION AND PERSPECTIVES Molecular imaging is an interdisciplinary research involving chemistry, physics, biology, engineering, medicine, nanoscience, and nanotechnology. Progress on the next generation of molecular probes will result from fusing multiple fluorescent dyes, drug carrier, and multiple magnetic NPs into a single nanoprobe that provides superior fluorescence, MR imaging, and drug delivery capabilities through the synergistic enhancement of its respective components. The development of technology in multimodal imaging is providing an alternative route to retrieve biological information in exquisite detail. Each of the most frequently used techniques, such as MRI, optical imaging, PET, and CT, have inherent advantages and disadvantages. Multifunctional NPs with simultaneous optical and MR imaging and carrier capability are being developed for high sensitivity and resolution of optical microscopy with the ability for MRI and therapy. These nanoplatforms enable the same biological structures of a specimen to be studied at dramatically different resolutions. The synthesis and surface chemistry of multifunctional nanoparticles are a rapidly developing area of research that focuses more attention on the in vivo requirements of biocompatibility and pharmacokinetics. The multimodal probes based on multifunctional NPs are expected to give accurate imaging and provide a research model for biological and medical application.

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Multifunctional Nanoparticles for Cancer Theragnosis SEULKI LEE, ICK CHAN KWON, and KWANGMEYUNG KIM Biomedical Research Center, Korea Institute of Science and Technology, Seoul, South Korea

Cancer theragnosis is a new concept in next-generation cancer treatment that combines early-stage diagnosis and efficient therapy. Recently, nanotechnology and molecular imaging have been combined to generate multifunctional nanoparticles that simultaneously facilitate cancer diagnosis and therapy (cancer theragnosis). These multifunctional nanoparticles show great promise in the emerging field of personalized medicine, because they allow detection as well as monitoring of an individual patient’s cancer at an early stage and delivery of anticancer agents over an extended period for enhanced therapeutic efficacy. Moreover, noninvasive and real-time monitoring of the multifunctional nanoparticles could enable clinical management of cancer for early diagnostic and response to drug treatment assessment. The application of cancer theragnosis is still in its infancy, but cancer theragnosis with multifunctional nanoparticles presents a promising new strategy in cancer treatment, permitting simultaneous early-stage diagnosis, drug delivery, and real-time monitoring for therapeutic efficacy. The hypothesis offered in this chapter is that cancer theragnosis, if properly integrated with established cancer research, provides extraordinary opportunities to meet these challenges.

22.1 INTRODUCTION Although many researchers have paid much attention to cancer treatment, no substantial progress has been made over the past 50 years. This is because current cancer treatments the lack selectivity of anticancer agents in cancer cells. The anticancer agents used in cancer treatment are randomly distributed in the body because they have no tumor selectivity, resulting in a relatively low therapeutic index. Therefore we need specific delivery systems designed to safely shuttle anticancer agents to the target cancer cells. For this purpose, various nanoparticles of polymers, metals, and semiconductors have been used as new drug delivery systems or imaging agents that have novel and unique properties, Nanoplatform-Based Molecular Imaging Edited by Xiaoyuan Chen C 2011 John Wiley & Sons, Inc. Copyright


