Biointegration of Medical Implant Materials: Science and Design

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Author: C. P. Sharma

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Biointegration of medical implant materials

© Woodhead Publishing Limited, 2010

Related titles: Cellular response to biomaterials (ISBN 978-1-84569-358-9) The response of cells to biomaterials is critical in medical devices. It has been realised that specific cell responses may be beneficial – encouraging adhesion, healing or cell multiplication. Cellular response to biomaterials will examine the response of cells with a wide range of materials, targeted at specific medical applications. The first section reviews cellular response to polymers and ceramics. A second group of chapters discuss cell responses and regenerative medicine for nerves, muscles and orthopaedic materials. The final chapters analyse the effect of surface chemistry and how it can be manipulated to provoke a useful cell response. Surfaces and interfaces for biomaterials (ISBN 978-1-85573-930-7) This book presents the current level of understanding on the nature of a biomaterial surface, the adaptive response of the biomatrix to that surface, techniques used to modify biocompatibility, and state-of-the-art characterisation techniques to follow the interfacial events at that surface. Tissue engineering using ceramics and polymers (ISBN 978-1-84569-176-9) Tissue engineering is rapidly developing as a technique for the repair and regeneration of diseased tissue in the body. This authoritative and wide-ranging book reviews how ceramic and polymeric biomaterials are being used in tissue engineering. The first part of the book reviews the nature of ceramics and polymers as biomaterials together with techniques for using them such as building tissue scaffolds, transplantation techniques, surface modification and ways of combining tissue engineering with drug delivery and biosensor systems. The second part of the book discusses the regeneration of particular types of tissue from bone, cardiac and intervertebral disc tissue to skin, liver, kidney and lung tissue. Details of these and other Woodhead Publishing materials books can be obtained by: • •

visiting our web site at www.woodheadpublishing.com contacting Customer Services (e-mail: [email protected]; fax: +44 (0) 1223 893694; tel.: +44 (0) 1223 891358 ext. 130; address: Woodhead Publishing Limited, Abington Hall, Granta Park, Great Abington, Cambridge CB21 6AH, UK)

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Biointegration of medical implant materials Science and design Edited by Chandra P. Sharma

Oxford

Cambridge


New Delhi

Published by Woodhead Publishing Limited, Abington Hall, Granta Park, Great Abington, Cambridge CB21 6AH, UK www.woodheadpublishing.com Woodhead Publishing India Private Limited, G-2, Vardaan House, 7/28 Ansari Road, Daryaganj, New Delhi – 110002, India www.woodheadpublishingindia.com Published in North America by CRC Press LLC, 6000 Broken Sound Parkway, NW, Suite 300, Boca Raton, FL 33487, USA First published 2010, Woodhead Publishing Limited and CRC Press LLC © Woodhead Publishing Limited, 2010 The authors have asserted their moral rights. This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. Reasonable efforts have been made to publish reliable data and information, but the authors and the publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from Woodhead Publishing Limited. The consent of Woodhead Publishing Limited does not extend to copying for general distribution, for promotion, for creating new works, or for resale. Specific permission must be obtained in writing from Woodhead Publishing Limited for such copying. Trademark notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation, without intent to infringe. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library. Library of Congress Cataloging in Publication Data A catalog record for this book is available from the Library of Congress. Woodhead Publishing ISBN 978-1-84569-509-5 (book) Woodhead Publishing ISBN 978-1-84569-980-2 (e-book) CRC Press ISBN 978-1-4398-3064-2 CRC Press order number: N10181 The publishers’ policy is to use permanent paper from mills that operate a sustainable forestry policy, and which has been manufactured from pulp which is processed using acid-free and elemental chlorine-free practices. Furthermore, the publishers ensure that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by Toppan Best-set Premedia Limited, Hong Kong Printed by TJ International Limited, Padstow, Cornwall, UK


Contents

Contributor contact details Preface

xi xv

1

Biointegration: an introduction C. K. S. Pillai and C. P. Sharma, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

1

1.1 1.2 1.3 1.4 1.5

Introduction Biointegration of biomaterials for orthopedics Biointegration of biomaterials for dental applications AlphaCor artificial corneal experience Biointegration and functionality of tissue engineering devices Percutaneous devices Future trends References

1 1 7 8

1.6 1.7 1.8

Part I Soft tissue biointegration 2

2.1 2.2 2.3 2.4 2.5 2.6 2.7 2.8

Biocompatibility of engineered soft tissue created by stem cells P. A. Clark, University of Wisconsin–Madison, USA; and J. J. Mao, Columbia University, USA Introduction Bone: from tissue to molecular organization Bone development Bone homeostasis Bone repair after injury Bone and joint disease Current treatment options and total joint replacements Current challenges of titanium implants

10 10 11 12

17

19

19 20 22 24 26 29 29 30 v


vi

Contents

2.9 2.10

Current titanium modifications for improved integration Mimicking nature toward achieving titanium ‘biointegration’: cytokines and implants Growth factor delivery: why is controlled and sustained release important? Future trends Acknowledgements Sources of further information and advice References

2.11 2.12 2.13 2.14 2.15 3

3.1 3.2 3.3 3.4

Replacement materials for facial reconstruction at the soft tissue–bone interface E. Wentrup-Byrne, Queensland University of Technology, Australia; L. Grøndahl and A. Chandler-Temple, The University of Queensland, Australia

32 34 35 37 39 39 40

51

Introduction Facial reconstruction Materials used in traditional interfacial repair Surface modification of facial membranes for optimal biointegration Future trends Acknowledgements References

51 55 60

4

Corneal tissue engineering Y.-X. Huang, Ji Nan University, China

86

4.1 4.2 4.3

Introduction Characteristics of the human cornea and its regeneration Special conditions for wound healing and tissue regeneration of the cornea Approaches to corneal tissue engineering Future trends References

86 87

3.5 3.6 3.7

4.4 4.5 4.6

73 78 78 79

90 95 108 109

5

Tissue engineering for small-diameter vascular grafts J. I. Rotmans, Leiden University Medical Centre, The Netherlands; and J. H. Campbell, University of Queensland, Australia

116

5.1 5.2

Introduction Required characteristics of tissue engineered blood vessels

116 118


Contents

vii

5.3 5.4 5.5 5.6

Approaches to vascular tissue engineering Future trends Conclusion References

120 136 138 138

6

Stem cells for organ regeneration K. D. Deb, Dayananda Sagar Institutions, India

147

6.1 6.2 6.3

Introduction Basic components of tissue engineering Tissue engineering and stem cells in organ regeneration Conclusions References

147 149

6.4 6.5

Part II Drug delivery 7

7.1 7.2 7.3 7.4

Materials facilitating protein drug delivery and vascularisation P. Martens, A. Nilasaroya and L. A. Poole-Warren, University of New South Wales, Australia

161 169 169

177

179

Introduction Hydrogel classification Factors influencing protein encapsulation and release Tissue engineering applications: vascularisation and protein delivery Conclusions Acknowledgements References

179 181 185

8

Inorganic nanoparticles for targeted drug delivery W. Paul and C. P. Sharma, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

204

8.1 8.2 8.3 8.4 8.5 8.6 8.7

Introduction Calcium phosphate nanoparticles Gold nanoparticles Iron oxide nanoparticles Conclusion Acknowledgements References

204 206 213 218 226 226 227

7.5 7.6 7.7


191 197 198 198

viii

Contents

9

Alginate-based drug delivery devices L. Grøndahl, G. Lawrie and A. Jejurikar, The University of Queensland, Australia

236

9.1 9.2 9.3 9.4 9.5 9.6

Introduction Alginate biopolymers Drug delivery using alginate matrices Future trends Acknowledgement References

236 237 247 258 259 259

10

Functionalised nanoparticles for targeted drug delivery S. Manju and K. Sreenivasan, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

10.1 10.2 10.3 10.4 10.5 10.6

Introduction Drug targeting Multifunctional nanocarrier systems Conclusion Acknowledgements References

Part III Design considerations 11

11.1 11.2 11.3 11.4

11.5 11.6 12

12.1 12.2

Biocompatibility of materials and its relevance to drug delivery and tissue engineering T. Chandy, 3M Drug Delivery Systems, USA Biocompatibility of materials and medical applications Biomaterials for controlled drug delivery Biomaterials for tissue engineering Role of the scaffold and loaded drug/growth factor in the integration of extracellular matrix and cells at the interface Future trends References Mechanisms of failure of medical implants during long-term use A. Kashi and S. Saha, SUNY Downstate Medical Center, USA Introduction Manufacturing deficiencies


267

267 269 282 289 289 290

299

301 301 307 311

315 320 321

326

326 327

Contents

ix

12.3 12.4 12.5 12.6 12.7 12.8 12.9 12.10

Mechanical factors (e.g. fatigue, overloading) Wear Corrosion Clinical factors for implant success and failure Failure mechanisms of non-load-bearing implants Failure analysis of medical implants Conclusion References

328 329 333 334 335 337 340 342

13

Rapid prototyping in biomedical engineering: structural intricacies of biological materials S. J. Kalita, University of North Dakota, USA

349

13.1 13.2 13.3 13.4 13.5 13.6 13.7 13.8 13.9 13.10 13.11

Introduction An overview of biomaterials Material properties of structural biomaterials Rapid prototyping – a novel manufacturing approach Designing structural implants Rapid prototyping in biomedical engineering – synopsis Rapid prototyping in mimicking structural intricacies of biological materials Patient-specific customized scaffolds via rapid prototyping Conclusion List of abbreviations References

377 387 387 389 390

Index

399


349 354 357 360 369 373

Contributor contact details

(* = main contact)

Chapter 2

Editor C. P. Sharma Division of Biosurface Technology Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences and Technology Poojappura Thiruvananthapuram 695 012 India Email: [email protected]

Chapter 1 C. K. S. Pillai and C. P. Sharma* Division of Biosurface Technology Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences and Technology Poojappura Thiruvananthapuram 695 012 India Email: [email protected]

P. A. Clark University of Wisconsin–Madison UW-Hospitals and Clinics Department of Neurological Surgery CSC K4/879 600 Highland Ave. Madison, WI 53792 USA Email: [email protected] J. J. Mao* Columbia University College of Dental Medicine Fu Foundation School of Engineering and Applied Sciences Department of Biomedical Engineering 630 W. 168 St. – PH7 East CDM New York, NY 10032 USA Email: [email protected]

xi © Woodhead Publishing Limited, 2010

xii


Chapter 3 E. Wentrup-Byrne Visiting Fellow Tissue Repair and Regeneration Program Institute of Health and Biomedical Innovation Queensland University of Technology 60 Musk Avenue, Kelvin Grove Brisbane, QLD 4059 Australia Email: [email protected]

Chapter 4 Y.-X. Huang Institute of Biomedical Engineering Ji Nan University Guang Zhou China 510632 Email: [email protected]

J. H. Campbell Australian Institute for Bioengineering and Nanotechnology Corner College and Cooper Rds (Bldg 75) The University of Queensland Brisbane, QLD 4072 Australia Email: [email protected]

Chapter 6 K. D. Deb Tissue Engineering and Regenerative Medicine Lab Department of Biotechnology Dr C. D. Sagar Center for Life Sciences Dayananda Sagar Institutions Shavige Malleswara Hills Kümarawamy Layout Bangalore 560078 India Email: [email protected], [email protected]

Chapter 5 J. I. Rotmans* Leiden University Medical Centre Department of Nephrology, C3-P Albinusdreef 2, 2333 ZA Leiden The Netherlands Email: [email protected]

Chapter 7 P. Martens*, A. Nilasaroya and L. A. Poole-Warren Graduate School of Biomedical Engineering University of New South Wales Sydney, NSW 2052 Australia Email: [email protected], [email protected]



xiii

Chapter 8

Chapter 11

W. Paul and C. P. Sharma* Division of Biosurface Technology Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences and Technology Poojappura Thiruvananthapuram 695 012 India

T. Chandy 3M Drug Delivery Systems Division 3M Center Building 260-03-A-06 St. Paul, MN 551344-1000 USA

Email: [email protected]

Chapter 12

Chapter 9

A. Kashi* SUNY Downstate Medical Center 450 Clarkson Avenue, Box 30 Brooklyn, NY 11203 USA

L. Grøndahl*, G. Lawrie and A. Jejurikar School of Chemistry and Molecular Biosciences The University of Queensland Cooper Rd, Brisbane, QLD 4072 Australia Email: [email protected], [email protected]

