Biomaterials for Cell Delivery: Vehicles in Regenerative Medicine 9781315151755, 1315151758, 9781351638272, 1351638270, 9781351647793, 1351647792, 9781498743174, 149874317X, 9781498743167

Table of contents :
Content: Cover
Half Title
Title Page
Copyright Page
Table of Contents
Series Preface
Preface
Editor
Contributors
Chapter 1: Natural Materials for Cell-Based Therapies
Chapter 2: Synthetic Polymers for Cell-Based Therapies
Chapter 3: Decellularized Tissues for Bioengineering of Whole Organs
Chapter 4: Cell-Based Approaches for Vascularized Tissue Formation
Chapter 5: Using Biomaterials to Deliver Cells In Vivo for Neural Tissue Engineering Applications
Chapter 6: Cell Carriers for Bone and Cartilage Repair In Vivo
Chapter 7: Delivering Therapeutic Cells to the Heart Chapter 8: In Vivo Cell Delivery: Pancreatic Islet TransplantationChapter 9: Cutaneous Wound Healing
Chapter 10: Incorporation of In Vitro Cell Conditioning for Enhanced Development of Tissue Engineered Skeletal Muscle Implants
Chapter 11: Conditioning Cells In Vitro to Facilitate Tendons and Ligament Regeneration
Chapter 12: In Vitro Cell Conditioning for the Development of Engineered Blood Vessels
Chapter 13: Perspectives for Clinical Translation: How Stem Cells and Biomaterials Affect Vasculogenesis and Neurogenesis in Preclinical and Clinical Models
Index

Citation preview

Biomaterials for Cell Delivery

Gene and Cell Therapy Series Series Editors

Anthony Atala & M. Graça Almeida-Porada PUBLISHED TITLES

Placenta: The Tree of Life, edited by Ornella Parolini Cellular Therapy for Neurological Injury, edited by Charles S. Cox, Jr. Regenerative Medicine Technology: On-a-Chip Applications for Disease Modeling, Drug Discovery and Personalized Medicine, edited by Sean V. Murphy and Anthony Atala Therapeutic Applications of Adenoviruses, edited by Philip Ng and Nicola Brunetti-Pierri Gene and Cell Delivery for Invertebral Disk Degeneration, edited by Raquel Madeira Gonçalves and Mario Adolfo Barbosa Bioreactors for Stem Cell Expansion and Differentiation, edited by Joaquim M. S. Cabral and Cláudia Lobato de Silva Biomaterials for Cell Delivery: Vehicles in Regenerative Medicine, edited by Aaron S. Goldstein

For more information about this series, please visit: https://www.crcpress.com/Gene-and-Cell-Therapy/book-series/CRCGENCELTHE

Biomaterials for Cell Delivery

Vehicles in Regenerative Medicine

Edited by

Aaron S. Goldstein

CRC Press Taylor & Francis Group 6000 Broken Sound Parkway NW, Suite 300 Boca Raton, FL 33487-2742 © 2019 by Taylor & Francis Group, LLC CRC Press is an imprint of Taylor & Francis Group, an Informa business No claim to original U.S. Government works Printed on acid-free paper International Standard Book Number-13: 978-1-4987-4316-7 (Hardback) This book contains information obtained from authentic and highly regarded sources. Reasonable efforts have been made to publish reliable data and information, but the author and publisher cannot assume responsibility for the validity of all materials or the consequences of their use. The authors and publishers have attempted to trace the copyright holders of all material reproduced in this publication and apologize to copyright holders if permission to publish in this form has not been obtained. If any copyright material has not been acknowledged please write and let us know so we may rectify in any future reprint. Except as permitted under U.S. Copyright Law, no part of this book may be reprinted, reproduced, transmitted, or utilized in any form by any electronic, mechanical, or other means, now known or hereafter invented, including photocopying, microfilming, and recording, or in any information storage or retrieval system, without written permission from the publishers. For permission to photocopy or use material electronically from this work, please access www.copyright. com (http://www.copyright.com/) or contact the Copyright Clearance Center, Inc. (CCC), 222 Rosewood Drive, Danvers, MA 01923, 978-750-8400. CCC is a not-for-profit organization that provides licenses and registration for a variety of users. For organizations that have been granted a photocopy license by the CCC, a separate system of payment has been arranged. Trademark Notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation without intent to infringe. Library of Congress Cataloging-in-Publication Data Names: Goldstein, Aaron S., editor. Title: Biomaterials for cell delivery : vehicles in regenerative medicine / editor(s): Aaron S. Goldstein. Other titles: Gene and cell therapy series. Description: Boca Raton : Taylor & Francis, 2018. | Series: Gene and cell therapy series | Includes bibliographical references. Identifiers: LCCN 2018029997 | ISBN 9781498743167 (hardback : alk. paper) | ISBN 9781315151755 (general) | ISBN 9781498743174 (pdf) | ISBN 9781351647793 (ePub) | ISBN 9781351638272 (Mobi/Kindle) Subjects: | MESH: Biocompatible Materials | Cell- and Tissue-Based Therapy | Regeneration | Tissue Engineering | Drug Delivery Systems | Drug Carriers Classification: LCC R857.M3 | NLM QT 37 | DDC 610.28--dc23 LC record available at https://lccn.loc.gov/2018029997 Visit the Taylor & Francis Web site at http://www.taylorandfrancis.com and the CRC Press Web site at http://www.crcpress.com

Contents Series Preface ...........................................................................................................vii Preface.......................................................................................................................ix Editor ........................................................................................................................xi Contributors ........................................................................................................... xiii

Chapter 1

Natural Materials for Cell-Based Therapies ........................................1 Marc Thompson and Mark Van Dyke

Chapter 2

Synthetic Polymers for Cell-Based Therapies ....................................25 Komal Prem, Cole H. Fuerste, and Guillermo A. Ameer

Chapter 3

Decellularized Tissues for Bioengineering of Whole Organs ........... 47 Tyler Callese, Andrea Peloso, Riccardo Tamburrini, Marcia Voigt, Joao Paulo Zambon, and Giuseppe Orlando

Chapter 4

Cell-Based Approaches for Vascularized Tissue Formation.............. 85 Banu Akar and Eric M. Brey

Chapter 5

Using Biomaterials to Deliver Cells In Vivo for Neural Tissue Engineering Applications ................................................................. 107 Laura de la Vega, Michaela Thomas, and Stephanie M. Willerth

Chapter 6

Cell Carriers for Bone and Cartilage Repair In Vivo ....................... 139 Dilip Thomas, Manus Biggs, Timothy O’Brien, and Abhay Pandit

Chapter 7

Delivering Therapeutic Cells to the Heart ....................................... 173 Maribella Domenech, Jaime E. Ramirez-Vick, and Donald O. Freytes

Chapter 8

In Vivo Cell Delivery: Pancreatic Islet Transplantation ................... 191 John Patrick McQuilling, Alan C. Farney, and Emmanuel C. Opara

Chapter 9

Cutaneous Wound Healing............................................................... 217 Yixiao Dong and Geoffrey C. Gurtner v

vi

Contents

Chapter 10 Incorporation of In Vitro Cell Conditioning for Enhanced Development of Tissue-Engineered Skeletal Muscle Implants ........ 241 David Remer, Juliana Passipieri, and George Christ Chapter 11 Conditioning Cells In Vitro to Facilitate Tendons and Ligament Regeneration ..................................................................................... 263 Chelsea E. Coffey, Zachary R. Mussett, and Vassilios I. Sikavitsas Chapter 12 In Vitro Cell Conditioning for the Development of Engineered Blood Vessels ................................................................................... 281 Chris A. Bashur and Mozhgan Shojaee Chapter 13 Perspectives for Clinical Translation: How Stem Cells and Biomaterials Affect Vasculogenesis and Neurogenesis in Preclinical and Clinical Models ....................................................... 297 Neel Sharma, Christina Ross, In Kap Ko, Yuan-Yuan Zhang, Shay Soker, and Tracy Criswell Index ...................................................................................................................... 333