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such as enhanced tumor targeting ability, prolonged in vivo biodistribution, and biocompatibility [1–3]. Various nanoparticles are known to accumulate at the tumor site by the unique vascular structural changes associated with the pathology of cancer, providing opportunities for the use of nanoscale delivery systems [4–7]. The blood vessels in tumors have several abnormalities, including a relatively high proportion of proliferating endothelial cells, increased tortuosity, deficiency in pericytes, and aberrant basement membrane formation. The defective vascular structure, created due to the rapid vascularization for serving fast-growing cancers, renders them permeable to macromolecules or nanoparticles. This passive targeting phenomenon discovered by Maeda and colleagues is called the “enhanced permeation and retention effect (EPR effect)” [8, 9]. It is now well established that macromolecules and nanosized particulates (e.g., polymer conjugates, polymeric micelles, dendrimers, and liposomes) accumulate passively in solid tumor tissues by the EPR effect, which increases the therapeutic index but decreases side effects. However, in vivo studies have shown that the tumor specificity of nanoparticles is not as high as anticipated. Nanoparticles introduced into the circulating blood are quickly removed from the body by the immune system. Indeed, a number of nanoparticles with different characteristics (e.g., different surface chemistry, size, surface charge, and molecular weight) have proved unsatisfactory in in vivo tests with regard to stability, biodistribution, and tumor targeting specificity [10, 11]. One promising approach that enhances the cancerspecific target efficacy is to develop nanoparticle constructs containing a targeting moiety, such as antibody, aptamer, and small molecule ligand, for optimized tumor homing with reduced nonspecific accumulation [12–14]. These targeting moieties showed selective delivery of nanoparticles into tumors or selective detection in tumors. Despite their improved selectivity, they still have to overcome an obstacle—a large percentages of nanoparticles continue to accumulate in the liver and spleen. Apparently, nanoparticulate delivery systems are not working well as expected. Furthermore, the final in vivo destination of nanoparticles is largely unknown, because it is difficult to acquire direct and noninvasive in vivo characteristics of nanoparticles. Recently, nanotechnology and molecular imaging have been combined to generate multifunctional nanoparticles that simultaneously facilitate cancer theragnosis [15]. Theragnostic nanoparticles containing therapeutic agents and imaging probes show great promise in the emerging field of cancer theragnosis, because they allow noninvasive imaging of in vivo characteristics of various nanoparticles as well as delivering anticancer agents over an extended period for enhanced therapeutic efficacy (Fig. 22.1) [16]. For these reasons, we believe that cancer theragnosis with multifunctional nanoparticles is a promising new strategy in cancer treatment, permitting simultaneous early-stage diagnosis, drug delivery, and real-time monitoring for therapeutic efficacy. The basic rationale is that nanometer-sized particles, such as polymeric micelles/particles, semiconductor quantum dots, and iron oxide nanocrystals, have theragnostic and multifunctional characteristics that are not available from one molecule or bulk solids. When nanosized particle structure is formed with or without tumor targeting ligands such as monoclonal antibodies, peptides, or small molecules, these nanoparticles can be used as a tumor-targeting drug delivery system as well as tumorspecific imaging agents, simultaneously. Moreover, real-time and noninvasive monitoring of the theragnostic nanoscale particles could enable researchers to rapidly decide whether the multifunctional nanoparticle is effective in cancer treatment. For these reasons, we believe that cancer theragnosis with theragnostic nanoparticles present a promising new strategy in cancer treatment, permitting simultaneous early-stage diagnosis, drug delivery, and real-time monitoring for therapeutic efficacy.

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FIGURE 22.1 Design of multifunctional nanoparticles with specific targeting moieties, therapeutic agents, and imaging probes in cancer treatment.

22.2 CONCEPT OF CANCER THERAGNOSIS 22.2.1 Cancer as Theragnostic Nanoparticle Target One of the current challenges in cancer treatment is tumor-targeting delivery of imaging agents and anticancer agents. In particular, cancer researchers have been actively exploring nanoparticles made of lipid-based micelles, natural/synthetic polymeric particles, and inorganic particles for cancer diagnosis and therapy [2, 17, 18]. This is because nanoparticles are known to accumulate at the tumor site by the so-called enhanced permeation and retention (EPR) effect, resulting in efficient passive targeting in solid tumor tissues (Fig. 22.2). Tumors show several abnormalities of blood vessel structure in comparison with physiological normal blood vessels. These include relatively high angiogenesis with increased proliferation of endothelial cells, increased tortuosity, and aberrant basement membrane formation [19, 20]. In general, the rapid formation of blood vessels in the tumor area leads to a discontinuous endothelium, with gaps several hundred nanometers in size between cells. The transportation of macromolecules from blood vessel to tumor occurs via discontinuous endothelium, vesicular vacuolar organelles, and fenestrations. However, it still remains debatable which pathways are predominantly responsible for tumor hyperpermeability and macromolecular transvascular transport. The optimum size of nanoparticles is not yet precisely known for effective accumulation in a tumor via the EPR effect. Some studies using liposomes and nanoparticles have shown that the cutoff size of the gap in the tumor endothelium is as large as 200 nm to