Chapter 10 S. Manju and K. Sreenivasan* Laboratory for Polymer Analysis Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences and Technology Poojappura Thiruvananthaparam 695 012 India Email: [email protected]

Email: [email protected]

Email: [email protected] S. Saha Department of Orthopaedic Surgery and Rehabilitation Medicine SUNY Downstate Medical Center 450 Clarkson Avenue, Box 30 Brooklyn, NY 11203 USA Email: [email protected]

Chapter 13 S. J. Kalita Engineered Surfaces Center School of Engineering and Mines University of North Dakota 4201 James Ray Drive, Suite 1100; Stop 8391 Grand Forks, ND 58202 USA Email: [email protected]. edu


Preface

The aim of this book is to enhance our understanding of the interfacial interaction and integration of implant materials with hard/soft tissue. The ultimate success of any implant inside the body is not only that it should be non-toxic and biocompatible with respect to its physico-chemical properties, including degraded products if any, for a desired application, but also that it integrates with the tissue as per the physiological requirements. Such issues have been discussed appropriately by experts in three separate sections on soft tissue biointegration, drug delivery and design considerations. The book will certainly be useful for academic faculty graduate students and the medical devices industry interested in understanding the concepts useful for enhancing the quality of their products. I thank all the authors who contributed the chapters in this book and Ms Lucy Cornwell for coordinating the communication link among all of us. I would also like to thank our former Director Prof. K. Mohandas, Dr K. Radhakrishnan Director SCTIMST and Dr G. S. Bhuvaneshwar Head BMT Wing SCTIMST Trivandrum for providing facilities to complete this project. Chandra P. Sharma

xv © Woodhead Publishing Limited, 2010

1 Biointegration: an introduction C. K. S. P I L L A I and C. P. S H A R M A, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

Abstract: This introductory chapter discusses the relevance of the basic aim of the book, reflecting on the importance of the interfacial concepts related to the biointegration of the implants with respect to their design and structure–function relationship. Various examples, including orthopedic implants, are discussed. Key words: biointegration, interface, medical implants, design, biomaterials.

1.1

Introduction

When biomaterials are designed, a set of properties are built in such a way as to ensure that, after implantation, they will help the body to heal itself. So it is of critical importance that these materials be integrated into organspecific repair mechanisms such as the physiologic process required for the biologization of implants (Amling et al., 2006). It should involve a direct structural and functionally stable connection between the living part and the surface of an implant. Although various materials have been developed in recent years with enhanced physical, surface and mechanical properties, the use of these materials in certain biological applications is often limited by poor tissue integration. So, the question is on how biomaterials can be converted to ‘living tissues’ after implantation. To cite an example, the bonding of hydroxyapatite to bone, which is considered as a true case of biointegration, is thought to involve a direct biochemical bond of the bone to the surface of an implant at the electron microscopic level and is independent of any mechanical interlocking mechanism (Meffert et al., 1987; Cochran, 1996). Several groups working on various aspects of the design, development and application of improved devices are concerned with how these materials become integrated with soft and hard tissues in the body and how these implanted systems have to match their physical–chemical and biological properties to those of their environment.

1.2

Biointegration of biomaterials for orthopedics

Biomaterials are defined as ‘materials intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ, or function 1 © Woodhead Publishing Limited, 2010

2


of the body’ (Williams, 1999). Orthopedic research is developing and advancing at a rapid pace as new techniques are applied to musculoskeletal tissues. The discovery of biologic solutions to important problems, such as fracture-healing, soft-tissue repair, osteoporosis, and osteoarthritis, continues to be an important research focus. At the same time, research on biomaterials and biomechanics is critical for advances in current areas such as tissue-engineering and cytokine delivery. In orthopedic applications, there is a significant need and demand for the development of a bone substitute that is bioactive and exhibits material properties (mechanical and surface) comparable with those of natural, healthy bone. Particularly, in bone tissue engineering, nanometer-sized ceramics, polymers, metals and composites have been receiving much attention recently. This is a result of current conventional materials not invoking suitable cellular responses to promote adequate osteointegration to enable implanted devices to be successful for long periods (Balasundaram and Webster, 2006; Barrère et al., 2008). Metallic materials are normally used for load-bearing members such as pins and plates, femoral stems, etc. Ceramics, such as alumina and zirconia, are used for wear applications in joint replacements, while hydroxyapatite is used for bone bonding applications to assist implant integration. Polymers, such as ultra high molecular weight polyethylene (UHMWP), are used as articulating surfaces against ceramic components in joint replacements. Porous alumina has also been used as a bone spacer to replace large sections of bone which have had to be removed due to disease (www.azom, 2004; http://academic.uprm.edu). In applications involving the loading phase, the best material has been titanium and its alloys, whereas calcium phosphate seems to be the best material to be used in joint replacement and osseointegration (the degree to which bone will grow next to or integrate into the implant). Titanium is used primarily for the loading faces, which include the pin structure, and the fabrication of plates and femoral stems. The integration of a biomaterial to bone involves, essentially, two processes: interlocking with bone tissue and chemical interactions with bone constituents. The direct bonding of orthopedic biomaterials with collagen is rarely considered; however, several non-collagenic proteins have been shown to adhere to biomaterial surfaces (Rey, 1998). Many studies have been reported on the biointegration of orthopedic devices. Hydroxyapatite (HA) films have been widely recognized for their biocompatibility and utility in promoting biointegration of implants in both osseous and soft tissue. In a study on hydroxyapatite-coated (by electroplating) cp-titanium implants, Badr and El Hadary (2007) showed the formation of recognizable osseointegration of bone regeneration with more and denser bone trabeculae, and concluded that electroplating provided a thin


Biointegration: an introduction

3

and uniform pure crystalline hydroxyapatite coating. The characterization of the precipitated film is promising for clinically successful long-term bone fixation. An AFM analysis of roughness on seven materials widely used in bone reconstruction carried out by Covani et al. showed that the biointegration properties of bioactive glasses can also give an answer in terms of surface structures in which chemical composition can influence directly the biological system (e.g. with chemical exchanges and development of specific surface electrical charge) and indirectly, via the properties induced on tribological behavior that expresses itself during the smoothing of the surfaces (Covani et al., 2007). The biological behavior of an implant, such as osseointegration, depends on both the chemical composition and the morphology of the surface of the implant. Irradiation with laser light (Nd:YAG (λ = 1064 nm, τ = 100 ns)) is used for the surface modification of Ti-6Al-4V – which is widely used in implantation to enhance biointegration (Mirhosseini et al., 2007). Conventional sputtering techniques have shown some advantages over the commercially available plasma spraying method for generating hydroxyapatite (HA) films on metallic substrates; however, the as-sputtered films are usually amorphous, which can cause some serious adhesion problems when post-deposition heat treatment is necessitated. Nearly stoichiometric, highly crystalline HA films strongly bound to the substrate were obtained by an opposing radio frequency (RF) magnetron sputtering approach. HA films have been widely recognized for their biocompatibility and utility in promoting biointegration of implants in both osseous and soft tissue (Hong et al., 2007). Oudadesse et al. (2007) studied the in vitro behavior of compounds in contact with simulated body fluid (SBF) and in vivo experiments in a rabbit’s thigh bones. The inductively coupled plasma–optical emission spectroscopy (ICP-OES) method was used to study the eventual release of Al from composites to SBF and to evaluate the chemical stability of composites characterized by the succession of SiO4 and AlO4 tetrahedra. The results obtained show the chemical stability of composites. At the bone– implant interface, the intimate links revealed the high quality of the biointegration and the bioconsolidation between composites and bony matrix. Histological studies confirmed good bony bonding and highlighted the total absence of inflammation or fibrous tissues, indicating good biointegration (Oudadesse et al., 2007). Zanotti and Verlicchi (2006) proposed a bioglass– alumina spacer that could perform an excellent arthrodesis by a mechanical stabilization (primary one) and a biological bio-mimetic stabilization (biointeraction, biointegration and biostimulation). Other advantages are easy use, even in the low somatic interspaces (C7-D1 type), reduction of convalescence, and reduced costs of this type of device. That makes it interesting, even in societies with low economic well-being (Zanotti and Verlicchi, 2006).


4


The process by which these materials are integrated into the organspecific repair cascade is named ‘bone remodeling’, which is the concerted interplay of two cellular activities: osteoclastic bone resorption and osteoblastic bone formation. The latter physiologic process not only maintains bone mass, skeletal integrity and skeletal function but is also the cellular process that determines structural and functional integration of bone substitutes. A molecular understanding of this process is therefore of paramount importance for almost all aspects of the body’s reaction to biomaterials and might help us to understand, at least in part, the success or the failure of various materials used as bone substitutes. Recent genetic studies have demonstrated that there is no tight cross-control of bone formation and bone resorption in vivo and that there is also a central axis controlling bone remodeling, radically enhancing our understanding of this process. Amling et al. (2006) have shown how an understanding of bone remodeling – the physiologic process required for the biologization of bone substitutes – has evolved during recent years, providing a platform for the design, development and application of improved biomaterials. Roughness was evaluated by measuring root mean square (RMS) values and RMS/average height (AH) ratio, in different dimensional ranges, varying from 100 microns square to a few hundreds of nanometers. The results showed that titanium presented a lower roughness than the other materials analyzed, frequently reaching statistical significance (Covani et al., 2007). Conversely, bioactive materials such as hydroxyapatite (HA) and bioactive glasses have demonstrated an overall higher roughness. In particular, this study focuses attention on AP40 and especially RKKP, which proved to have a significant higher roughness at low dimensional ranges. This determines a large increase in surface area, which is strongly connected with osteoblast adhesion and growth, and also with protein absorption (Fig. 1.1). One should mention here the famous osseointegration concept evolved by Per Ingvar Brånemark, closely coupled with the design of a cylindrical titanium screw (Fig. 1.2) (Albrektsson et al., 1981; Brånemark et al., 1985) having a specific surface treatment to enhance its bioacceptance (Adell et al., 1990). The titanium screw (Fig. 1.2) underwent many animal and, subsequently, human clinical trials to test the success rate, the concept and the design of this implant. A fixture is osseointegrated if it provides a stable and apparently immobile support of a prosthesis under functional loads, without pain, inflammation, or loosening. Titanium’s ability to be integrated in the bone has been known for more than 25 years of experience and research that form the basis of the knowledge and use of implant technology today. Osseointegration of an implant is a direct structural and functionally stable connection


Biointegration: an introduction Titanium

Bone

Osseointegrated (a)

Titanium Connective tissue

5 Bone

Non-integrated (b)

1.1 (a) and (b) Biointegration of titanium showing a lower roughness (reproduced from Covani et al., 2007 with permission from Global Rights, John Wiley & Sons Inc.).

1.2 The titanium screw (reproduced with permission from Brånemark et al., Tissue-integrated Prostheses: Osseointegration in Clinical Dentistry, Quintessence Publishing Co., Chicago, 1985).


6


between the living bone and the surface of an implant that is exposed to mechanical load (http://www.dental-oracle.org). Specialized coatings are being developed for orthopedic implants, biomedical use, and other scientific applications. A HA coating for surface improvement gave rapid osseointegration and biointegration in four weeks with 90% of implant–bone contact at ten months in contrast to titanium alone, which required ten weeks to be osseointegrated with 50% implant– bone contact at ten months. A very recent material is osteopontin, an extracellular glycosylated bone phosphoprotein with a polypeptide backbone that makes dead metal ‘come alive’. Surrounding cells ‘don’t see an inert piece of metal, they see a protein and it’s a protein they know’. Other novel innovations include the development of nanostructure materials and diagnostic techniques for both in vitro and in vivo applications (Namavar et al., 2007). Typically, in regards to orthopedic devices, the primary concerns are wear, infection and failure of biointegration (Harris and Richards, 2006; Viceconti et al., 2004). The survival rate of an implant under optimal conditions is at least 96% after five years. Zim reports development of a high-porosity expanded polytetrafluoroethylene that has been fabricated to provide a softer feel with less shrinkage and migration because of better biointegration and cellular ingrowth (Zim, 2004). Long-term results with porous polyethylene have demonstrated superior biocompatibility and minimal complications. Hydroxyapatite cement has been associated with an immunoguided delayed inflammatory reaction that leads to thinning of the overlying skin and exposure of the implant. Applications of distraction osteogenesis are rapidly expanding and include deformities of the mandible, midface, and cranium. There has been a trend toward the use of internal hardware, and internal devices are being developed to deliver a greater degree of vector control. Biodegradable devices have been developed to eliminate the second surgical procedure necessary for hardware removal. In the future, successful tissue engineering could eliminate many of the drawbacks associated with implants and osteotomies. The ability to stimulate stem cells to generate autogenous bone has been demonstrated. Computer technology has been successfully used to integrate laser surface scanning and digitizing with computer-aided design and manufacturing to produce facial prostheses. Technologic advances in biomaterials, distraction hardware, computer modeling, and tissue engineering will continue to supply the surgeon’s repertoire with improved methods to augment and restore the craniomaxillofacial skeleton. Whenever metallic devices are implanted in vivo, successful biointegration requires that host cells colonize the highly reactive implant surface (Schmidt and Swiontkowski, 2008). Bacteria such as staphylococci can also become adherent to metallic or polymeric implants and will compete with



7

host cells for colonization of the implant surface. It has been demonstrated in animal models that contaminated fractures without internal fixation develop clinical infection more commonly than similar fractures treated with internal fixation at the time of colonization. Because of the potential for infection whenever internal fixation is utilized, appropriate prophylactic antibiotic coverage for staphylococci and Gram-negative organisms should be provided.