Series Preface Gene and cell therapies have evolved in the past several decades from a conceptual promise to a new paradigm of therapeutics, able to provide effective treatments for a broad range of diseases and disorders that previously had no possibility of cure. The fast pace of advances in the cutting-edge science of gene and cell therapy, and supporting disciplines ranging from basic research discoveries to clinical applications, requires an in-depth coverage of information in a timely fashion. Each book in this series is designed to provide the reader with the latest scientific developments in the specialized fields of gene and cell therapy, delivered directly from experts who are pushing forward the boundaries of science. In this volume of the Gene and Cell Therapy book series, Biomaterials for Cell Delivery, the editor has assembled a remarkable team of outstanding investigators across disciplines to present a series of riveting articles on the intersection of biomaterials and technologies that play an integral role in enhancing the effectiveness of cell therapies and regeneration and foretelling major conceptual advances in the field. The chapters describe the state-of-the-art technologies and applications of biomaterials, scaffolds, and decellularized organs, as well as provide cutting-edge information on angiogenesis and cell delivery for the regeneration of peripheral nerve, bone, muscle, cartilage, skin, cardiac, and pancreatic islet tissues. In addition, the fundamental issues of cell and scaffold choice are presented under a clinical perspective for the treatment of cardiovascular and neurodegenerative pathologies. We would like to thank the volume editor, Aaron Goldstein, and the authors, all of whom are remarkable experts, for their valuable contributions. We would also like to thank our senior acquisitions editor, C.R. Crumly, and the Taylor & Francis/ CRC Press staff for all their efforts and dedication to the Gene and Cell Therapy book series. Anthony Atala Wake Forest Institute for Regenerative Medicine M. Graça Almeida-Porada Wake Forest Institute for Regenerative Medicine

vii

Preface Tissue engineering can be envisaged as an offshoot of biomaterials, whereby a temporary, porous three-dimensional scaffold can be combined with cells and biologicals to achieve a biologically active construct to guide tissue regeneration. Efforts over the past two decades to employ this paradigm have revealed the complexity of the in vivo host response to such constructs and how seemingly subtle biomaterial properties of the scaffold can affect functional outcomes. For example, stiffer materials are more conducive to the regeneration of skeletal tissues, whereas softer materials are more effective for developing capillary networks and nervous and fat tissues. In reality, these biomaterials should not be perceived as wholly inert, but as biologically active entities that, upon implantation, can elicit host responses such as inflammation, angiogenesis, and tissue regeneration. Hence, the chemistry and biochemistry of the biomaterial can be designed—through the incorporation of cytokines, morphogens, and miRNAs—to stimulate regenerative processes, while the incorporation of labile bonds permits the release of payloads or the disintegration of the scaffold in concert with host tissue infiltration. On the other side of the coin, host tissue infiltration and regeneration requires recruitment, proliferation, and maturation of the appropriate cell types. This can be a limiting factor for regeneration if the particular cell type is highly specialized (e.g., neurons, liver hepatocytes, insulin-secreting beta cells, cardiomyocytes), if the target tissue is poorly vascularized (e.g., ligament, cartilage), or if the tissue deficit is large. Here, regenerative strategies have been built around the delivery of tissue repair cells: differentiated cells that exhibit the target phenotype but do not necessarily proliferate rapidly; undifferentiated (e.g., stem) cells that have the potential to proliferate extensively and mature into the target tissue phenotype; or a mixture of different cell types. However, cell-based therapies often suffer from poor cell adhesion, viability, and retention at the target site. This is unsurprising as the target site is often marked by inflammation and hypoxia. In this case, a biomaterial vehicle can provide a supportive microenvironment to facilitate cell viability, proliferation, and tissue regeneration. In general, no single strategy exists that can be applied broadly to the regeneration of many tissues and organs: each has its own unique structure, function, and requirements. Likewise, for any given tissue or organ, no single regenerative strategy exists. Rather, a continuum of strategies exists: from the simple to the complex; from the novel, which may not yet be ready for testing in vivo; to the mature, which are already in clinical trials. Thus, the purpose of this book is to summarize key strategies and recent accomplishments in the area of developing cell–biomaterial constructs for regenerative medicine. To this end, Chapters 1 through 3 review the state of the art of biomaterial scaffolds, which include natural and synthetic materials as well as decellularized organs. Because tissue regeneration hinges on a vascular network to provide cells, oxygen, and nutrients, Chapter 4 is devoted to the fundamental problem of angiogenesis. Thereafter, the bulk of this book is devoted to unique problems associated with particular tissues and organs. These are organized into two sections. ix

x

Preface

First, Chapters 5 through 9 concern cell delivery for the regeneration of peripheral nerve, bone, cartilage, skin, cardiac, and pancreatic islet tissues. Next, Chapters 10 through 12 examine the value-added impact of culturing cells within a bioreactor in vitro prior to implantation for the purpose of enhancing muscle, tendon/ligament, and vascular tissue function in vivo. Finally, Chapter 13 revisits the central issues of cell and scaffold choice from the clinical perspective for the treatment of cardiovascular diseases and neurodegenerative pathologies. While no singular volume can be either comprehensive or completely up to date, this book highlights modern strategies and accomplishments of the past decade in the areas of biomaterial vehicles for cell delivery, as well as the various types of cells and pharmaceutical agents that can facilitate tissue regeneration. Aaron S. Goldstein

Editor Aaron S. Goldstein, PhD, designs and evaluates two- and three-dimensional tissue microenvironments to guide stem cell differentiation into orthopedic tissue phenotypes. His research involves the combination of biocompatible materials and materials processing techniques to systematically and spatially vary the chemistry, topography, and mechanical properties of the biomaterial surfaces that are presented to cells. His interests also include the use of perfusion and mechanical stretch bioreactors to stimulate cell phenotypes through the activation of mechanotransductive signaling pathways. He is the author of more than 40 peer-reviewed research articles in the areas of biomaterials, cell adhesion, and tissue engineering. He earned a BS in chemical engineering from the University of California and a PhD in chemical and biomedical engineering from Carnegie Mellon University. He was a postdoctoral research fellow in the Department of Bioengineering at Rice University before joining the Department of Chemical Engineering at Virginia Tech in 1999.

xi

Contributors Banu Akar Pritzker Institute of Biomedical Science & Engineering and Department of Biomedical Engineering Illinois Institute of Technology and Hines Veterans Hospital Chicago, Illinois Guillermo A. Ameer Department of Biomedical Engineering Northwestern University Evanston, Illinois Chris A. Bashur Department of Biomedical Engineering Florida Institute of Technology Melbourne, Florida Manus Biggs Centre of Research in Medical Devices (CÚRAM) National University of Ireland Galway Galway, Ireland Eric M. Brey Pritzker Institute of Biomedical Science & Engineering and Department of Biomedical Engineering Illinois Institute of Technology and Hines Veterans Hospital Chicago, Illinois