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FIGURE 22.2 Schematic representation of tumor-specific accumulation via “active targeting,” which is achieved by functionalizing the surface of nanoparticles with ligands that promote cellspecific recognition, and “Passive targeting,” which is achieved by extravasation of nanoparticles through increased permeability of the tumor vasculature.

1.2 mm [21, 22]. In addition, the direct observation of tumor vasculature has demonstrated a tumor-dependent gap size between the endothelium cells from 200 nm to 2 mm [23, 24]. These size ranges appear to be an indication that drug-loaded nanoparticles may accumulate in the tumor area. Consistent with this, administered liposomal formulations with DOX entrapment have been described to exhibit favorable pharmacokinetics due to EPRmediated tumor targeting, as compared with free DOX [25]. Furthermore, DOX-loaded polymeric nanoparticles showed long circulation in the blood for more than 3 days, and gradual accumulation in tumors via the EPR effect [26, 27]. In theory, genes and proteins could be delivered to primary or metastasized tumors via the EPR effect. This suggests that a wide variety of polymeric nanoparticles may be used for tumor targeting of anticancer drugs. However, it should be noted that the vessel permeability that forms a cornerstone of the EPR effect varies during tumor progression. In addition, transvascular transportation of polymeric nanomedicines will depend on the tumor type and anatomical location, as well as the physicochemical properties of the utilized polymer in nanomedicines. Moreover, specific tumor accumulation of various nanoparticles via active targeting is achieved by chemical attachment of targeting moieties that strongly interact with tumor-associated or tumor-specific antigens located on the target cancer cells. Myriad nanoparticles have several advantages such as large surface area (ideal for efficient modification with a wide range of imaging moieties), prolonged plasma circulation time, enhanced stability, improved targeting ability, and reduced nonspecific binding. For specific targeting to tumors, these nanoparticles could be chemically modified with tumor-targeting proteins (mainly antibodies and their fragments), nucleic acids (aptamer), or other receptor ligands such as peptides, vitamins, and carbohydrates [28–31]. The use of tumor-targeting moieties