1.3

Biointegration of biomaterials for dental applications

The biomaterials community is producing new and improved implant materials and techniques to meet the growing demands for biomaterials in dental applications. The main property required of a biomaterial is that it does not illicit an adverse reaction when placed into service (www.azom.com). Thus, a variety of materials – metallic (pins for anchoring tooth implants and as parts of orthodontic devices), ceramics (tooth implants including alumina and dental porcelains) and polymeric (orthopedic devices such as plates and dentures) – with excellent biointegration have been developed and have been placed into service. Starting with the ‘Integral Biointegrated Dental Implant System’, consisting of a titanium implant cylinder coated with calcitite that permitted the bone to actually bond with the implant surface in a jaw restoration (Roling, 1989), dental surgery has undergone a revolution in both implant techniques and materials technology. It was shown early on that the surface oxide of titanium appears to be central to the ability of titanium implants to achieve osseointegration, and ceramic coatings appear to improve the ingrowth of bone and promote chemical integration of the implant with the bone (Wataha, 1996). Badr and El Hadary (2007) have reported development of osseointegration of the regenerated bone when hydroxyapatite was coated onto the surface of commercially pure titanium (cpTi) implants using an electroplating technique. Pelsoczi et al. obtained more effective osseointegration on surface modifications of titanium implants with an excimer laser. In this case, it was easy to achieve the desired morphology (microstructure) and physical-chemical properties that control the biointegration process (Pelsoczi et al., 2004). X-ray photoelectron spectroscopy (XPS) studies show that laser treatment, in addition to micro-structural and morphological modification, results in a decrease of surface contamination and thickening of the oxide layer. Thin film deposition of ceramic oxides onto titanium by excimer lasers and pulsed lasers has been successfully employed by other groups to improve the surface characteristics for facilitating biointegration, e.g. pulsed laser deposition of bioceramic thin films from human teeth (Smausz et al., 2004) and surface modifications induced


8


by ns and sub-ps excimer laser pulses on titanium implant material (Bereznai et al., 2003).

1.4

AlphaCor artificial corneal experience

AlphaCorTM is a biocompatible, flexible, one-piece artificial cornea (keratoprosthesis) designed to replace a scarred or diseased native cornea. It is a one piece convex disc consisting of a central transparent optic and an outer skirt that is entirely manufactured from poly(2-hydroxyethyl methacrylate) or PHEMA. AlphaCor’s material and patented features are designed to promote retention and optimize patient outcome. The outer skirt is an opaque, high water-content, PHEMA sponge. The porosity of the sponge encourages biointegration with host tissue and thus promotes retention of the implanted device. The central optic core is a transparent PHEMA gel, providing a refractive power similar to that of the human cornea. The optic core is designed to allow the patients’ visual potential to be achieved. The junctional zone between the skirt and central optic is the interpenetrating polymer network or IPN. This is a permanent bond formed at the molecular level and is designed to prevent the down-growth of cells around the optic, which can lead to the formation of retroprosthetic membranes, one of the major complications historically associated with artificial corneas (http://www.medcompare.com; http://www.pricevisiongroup.com). The World Health Organization (WHO) reports that corneal blindness affects more than 10 million people worldwide; however, only 100 000 people received corneal transplants each year. This shortfall is due to a combination of inadequate supply of donor corneas and the unsuitability of some patients to receive a corneal graft. AlphaCor is designed for use in patients who have had multiple failed corneal transplants or in those patients in whom a donor graft is likely to fail. Its patented design features are aimed to promote retention, minimize post-operative complications and restore vision in patients who cannot receive, or are unlikely to have, a beneficial outcome from a human donor graft. AlphaCor is available in two versions, to suit those with a natural lens (phakic) or artificial lens (pseudophakic) and for those without a lens (aphakic). Keratoprosthesis for artificial cornea surgery is a procedure for restoring the sight of patients suffering from a severely damaged anterior segment due to trauma, chemical burns, infections, etc. The ideal keratoprosthesis would be inert and not be rejected by the patient’s immune system, be inexpensive, and maintain long-term clarity. In addition, it would be quick to implant, easy to examine, and allow an excellent view of the retina. Coassin et al. (2007) reported histopathologic and immunologic characteristics of late artificial corneal failure in a small series of patients who



9

1.3 AlphaCorTM artificial cornea (reproduced with permission from M/s Addition Technology Inc, 950 Lee Street, Des Plaines, IL 60016, USA).

underwent AlphaCor implantation, but light microscopic examination of the specimens disclosed adequate biointegration with no foreign body response. Immunofluorescence studies of the skirt exhibited expression of inflammatory cytokines such as interleukin-1β (IL-1β) and tumor necrosis factor α (TNF-α), and some interferon γ (IFN-γ). The keratocytes stained positively for Thy-1 and smooth muscle actin, but negatively for CD34. The AlphaCor implant (Fig. 1.3) is a viable method of treatment for multiple failed PKPs, but it may be associated with unique complications, including corneal stromal melting, focal calcification, and retroprosthetic membrane formation. Infectious keratitis may be a risk factor for corneal stromal melting and needs to be managed aggressively. Explantation of the implant is essential if the skirt is exposed (Chow et al., 2007). Hicks et al. (2005) showed that histologic findings of the AlphaCor skirt in humans are consistent with earlier animal studies. Their study confirmed that biointegration by host fibroblastic cells, with collagen deposition, occurred after AlphaCor implantation in humans. In cases in which stromal melting had occurred, biointegration was seen to be reduced. On correlating preoperative clinical factors with biointegration observed histologically, preoperative vascularization appears not to be required for AlphaCor biointegration (Hicks et al., 2005). Another study demonstrated that systemic factors


10


affected the risk of retroprosthetic membrane formation with AlphaCor (Hicks and Hamilton, 2005). Hicks et al. (2005) further showed that device biointegration gets reduced in the cases of patients with a history of ocular herpes simplex virus (HSV) because the extensive lamellar corneal surgery involved in AlphaCor implantation may precipitate reactivation of latent HSV, such that reactivation and resultant inflammation could facilitate melting of corneal stromal tissue anterior to the device (Hicks et al. 2002). Chirila (2006), in a recent publication, claims that the first keratoprosthesis based on polyurethane was made in 1985 by Lawrence Hirst, an Australian ophthalmologist then working in St Louis, USA. This keratoprosthesis, which also had a porous skirt, was inserted intralamellarly in a monkey cornea and was followed up clinically for about three months. There were no significant postoperative complications, and the histology of the explant indicated proper biointegration of the prosthetic skirt within the host stromal tissue (Chirila, 2006; Chirila et al., 1998). Hydrogel lenses may even make their way deeper into the eye, as replacements for inner-eye lenses damaged by cataracts.

1.5

Biointegration and functionality of tissue engineering devices

Experiments on animals have underlined the importance of vascularization for biointegration and functionality of any given tissue engineering device. Polykandriotis et al. recently showed that the presence of a vascular bed prior to cell transplantation might protect against hypoxia-induced cellular death, especially at central portions of the matrix, and therefore ensure physiological function of the device. The generation of vascularized bioartificial tissue substitutes might offer new modalities of surgical reconstruction for use in reparative medicine (Polykandriotis et al., 2006).

1.6

Percutaneous devices

Percutaneous devices play an essential role in medicine; however, they are often associated with a significant risk of infection. One approach to circumvent infection would be to heal the wound around the devices by promoting skin cell attachment (Fukano et al., 2006). Biointegration through human fibronectin (FN) plays a key role in the biointegration of implants, as the success depends on adsorption of proteins like FN. Indeed FN can be an intermediary between the biomaterial surface and cells (Sousa et al., 2005). Isenhath et al. (2007) developed an in vivo model that permits examination of the implant/skin interface and that will be useful for future studies designed to facilitate skin cell



11

attachment where percutaneous devices penetrate the skin. The space created between the skin and the device becomes a haven for bacterial invasion and biofilm formation, and this results in infection. Sealing this space via integration of the skin into the device is expected to create a barrier against bacterial invasion. Porous poly(2-hydroxyethyl methacrylate) (PHEMA) rods were implanted for seven days in the dorsal skin of C57 BL/6 mice. The porous PHEMA rods were surface-modified with carbonyldiimidazole (CDI) or CDI plus laminin 5, with unmodified rods serving as control. Implant sites were sealed with 2-octyl cyanoacrylate; corn pads and adhesive dressings were tested for stabilization of implants. All rods remained intact for the duration of the study. There was histological evidence of both epidermal and dermal integration into all PHEMA rods, regardless of treatment. The effects of hyperbaric oxygen (HBO) therapy on biointegration of porous polyethylene (PE) implanted beneath dorsal burn scar and normal skin of Sprague-Dawley rats were microscopically examined and the ratio of fibrovascular ingrowth (FVI) was determined for each rat (Dinar et al., 2008). The results showed that HBO therapy enhanced biointegration of porous PE in hypoxic burn scar areas via improving collagen synthesis and neovascularization; otherwise, it apparently delayed tissue ingrowth into a porous structure implanted in normal healthy tissues.

1.7

Future trends

The future of biointegration and the future of implants are considered to be bright, as advancements in frontier biomaterials are advancing rapidly to unravel the physiologic process required for the biologization of the implants and developing materials that become intrinsically integrated into the organ-specific repair mechanisms. One example is the ‘designer implant’, which could carry different types of proteins, one set to spur soft tissue healing, another to encourage hard tissue growth on another front. According to Rush, future devolvement of osteobiologic materials will no doubt replace materials currently being used (Rush, 2005). Techniques to improve biointegration and manipulation of the healing environment will be developed such that future graft substitutes may exceed even autogenous bone in their reliability. An understanding of the cascade of events that occurs with bone healing and graft incorporation will enhance the chances to augment or manipulate the grafting process. The biomaterials community is producing new and improved implant materials and techniques to meet this demand. A counter force to this technological push is the increasing level of regulation and the threat of litigation. To meet these conflicting needs it is necessary to have reliable methods of characterization of the material and material/host tissue interactions (www.azom.com). In addition,


12


progress on the road to regenerating major body parts, salamander-style, could transform the treatment of amputations and major wounds (Muneoka et al., 2008). At the same time, the newly emerging nanobiotechnology is revolutionizing capability to resolve biological and medical problems by developing subtle biomimetic techniques. The use of nanoscale materials is expected to increase dramatically in many applications of medicine and surgery. It is interesting to note that the size of these nanomaterials is comparable to many biological systems and so there is large scope for biointegration. Additionally, nanomaterials exhibit fundamentally different properties from their properties in bulk, such that they can be tailor-made to provide properties to suit a specific application (Namavar, 2003; Christenson et al., 2007). Significant advancements are expected to achieve the desired goals and their clinical use, especially in areas such as nanostructured coatings, nanostructured porous scaffoldings and other nanobiomaterials. The current trends in nanobiotechnology, thus, offer a bright future through the use of nanobiomaterial in achieving biointegration.