Tyler Callese Wake Forest School of Medicine Winston-Salem, North Carolina George Christ Department of Biomedical Engineering and Department of Orthopedic Surgery University of Virginia Charlottesville, Virginia Chelsea E. Coffey Stephenson School of Biomedical Engineering University of Oklahoma Norman, Oklahoma Tracy Criswell Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Winston-Salem, North Carolina Laura de la Vega Department of Mechanical Engineering University of Victoria Victoria, British Columbia, Canada Maribella Domenech Department of Chemical Engineering Universidad de Puerto Rico–Mayagüez Mayagüez, Puerto Rico

xiii

xiv

Contributors

Yixiao Dong Department of Surgery Stanford University School of Medicine Stanford, California

John Patrick McQuilling Wake Forest Institute for Regenerative Medicine Winston-Salem, North Carolina

and

and

Shanghai Institute for Advanced Immunochemical Studies (SIAIS) ShanghaiTech University Shanghai, China

Virginia Tech–Wake Forest School of Biomedical Engineering & Sciences (SBES) Blacksburg, Virginia

Alan C. Farney Wake Forest Institute for Regenerative Medicine and Department of Surgery Wake Forest School of Medicine Winston-Salem, North Carolina

Zachary R. Mussett Stephenson School of Biomedical Engineering University of Oklahoma Norman, Oklahoma

Donald O. Freytes Comparative Medicine Institute North Carolina State University and Joint Department of Biomedical Engineering North Carolina State University/ University of North Carolina at Chapel Hill Raleigh, North Carolina Cole H. Fuerste Department of Biomedical Engineering Northwestern University Evanston, Illinois Geoffrey C. Gurtner Department of Surgery Stanford University School of Medicine Stanford, California In Kap Ko Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Winston-Salem, North Carolina

Timothy O’Brien Regenerative Medicine Institute National University of Ireland Galway Galway, Ireland Emmanuel C. Opara Wake Forest Institute for Regenerative Medicine Winston-Salem, North Carolina and Virginia Tech–Wake Forest School of Biomedical Engineering & Sciences (SBES) Blacksburg, Virginia Giuseppe Orlando Wake Forest School of Medicine Winston-Salem, North Carolina Abhay Pandit Centre of Research in Medical Devices (CÚRAM) National University of Ireland Galway Galway, Ireland Juliana Passipieri Department of Biomedical Engineering University of Virginia Charlottesville, Virginia

xv

Contributors

Andrea Peloso Wake Forest School of Medicine Winston-Salem, North Carolina Komal Prem Department of Biomedical Engineering Northwestern University Evanston, Illinois Jaime E. Ramirez-Vick Department of Biomedical, Industrial & Human Factors Engineering Wright State University Dayton, Ohio David Remer Department of Biomedical Engineering University of Virginia Charlottesville, Virginia Christina Ross Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Winston-Salem, North Carolina Neel Sharma Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Winston-Salem, North Carolina Mozhgan Shojaee Department of Biomedical Engineering Florida Institute of Technology Melbourne, Florida

Vassilios I. Sikavitsas Stephenson School of Biomedical Engineering and School of Chemical, Biological, and Materials Engineering University of Oklahoma Norman, Oklahoma Shay Soker Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Winston-Salem, North Carolina Riccardo Tamburrini Wake Forest School of Medicine Winston-Salem, North Carolina Dilip Thomas Regenerative Medicine Institute National University of Ireland Galway and Centre of Research in Medical Devices (CÚRAM) National University of Ireland Galway Galway, Ireland and Cardiovascular Institute Stanford University Palo Alto, California Michaela Thomas Department of Mechanical Engineering University of Victoria Victoria, British Columbia, Canada

xvi

Marc Thompson Department of Biomedical Engineering and Mechanics Virginia Tech Blacksburg, Virginia Mark Van Dyke Department of Biomedical Engineering and Mechanics Virginia Tech Blacksburg, Virginia Marcia Voigt Wake Forest School of Medicine Winston-Salem, North Carolina

Contributors

Stephanie M. Willerth Department of Mechanical Engineering University of Victoria Victoria, British Columbia, Canada Joao Paulo Zambon Wake Forest School of Medicine Winston-Salem, North Carolina Yuan-Yuan Zhang Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Winston-Salem, North Carolina

1

Natural Materials for Cell-Based Therapies Marc Thompson and Mark Van Dyke

CONTENTS 1.1

Introduction ...................................................................................................... 2 1.1.1 Hydrogels ..............................................................................................3 1.1.2 Foam Biomaterials ................................................................................3 1.1.3 Films .....................................................................................................4 1.1.4 Cell Encapsulation ................................................................................ 4 1.1.5 Composites............................................................................................5 1.2 Protein-Based Biomaterials ..............................................................................5 1.2.1 Keratin .................................................................................................. 5 1.2.1.1 Structure.................................................................................5 1.2.1.2 Fabrication ............................................................................. 6 1.2.1.3 Applications ........................................................................... 7 1.2.2 Fibrin .................................................................................................... 7 1.2.2.1 Structure.................................................................................8 1.2.2.2 Fabrication ............................................................................. 8 1.2.2.3 Applications ........................................................................... 9 1.2.3 Collagen ................................................................................................9 1.2.3.1 Structure............................................................................... 10 1.2.3.2 Fabrication ........................................................................... 10 1.2.3.3 Application ........................................................................... 11 1.3 Polysaccharide Biomaterials ........................................................................... 13 1.3.1 Alginates ............................................................................................. 13 1.3.1.1 Structure............................................................................... 13 1.3.1.2 Fabrication ........................................................................... 14 1.3.1.3 Application ........................................................................... 14 1.3.2 Chitosan .............................................................................................. 15 1.3.2.1 Structure............................................................................... 15 1.3.2.2 Fabrication ........................................................................... 16 1.3.2.3 Application ........................................................................... 16 1.3.3 Hyaluronic Acid/Hyaluronan ............................................................. 17 1.3.3.1 Structure............................................................................... 17 1.3.3.2 Fabrication ........................................................................... 18 1.3.3.3 Application ........................................................................... 18 1.4 Summary ........................................................................................................ 19 References ................................................................................................................ 19 1

2

1.1

Biomaterials for Cell Delivery

INTRODUCTION

A biomaterial can be described as any material designed to interact in concert with a biological system. Biomaterials past and present have long stood as a symbol of scientists’ attempts to interact with, recreate and improve upon human anatomy and function. From surgical sutures to organ replacement, investigators have strived to better understand the details of human biology and apply this knowledge to numerous applications using a myriad of conventional and novel biomaterials. Today, that work has resulted in innovations correlating to innumerable lives improved and saved. Improvement of the synthesis and application of materials in particular have spurred the growth of an entire field of biomaterials research, into which this chapter will take an in-depth look, specifically, the impact of natural biomaterials used for delivering cells for tissue regeneration purposes. In terms of tissue repair and regeneration, biomaterials have increased in complexity and robustness while their ability to support the repair or replacement of damaged tissue has become more probable. Today, clinically applied tissue regeneration procedures are in large part performed using either autografts (tissue originating from the same individual) or allografts (tissue originating from another human individual) (Khan et  al. 2005). Autografting, which has been used extensively with satisfactory results, has several limitations, including patient pain, medical costs and a finite supply of suitable donor tissue (Oryan et  al. 2014). Allografts, while more abundant than autograft tissue, carry the uncertainty of biocompatibility and the potential for disease transmission. The delivery of cells for tissue regeneration by means of natural biomaterial vehicles that are more accessible than autograft tissue, and elicit reduced complications compared with allograft tissue, hold great promise as a means to overcome current limitations in repairing and regenerating defective or damaged tissues (Kanbe et al. 2007; Naderi et al. 2011). A prerequisite of classical tissue engineering approaches is a suitable biomaterial scaffold or substrate with an architectural design, chemical, mechanical and physical makeup comparable to that of the native tissue. Prior clinical practices of delivering cell suspensions directly into a defect site have proven to be problematic—if not ineffective—due to insufficient retention within the defect and subsequent flushing into the surrounding tissue (Endres et al. 2010). Consequently, the implementation of naturally derived vehicles, composed of more durable materials found readily in the body and more specifically within the tissue region of interest, are an attractive option for mimicking the native host tissue in the form of three-dimensional (3D) prefabricated scaffolds that retain the delivered cell population (Kretlow and Mikos 2007; Wang et  al. 2015). Herein we focus on scaffolds derived from natural polymers (i.e., proteins, polysaccharides) that can then be processed into hydrogels, foams and films, as just a few examples. These materials are capable of conforming to irregular shapes at the site of a defect or injury for complete access and infiltration of the wound and of delivering cells and pharmaceutical agents to guide tissue regeneration.