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allows facilitated cellular uptake of the drug by receptor-mediated endocytosis, which is an active process requiring a significantly lower gradient of internalized substance across the plasma membrane when compared with simple endocytosis, as well as delivering the multifunctional nanoparticles containing therapeutic agents or imaging probes in cancer theragnosis [32]. 22.2.2 Molecular Imaging in Cancer Theragnosis The physicochemical properties can determine nanoparticles’ behavior in a biological system, such as biodistribution, safety, targeting efficacy, and multifunctional activity in drug delivery systems and molecular imaging. General traditional techniques (electron microscopy, dynamic light scattering, energy dispersive X-ray, zeta potential, etc.) can be used to characterize the physicochemical properties of nanoparticles. However, the in vitro physicochemical properties of nanosystems do not always match with in vivo behaviors of nanoparticles, due to the in vivo biological microenvironment. Therefore characterization of nanoparticles must be carried out under in vivo conditions, not only in vitro. Unfortunately, there is only limited information about the interaction between nanosystems and the biological microenvironment. The physicochemical properties of optimum nanoparticles for drug carriers and imaging probes have not been fully characterized. Despite the benefits that nanosystems have contributed to cancer treatment, some applications remain to be improved; for example, specific targeting to the active site, efficient drug delivery inside the target cells or tissues, and early-stage diagnosis. So resonable and well-established approaches for in vivo characterization of nanosystems are required for novel nanocarriers to be used in drug delivery and molecular imaging. Rationally designed nanosystems, based on our understanding of their interactions with the biological environment and molecular mechanisms in vivo, are necessary for achieving efficient drug delivery and molecular imaging. Consequently, it is very important to understand how multifunctional nanoparticles directly or indirectly interact with target cancer cells. It is necessary to learn how they interact with cancer cells, how the particles are targeted, how they detect early-stage cancer, and how they efficiently treat the tumor. All of theses processes can now be visualized via “molecular imaging” techniques at the molecular level of resolution. Molecular imaging has recently emerged and is a rapidly growing field; it is providing new opportunities to directly visualize in vivo characteristics of multifunctional nanoparticles. Currently, in vivo molecular imaging techniques, including magnetic resonance imaging (MRI) [33, 34], positron emission tomography (PET) [35, 36], computed tomography (CT) [37], and near-infrared (NIR) fluorescence imaging [38, 39], are extensively used to characterize the in vivo behaviors of nanosized drug carriers and imaging probes (Table 22.1). Positron emission tomography (PET) and single photon emission computed tomography (SPECT) are two important nuclear imaging techniques for clinical applications that show high sensitivity and sufficient tissue penetration. In the case of small laboratory animals, microPET and microSPECT are also designed for high-resolution imaging at the molecular level [35, 36]. For target-specific imaging, antibodies, peptides, and nucleic acids are generally labeled for nuclear imaging. Radionuclides for SPECT such as 99m Tc, 123 I, and 125 I have relatively long half-lives compared to those for PET, in which the radiotracers are labeled with positron emitting radioactive atoms such as 11 C, 13 N, 15 O, and 18 F. Also, MRI provides both functionality as well as anatomical information. Recently, a novel noninvasive molecular magnetic resonance imaging (mMRI) technology was developed for

546 Pseudoquantification

Yes Yes Yes

<10 cm

1.2 mm

SPECT

Fluorescence 1 mm imaging

No limit cm No limit

50 ␮m 50 ␮m 1.2 mm

CT Ultrasound PET

Yes

Yes

No limit

10–100 ␮m

MRI

Quantification

No limit

Depth

Resolution

mTc- or 11 In-labeled compounds

Near-infrared fluorochromes, QDs

99

Paramagnetic chelates, magnetic particles Iodinated molecules, Gold Microbubbles 18 F-, 64 Cu-, or 11 Clabeled compounds

Imaging agents

Tumor, liver Angiography Tumor, pharmacokinetics, pharmacodynamics, biodistribution Tumor pharmacokinetics, pharmacodynamics, biodistribution Tumor, biodistribution, molecular events, physiological events

Tumor, liver

Imaging Targets

Photodynamic therapy

No

Tumor-targeting drug delivery system Hyperthermia therapy US responsive drug delivery No

Therapy

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TABLE 22.1 Characteristics of Currently Available Imaging Techniques

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visualization of the biological process at the cellular and the molecular levels based on advances in MRI physics and target-specific contrast imaging agents [33, 34]. Although it has limited tissue penetration (a few centimeters), optical imaging is a very powerful and noninvasive technique that provides excellent spatial and temporal resolution. The light in the visible wavelength range is used for conventional and intravital microscopic observation. The near-infrared (NIR) region around 650–900 nm has been used to image deep tissue, since there is lower light absorption from hemoglobin, water, and lipids [37, 40]. In general, various nanoparticles have been directly labeled with a wide range of imaging probes according to the imaging modality via conventional chemistry. For example, fluorescent probes for optical imaging, paramagnetic agents for MR imaging, and radiolabeled probes for PET imaging were labeled to nanoparticles. Novel nanosystems with imaging probes can be used in real-time monitoring and visualization in a noninvasive way, allowing for clinical uses in animals and humans. With the help of various imaging systems, the most important characteristics of nanosystems in vivo were studied, including in vivo ADME (absorption, distribution, metabolism, and excretion), targeting efficacy, and their multifunctional properties as drug carriers and imaging probes. Also, in vivo experimental uncertainties arising from interanimal variations are greatly reduced, because each animal serves as its own “control” for consecutive analyses at the same condition. The different imaging technologies revealed the information on circulating or localized nanoparticles, which has close relevance with the in vivo fate of different nanoparticles in small animals or humans, in that drug carriers and imaging probes are passively targeted or nonspecifically localized.