1.8

References

adell r, eriksson b, lekholm u, brånemark p i and jemt t (1990), ‘Long-term follow-up study of osseointegrated implants in the treatment of totally edentulous jaws’, Int J Oral Max Impl, 5, 347–359. albrektsson t, brånemark p i, hansson h a, lindstrom j (1981), Osseointegrated titanium implants: Requirements for ensuring a long-lasting, direct bone-toimplant anchorage in man. Acta Orthop Scand, 52, 155–170. amling m, schilling a f, pogoda p, priemel m and rueger j m (2006), ‘Biomaterials and bone remodeling: The physiologic process required for biologization of bone substitutes’, Eur J Trauma, 32, 102–106. Doi: 10.1007/s00068-006-6049-6. badr n a and el hadary a a (2007), ‘Hydroxyapatite-electroplated cp-titanium implant and its bone integration potentiality: An in vivo study’, Impl Dent, 16, 297–308. Doi: 10.1097/ID.0b013e31805d7dc4. balasundaram g and webster t j (2006), ‘Nanotechnology and biomaterials for orthopedic medical applications’, Nanomedicine, London, 1, 169–176. barrère f, mahmood t a, de groot k and van blitterswijk c a (2008), ‘Advanced biomaterials for skeletal tissue regeneration: Instructive and smart functions’, Mat Sci Eng R, 59, 38–71. bereznai m, pelsöczi i, tóth z, turzó k, radnai m, bor z and fazekas a (2003), ‘Surface modifications induced by ns and sub-ps excimer laser pulses on titanium implant material’, Biomaterials, 24, 4197–4203. Doi: 10.1016/ S0142-9612(03)00318-1. brånemark p i, zarb g a and albrektsson t (1985) Tissue-integrated Prostheses: Osseointegration in Clinical Dentistry, Chicago, Quintessence Publishing Co., pp. 1–356.



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chirila t v, hicks c r, dalton p d, vijayasekaran s, lou x, hong y, clayton a b, ziegelaar b w, fitton j h, platten s, crawford g j and constable i j (1998), ‘Artificial cornea’, Prog Polym Sci, 23, 447–473. chirila t v (2006), ‘First development of a polyurethane keratoprosthesis and its Australian connection: An unbeknown episode in the history of artificial cornea’, Clin Exp Ophthalmol, 34, 485–488. Doi: 10.1111/j.1442-9071.2006. 01251.x. chow c c, kulkarni a d, albert d m, darlington j k and hardten d r (2007), ‘Clinicopathologic correlation of explanted AlphaCor artificial cornea after exposure of implant’, Cornea, 26, 1004–1007. Doi: 10.1097/ICO.0b013e3180e799f0. christenson e m, anseth k s, van den beucken j j j p, chan c k, ercan b, jansen j a, laurencin c t and mikos a g (2007), ‘Nanobiomaterial applications in orthopedics’, J Orthopaed Res, 25, 11–22. coassin m, zhang c, green w r, aquavella j v and akpek e k (2007), ‘Histopathologic and Immunologic Aspects of AlphaCor Artificial Corneal Failure’, Am J Ophthalmol, 144, 699–704 e4. Doi: 10.1016/j.ajo.2007.07.025. cochran d (1996), ‘Implant therapy I’, Ann Periodontol, 1, 707–791. covani u, giacomelli l, krajewski a, ravaglioli a, spotorno l, loria p, das s and nicolini c (2007), ‘Biomaterials for orthopedics: A roughness analysis by atomic force microscopy’, J Biomed Mater Res A, 82, 723–730. dinar s, agir h, sen c, yazir y, dalcik h and unal c (2008), ‘Effects of hyperbaric oxygen therapy on fibrovascular ingrowth in porous polyethylene blocks implanted under burn scar tissue: An experimental study’, Burns, 34, 467–473. Doi: 10.1016/j.burns.2007.04.014. fukano y, knowles n g, usui m l, underwood r a, hauch k d, marshall a j, ratner b d, giacelli c, carter w g, fleckman p and olerud j e (2006), ‘Characterization of an in vitro model for evaluating the interface between skin and percutaneous biomaterials’, Wound Repair Regen, 14, 484–491. Doi: 10.1111/j.1743-6109.2006. 00138.x. harris l g and richards r (2006), ‘Staphylococci and implant surfaces: A review’, Injury, 37, S3–S14. Doi:10.1016/j.injury.2006.04.003. hicks c r, crawford g j, tan d t, snibson g r, gondhowiardjo t d, lam d s c and downie n (2002), ‘Outcomes of implantation of an artificial cornea, AlphaCor: Effects of prior ocular herpes simplex infection’, Cornea, 21, 685–690. Doi: 10.1097/00003226-200210000-00010. hicks c r and hamilton s (2005), ‘Retroprosthetic membranes in AlphaCor patients: Risk factors and prevention’, Cornea, 24, 692–698. Doi: 10.1097/01.ico. 0000154380.13237.ea. hicks c r, werner l, vijayasekaran s, mamalis n and apple d j (2005), ‘Histology of AlphaCor skirts: Evaluation of biointegration’, Cornea, 24, 933–940. Doi: 10.1097/01.ico.0000160969.50706.7f. hong z, luan l, paik s b, deng b, ellis d e, ketterson j b, mello a, eon j g, terra j and rossi a (2007), ‘Crystalline hydroxyapatite thin films produced at room temperature – An opposing radio frequency magnetron sputtering approach’, Thin Solid Films, 515, 6773–6780. Doi: 10.1016/j.tsf.2007.02.089. http://academic.uprm.edu/~mgoyal/materialsmay2004/k04orthopaedic.pdf. http://www.dental-oracle.org/uk/implant/Pages/c.html. http://www.karger.com/gazette/65/lidgren/art_5_0.html.


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http://www.medcompare.com/details/40895/AlphaCorAndtrade-Artificial-Cornea. html. http://www.pricevisiongroup.com/cornea_transplant.html. isenhath s n, fukano y, usui m l, underwood r a, irvin c a, marshall a j, hauch k d, ratner b d, fleckman p and olerud j e (2007), ‘A mouse model to evaluate the interface between skin and a percutaneous device’, J Biomed Mater Res Part A, 83, 915–922. Doi: 10.1002/jbm.a.31391. meffert r m, block m s and kent j n (1987), ‘What is osseointegration?’, Int J Periodontics Restorative Dent, 7, 9–21. mirhosseini n, crouse p l, schmidth m j j, li l and garrod d (2007), ‘Laser surface micro-texturing of Ti-6Al-4V substrates for improved cell integration’, Appl Surf Sci, 253, 7738–7743. muneoka k, han m and gardiner d m (2008), ‘Regrowing human limbs’, Sci Amer 298, 56–63. namavar f, jackson j d, sharp g, mann e e, bayles k, cheung c l b, feschuk c a, varma s, haider h and garvin k l (2007), Searching for Smart Durable Coatings to Promote Bone Marrow Stromal Cell Growth While Preventing Biofilm Formation, Mater Res. Soc. Symp Proc., Vol. 954 © Materials Research Society 0954-H04-04. namavar f (2003), ‘Applications of nanotechnology for alternative bearing surfaces in orthopaedics’, Proceedings of the 8th Ceramics, Cells and Tissues MeetingSeminar, Faenza, Italy March, 2003, Volume edited by A Ravaglioli and A Krajewski, ISTEC-CNR (December 2003). oudadesse h, derrien a c, mami m, martin s, cathelineau g and yahia l (2007), ‘Alumino silicates and biphasic HA-TCP composites: Studies of properties for bony filling’, Biomed Mater, 2, art.no.S09, S59–S64. Doi: 10.1088/1748-6041/2/1/ S09. pelsoczi k i, bereznai m, tóth z, turzó k, radnai m, bor z and fazekas a (2004), ‘Surface modifications of titanium implant material with excimer laser for more effective osseointegration’ (Titán-minták felületének módosítása excimer lézerrel a hatékonyabb osszeointegráció erdekében), Fogorvosi Szemle, 97, 231–237. polykandriotis e, arkudas a, euler s, beier j p, horch r e and kneser u (2006), ‘Prevascularisation strategies in tissue engineering’ (Prävaskularisationsstrategien im tissue engineering), Handchirurgie Mikrochirurgie Plastische Chirurgie, 38, 217–223. rey c (1988), ‘Orthopedic biomaterials, bioactivity, biodegradation: A physicalchemical approach’, J Biomech, 31, Supplement 1, July 1998, 182. roling t (1989), ‘Biointegration revolutionizes dental surgery’, Sulzer Technical Review, 71, 7–10. rush s m (2005), ‘Bone graft substitutes: Osteobiologics’, Clinics in Podiatric Medicine and Surgery, 22, 619–630. Doi: 10.1016/j.cpm.2005.07.004. schmidt a h and swiontkowski m f (2008), ‘Pathophysiology of infections after internal fixation of fractures’, J Am Acad Orthop Surg, 285–291. smausz t, hopp b, huszár h, töth z and kecskeméti g (2004), ‘Pulsed laser deposition of bioceramic thin films from human tooth’, Appl Phys A-Mater, 79, 1101–1103. sousa s r, moradas-ferreira p and barbosa m a (2005), ‘TiO2 type influences fibronectin adsorption’, J Mater Sci-Mater M, 16, 1173–1178. Doi: 10.1007/ s10856-005-4725-4.



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viceconti m, davinelli m, taddei f and cappello a (2004), Automatic generation of accurate subject-specific bone finite element models to be used in clinical studies, J Biomech, 37, 1597–1605. Doi:10.1016/j.jbiomech.2003.12.030. wataha j c (1996), ‘Materials for endosseous dental implants’, J Oral Rehabil, 23, 79–90. williams d f (1999), ‘The Williams Dictionary of Biomaterials’, Liverpool University Press, Liverpool. www.azom.com, Biomaterials: an overview. zanotti b and verlicchi a (2006), ‘Is one cervical prosthesis equal to another?’ (Una protesi cervicale vale l’altra?), Rivista Medica, 12, 125–130. zim s (2004), ‘Skeletal volume enhancement: Implants and osteotomies’, Current Opinion in Otolaryngology and Head and Neck Surgery, 12, 349–356. Doi: 10.1097/01.moo.0000130576.04818.55.


2 Biocompatibility of engineered soft tissue created by stem cells P. A. C L A R K, University of Wisconsin–Madison, USA; and J. J. M AO, Columbia University, USA

Abstract: Orthopedic and dental implants replace millions of arthritic, traumatic or lost skeletal or dental structures. All orthopedic or dental implants can fail, and the central reason for failure is that metallic implants do not remodel with host tissue. Current implants rely primarily on tissue growth onto implants or an ‘outside-in’ strategy. This chapter discusses an ‘inside-out’ strategy to induce tissue ingrowth by cytokine delivery. Drug-eluting porous implants have the advantage not only of reducing bulk metal mass, but also of harboring cytokines that are programmed to release into surrounding tissue. This coupled inside-out and outside-in strategy improves bone ingrowth. Key words: orthopedic implants, medical implants, cardiac implants, dental implants, growth factors, cytokines, controlled release, microencapsulation.

2.1

Introduction

Tissue and organ defects resulting from trauma, chronic diseases, tumor resection or congenital anomalies necessitate the restoration of the lost anatomical structures. Due to a lack of biological replacements for skeletal structures, implantation of biocompatible metals such as titanium is currently the preferred treatment (Ratner et al., 1996; Misch, 1993; Kienapfel et al., 1999). Despite high success rates of initial anchorage (over 90%) (Ashley et al., 2003), titanium implants require long healing times before functional loading, and are subject to failure from inadequate initial bone ingrowth or long-term osteolysis at the bone–implant interface. Current approaches in modifying titanium implants to overcome these limitations focus on biomaterial composition and processing, surface roughening, and chemical surface modification, among others. Taking cues from biology and tissue engineering has led to the idea of biointegration, entailing the use of biologically active agents to modulate the bone ingrowth process and improve implant anchorage. Biointegration of orthopedic implants represents a daunting task considering the complex environment of healing and homeostatic bone, but fortunately the fields of tissue engineering and drug 19 © Woodhead Publishing Limited, 2010

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delivery have developed micron- and nano-scale systems for controlled release of various biologically active agents. This chapter will first address the key biological considerations of biointegration of orthopedic implants. Mimicking these processes using drug delivery systems toward improved short- and long-term efficacy in these implants will then be discussed, concluding with future trends in this emerging field.