Natural Materials for Cell-Based Therapies

3

1.1.1 HYDROGELS Hydrogels are a class of biomaterials with high water content and soft tissue-like mechanical properties (Matricardi et al. 2013, Dubbini et al. 2015) that make them attractive as scaffolds for a variety of tissue engineering applications. Moreover, highly moldable hydrogels can be injected into irregular-shaped defects, enabling targeted delivery of cells for in vivo tissue engineering (Park et al. 2012) with limited surgical invasion (Zhang et al. 2014). Further, injectable biomaterials may be designed to crosslink in situ to facilitate integration and tissue regeneration at the region of interest (Montanari et al. 2015). In situ crosslinking is generally accomplished by chemical polymerization, photopolymerization or thermal crosslinking (Berger et al. 2004; Hennink and van Nostrum 2012; Montero-Rama et al. 2015). Chemical polymerization and photopolymerization have been applied extensively for gelation of injectable hydrogels, although the use of chemical initiators risks cytotoxicity and introduces undesired complexity to the delivery systems (Elias et al. 2015). Thermal crosslinking is applicable for thermosensitive hydrogels that undergo phase transitions in response to temperature change; however, these systems are consequently limited to temperature ranges that are well tolerated by living cells.

1.1.2 FOAM BIOMATERIALS Instances in which a solid or microporous hydrogel scaffold does not provide the desired level of nutrient, fluid or waste exchange may warrant the application of hydrogel foams or sponges. Implantation of interconnected, macroporous foams, commonly fabricated from alginate, collagen or chitosan (Anderson et al. 2014; Mi et al. 2001; Ranucchi et al. 2000), allows for more complete cellular invasion and improved mass transport of oxygen, nutrients and waste. Increases in porosity also provide greater cavity space for new tissue ingrowth and formation without having to compete with the implanted biomaterial, which may not degrade at the same rate that tissue is formed (Anderson et al. 2014; Prieto et al. 2014). These foams may be delivered prewetted; however, an added benefit of non-wetted foams is that improved fluid infiltration within the dry material may result in further adsorption of cells, nutrients and factors from the surrounding milieu that are beneficial to tissue regeneration. Depending on the intended use, foams have exhibited highly variable mechanical properties, with more elastic foams exhibiting Young’s moduli of 4 kPa and more rigid foams displaying compressive strengths of up to 200  kPa, supporting their use in soft tissue as well as bone regeneration, respectively (Karashima et al. 2009; Yu et al. 2013). Macroporosity in a biomaterial foam is not naturally occurring, as in the microporousness of similar biomaterial constructs. Instead, pores of larger and more consistent diameter are produced using techniques such as gas foaming, porogen leaching and phase separation (Christenson et al. 2007; Moglia et al. 2011; Zhang et al. 2005). While successful material synthesis has been accomplished and tested clinically, incomplete

4

Biomaterials for Cell Delivery

leeching of the sacrificial pore-forming material risks cytotoxicity, and potentially deleterious consequences for the regeneration of native tissue (Prieto et al. 2014). Foams also experience occasional difficulty during implantation. Foams are often cast to a specific size and shape that is not necessarily an exact fit for a defect of a particular size and shape, which can vary in cases such as broken bones. In addition, foams can undergo expansion or contraction during fabrication (Karashima et al. 2009). Consequently, foams may need to undergo a shaping step, which can lead to significant chipping and breaking, especially with harder and more brittle constructs.

1.1.3

FILMS

In line with improving transport through biomaterials, thin films or biomaterial sheets have been employed when the environment calls for rapid transport of materials. Previously, collagen and chitosan-based biomaterial films have been synthesized as corneal implants or wound dressings, respectively (Lee et al. 2001; Lahooti et al. 2016). In both instances, these biomaterials proved effective, allowing for sufficient oxygen, fluid and waste transport. In most cases, these films are easily biodegraded or removed (several uses include topical or superficial applications). In scenarios in which films will not experience large mechanical strains (as is the case with corneal implants) where materials can only withstand tens of kilopascals in tensile forces at most, softer and more flexible films are preferred. On the other hand, stiffer and tougher materials have been processed into films for applications that experience thousands of kilopascals of tensile stress (Vrana et al. 2007; Kroustalli et al. 2013).

1.1.4

CELL ENCAPSULATION

As an emerging methodology in tissue engineering, complete cell encapsulation has attracted increasing interest in recent years (Hunt and Grover 2010; Man et al. 2012; Uludag et al. 2000). It promotes tissue regeneration by the localized retention of cells and delivery of therapeutic trophic factors over a desirable period of time. In addition, the biomaterial encapsulant can mitigate the risks associated with direct cell injection such as apoptosis within the first few days of implantation (Abbah et al. 2006). To date, a broad variety of multipotent progenitor cell types have been examined for the regeneration of musculoskeletal, nerve, hepatic, and cardiac tissues, including mesenchymal stem cells (MSCs) (Niemeyer et al. 2007), bone marrow progenitor cells (Steadman et al. 2002), adipose tissue stromal cells (ATSCs) (Cowan et al. 2004) and muscle-derived stem cells (Prigozhina et al. 2008). These primary cell types exhibit a variety of favorable properties for potential delivery. For example, MSCs from bone marrow are negative for immunologically relevant surface markers and inhibit proliferation of allogenic T-cells in vitro (Hennink and van Nostrum 2012) Although concerns exist that MSCs from bone marrow may lose their immunosuppressive potential after expansion in vitro or implantation into allogenic recipients (Prigozhina et al. 2008). Alternatively, ATSCs are easily harvestable with lower

Natural Materials for Cell-Based Therapies

5

donor site morbidity when compared with other pluripotent stem cell sources (Cowan et al. 2004). Additionally, reports indicate that ATSCs possess phenotypic and functional characteristics similar to MSCs derived from bone marrow, and attach and proliferate easily in culture, making them potentially available on a large scale.

1.1.5 COMPOSITES To date, a wide variety of natural biomaterials have been identified and employed for tissue engineering applications. Individually, these materials exhibit similar properties to a wide range of biological tissues. However, in most cases, no singular biomaterial perfectly mimics the biological system it is meant to recreate. As a result, composite materials at varying concentrations and ratios are often employed to modulate factors such as degradation rate and better mimic the chemical, biochemical, and mechanical properties of the system of interest. In this chapter, the focus is mainly on individual natural biomaterials for cell delivery; however, in many cases composites are necessary to better achieve and maintain tissue regeneration. Within biological systems, three major classes of natural polymers exist: polysaccharides (e.g., cellulose, alginate), polynucleic acids (e.g., DNA, RNA), and proteins. However, in reality, proteoglycan and glycoprotein hybrids also exist. In the following sections, we describe a few of the most popular and promising natural proteins and polysaccharides for cell delivery and tissue regeneration.