22.3 MULTIFUNCTIONAL NANOPARTICLES FOR CANCER THERAGNOSIS 22.3.1 Polymeric Multifunctional Nanoparticles for Cancer Theragnosis Polymeric nanoparticles have been applied as effective theragnostic nanoparticles capable of enhancing the diagnostic and therapeutic efficacy of imaging probes and therapeutic agents in cancer treatment. Modern polymer chemistry has produced an increasing number of polymeric nanoparticles for cancer theragnosis, including dendrimers, natural/synthetic amphiphilic polymers, and grafted or branched polymers (Fig. 22.3). When they form a self-assembled nanoparticle, they could exhibit unique physicochemical characteristics in the aqueous condition, possibly having unusual rheological features, a small hydrodynamic radius, or thermodynamic stability [41]. In particular, polymeric nanoparticles exhibited ideal in vivo characteristics, such as biocompatibility and biodegradability, and prolonged circulation time in the bloodstream. These unique in vivo characteristics of polymeric nanoparticles have been extensively exploited in cancer treatment. This is because biocompatible and prolonged circulating polymeric nanoparticles in vivo allow them to accumulate and extravasate into the tumor tissue, where a disorganized vasculature and defective vascular architecture is usual. Recently, a fusion technology of polymer chemistry and molecular imaging approaches has led to the generation of cancer theragnosis based on novel polymeric nanoparticles. Most imaging probes currently used in clinical applications are physically/chemically introduced into the polymeric nanoparticles. For example, PET/SPECT radionuclides (11 C, 18 F, and 64 Cu), MRI imaging probe (such as Gd3+ ), and optical imaging probes (fluorochromes,

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FIGURE 22.3 Various polymeric nanoparticles for cancer theragnosis. Physical loading or chemical conjugation of therapeutic agents and imaging probes on polymeric nanoparticles were used for cancer theragnosis.

quantum dot, and photoproteins) were simply introduced into the polymeric nanoparticles by physical complex or chemical conjugation. Polymeric nanoparticles containing imaging probes could be noninvasively visualized in animals and humans and the in vivo characteristics of polymeric nanoparticles—biodistribution, tumor-targeting efficacy, and therapeutic efficacy—are monitored by various imaging modalities. In addition, different imaging probe-labeled polymeric particles have presented distinguishing properties, such as prolonged plasma half-lives, enhanced stability, reduced toxicity, and improved target specificity, compared to small molecular weight imaging probes. In order to make theragnostic and multifunctional nanoparticles, various polymeric nanoparticles were chemically or physically modified with radionuclides (11 C, 18 F, and 64 Cu), Mn2+ , Gd3+ , and optical imaging probes (fluorochromes, quantum dots, and photoproteins). Simultaneously, anticancer agents of hydrophobic chemical drugs, peptides, proteins, and DNAs were encapsulated or chemically grafted to the polymeric nanoparticles. After intravenous injection, multifunctional polymeric nanoparticles exhibited tumortargeting ability with low nonspecific uptake by other tissues. Simultaneously, they could be monitored in real-time and visualized in a noninvasive way, allowing for clinical uses in animals and humans. Radionuclide imaging is commonly divided into two general modalities: single photon emission computed tomography (SPECT) and positron emission tomography (PET) [42]. For PET applications, first-choice positron-emitting radionuclides are 18 F, 11 C, 13 N, and 15 O. These elements are able to emit a positive electron from the nucleus. In particular, the principal applied radionuclide for PET is Fluor-18 (18 F), which is known for its ideal