2.2

Bone: from tissue to molecular organization

Bone is the main weight-bearing tissue of the body, with varying and complex macroscopic designs due to its distinct functions in different regions of the body. In general, bone exists either as cortical with low porosity and high density, or cancellous (or trabecular) with microscopic interconnecting bony trabecula giving macroscopically high porosity and low density (Marks and Odgren, 2002). Biochemically, bone is composed of about 35% organic matrix (osteoid), mainly Type I collagen fibers along with proteoglycans and noncollagenous proteins, and about 65% inorganic mineral, mainly calcium and phosphate in the form of hydroxyapatite (Lind, 1998; Misch, 1993). This general composition gives bone marked rigidity while retaining some elasticity (Marks and Odgren, 2002), with the collagen fibers of the organic matrix providing high tensile strength to resist pulling forces and the inorganic mineral providing high compressive strength to resist crushing forces (Marks and Odgren, 2002; Misch, 1993; Alberts et al., 2002). A key facet of bone tissue during development and maintenance is the constant re-organization of the extracellular matrix to satisfy local loadbearing requirements. This process is driven by the two main cell phenotypes of bone: bone-forming osteoblasts and bone-resorbing osteoclasts (Marks and Odgren, 2002; Hole and Koos, 1994). Acting as possible sensors and signaling agents for the osteoblasts and osteoclasts are the osteocytes, post-mitotic terminally differentiated osteoblasts encased in bone matrix that communicate via long processes known as canaliculi (Marks and Odgren, 2002). Located in the cavities of long bones and among trabecula in cancellous bone is the bone marrow. This tissue contains both red marrow, the site of new blood cell production or hematopoiesis throughout life, and yellow marrow, which is mostly fat cells (Hole and Koos, 1994). The bone marrow generally transitions from red to yellow with age, although this trend can be reversed in injurious or other special instances (Hole and Koos, 1994). The marrow contains a milieu of cells, including red and white blood cells, osteoblasts, fibroblasts, adipocytes (fat cells), and blood vessel cells (Hole and Koos, 1994). Fibroblast-like cells residing within the bone marrow stroma, or connective tissue of the marrow, have also been isolated that


Biocompatibility of engineered soft tissue created by stem cells

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possess extensive proliferative potential and differentiation ability to multiple mesenchymal lineages, including osteoblasts, chondrocytes (cartilage cells), and adipocytes (Pittenger et al., 1999; Caplan, 1991a,b; Alhadlaq and Mao, 2004; Cassiede et al., 1996). These multipotential cells, termed mesenchymal stem cells (MSCs) or bone marrow stromal cells, likely play a key role in repair after injury, and are present throughout life (Caplan, 1991b; Alhadlaq and Mao, 2004). Lining the outer wall of the marrow cavity and the outer surface of bones are thin linings of tissue called the endosteum and periosteum, respectively. These tissues are similar in composition and morphology, being composed of flattened cells (Marks and Odgren, 2002). Recently, the endosteal lining has been identified as an important hematopoietic stem cell (HSC) ‘niche’, the specialized compartment in which stem cells reside (Scadden, 2006; Taichman, 2005). Through strict control of the microenvironment, the endosteal cells maintain the HSCs, which can differentiate to every blood lineage, until they are needed (Taichman, 2005; Scadden, 2006). The endosteum and periosteum may also contain osteoprogenitor cells that can mobilize after injury (Hutmacher and Sittinger, 2003; Hanada et al., 2001). Not to be forgotten, like any tissue, bone and its marrow require a rich vascular supply for oxygen and nutrients and for disposal of waste products. As will be discussed later in the chapter, vessel formation or ingrowth is critical for bone formation during development and after injury. Mural cells associated with blood vessels, particularly the pericytes, have demonstrated multilineage potential and may also participate in bone repair after injury (Collett and Canfield, 2005; Doherty and Canfield, 1999). The tissue and cellular processes that organize and maintain bone are molecularly coordinated and controlled largely by bioactive chemicals termed cytokines or growth factors (Gilbert, 1997). During development, homeostasis, and after injury, a multitude of skeletal growth factors act as both temporal and spatial coordinating molecules to induce chemotaxis, mitosis, differentiation, changes in extracellular matrix production, and even apoptosis (Roberts, 2000; Alliston and Derynck, 2000). These cytokines can exert their effects on local cells (paracrine), on the same cells that released them (autocrine), or after absorption and transport via the bloodstream (endocrine) (Lind, 1998). Some cytokines exert their effects on very specific lineages of cells while others exhibit context-dependent effects on multiple cell phenotypes. Growth factor effects on cells depend heavily on the dosage, with most cells demonstrating biphasic responses. Some of the most well-studied cytokines in skeletal biology are members of the transforming growth factor β (TGFβ) superfamily, which include multiple TGFβ isoforms and the bone morphogenetic proteins (BMPs). Members of this superfamily play important and oftentimes critical roles in the growth and


22


maintenance of bone and cartilage. Perhaps most strikingly, the BMPs were discovered by Marshall Urist in 1965 by their ability to induce ectopic bone formation in skeletal muscle (Lou, 2001; Linkhart et al., 1996; Lind, 1998; Urist, 1965; Urist et al., 1979). Fibroblast growth factors (FGFs) and the insulin-like growth factors (IGFs) also participate in many skeletal development and repair processes (Marie et al., 2002; Linkhart et al., 1996). Important to vasculogenic and angiogenic processes, and therefore to bone development and homeostasis, are vascular endothelial growth factor (VEGF) and platelet-derived growth factor (PDGF) (Lind, 1998; Franceschi, 2005; Gerstenfeld et al., 2003). Details of these, as well as other cytokines, in skeletal biology will be expounded upon later in the chapter, but their early introduction is warranted as their appearance is a common theme throughout skeletal development, homeostasis, and repair.

2.3

Bone development

Becoming apparent in recent years is the uncanny resemblance that repair processes have to embryonic development (Gerstenfeld et al., 2003; Caplan, 2003), and therefore it is important to confer at least a general understanding of the mechanisms of bone formation during embryogenesis. Embryonic limb bud formation begins by budding from the lateral surface at specific levels (Tuan, 2004). Undifferentiated mesenchyme begins to pocket out as a limb field, expressing FGF10. Thickening of the edge of the bud forms the apical ectoderm ridge, and begins to express FGF8, BMP2, BMP4, and Msx2. At this point, spatial patterning (ventral vs. dorsal, anterior vs. posterior) is initiated, as determined by gradients of Hox genes as well as by Wnt (related to the developmental molecules wingless in Drosophila) and FGF4. The mesenchyme then condenses, with these cells secreting a variety of signaling factors, including growth and differentiation factor (GDF) 5, BMP2, BMP4, BMP7, and FGF9 (Tuan, 2004; Ornitz, 2005). Condensation of mesenchymal cells to trigger lineage specification is common throughout development of bone and cartilage, suggesting critical roles for extracellular matrix (ECM) components (Gilbert, 1997) and cell shape. Lineage specification of adult MSCs isolated and expanded ex vivo also require proper cell shape and ECM (Cassiede et al., 1996; Pittenger et al., 1999; McBeath et al., 2004), exemplifying the importance of basic developmental concepts in regenerative medicine. Condensing mesenchymal cells begin to differentiate to chondrocytes and express noggin, a potent inhibitor of BMP signaling (Tuan, 2004). The chondrocytes then begin forming cartilaginous tissue matrices that will be the model for future bones. After the cartilaginous scaffolds of the bones form, the skeletal structures begin to mature through endochondral ossification, one of two mechanisms



23

by which bone forms during development. Proliferating cartilage cells in the center of the bone models begin expressing Indian hedgehog (IHH). IHH, parathyroid hormone-related protein (PTHrP), and multiple FGFs coordinate the proliferation of chondrocytes as well as their further maturation into hypertrophic chondrocytes (St-Jacques et al., 1999; Day and Yang, 2008; Ornitz, 2005; Kronenberg, 2006). As the hypertrophic chondrocytes enlarge, they begin to both degrade and mineralize the surrounding matrix before dying and degenerating. Triggering bone formation is the invasion of blood vessels, which bring along osteoclasts to dispose of disintegrating tissue and osteoprogenitors that differentiate into osteoblasts and begin building the bone (Alberts et al., 2002; Hole and Koos, 1994). This close relationship between angiogenesis and osteogenesis has implicated VEGF as playing a major role in the coordination of endochondral ossification (Dai and Rabie, 2007). Osteoblasts deposit a bed of collagenous osteoid that is subsequently mineralized to mature bony tissue, working from the center of the developing bone forming an ossification front that continues to push the chondrocytes outward (Gilbert, 1997). In long bones, the chondrocytes, begin bulging to form heads, and the progression of the ossification front slows. As blood vessels penetrate the heads, a secondary ossification center forms that again ossifies, pushing the chondrocytes outward. Where the secondary and primary ossification fronts meet remains cartilaginous and forms the epiphyseal growth plate, making it possible for further bone extension, until adulthood when bone growth is completed (Gilbert, 1997). Chondrocytes pushed to the edge of the long bones form the articular cartilage, driven in part by TGFβ-induced ECM formation (Eames et al., 2003), and cells remaining outside the bone differentiate to form the periosteum (Hole and Koos, 1994). The second mechanism of bone development, exemplified in the plates of the skull, is intramembranous ossification. Membrane-like layers of primitive connective tissues first appear at the site where bones are to be grown. Mesenchymal cells derived from the neural crest interact with the extracellular matrix of the head epithelial cells to form the bones (Gilbert, 1997). Signaling during these mesenchymal–epithelial interactions includes the BMPs (Gilbert, 1997), TGFβs (Kanaan and Kanaan, 2006), and high Wnt signaling within the mesenchymal condensates to induce osteoblast differentiation (Day and Yang, 2008). Again, vessel formation is critical, as the mesenchymal cells condense around and begin bone formation immediately adjacent to capillaries (Gilbert, 1997). Coordinating these mesenchymal cells through proliferation to differentiation to osteogenesis are members of the FGF (FGF18 and FGF2) and BMP families. Bone formation from differentiating osteoblasts commences the same as in endochondral ossification, forming bony islands that eventually connect to form the plates of the skull.


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During skeletal development, it is interesting that separate embryonic lineages differentiate to form the long bones and craniofacial bones. The long bones originate from the lateral plate mesoderm of the developing embryo, giving these cells a mesenchymal lineage, while craniofacial bones are generated from neural crest cells of ectodermal lineage (Alberts et al., 2002). Although not determined, the different heritage of these cells may influence their response to various growth factor stimuli in adult life, and therefore location must be considered during design and testing of implants. As mentioned above, bone growth occurs post-natally at epiphyseal growth plates. Growth at the plates resembles that of endochondral ossification. Proliferating chondrocytes extend the bone at the leading edge of an ossification front, leaving behind hypertrophic chondrocytes forming calcified cartilage. The calcified cartilage is removed by invading osteoclasts and replaced with mature bone by osteoblasts (Gilbert, 1997). Critical in stimulating the growth of the epiphyseal growth plates are factors including growth hormone (GH) and IGF1 (Gilbert, 1997). As during development, IHH and PTHrP constitute a feedback loop regulating chondrocyte proliferation and differentiation (van der Eerden et al., 2003). Multiple BMPs, including BMP2, BMP4, and BMP7, also play a role, as well as members of the FGF family (van der Eerden et al., 2003). Although a daunting list, it is important to be aware of the multitude of cytokines and the processes they coordinate during development. As the molecular mechanisms underlying bone repair after injury are elucidated, the number of similarities with embryonic developmental processes continues to compound (Gerstenfeld et al., 2003; Dimitriou et al., 2005).

2.4

Bone homeostasis

The structure of bone tissue at any given anatomical location is no mistake, and reflects the optimal mass and morphology to develop the strength required, as well as the optimal shape to satisfy its local load-bearing requirements (Guyton and Hall, 1996; Frost, 1987). These changes in structure, as well as repair of bone microdamage, are accomplished through highly coordinated processes involving the osteoblasts and osteoclasts, and disruptions in this delicate balance is attributed to many bone diseases (Marks and Odgren, 2002). This remodeling occurs via cooperation of many cell types in a basic multicellular unit (BMU), morphologically resembling a cutting cone – osteoclasts leading the way in resorbing bone, osteoblasts following laying down new bone, and blood vessels at the end providing essential nutrients to the newly formed bone (Jilka, 2003). The exact cellular and molecular signals that initiate bone remodeling are unknown, but either resident osteocytes or bone lining cells are likely responsible (Jilka, 2003).