1.2

PROTEIN-BASED BIOMATERIALS

1.2.1 KERATIN Keratins are a family of structural proteins, and can be found within a broad variety of animal tissues including: skin, hair, claws, horns, hooves, whale baleen, and bird feathers. The toughness and hardness of these proteinaceous materials derive from the organization and extensive crosslinking of the individual proteins. 1.2.1.1 Structure Two types of mammalian α-keratins have been characterized and denoted as the “soft” α-keratins found in skin (epidermal keratins) and the “hard” α-keratins (trichocytic keratins) stemming from epidermal appendages, such as hair, claws, and quills. Both forms of keratin are stabilized around the time of cell death by disulfide bonds between pairs of cysteine residues. The number of disulfide bonds in trichocytic keratin is greater than in epidermal keratin. Electron micrographs of cross-sections reveal that both types of keratin exhibit a filament and matrix-like texture (Bragulla and Homberger 2009). In the case of the epidermal keratins, the matrix is believed to comprise some or all of the N- and C-terminal domains of the molecules. The cores of the filaments are formed from rod domains that are approximately 40–50 nm in length. A comparable structure exists in the keratins of birds and reptiles except for the fact that the secondary structure of the filaments is predominantly the β-sheet (Wang et al. 2016). A major difference between epidermal and trichocytic keratins

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Biomaterials for Cell Delivery

is that the matrix in trichocytic keratins contains considerable amounts of sulfur-rich and glycine–tyrosine-rich proteins. Also, during the transition from the reduced to the oxidized state, the framework of the intermediate filaments undergoes both molecular slippage and compaction, a process that is much more apparent under chemically induced processes, but has been shown to occur in natural settings as well (Grune et al. 1997; Thiele et al. 1999). This brings the cysteine residues of neighboring molecular segments into axial alignment and enables the formation of disulfide linkages (Hill et al. 2010). By comparison, epidermal proteins contain very little cysteine and particular regions contain none at all, which indicates that any change in the framework of epidermal intermediate filaments on oxidation will not depend on changes in free energy associated with the formation of disulfide linkages. There is a clear functional relationship between the concentrations of disulfide bonds in the two types of keratin and their mechanical properties. Epidermal keratins have a sufficient concentration to ensure insolubility without sacrificing flexibility, whereas trichocytic keratins, with their extensive crosslinking, are tailored to strengthen epidermal appendages to a degree consistent with their function (Asquith 1977). Prior research suggests keratin and keratin derivatives exhibit several attractive mechanical and biological properties that support their use as a cell delivery vehicle. Upon hydration, keratin hydrogels mechanically perform as a flexible viscoelastic material. Data on the compressive modulus of keratin hydrogels ranges from approximately 4 MPa in dry scaffolds to almost 0.01 MPa after just 5 minutes of hydration (de Guzman et al. 2011). Tensile tests on keratin from various sources have shown that, once isolated, keratinized structures can display an elastic modulus anywhere between 0.01 and 9 GPa (Papir et al. 1975). In terms of biological cues, keratin contains intrinsic amino acid motifs in the form of RGD, LDV, and EDS amino acid– binding sites, which are integrin specific, as well as other classes of receptors that are non-integrin specific based on antibody blocking of the β2 integrin subunit. Both groups support cellular adhesion specifically (Hill et  al. 2010; Rouse et  al. 2010; Sando et al. 2010). Along with a natural propensity for cell adhesion, keratin proteins exhibit a uniquely low inflammatory response upon implantation. The mechanisms of the process are still in question, but the validity of this low response has been documented via analyses of macrophage activation when in contact with keratin, where keratin increased anti-inflammatory macrophage cytokines and decreased pro-inflammatory macrophage cytokines (Fearing et al. 2014). 1.2.1.2 Fabrication Keratins are generally extracted through chemical means that first break the disulfide bonds prevalent in keratinized tissues. The alpha- and less common gammakeratins are converted to their non-crosslinked forms via oxidation or reduction, as cysteine is converted to either cysteic acid or cysteine, respectively. Next, free proteins are extracted with solvents capable of denaturing the proteins for subsequent purification (by filtration and dialysis). Here, oxidizing denaturants produce derivatives referred to as keratoses, while reducing denaturants produce cysteine-containing proteins called kerateines. While essentially homologous, the sulfur-containing cysteine groups primarily define the differences between keratoses and kerateines. Keratoses are non-disulfide crosslinkable, but are water soluble and susceptible to

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hydrolytic degradation due to the polarized protein backbone. The higher polarized nature of keratoses causes them to degrade relatively quickly in vivo (in days to weeks). Kerateines are less polar and therefore slightly less soluble in water. They are more stable and can be re-crosslinked through oxidative coupling of cysteine groups. This results in biomaterials that can persist in vivo for weeks to months (Noishiki et al. 1982). 1.2.1.3 Applications Investigations of the structure and properties of soluble keratins has continued, but beginning in the 1970s, the processing of these protein solutions into derivative physical states such as gels and films began appearing in the literature (Van Dyke and Nanney 2002). Early keratin biomaterials appeared in the form of a wool derivative applied as a vascular graft coating, which was successfully implanted into a canine model for approximately 200 days (Noishiki et al. 1982). Since then, keratins have been used in several in vitro and preclinical models for wound healing (Burnett et al. 2013), bone regeneration (Zhao et al. 2015), hemostasis (Burnett et al. 2013) and peripheral nerve repair (Sierpinski et al. 2008). Preclinical studies have suggested the viability of a keratin-based nerve repair biomaterial. Keratin hydrogels introduced in vivo have resulted in significant increases in Schwann cell proliferative and migratory ability, suggesting a potentially improved nerve regenerative response (Nakaji-Hirabayashi et  al. 2008; Sierpinski et al. 2008). These materials were mechanically weaker compared with other hydrogels but maintained their structure after hydration. In other previous studies, keratin-based hydrogels were evaluated for their ability to mitigate further tissue damage after spinal cord injury. (Fearing et al. 2014). However, in some cases the requirement of strong denaturants to dissolve keratin molecules makes it difficult to prepare cell-seeded hydrogels for transplantation given that residual denaturants in the hydrogel can disrupt cellular processes or act in a cytotoxic manner against seeded cells. Keratin has also been previously exploited for its ability to induce cell differentiation. Previously, human cardiac stem cells (hCSCs) have been shifted to a smooth muscle cell lineage in vitro (Ledford et  al. 2017). Other studies have displayed keratins ability to induce macrophage differentiation, suggesting that it be exploited as a potential wound healing therapy (Fearing et al. 2014). Results from other studies suggest that keratin biomaterials can in fact promote a greater production of anti-inflammatory cytokines and suppress production of pro-inflammatory cytokines. This is thought to be a function of the ability of keratins to modulate macrophage phenotype during the wound healing process, leading to decreased inflammation and an improved environment for regeneration as a whole (Fearing et al. 2014).