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half-life (1.83 h) [43]. Recently, Weissleder and co-workers reported the synthesis and in vivo characterization of an 18 F modified trimodal nanoparticle (18 F-CLIO) [44]. First, 18 FPEG3 was prepared using Synthra RN Plus automated synthesis. Second, 18 F-PEG3 was added to azido-CLIO. The reaction mixture was incubated for 40 min via a “Click” reaction and filtered to remove the 18 F-CLIO from unreacted 18 F-PEG3 [45].Purified 18 F-CLIO was injected into the tail vein of three mice, and dynamic PET imaging was performed to determine organ biodistribution and blood clearance over time. Figure 22.4 demonstrates one imaging series in which data were collected over the course of 16 h after intravenous injection of 18 F-CLIO and shows images acquired at 2, 7, and 16 h. Initially, high sustained activity was detected in the blood pool, demonstrated by high PET signal in the cardiac chambers. Over time, this signal declined, while an increase of activity was observed in the liver and spleen. The low renal and bladder uptake emphasizes that 18 F-CLIO is stable in vivo, and the tracer is not excreted via the kidneys. Taken together, these PET data follow the typical pharmacokinetics of CLIO, showing redistribution of the blood pool signal to the liver and lymphatic organs over the course of several hours (Fig. 22.4, d–f). This nanoparticle is biodegradable and had previously been shown to be broken down into elemental components in a weeks to months time frame. A major future application of the above approach is to use nanoparticle platforms for targeted molecular imaging. Nanoparticles allow multivalent attachment of small affinity ligands, vastly increasing avidity. Kwon and co-workers recently reported hydrophobically modified glycol chitosan with hydrophobic bile acids which stably formed self-assembled nanoparticles in aqueous condition. Various anticancer drugs, such as Taxol, cisplatin, camptothecin (CPT), docetaxel, and therapeutic peptide drugs were easily encapsulated into the hydrophobic cores of glycol chitosan nanoparticles due to their stable nanoparticle structure. After encapsulation of anticancer therapeutic agents, these glycol chitosan nanoparticles were modified with near-infrared (NIR) fluorescence imaging probe, Cy5.5, for optical imaging simultaneously [26, 46–49]. In vivo NIR fluorescence images showed that anticancer-agent-encapsulated glycol chitosan nanoparticles presented a prolonged circulation time of 1 week and preferential accumulation in the target tumor area compared to normal tissues. From the optical imaging data, tumor-targeting efficacy of glycol chitosan nanoparticles was four to sevenfold higher than that of other organs, providing decisive evidence that the nanoparticles specifically targeted tumors. It can be emphasized that the glycol chitosan nanoparticles progressively accumulated in the tumor tissues only based on the EPR effect within 24 h and lasted up to 10 days with considerably low uptake in the liver and spleen as compared with that of conventional nanoparticle-based formulations. Therefore it is deduced that cancer theragnosis with multifunctional polymeric nanoparticles presents a promising new strategy in cancer treatment, permitting simultaneous early-stage diagnosis, drug delivery, and real-time monitoring for therapeutic efficacy. Photodynamic therapy (PDT) is a method for cancer treatment that is a combination of photosensitizers and light. When irradiated with light, the photosensitizers produce singlet oxygen or other reactive oxygen species via a transfer of energy to molecular oxygen, which damages adjacent tissues. However, PDT has barriers: limited tumor selectivity of PDT agents and phototoxicity to skin and eyes. Choi et al. [50] recently proposed a proteasemediated PDT (PM-PDT) to avoid unexpected toxicities. Based on the concept of enzymatic cleavable imaging probe, they used PGC polymer conjugated with chlorine e6 (Ce6), as a photosensitizer that exhibits fluorescence signal and singlet oxygen generation (SOG). Normally, they show photosensitizer–photosensitizer self-quenching (Fig. 22.5a). When the

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(a)

(b)

(c)

(d)

FIGURE 22.4 Dynamic PET/CT imaging of BALB/c mouse after injection with 18 F-CLIO. Fused PET/CT coronal images (a) at 2 h and (b) 16 h after injection of 18 F-CLIO. Three-dimensional fused PET-CT images at (c) 2 h and (d) 16 h postinjection. Arrow (green) indicates blood pool region of interest (ROI) and asterisk indicates liver ROI. (Reproduced with permission from Devaraj et al. [44]. Copyright © 2006 American Chemical Society.)