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The signaling cells first recruit osteoclast progenitors to the site of remodeling by releasing chemokines, including monocyte chemoattractant protein-1 (MCP-1, also known as CCL2) and stromal cell-derived factor (SDF-1, also known as CXCL12) (Matsuo and Irie, 2008). Initial differentiation of osteoclasts is induced by the same osteocytes or stromal cells by cytokines, including macrophage-colony stimulating factor (M-CSF, also known as CSF-1) and receptor activator for nuclear factor κB ligand (RANKL) that bind to their correlative receptors on osteoclast progenitors (Jilka, 2003; Matsuo and Irie, 2008). Allowing for fine control of the process, signaling cells also release osteoprotegerin (OPG), which binds to RANKL and prevents its signaling to osteoclast progenitors (Jilka, 2003). Systemic factors such as parathyroid hormone (PTH) and 1,25-vitamin D3 as well as paracrine factors such as interleukin (IL)-1, IL-6, IL-11, and tumor necrosis factor (TNF) can also influence the extent of osteoclast differentiation. Osteoclasts then initiate the process of ‘coupling’ with osteoblasts to form the BMU, through direct cytokine release, cell-to-cell signaling, or liberation of cytokines from resorbed bone matrix (Matsuo and Irie, 2008). Critical in this process are the familiar skeletal cytokines including BMPs, IGFs, and TGFβ (Jilka, 2003; Matsuo and Irie, 2008), as well as a few other candidates including PDGF, hepatocyte growth factor (HGF), and Wnts (Matsuo and Irie, 2008). TGFβ1 is important in the chemotaxis and proliferation of progenitor cells as well as initial extracellular matrix deposition (Dimitriou et al., 2005), but seems to exert inhibitory effects on final cell differentiation and maturation (Alliston and Derynck, 2000). Reinforcing the idea that proper temporal expression of cytokines is critical, TGFβ1 promotes the production of critical initial extracellular matrix components, such as collagen Type I, osteopontin, and osteonectin (Alliston and Derynck, 2000; Lu et al., 2001), while possibly inhibiting some of the late stage markers of osteoblast differentiation. BMPs seem to induce differentiation of immature mesenchymal cells (Lou, 2001), but the BMPs, especially the popular BMP-2, have been shown to exhibit chemotactic and proliferative effects as well (Dimitriou et al., 2005; Lind, 1998). The IGFs work throughout the bone remodeling process, acting potently as survival factors to prevent apoptosis of differentiating and mature osteoblasts (Jilka, 2003). When the osteoblasts complete bone formation and the new bone is properly vascularized, the bone remodeling process is terminated by either apoptosis of the osteoblasts or their terminal differentiation to osteocytes or bone lining cells (Matsuo and Irie, 2008). It is interesting to note the two completely different lineages from which osteoblasts and osteoclasts arise. The mononucleated osteoblasts originate from progenitor cells, likely located in the bone marrow or periosteum (Marks and Odgren, 2002; Gerstenfeld et al., 2003; Caplan, 1991b; Pittenger et al., 1999; Alhadlaq and Mao, 2004). Osteoclasts originate from the same


26


precursors as blood cells, and are formed by fusion of these mononuclear precursors (Gilbert, 1997; Jilka, 2003). The involvement of such seemingly unrelated lineages (blood vs. bone) exemplifies how complex and orchestrated bone homeostasis is.

2.5

Bone repair after injury

The events of bone repair after injury (Fig. 2.1) as alluded to above share multiple cellular and molecular events with development and homeostasis, and the process has been referred to as a regeneration that progresses like a post-natal developmental process (Gerstenfeld et al., 2003; Dimitriou et al., 2005). Multiple tissues are required for successful repair and progress in overlapping phases including the inflammatory, stabilization, repair, and remodeling phases (Fig. 2.1). Although there is evidence for key involvement of other lineages, we will focus on the cellular response of the bone cells after injury, which include five key cellular steps that must occur for rapid, successful regeneration of bone tissue: cell recruitment and chemotaxis to injury site, cell proliferation, extracellular matrix deposition, cell differentiation, and mineralization and maturation (Puleo and Nanci, 1999; Greenhalgh, 1996; Dimitriou et al., 2005). The inflammatory phase is the first phase of bone repair after injury and initiates many critical processes to set in motion the complex events to follow (LeGeros and Craig, 1993; Puleo and Nanci, 1999; Kienapfel et al., 1999; Probst and Spiegel, 1997), as demonstrated by the ability of antiinflammatory drugs to significantly decrease bone ingrowth in orthopedic implants (Cook et al., 1995; Goodman et al., 2002). The disruption of many blood vessels results in the activation and aggregation of platelets, leading to the formation of a blood clot, called a hematoma, which is the initial stabilization of the site and repository of many cytokines (Probst and Spiegel, 1997). White blood cells, leukocytes and neutrophils, soon reach the injury site, along with monocytic phagocytes, which differentiate into macrophages. Aside from fighting off infection, these macrophages act as growth factor factories, manufacturing and releasing approximately 100 biologically active substances (Probst and Spiegel, 1997). Apart from the immune phenotypes, disruption of the bone matrix itself leads to release of cytokines, such as TGFβ1 that exists in a latent form in high quantities in bone until released after injury (Roberts, 2000). Taken together, a plethora of coordinating cytokines and growth factors are present at the initial wound site in large quantities, among them the skeletal cytokines PDGF, BMPs, TGFβs, FGFs, and IGFs (Probst and Spiegel, 1997). TGFβ1 seems especially important, as it is synthesized and released in high amounts by the immune cells, is released from the disrupted bone matrix, and induces positive regulation by resident osteoblasts and



27

Progenitor Oc

Host bone

Woven bone

Injury site ECM

Ob

Oy Lamellar bone

Titanium implant Time

Tissue response

Inflammatory response Stabilization Repair Remodel

Bone cell response

Recruitment/ chemotaxis Proliferation ECM deposition Differentiation Mineralization/maturation

2.1 Bone repair after titanium implant placement. Initial placement of a titanium implant initiates a tissue cascade that includes the inflammatory phase, repair phase, and remodeling phase. After the inflammatory phase is nearly complete, these tissue responses are regulated largely by bone cells. Recruitment and chemotaxis of progenitor cells from various areas of bone, including periosteum, endosteum, and bone marrow, to the injury site is the first cellular phase. The cells then begin to proliferate and deposit extracellular matrix (ECM) to stabilize the injury site. Upon stabilization and blood vessel ingrowth, the progenitor cells differentiate to osteoblasts (Ob) and begin to lay down immature, unorganized woven bone. During the final remodeling phase, osteoclast (Oc) progenitors are recruited that subsequently differentiate and begin to digest the immature bone. Osteoblasts (Ob) follow closely behind, depositing mature, organized lamellar bone. Some osteoblasts that become entrapped in bone matrix differentiate to osteocytes (Oy). Various growth factors, including TGFβs, BMPs, and IGFs, coordinate this process and can exert different spatial and temporal effects.


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osteoprogenitors (Probst and Spiegel, 1997; Linkhart et al., 1996; Roberts, 2000; Alliston and Derynck, 2000). Once the initial inflammatory phase is completed, stabilization of the injury site continues in the proliferative phase. In fractures, granulation tissue is often first laid down, then replaced by cartilage, and finally is replaced by bone (Probst and Spiegel, 1997). The formation of connective tissue is accomplished by the migration, proliferation, extracellular matrix deposition, and differentiation of progenitor cells (Davies, 2003; Caplan, 1991a). The origin of these progenitors is currently unknown, but many cellular pockets are culprit (Dimitriou et al., 2005) – multipotent mesenchymal stem cells found in bone marrow (Caplan, 1991b; Pittenger et al., 1999; Alhadlaq and Mao, 2004), osteoprogenitor cells of the periosteum and endosteum covering the outer and inner surfaces of bone (Hutmacher and Sittinger, 2003; Gerstenfeld et al., 2003), bone lining cells on the surfaces of bone and in bone pockets (Marks and Odgren, 2002), or multipotent pericyte cells surrounding blood vessels (Collett and Canfield, 2005; Doherty and Canfield, 1999). The periosteum seems to be the major provider of cells for the external callus, and multipotent mesenchymal stem cells of the marrow appear to be the primary cell source for the internal callus (Gerstenfeld et al., 2003; Probst and Spiegel, 1997; Hole and Koos, 1994). Besides providing cellular help, the periosteum cells release cytokines that stimulate differentiation of multipotent cells to osteoblastic phenotypes such as BMP-2, as well as driving the ingrowth of new blood vessels by VEGF and PDGF (Gerstenfeld et al., 2003). After the initial inflammation phase, the levels of TGFβ1, among other factors, remain elevated (Gerstenfeld et al., 2003), suggesting continued and varied roles throughout the fracture healing process. Of critical importance in the proliferative phase is the growth of new blood vessels, angiogenesis, at the wound site. Bone can form only where there is an adequate blood supply available (Probst and Spiegel, 1997). Angiogenesis also depends on the many cytokines and other biologically active substances released by inflammatory cells. Peptides such as TGFβ, FGF2, PDGF, and VEGF have all been shown to play a part in stimulating angiogenesis (Probst and Spiegel, 1997). The mineralization and maturation phase is the final step in bone repair. After suitable mechanical stabilization is achieved from extracellular matrix deposition, osteoprogenitor cells differentiate into osteoblasts and form a bone matrix at the injury site. Differentiation factors such as TGFβ and BMPs are elevated and critical extracellular matrix components such as osteocalcin and osteopontin begin to be expressed. Much of this bone formation is uncoordinated, leading to the formation of mainly woven bone at the injury site. Transition from woven to organized lamellar bone follows the pattern of homeostatic bone remodeling, and results in bone



29

strong enough to restore biomechanical competence (Probst and Spiegel, 1997).

2.6

Bone and joint disease

The skeletal system is subject to many disorders that diminish its function, such as trauma, chronic diseases, tumor resection, or congenital anomalies. These disorders compromise and, if left untreated, usually incapacitate the multiple functions of the skeleton, exhibited clinically as increased fractures or debilitating pain. Musculoskeletal injury affects more than 28 million patients in the United States every year and has a $254 billion impact on the US economy per year (USBJD, 2006). Bone homeostasis is a delicate balance between bone formation and resorption, and if disrupted can lead to many disease states (Marks and Odgren, 2002). Osteoporosis is clinically defined as a symptomatic, generalized decrease in bone mass (McCarthy and Frassica, 1998). Osteoporosis results in increased fracture risk, mostly in the vertebrae and hip (McCarthy and Frassica, 1998), and because of impaired bone repair in these patients, fractures often heal slowly, if at all. Like almost every other tissue of the body, bone is subject to cancerous growth of its cells, although to a much less degree than other tissues, with malignant bone tumors being rare (McCarthy and Frassica, 1998). Osteosarcomas are the most common primary malignant bone tumor, usually affecting children and young adults (McCarthy and Frassica, 1998). Luckily, osteosarcomas can usually be reliably removed from the body with a longterm survival rate afterwards. However, complete resection can sometimes require removal of large portions of bone that cannot heal on its own. In human joints, degenerative diseases can progress to severe states leading to pain and debilitation, necessitating treatment (Hayes et al., 2001; Buckwalter, 2002; Gay et al., 2002; Gravallese, 2002). Osteoarthritis, a degenerative disease characterized by loss of articular cartilage (McCarthy and Frassica, 1998), and rheumatoid arthritis, an inflammatory autoimmune disease that destroys a patient’s own cartilage, are both joint diseases that usually lead to debilitating pain or physical disability (McCarthy and Frassica, 1998). Although normally resulting from problems with the articular cartilage, treatment of these diseases often requires total joint replacements that engage bone sites for strength and stability.