1.2.2 FIBRIN Fibrin falls within the category of water-soluble biopolymers that are readily found in the body and have key applications in tissue regeneration. As one of the key components to the physical contribution of both wound healing and tissue formation, fibrin

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contains intrinsic factors to induce blood clotting, fibrinolysis and cell-to-extracellular matrix (ECM) interactions pertaining to tissue remodeling (Bensaid et al. 2003). These properties along with the body’s ability to recognize fibrin as a naturally present material, contribute to a reduced inflammatory response upon implantation, and the prevalence of fibrin in the blood suggests it has potential as an autologous material for transplant (Ahmed et al. 2008). As such, fibrin biomaterials have been produced in the form of microbeads, glues, and hydrogels. Fibrin has possibly experienced more support than any other biomaterial in terms of Food and Drug Administration (FDA)approved clinical applications, given that fibrin materials have been approved in the form of sealants and hemostatic formulations such as Evicel® or Crosseal® (Ahmed et al. 2008; Ulery et al. 2011). 1.2.2.1 Structure Fibrin is derived from its precursor, fibrinogen, via enzymatic cleavage by thrombin. Fibrinogen consists of three polypeptide chains (Aα, Bβ, and γ) that are joined together by six disulfide bridges and contain E and D domains (Mosesson et al. 2001). Thrombin-mediated cleavage of fibrinopeptide A from the Aα chains and similar cleavage of fibrinopeptide B from the Bβ chains leads to conformational changes that expose polymerization sites. Cleavage of A and B peptides results in exposed E domains (EA/EB) that can subsequently crosslink with exposed D domains (DA/DB). Concomitantly, the blood coagulation factor XIIIa, a transglutaminase, rapidly crosslinks γ chains by introducing intermolecular ζ-(γ-glutamyl) lysine covalent bonds between the glutamine of one γ-chain and lysine of the other γ-chain (Shainoff et al. 1990; Mosesson 2003). These covalent crosslinking mechanisms produce a stable form of fibrin that is resistant to proteolytic degradation. Stability can be further reinforced through the addition of crosslinkers such as genipin. 1.2.2.2 Fabrication Fibrin hydrogels can be synthesized from commercially purified allogeneic fibrinogen and purified thrombin. One clear advantage of fibrin gels over other biomaterial gels, however, is that they can be obtained autologously. Fibrin gels can be produced via a subject’s blood and used as an autologous scaffold for seeded cells to create a 3D construct that exhibits little to no toxicity or inflammatory responses (Ye et  al. 2000). Furthermore, cells entrapped within fibrin gels have been shown to produce more collagen and elastin (Long and Tranquillo 2003; Neidert et al. 2002), the major building blocks for the extracellular matrix environment, compared with cells entrapped in collagen gels, for example. Fibrin gels have shown a tendency toward shrinkage that may, in some cases, be a desired property since it results in tension within the developing tissue, resulting in a greater production of collagen when the gel is remodeled (Eyrich et al. 2007). This contraction improves mechanical stability under loading and mitigates the loss of cells and freshly formed ECM.

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1.2.2.3 Applications Fibrinogen and subsequent fibrin have major biological impacts on the efficiency of fibrinolysis, blood clotting, inflammatory response, wound healing, and general cell and matrix interactions (Mosesson et  al. 2001). The clinical applications of fibrin have largely remained in the realm of hemostatic agents for invasive surgeries related to the heart, spleen, and liver, and in patients suffering from hemophilia. As a natural component in blood clotting, fibrin has also seen significant application as a sealant in soft tissue dissection procedures and as an agent to reduce bleeding in vascular, intestinal, and colonic anastomosis. However, for the purposes of this chapter, the focus is on the use of fibrin hydrogels to deliver cells. Fibrin hydrogels have been used extensively in the last decade in a variety of tissue engineering applications, including engineering of adipose, cardiovascular, ocular, muscle, liver, skin, cartilage, and bone tissues. In addition, they have been used for promoting angiogenesis (Janmey et al. 2009). These gels could have many advantages as a cell delivery vehicle in terms of biocompatibility, biodegradation and hemostasis (Janmey et al. 2009; Mol et al. 2005). Currently, mechanical conditioning of fibrin hydrogel scaffolds loaded with cells is thought to be particularly relevant in the field of cardiovascular tissue engineering to enhance tissue formation and organized contractibility (Mol et al. 2005). However, as a potential cell delivery vehicle, this material has three major disadvantages: shrinkage of the gel during the formation of flat sheets, low mechanical stiffness, and rapid degradation in vivo (over several days due to cell-associated enzymatic activities) (Ahmed et  al. 2008). Nevertheless, these limitations can often be overcome by modifying the fibrin. Gel shrinkage can be prevented by incorporating a fixative such as poly-l-lysine into the fibrin gel. Mechanical stiffness can be increased by combining fibrin hydrogels with other scaffold materials (e.g., polyurethanes, polycaprolactone-based polyurethanes, polycaprolactone, β-tricalciumphosphate, polyethylene glycol [Buehler 2006]) to obtain constructs with desired mechanical strength. Lastly, rapid degradation can be overcome by the incorporation of chemical inhibitors into the fibrin gels to delay enzymatic degradation (Ahmed et al. 2008).

1.2.3 COLLAGEN Collagens and similarly sourced gelatins are natural proteins found abundantly in organisms to strengthen and support structural units within the body. While at least 15 types of collagen exist, types I, II, and III constitute at least 80% of the collagen in humans and are fibrillar in structure. Type I collagen is the most abundant and is found in most connective tissues, including bone. Type II is found largely in cartilage, while type III collagen is present in extensible tissues (e.g., lungs, skin, vascular system). Type IV collagen is less common but can be found in basement membranes. Collagens of type V through XV are found in smaller amounts in skin and in various connective and non-connective tissues. The prevalence of collagens

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throughout the body incurs an overall weak antigenicity, making collagen one of the most appropriate biomaterials of choice in terms of mitigating an immunological response. 1.2.3.1 Structure The fundamental structural unit of collagen is called tropocollagen: a 300-nm long, 1- to 5-nm diameter protein that consists of three subunits arranged in a triple helical structure (Ramakrishna et al. 2001). In the case of the fibrillar collagens (types I, II, III, V), tropocollagen molecules are arranged side-by-side to form fibrils (and in some cases ultrastructure fibers) that can be subsequently stabilized by crosslinking lysine or hydroxylysine. Other types of collagen, like type IV, do not form tightly bound fibrillar networks. This type arranges fibers head-to-head and lacks a regular glycine, thus preventing a collagen helix from forming and resulting in a thin formation of disorganized kinks. Type IV collagen is primarily found in the basal lamina, a thin amorphous membrane, the purposes of which include support and facilitating filtration. The distance between two helices within a fibril is approximately 67  nm, and the diameter of the fibrils ranges from 50 to 200  nm (Cicchi et al. 2013). Aside from its abundance within the body, the simplistic methods of blending collagen with other natural materials—to alter mechanical and physical properties— make collagen one of the more highly developed biomaterials for cell delivery (Li et al. 2013). The subunits of type I collagen, the most commonly employed form of collagen, are two α1 chains and one α2 chain. Each chain contains 1,050 amino acids that are wound around one another in a right-handed triple helix. These α chains are made of a repeating 3 amino acid motif, with the most common sequence being a glycine, proline, and hydroxyproline (Gly-Pro-Hyp) (Ramakrishna et  al. 2001). The internally positioned glycine residues permit tight packing of the chains, which along with hydrogen bonding between different α chains, stabilize the triple helix and prevent degradation by chemical or enzymatic means. Individual collagen fibrils arrange together to form collagen fiber bundles with a diameter varying between 0.5 and 3 mm in the skin or in the lamellae of the cornea. The organization of fiber bundles into a dense network at a larger scale creates the architectural substrate of the connective tissue, which affects the mechanical properties of the tissue as well as the physiological conditions of the overlying cellular epithelium (Brodsky and Persikov 2005). 1.2.3.2 Fabrication The nature of the crosslinked collagen in connective tissue causes it to dissolve very slowly, even in boiling water. Nevertheless, collagen extraction from tissue sources can be achieved by partial hydrolysis with a dilute acid or base, cleaving the crosslinks but keeping the collagen chains intact (Schrieber and Gareis 2007; Schmidt et al. 2016). In acidic extraction, cleavage of noncovalent inter- and intramolecular bonds occurs as the solution penetrates the structure and swells. This is more suitable for extracting collagens from fragile tissues with less intertwined collagen fibers. In contrast, the alkaline process, typically involving sodium hydroxide, is