PGC polymer conjugated with Ce6 contacts with protease CaB, Ce6-conjugated-PGC (Ce6PGC) can be cleaved and induces stronger fluorescence activation. In its native state, SOG of Ce6-PGC was only 13% compared to free Ce6, however, its SOG dramatically increased up to 79% with CaB treatment. In addition, Ce6-PGC showed high accumulation in an HT1080 tumor at 24 h postinjection (Fig. 22.5b), and its fragments were mainly distributed in the cellular cytoplasm. Apoptosis and significant tissue loss were observed in large areas of the tumor, after light illumination. The proposed PM-PDT strategy demonstrates visualization of the target and local drug concentration before therapy, and could be used as a primary anticancer treatment or as an adjuvant to other therapeutic options. However, this PDT approach based on protease cleavage is restricted by the limited tissue penetration of light. However, it illustrates key issues of concern, including a defined biological rationale, safety, pharmacokinetics, more information in preclinical trials, and manufacture on a large scale for commercial challenges. Once optimized, we hope that the multifunctional nanoparticle systems will provide improved treatment for life-threatening and debilitating diseases.

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(a)

(b)

(c)

(d)

FIGURE 22.5 (a) Schematic representation of PM-PDT strategy. Without protease, Ce6 molecules are self-quenched. After proteolytic cleavage of the peptide backbone, Ce6 and SOG were released and fluorescent properties were exhibited after light excitation. (b) Biodistribution of Ce6 in bilateral flank tumors. (c) Fluorescence images of tumor tissues that were stained after treatment with Ce6conjugated L-PGC and light. (d) TUNEL staining of Ce6-conjugated L-PGC-treated tumors 24 h after light illumination. (Reproduced with permission from Choi et al. [50], Copyright © 2006 American Association for Cancer Research.)

22.3.2 Multifunctional Magnetic Nanoparticles for Cancer Imaging and Therapy Magnetic nanoparticles (MNPs) have become important tools for magnetic resonance imaging (MRI) contrast enhancement in many cancers [51, 52] and have potential in current clinical diagnostic and therapeutic methods. In order to impart stability, biocompatibility, and functionality for targeting, MNPs require surface coating or conjugation of various materials [53]. Commonly, lipids [54], proteins [55], dendrimers [56], PEG [57], and ® ® polysaccharides [58] were used for the MNP’s surface coating. Ferridex , Combidex , ® Resovist , and AMI-288/gerumoxytrol, are clinically used as MNPs, which have been based on dextran or a similar carbohydrate coating. However, these MNPs still have limitations; they are fairly nonspecific for cancer diagnosis, and their signal enhancement is still unsatisfactory. To improve specificity and signal enhancement of MR contrast agents, recently, there has been a growing interest in the development of targeted MNP-based probes for tumor diagnostics. Furthermore, MNPs have been reported as drug carriers for delivering and releasing drugs [59]. Surface modified multifunctional MNPs with targeting moieties and/or therapeutic drugs can be targeted to tumors or other lesions at molecular or cellular events [18, 56, 60–63], and further allow simultaneous diagnosis and treatment of the diseased tissues with more efficient targeting, precise diagnosis, and therapeutic applications. As previously described, MNPs can serve both as contrast enhancement agents in MRI and as drug carriers in controlled drug delivery, targeted at cancer diagnosis and therapy,