2.7

Current treatment options and total joint replacements

Diagnosed and treated early, bone and joint diseases often can be controlled or even reversed with non-invasive methods such as drug or hormone


30


treatments (McCarthy and Frassica, 1998). Unfortunately, these diseases often progress to a late stage, characterized by severe pain and debilitation that inhibits or severely limits normal daily activity, requiring more radical and invasive strategies to restore an acceptable quality of life. At this point of disease progression, current medical approaches dictate full replacement. Aside from total joint replacements, load-bearing plates are often used to stabilize bone after fracture or tumor resection. In comparison with donor site morbidity and pain in association with autologous tissue grafting, synthetic materials have the advantage of ready and endless supply without any sacrifice of donor tissue. Therefore, replacement using synthetic biomaterials such as metals or ceramics has become the preferred treatment for total joint replacement. The current ‘gold standard’ in metals for treatment of skeletal disorders is the titanium implant. Titanium combines high strength and excellent biocompatibility (Long and Rack, 1998; Ratner et al., 1996), arising from a highly inert passivating layer of titanium oxide (TiO2) that forms instantly on exposure to air (Ratner et al., 1996). This passivating layer also is responsible for titanium implants’ remarkable capacity for integration with host bone (Albrektsson et al., 1981; Kienapfel et al., 1999), which occurs first by bone ingrowth in which bone forms within the irregular surfaces of the titanium (Kienapfel et al., 1999) and subsequent osseointegration, defined as ‘direct contact between living bone and an implant on the light microscope level’ (Brånemark et al., 1969). Titanium has high success rates of initial anchorage (over 90%) and has been shown to co-exist with host bone for the life of the patient (Ashley et al., 2003). After insertion, the host bone response to the titanium implant is paramount for a successful outcome (Misch, 1993). Implant integration proceeds almost identically to that of bone repair (Fig. 2.1), practically down to the molecular level (Kojima et al., 2008; De Ranieri et al., 2005a; Schierano et al., 2005). Much like fracture repair, titanium implants can integrate to form a strong and stable interlock with host bone. However, unlike fracture repair, a high degree of initial fixation at surgery is critical for successful performance of the implanted device (Puleo and Nanci, 1999), as any relative motion between implant and bone can result in permanent fibrous encapsulation and failure of the implant (Szmukler-Moncler et al., 1998; Brunski, 1999).

2.8

Current challenges of titanium implants

Titanium implants predictably restore mechanical function, yet failures still remain in the area of initial and long-term integration of the implant into host bone (Merickse-Stern et al., 2001; Greenfield et al., 2002). Inadequate



31

bone ingrowth and failure to maintain adequate peri-implant bone density because of osteolysis at the bone–implant interface still lead to many failures of titanium implants, requiring painful revision surgeries (Saleh et al., 2004). Although the precise mechanisms of implant failure are not well understood, the ultimate direct cause is either initial or long-term failure of implant integration with host bone, resulting in the implant loosening and ultimately secondary surgery. Aseptic loosening of the implant is the most common cause for long-term revision surgery (McCarthy and Frassica, 1998). Loosening occurs through multiple mechanisms. Like any artificial biomaterial, the implant is subject to wear and tear from the constant mechanical loading that is applied to it, leading to a gradual loosening (McCarthy and Frassica, 1998). A biologic response occurs via one of two mechanisms. During loading, particles of varying sizes are released both from the titanium implant itself and from the plastic (polyethylene) acetabular cup that usually accompanies the implant. This particle wear from polyethylene is taken up by both macrophages for small particles and giant cells for the larger particles, which invokes macrophages to release osteoclast-specific cytokines, inducing osteolysis at the bone–implant interface (McCarthy and Frassica, 1998; McKoy et al., 2000; Ingham and Fisher, 2000). Secondly, the mis-match of biomechanical properties between titanium and bone (the modulus of elasticity of titanium is about ten times greater than that of cortical bone and one hundred times greater than that of cancellous (An, 2000; Ratner et al., 1996)) can lead to abnormally low stress transfer to surrounding bone, initiating a remodeling response from osteoclasts and bone disuse atrophy (McKoy et al., 2000). This condition is known as stress shielding (McKoy et al., 2000; Sumner and Galante, 1992). The most important factors in the relief of stress shielding are the design of the implant to distribute load equally to surrounding bone and the modulus of elasticity of the material used in the implant (McKoy et al., 2000). Because of these long-term issues with titanium implants, it may seem that failure of a titanium implant for joint replacement is inevitable. However, these processes are gradual, occurring over many years (McCarthy and Frassica, 1998). A strong initial bone formation response providing good fixation of the implant, combined with long-term delivery of osteogenic stimuli, such as cytokines, to counteract the bone resorption response could significantly extend the life of titanium implants. Therefore, modulating the complex tissue and cellular bone repair responses at the bone–implant interface to improve the implant fixation process is desirable to improve both short- and long-term efficacy of titanium implants.


32


2.9

Current titanium modifications for improved integration

Modification of titanium implants to improve the integration process is by no means a new idea, and much research using various approaches has been reported. The performance of an implant depends on two major factors: the behavior of the material in the host and the response of the host to the implant (Puleo and Nanci, 1999). The behavior of the material in the host depends on the biocompatibility of material implanted. In the case of titanium, once exposed to oxygen in ambient air or in the body, titanium forms a passivating layer of titanium oxide (TiO2) almost instantly, making the surface bioinert and allowing bone ingrowth (Ratner et al., 1996). Different treatments of the surface (Sul et al., 2001; Pan et al., 1998) and even different methods of sterilization (Binon et al., 1992) can affect the thickness of this oxide layer, which in turn affects cellular response to the titanium implants (Pan et al., 1998). Tailoring the response of the host to the implant has focused mainly on surface modification of implanted titanium (Puleo and Nanci, 1999). Perhaps the simplest and most successful methods have been morphological alterations of surface morphology and roughness of the implant. Surface roughening to create a porous titanium surface on the macro- and micro-scales has been accomplished by many methods, including machine-smoothing, grit-blasting, sand-blasting, acid etching, and plasma-spraying (Boyan et al., 1996; Pilliar, 2005). Use of these surface roughened implants has become commonplace, and the idea of ‘bone ingrowth’ into the porous spaces of titanium implant surfaces has become an important first step in the integration of titanium implants with host bone (Kienapfel et al., 1999; Pilliar, 2005). Varying degrees of surface roughness have been shown to effect cellular response to titanium implants (Jayaraman et al., 2004; Brett et al., 2004; Mante et al., 2003; Sittig et al., 1999; Schwartz et al., 1997; Boyan et al., 1996) and ultimately bone ingrowth (Boyan et al., 1996) in both dentistry and orthopedics (Puleo and Nanci, 1999; Kienapfel et al., 1999), with increasing surface roughness correlating to increased cell proliferation (Mante et al., 2003) and bony ingrowth (Frenkel et al., 2002; Groessner-Schreiber and Tuan, 1992; Wong et al., 1995; Thomas and Cook, 1985) up to an optimal value (Ronold et al., 2003). On the other hand, macroscopic or microscopic roughening of the titanium allows development of stress and strain concentrations, as is the case for some macroscopic screw threads (Van Oosterwyck et al., 1998), and if these grow too large, bone ingrowth can be impeded (Qin et al., 1996). Physiochemical methods of surface modification refer to increasing surface free energy in order to increase tissue adhesion; methods used



33

include glow discharge and charging the surface of the implant (Puleo and Nanci, 1999). An important approach in this category is coating implants with osteoinductive and osteoconductive materials, the most established being calcium phosphate derivatives (Lavelle et al., 1981; Kay, 1993) such as hydroxyapatite (HA) and tricalcium phosphate (TCP). These coatings have been prepared using various techniques, including precipitation, plasma spraying, and sputtering (Hulshoff et al., 1996; Jansen et al., 1993; Wen et al., 1998), and regardless of coating technique have consistently shown increased implant fixation when compared to non-coated titanium (Soballe et al., 1992; Maxian et al., 1993; Tisdel et al., 1994; Dean et al., 1995). In vitro assays have shown improved mineralization (Hulshoff et al., 1996; Morgan et al., 1996; ter Brugge et al., 2002) and increased osteoprogenitor differentiation (ter Brugge et al., 2002; Hulshoff et al., 1996) of calcium phosphate surfaces, with the degree of crystallinity of calcium phosphate playing a major role in cell response (Hulshoff et al., 1996; Berube et al., 2005). However, calcium phosphate coatings are not perfect, in that the mechanical strength between the calcium phosphate and titanium is often of questionable and variable strength (Filiaggi et al., 1991; Ergun et al., 2003) and the calcium phosphate may also stimulate osteoclast mediated bone resorption (Gottlander et al., 1997). Many newer approaches of surface modifications fall under the category of biochemical methods, in which proteins, enzymes, or peptides are immobilized on biomaterials for the purpose of inducing specific cell and tissue responses (Puleo and Nanci, 1999). Coatings in this category include collagen (Kim et al., 2005; Park et al., 2005), the RGD peptide (Ferris et al., 1999; Huang et al., 2003), and fibronectin (Yang et al., 2003), and have resulted in improvements for both in vitro assays and short-term in vivo implantations (Ferris et al., 1999; Schliephake et al., 2005; Morra et al., 2005; Rammelt et al., 2004). These surface modifications show promise for improved implant fixation, although long-term and human studies need still to be performed. During past decades, the premise of the design of synthetic tissue implants has been to use inert and bulk materials that permit the integration of host tissue. This approach allows titanium implants to retain their high strength; however, this high strength diverts functioning mechanical stress necessary for healthy peri-implant bone and leads to stress shielding related osteoclastogenesis and osteolysis (McKoy et al., 2000; Sumner and Galante, 1992). Mechanics of materials theory dictates that the removal of material from the bulk decreases the apparent modulus of elasticity (Hibbeler, 1997), reducing the disparity of the bone to implant modulus mis-match (Thelen et al., 2004). Removal of pieces of the bulk of titanium implants would also potentially lead to a greater degree of interlock between the implant and host, presumably leading to better implant fixation.


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Because of the aforementioned reasons, the use of a hollow implant design seems advantageous. Hollow designs such as meshes and cages have been tested for their integrative potential, but largely in non loadbearing applications (Vehof et al., 2002; Aytac et al., 2005). Load-bearing hollow implants in dentistry have existed almost as long as solid implants (Schroeder et al., 1976; Merickse-Stern et al., 2001; Misch, 1993). Although mechanical failures of these implants have been reported (Schwarz, 2000; Levine et al., 1999), the implants have performed at least as well as other implant designs (Buser et al., 1997). New tissue engineering technologies in the areas of drug and cell delivery have miniaturized delivery vehicles to the point where the hollow core of a titanium implant could also be used as a storage space for osteogenic drugs and/or cells.

2.10

Mimicking nature toward achieving titanium ‘biointegration’: cytokines and implants

With further biological and molecular understanding of bone development, growth, and repair, the importance of soluble cytokines has been realized in orthopedics. Identification of cytokines that regulate the key steps of titanium implant integration (cell recruitment and chemotaxis to injury site, cell proliferation, extracellular matrix deposition, cell differentiation, and mineralization and maturation) (Alliston and Derynck, 2000; Probst and Spiegel, 1997; Schwartz et al., 1997; Puleo and Nanci, 1999; Steinbrech et al., 2000; Hunziker and Rosenberg, 1996) has opened the door for powerful molecular based approaches that mimic the natural healing process. Although not a defined cocktail, platelet-rich plasma contains a multitude of cytokines expressed during the inflammation phase, and has been used to increase the volume of peri-implant bone in rat tibiae (Fontana et al., 2004) and increase bone ingrowth when delivered to peri-implant defects via a demineralized freeze-dried bone carrier (Sanchez et al., 2005). Injection of FGF2 increased new bone formation and connection to titanium implants implanted in rat medullary cavities (Takechi et al., 2008). Osteogenic factors such as bone morphogenetic proteins (BMPs), transforming growth factor β (TGFβ), and platelet-derived growth factor (PDGF) have repeatedly led to improvements in many bone-implant integration parameters (Wikesjo et al., 2005; Soballe et al., 2004; Lynch et al., 1991; De Ranieri et al., 2005b; Jansen et al., 2005). Use of TGFβ1 has been shown to improve bone repair and titanium implant healing (Sumner et al., 1995; Hong et al., 2000a,b; Yamamoto et al., 2000; Jansen et al., 2005; Ehrhart et al., 2005). In applications where there are existing gaps, such as augmentation of bone at implant sites or peri-implant defects, increased regeneration of bone, often at comparable levels to bone grafts, has been accomplished with delivery of TGFβ and BMP-2 via various scaffolds (Cochran et al., 1999;



35

Clokie and Bell, 2003; Clarke et al., 2004; Yamamoto et al., 2000; Vehof et al., 2002; Jansen et al., 2005). Adsorption of TGFβ and/or BMP to tricalcium phosphate or hydroxyapatite coatings on titanium implants leads to enhancement of bone ingrowth in gap models (Sumner et al., 1995, 2006; Lind, 1998; Zhang et al., 2004). Obviously, cytokines possess great potential to improve titanium–bone integration.

2.11

Growth factor delivery: why is controlled and sustained release important?