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commonly used for extracting collagens from thicker tissue samples over periods of days to weeks. It should be noted that the extracted collagen is inherently not tropocollagen because the initial cleavage introduces opportunities for further cleavage and ultimately more degradation of the triple-helix conformation, thus dysregulating the molecular structure. Inhibiting endogenous proteases are currently the most commonly applied method of preventing collagen denaturation, but there have also been attempts to retain the tropocollagen structure by means of pretreating the source material. The introduction of more carboxyl groups with acid chlorides, anhydrides or chloroanhydrides have been employed previously to improve acid and denaturant infiltration while retaining the triple-helix structure of the tropocollagen (Parenteau-Bareil et al. 2010; Xu et al. 2017). Crosslinking post-fabrication is generally necessary to achieve relatively stable collagen scaffolds. The extent of this crosslinking (and subsequent degradability of the resultant scaffolds) can be controlled with various crosslinking agents such as glutaraldehyde or formaldehyde (Oliver et al. 1980). One attractive feature of collagen, like many biomaterials discussed here, is that it can be used for cellular encapsulation because the biomaterial self-assembles from liquid monomers to a solid polymer meshwork. This gelation usually occurs when collagen solutions are switched from an acidic pH at a low temperature to a neutral pH at body temperature (Schmidt et al. 2016). However, gelation can also be affected by ionic strength and light. In practice, gelation is relatively easy to initiate but may diminish the viability of cells that are sensitive to environmental changes. 1.2.3.3 Application In practice, collagen has been fabricated into sponges, microparticles, films, and viscoelastic hydrogels. Collagen sponges are commonly used in the topical treatment of burn wounds and ulcers (Ruszczak 2003) and are one of the more popular biomaterials for in vitro model systems. The main advantages of collagen sponges for these applications are their ability to absorb tissue exudate, tightly stick to wet wound surfaces, preserve moisture in the wound, protect the wound from deleterious mechanical exposure and prevent secondary bacterial infections. Furthermore, collagen favors infiltration and growth of cells responsible for tissue regeneration as well as infiltration of inflammatory cells (Chattopadhyay and Raines 2014). Collagen sponges are prepared by lyophilization of aqueous solutions of the acid- or baseextracted forms. Porosity of lyophilized sponges varies based on the concentration of collagen, the rate of freezing before lyophilization, as well as the concentration of solid or alternate materials within the construct. In testing, collagen gels exhibit tailorable mechanical properties with elastic moduli ranging from 5 to almost 25 kPa, along with compressive mechanical failures up to 8 kPa; these properties can be easily tuned based on factors like collagen concentration and the pH of solution used (Roeder et  al. 2002; Yamamura et  al. 2007). Collagen films have exhibited even greater mechanical strength when further crosslinked with additives such as glutaraldehyde or increasing molecular crosslinking via UV irradiation. Films have remained intact while undergoing tensile stresses over 40 MPa after glutaraldehyde crosslinking (Koide et al. 1997).

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Collagen sponges have also been combined with growth factors and proteins for repair of bone tissue injuries. These bioactive materials have induced differentiation of mesenchymal stem cells into chondrogenic and osteogenic cells, as well as favoring osteoblast proliferation (Takahashi et al. 2005). Further, preclinical and clinical trials of recombinant human bone morphogenetic protein-2 incorporated into collagen sponges indicate the high efficiency of this morphogen in the repair of criticalsize bone defects, spinal fusions, and bone fractures as well as its application in dental practice (Osidak et al. 2014). It should be noted that while collagen biomaterials have been deemed relatively biocompatible in most cases that call for collagen, it is still associated with an inflammatory response in some cases. This is largely assumed to be a consequence of the small fraction of non-collagenous proteins or denatured collagen proteins, but can often be minimized through crosslinking with glutaraldehyde. Collagen microparticles are commonly generated from water-in-oil emulsions. As with most other polymer constructs, crosslinking improves stability and durability of the particles but can inhibit immunogenicity. To improve retention in the body, collagen microparticles can be combined with other polymer materials such as poly(lactic-co-glycolic) acid and polycaprolactone (Schlapp and Friess 2003), but these products have not reached clinical application. Previous studies have successfully encapsulated human mesenchymal stem cells using collagen, but more often than not these delivery systems require a composite consisting of collagen and another material, or do not meet the mechanical requirements for the system of interest, both due to the difficulties of balancing mechanical properties for delivery with cell survival (Batorsky et al. 2005). Collagen has predominantly been utilized in the form of thin collagen hydrogels or dried into films that are rehydrated prior to use. Some of the earliest FDAapproved collagen hydrogels include Apligraf TM, one of the earliest bioengineered skin substitutes consisting of a bilayered collagen hydrogel with keratinocytes on one side and fibroblasts on the other (Chattopadhyay and Raines 2014). Similar products were conceived with intrinsic antimicrobial properties. Fortaderm®, for example, is a collagen-based polyhexamethylene biguanide (PHMB) antimicrobial wound dressing that gained FDA approval in 2001. An attractive area of research is for a flexible and translucent biomaterial for corneal implants. Here, collagen films can reproduce the corneal shape, but require crosslinking to improve mechanical strength (Mi et al. 2010). Among other uses, collagen hydrogels have been employed as a substrate for vascular implants and seeded with endothelial cells. Collagen sponges have been loaded with chondrocytes as well as alveolar osteoblasts, in separate studies, to replace damaged cartilage and bone, respectively (Fujisato et  al. 1996; Xiao et  al. 2003; Yamamura et al. 2007). In summary, collagen is an attractive biomaterial because of its biocompatibility and broad availability. While it is well recognized that collagen gels have poor tensile mechanical properties, are too weak for surgical manipulations, and are susceptible to rapid enzymatic degradation (Hoffman 2012), these issues can often be overcome by crosslinking methodologies as well as by combining collagen with other biomaterials.

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1.3 POLYSACCHARIDE BIOMATERIALS Polysaccharides (literally chains of sugar molecules) constitute the single largest fraction of biomass worldwide. These polymers include cellulose, alginates (from algae and seaweed), chitin (the exoskeleton of insects and crustaceans), as well as the glycosaminoglycans (GAGs) found in animal tissues. In general, polysaccharides are well tolerated in vivo, but vary extensively in their properties. Celluloses are highly crystalline and nondegradable in vivo, while GAGs are amorphous and readily degradable. Polysaccharide biomaterials are generally not as bioactive in terms of cellular or matrix signaling, compared with collagens and keratins. However, there are benefits to their simplicity in that the mechanical and biologic properties found in these biomaterials are the same or relatively well conserved, whether they are used on a micro- or a macro-scale, resulting in a predictable or more tailorable product. For example, alginate biomaterials are composed of what are termed M and G residues. They produce tailorable material properties due to the fact that G residues are readily associated with crosslinking; thus, a stiffer material can predictably be fabricated by simply increasing the G residue fraction. Similarly, the basic units of alginate are relatively bioinert and can form a gel in the presence of divalent cations, such as calcium. As long as the concentration of cations involved does not produce a cytotoxic effect, alginate gels produce minimal toxicities whether they are encapsulating a single cell or are acting as a bulk hydrogel encapsulating a large cell payload. In the following sections, a few of the most common polysaccharide materials for cell encapsulation and tissue regeneration, and their uses, are described.