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simultaneously [64, 65]. Conjugation MNPs with therapeutic drugs or targeting moieties are an effective method of theragnosis, which combines diagnosis and therapeutic applications. For example, Kohler et al. [66, 67] fabricated MNPs covalently bound with methotrexate (MTX), a therapeutic drug that can target folate receptor-overexpressing (FR+) cancer cells. The conjugation of MNPs was performed with MTX or MTX-PEG through amidation between COOH of MTX or PEG-MTX and the amine groups on the MNP surface. In addition to a therapeutic effect, MTX was conjugated to PAMAM nanoparticles with folic acid analog to enhance cytotoxicity to FR+ cancer cells [68]. After a 2-h incubation, MTXconjugated MNPs showed higher internalization into FR+ HeLa (ca. 10-fold) and MCF-7 (ca. 20-fold) cells in comparison with normal cells. The chemical conjugation of MTX on MNPs provides stable linkage in the physiological condition. However, after internalization of MTX-conjugated MNPs into FR+ cancer cells, MTX could be released by low pH and intracellular enzymes; the released MTX induces apoptosis in cancer cells. When MTXconjugated MNPs were introduced into normal cells, they greatly reduced the toxicity of MTX, suggesting MTX-conjugated MNPs exhibited severe toxicity against FR+ cancer cells, but did not exhibit side effects in normal tissues. Polymer coating on the surface of MNPs has enhanced stability and biocompatibility in vivo. However, it is possible that the coated polymers can be removed under harsh conditions in vivo and MNPs could aggregate [69, 70]. Recently, thermally crosslinked superparamagnetic iron oxide nanoparticles (TCL-SPIONs) have been developed by using poly(3-(trimethoxysilyl)propyl methacrylate-r-PEG methyl ether methacrylate-r-Nacryloxysuccinimide), which is an antibiofouling PEG copolymer that prevents nonspecific adsorption of plasma protein [71]. The antibiofouling PEG copolymers have a COOH group on TCL-SPIONs, which can be further converted to an amine group as the functional group. Finally, Cy5.5 is conjugated to amine modified TCL-SPION to obtain Cy5.5 dye-labeled TCL-SPIONs for use as a multimodal (MR/fluorescence) imaging probe [72]. Without any targeting moieties, TCL-SPIONs showed high accumulation in the tumor area and demonstrated MRI and fluorescence images. More recently, researchers developed a novel multifunctional drug carrier based on TCL-SPIONs as the imaging agent, DOX@TCL-SPIONs [73]. DOX is a well-known anticancer agent that was successfully incorporated in the polymeric shell of TCL-SPIONs via electrostatic interactions between positively charged DOX and the negatively charged polymer. Loaded DOX was released quicker at mildly acidic pH of tumor areas than at neutral pH of vascular compartments. Thus when DOX@TCL-SPIONs were systemically administered, DOX@TCL-SPIONs showed much lower toxicity in normal organs than free DOX, but showed a therapeutic response in cancer cells. In vivo MRI and fluorescence images indicated that DOX@TCLSPIONs could efficiently detect a Lewis lung carcinoma (LLC) and sufficiently deliver DOX to the tumor tissue—thereby showing excellent anticancer activity. Coupled with drug delivery to tumor areas, antisense oligonucleotides have become a powerful tool for effective antisense therapy. The problems, however, are associated with rapid degradation by exonuclease and poor diffusion across the cell membrane, which must be solved for clinical antisense therapy [74, 75]. Dendrimers are excellent candidates for antisense oligonucleotide delivery agents as nonviral vectors since they are safer, simpler to use, and more easily mass-produced than other viral vectors [76, 77]. To solve the problems in clinical antisense therapy, Pan et al. [78] fabricated polyamidoamine (PAMAM) dendrimer-conjugated MNPs as a gene delivery carrier. The MNP’s surface was coated with 3-aminopropyl-trimethoxysilane (APTS) to make G0 dMNP, and then further added to excessive methacrylate. After rinsing, the methacrylate group of G0 dMNP was further

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converted to amine by using excessive ethylenediamine. Stepwise growth using methacrylate and ethylenediamine was repeated until the desired number of generations from 1.0 to 5.0 (G1.0 -G5.0 ) was achieved. As the generation of dendrimer increased, a