1.3.1

ALGINATES

Alginate, a natural polysaccharide extracted from seaweed and other algae, is one of the most common cell and drug delivery vehicles. The ease with which nontoxic and nonimmunogenic gel can be formed has resulted in some of the more successful studies in terms of cell encapsulation and delivery, while its broad availability makes it one of the most cost-effective biomaterials employed today. 1.3.1.1 Structure Alginate describes a whole family of linear copolymers containing blocks of (1,4)-linked β-d-mannuronate (M) and α-l-guluronate (G) residues. The blocks are composed of consecutive G residues, consecutive M residues or alternating M and G residues. The length and order of these blocks are found to be dependent on the source. More than 200 different alginates are currently being manufactured (Tonnesen and Karlsen 2002). It is believed that the G-blocks are solely responsible for in intermolecular crosslinking via divalent cations (e.g., Ca2+) to form hydrogels. The composition of M and G block ratios, their sequence, the G-block length, and molecular weight are critical factors affecting the physical properties of alginate and its resultant hydrogels (Cao and Mooney 2007), with higher gel stiffness associated with increased G-block content and polysaccharide molecular weight (Lee and Mooney 2012). Different alginate sources offer a range of chemical structures. For example, bacterial alginates with high concentrations of G-blocks form gels with

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relatively high stiffness. These physical properties, in turn, control the stability of the gels, the rate of drug release from gels, and the phenotype and function of cells encapsulated within the gels (Hay et al. 2010). 1.3.1.2 Fabrication Alginate biomaterials have been produced in the form of hydrogels, microspheres, foams, films, and sponges. Alginates are used in commercial products and clinical evaluations as cell carriers and drug delivery vehicles. Primarily used as a gel, the characteristics depend on the alginate type, concentration, and crosslink density. The concentration of alginate in solution, which can reach as high as tens of milligrams per milliliter, greatly determines the type and size of resulting pores (Malafaya et  al. 2007). Increased solution concentration and crosslinking density results in fewer or smaller pores and an overall decrease in permeability. This is relevant for the diffusion of solutes, nutrients, oxygen and soluble factors. Concentration and crosslinking of alginate also determine the mechanical strength and therefore the stability of the material encapsulating the cells. Alginate gels have previously withstood tensile strains of up to 70  kPa before failure (Drury et al. 2004). It is still a challenge to effectively fabricate cell-encapsulating alginate constructs because these largely form contiguous gels that sufficiently hold cells in place, but limit the exchange of oxygen, nutrients, and wastes between the encapsulated cells and the surrounding environment (Hay et al. 2010). This is especially difficult due to the fact that alginate is inherently nondegradable in mammals, given that they lack the necessary enzymes to cleave the polymer chains. However, ironically, crosslinked alginate gels can dissolve through the release of their divalent ions into the surrounding microenvironment. Nevertheless, the average molecular weights of many commercially available alginates are higher than the renal clearance threshold of the kidneys. Therefore, despite the ability of alginate hydrogels to dissolve in vivo, the alginate molecules will likely not be completely removed from the body (Augst et al. 2006). While sufficient in cellular encapsulation, it should be noted that alginates do not possess cell adhesion sites (Bidarra et al. 2014). To circumvent this, alginates have been combined with proteins such as collagen or fibronectin. Alternatively, short amino acid sequences found in the ECM, rather than full-length proteins, can be employed. 1.3.1.3 Application Microencapsulation of cells within alginate hydrogels is an attractive approach because this material can provide a uniquely protective shell for cells and act as a controlled release system for growth factors or drugs. When exposed to divalent cations (e.g., Ca2+) at molar ratios of approximately 0.5, alginate solutions quickly crosslink to form hydrogels under mild and physiologic conditions and can be readily injected into tissues with minimal invasion (Gu et  al. 2004; Serra et  al. 2011). Consequently, previous and current research supports the idea that a mixture of growth factors contained in alginate gels can promote maintenance of stem cell characteristics in cells seeded on or encapsulated within alginate hydrogels (Gu et al. 2004).

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Prior in vitro studies have evaluated the key cellular functions associated with wound healing, tissue regeneration, cellular migration, proliferation and differentiation within alginate gels. In vivo studies further examined whether calcium crosslinked alginate systems loaded with autologous chondrocytes could stimulate cartilage formation when subcutaneously implanted into a porcine model (Ren et al. 2015). Cell encapsulation in alginate has thus far proven to be one of the more promising methodologies for sustained cell viability. When delivering transplanted pancreatic islets, alginate provides sustainable shielding from immune rejection and forms a matrix to increase the insulin secretion of islet progenitor cells, a potential treatment for diabetes. Since the islets used to treat one patient with diabetes are usually acquired from two or more donors, widespread access to direct delivery of islet cells would require an approximately 100-fold increase in tissue availability (Ricordi 2003). The concept of immune protection via the alginate encapsulation method becomes more relevant when considering the use of insulin-producing cells from stem cells or animal donors (i.e., xenotransplant). In other disease areas, previous studies have succeeded in the delivery of alginate-encapsulated mesenchymal stem cells that secrete vascular endothelial growth factor to treat peripheral arterial disease (Tonnesen and Karlsen 2002). The concept of maintaining cell viability for an extended period therefore makes cell encapsulation a highly sought-after method of therapy.

1.3.2

CHITOSAN

Chitosan, an amino-polysaccharide obtained by alkaline deacetylation of chitin, is a natural and abundant polymer that is considered an attractive candidate material for tissue engineering. Chitosan-based scaffolds have been shown to be biodegradable, nonimmunogenic and biocompatible, and thus are used as potential therapeutic scaffolds for tissue engineering processes such as cell encapsulation and cell culture (Di Martino et al. 2005). 1.3.2.1 Structure Depending on its source, chitin occurs as two allomorphs, namely the α and β forms. A third allomorph, γ-chitin, also exists, but is believed to simply be a variant of the α family. The most abundant allomorph is α-chitin, existing in fungal and yeast cell walls, krill, lobster and crab tendons and shells, as well as in insect cuticle. These α-chitins have proven to be particularly interesting because many of them present remarkably high crystallinity and high purity (Rinaudo 2006). The rarer β-chitin is found in association with proteins in squid pens as well as in the lorica built by some seaweeds or protozoa. To date, β-chitin is less commonly used because it is not as readily available and tends to be more unstable compared with α-chitin (Jang et al. 2004; Revol and Chanzy 1986; Rudall 1969). Chitosan is a linear polysaccharide composed of glucosamine linked to N-acetyl glucosamine, the ratio of the two being referred to as the degree of deacetylation. Depending on the source and preparation procedure, its molecular weight may range from 300 to over 1,000 kD, with a respective degree of deacetylation of 30 to 95%. Chitosan is normally insoluble in aqueous solutions above pH 7; however, it is soluble

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at dilute concentrations under acid conditions (pH