As discussed, the interplay of soluble cytokines is an elegant process, both spatially and temporally. One may question why improved delivery systems are required for cytokines when such positive outcomes as those above are achieved through simple injection to the implant site or adsorption to the titanium implant. Unfortunately, several limitations and areas for improvements are evident in these systems. First, cytokines traditionally exhibit a short half-life in vivo before diffusion from the injury site or denature from enzymes (Moioli et al., 2007; Alliston and Derynck, 2000). Therefore, large amounts of TGFβ are required for improvement of key bone ingrowth parameters (De Ranieri et al., 2005b; Jansen et al., 2005; Ehrhart et al., 2005; Sumner et al., 1995). From an application perspective, the high expense of most cytokines makes this a major drawback for widespread use. More importantly, large amounts of delivered skeletal growth factor decreases safety and efficacy, potentially leading to side effects such as ectopic bone formation (Wildemann et al., 2004) or oncogenic transformation. Second, mimicking nature through controlled temporal and spatial release will most likely produce more optimal and efficacious effects. Injected or adsorbed proteins are rapidly exhausted whereas levels of many cytokines, including TGFβ and BMP2, remain elevated throughout the fracture repair process (Gerstenfeld et al., 2003). Designed delivery would take advantage of individual cytokines for optimal use; for example, early delivery of TGFβ for mobilization and expansion of osteoprogenitors followed by BMP2 for maturation and mineralization of the resulting pool. Fortunately, biomaterials research has provided a multitude of growth factor delivery systems to overcome these challenges, focused mainly on the use of biodegradable polymers. Many of the most common systems incorporate poly-L-lactic acid (PLLA), poly-glycolic acid (PGA), or the co-polymer poly-lactic-co-glycolic acid (PLGA) (Moioli et al., 2006; Cohen et al., 1991; Lu et al., 2000; Wildemann et al., 2004). As these polymers hydrolyze in the in vivo environment to their biocompatible byproducts, encapsulated cytokines are slowly released to surrounding tissue. However, as


36


plastics, these systems lack the required mechanical strength for functional loading of the skeletal system (Ratner et al., 1996). Using special techniques, multiple polymers have been prepared as microparticles, or microspheres, that can encapsulate water soluble proteins (Cohen et al., 1991; Iwata and McGinity, 1992) including cytokines (Lu et al., 2001; Oldham et al., 2000; Moioli et al., 2006; Clark et al., 2008). The small size of polymer microspheres allows for their incorporation into stronger materials, such as titanium, while still exhibiting slow release over an extended period of time. Injection of these microparticles has also been shown to be feasible (Jiang et al., 2005), which could be a potential route to deliver osteogenic cytokines if maintenance of implant fixation was required in the long-term. Using these approaches, we have fabricated titanium implants with a hollow core and interspersed macropores; these hollow implants maintain the biomechanical strength required for load-bearing applications but provide space for tissue engineered scaffolds delivering bioactive molecules or cells (Fig. 2.2a) (Clark et al., 2008). For incorporation into the implant, PLGA microspheres (MS) averaging 64 μm in diameter (Fig. 2.2b) were fabricated to release approximately 100 ng TGFβ1 in a controlled fashion up to four weeks (Fig. 2.2c). Dosage was calculated from release kinetics and corresponded to 0.005 μg/mm3 of injury volume, equivalent to that found using controlled release of TGFβ1 to repair skull defects (Ueda et al., 2002; Yamamoto et al., 2000). After injection into a gelatin carrier, the PLGA microspheres could be subsequently packed into the hollow core of the titanium implant. The hollow implants were placed unicortically into the humeri of adult New Zealand white rabbits (Fig. 2.2d) and allowed to integrate for four weeks. As controls, placebo microspheres (no TGFβ1) as well as 1 μg and 100 ng adsorbed to gelatin carrier (rapid release) were used. After analysis using scanning electron microscopy (SEM), histology, image-assisted histomorphometry, and microcomputed tomography (μCT) imaging, it was demonstrated that incorporation of TGFβ1 in the hollow implant led to significantly increased osseointegration parameters (Figs 2.2e–l). Increased bone apposition, measured as bone-to-implant contact (BIC), to the titanium surface was observed in the 1 μg TGFβ1 gelatin and 100 ng TGFβ1 controlled release MS groups, as compared to 100 ng gelatin and placebo MS groups (Figs 2.2e–h, arrows). Concurrently, increased woven bone (WB) density, measured as bone volume/tissue volume (BV/TV), was seen within the macropores of the 1 μg TGFβ1 gelatin and 100 ng TGFβ1 controlled release MS groups, as compared to 100 ng gelatin and placebo MS groups (Figs 2.2i–l). The measured increases of 96% for BIC and 50% for BV/TV via control-released TGFβ1 over placebo MPs were comparable to bone ingrowth studies using growth factor adsorption, with the important difference that the control-release approach reduced the required drug


(b)

(a) 2.8 mm

4 mm

200 μm

No TGFβ1 (placebo MS) (e)

Ti

0

0

7

14 21 28 Day

1 ng/mL TGFβ1 controlled release (h) Ti

Ti 2 mm (k)

WB 500 μm

Ti 500 μm

Ti

25

2 mm (j)

WB

(d)

37

50

(g)

Ti

2 mm

(i)

(c)

75

1 μg TGFβ1 gelatin

100 ng TGFβ1 gelatin (f)

Ti

Cumulative release (%± S.D.)


2 mm (l)

WB

Ti 500 μm

WB

Ti 500 μm

2.2 Hollow titanium implants with a hollow core and macropores were custom fabricated (a). Poly-lactic-co-glycolic acid (PLGA) microspheres encapsulating TGFβ1 with an average diameter of 64 μm were constructed (b) that demonstrated release kinetics of a quick initial burst at three days and sustained release up to four weeks (c). To compare the outcome to a rapid release system, 1 μg or 100 ng of TGFβ1 was also adsorbed to a gelatin sponge carrier (f, g, j, k). After four weeks of implantation in rabbit humeri (d), controlled delivery of TGF-β1 led to increases in bone-to-implant contact (BIC) (e–h, arrows) and woven bone (WB) volume within macropores (BV/TV) (i–l) as compared to placebo control spheres (e and i). Approximately ten times more TGFβ1 was required for rapid release from the gelatin sponge (g and k) to obtain results comparable with controlled release using PLGA microspheres (h and l). Scanning electron microscopy (SEM) imaging. Ti, titanium implant; MS, microspheres.

dose by ten-fold (Sumner et al., 1995; Ehrhart et al., 2005; Schmidmaier et al., 2001). These results could have significant implications in potential reductions of cost and toxicity of in vivo delivered biological cues, and established a proof-of-concept for porous implants as delivery vehicles for controlled release of microencapsulated bioactive cues (Clark et al., 2008).

2.12

Future trends

Growth factor delivery for enhancement of orthopedic implant integration is a field just beginning to emerge, part of the larger fields of tissue engineering and regenerative medicine incorporating materials science, cell and


38


molecular biology, biochemistry, and bioengineering. Success in enhancing orthopedic implant integration and efficacy requires a concerted effort from all these areas in refining and improving the above discussed approaches. Biomaterials delivery systems are becoming more and more elegant every day. Poly-lactic-co-glycolic acid (PLGA) microspheres represent a standard in growth factor delivery. Depending on the formulation of PLGA microspheres, especially the ratio of PLA to PGA, controlled release at various rates from weeks to months can be achieved (Moioli et al., 2006; Cohen et al., 1991), and the polymer has already been approved as safe and efficacious for human use by the United States Federal Drug Administration (FDA). Further refinements to PLGA as well as other materials for drug delivery are being actively investigated, including but not limited to poly-ethylene glycol (PEG) derivatives (Stosich et al., 2007; Barralet et al., 2005), oligo(poly(ethylene glycol) fumarate) (Temenoff et al., 2004; Holland et al., 2005), chitosan microspheres (Cho et al., 2004), and hyaluronic acid (Kim and Valentini, 2002; Angele et al., 1999; Bulpitt and Aeschlimann, 1999). More advanced load-bearing biomaterials are also becoming increasingly available. New processing procedures for titanium have led to the advent of microporous titanium and titanium foams (Thelen et al., 2004; Wen et al., 2002), which have porosities conducive for bone formation while still retaining weight-bearing strength. Similar processes exist for ceramics as well (Kim et al., 2004; Baran et al., 2004). Materials such as these may provide hollow implants with more stable biomechanical loading properties, while still maintaining porosity for the delivery of bone enhancing cytokines. Numerous cytokines have been implicated at different spatial and temporal locations during bone growth and repair after injury, including the popular skeletal factors (TGFβs, FGFs, BMPs) (Gerstenfeld et al., 2003; Dimitriou et al., 2005; Kuroda et al., 2005) as well as new candidates (Wnts and Hedgehog) (Day and Yang, 2008; St-Jacques et al., 1999; Franceschi, 2005; van der Eerden et al., 2003). It is unlikely that any single growth factor is optimal in inducing bone repair and integration after titanium implantation. Adaptation of the PLGA microsphere system or use of other systems are beginning to be used for multiple growth factor delivery (Richardson et al., 2001; Simmons et al., 2004; Sumner et al., 2006; Holland et al., 2007). Not to be forgotten, vascularization is critical for successful bone repair, and angiogenic strategies are also being developed to be used alongside osteogenic ones (Richardson et al., 2001; Patel et al., 2008). Incorporation of various cell types alongside medical implants is quickly emerging as a viable treatment option. Osteoprogenitors or mature osteoblasts have been demonstrated to accelerate and enhance the osseointegration of titanium implants (Frosch et al., 2003). The discovery of mesenchymal stem cells capable of multipotential differentiation has been



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a major breakthrough for regenerative medicine, providing a source of highly expandable cells for cartilage, bone, and adipogenic applications (Caplan, 1991a; Alhadlaq et al., 2004; Williams et al., 2003). Mesenchymal stem cells (MSCs) could easily be incorporated into hollow implants alongside cytokines, possibly leading to inside-out bone formation by delivered MSCs and outside-in bone formation by MSCs recruited by delivered cytokines from the external environment (Clark et al., 2008). Although there is still much to be learned on how to control their osteoblastic differentiation, human embryonic stem cells (hESCs) (Thomson et al., 1998; Sottile et al., 2003) and the recently discovered induced pluripotent stem (iPS) cells (Yu et al., 2007; Takahashi et al., 2007) also represent candidates for bone regeneration applications.

2.13

Acknowledgements

This work is supported by NIH grants DE15391 and EB02332 to J.J.M. P.A.C is supported by a post-doctoral fellowship through the UW–Madison Stem Cell Training Program funded by NIH (5T32AGO27566-03).

2.14

Sources of further information and advice

Further information in the fields discussed can be obtained from various locations. Further exploration into biology can be found through various textbooks, such as for developmental biology, Developmental Biology by S. F. Gilbert, 1997, or newer editions (Gilbert, 1997), molecular and cellular biology, Molecular Biology of the Cell by B. Alberts, A. Johnson, J. Lewis, M. Raff, K. Roberts, and P. Walter, 2002, or newer editions (Alberts et al., 2002), bone biology, Principles of Bone Biology by J.P. Bilezikian, L.G. Raisz, and G.A. Rodan, 2002 or newer editions (Bilezikian, 2002), and growth factor biology, Skeletal Growth Factors by E. Canalis, 2000, or newer editions (Canalis, 2000). Further information on biomaterials for medical applications can be found through textbooks such as Biomaterials Science: An Introduction to Materials in Medicine by B. D. Ratner (Ratner et al., 1996) or newer editions. Principles of Tissue Engineering by R. Lanza, R. Langer, and J. P. Vacanti (Lanza et al., 2007) gives a thorough overview of the tissue engineering and regenerative medicine fields. Translational Approaches in Tissue Engineering and Regenerative Medicine by J. J. Mao, G. Vunjak-Novakovic, A. Mikos and A. Atala (Mao et al., 2007) provides a comprehensive coverage of cutting-edge science and technologies regarding implant integration, tissue engineering and translational approaches in regenerative medicine. Many organizations are rich with pertinent information. The National Institutes of Health (NIH, www.nih.gov) and its associated institutes, such


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as the National Institute of Arthritis and Musculoskeletal and Skin Diseases (NIAMS, www.niams.nih.gov), provide background on bone diseases as well as current therapies and approaches. The National Institute of Dental and Craniofacial Research (NIDCR) has a wealth of information regarding dental and craniofacial implants (www.nidcr.nih.gov). Many societies and foundations have been founded to further research in bone diseases and repair that provide background as well as recent advances in various fields – the American Academy of Implant Dentistry (AAID, www.aaid-implant. org), American Academy of Orthopedic Surgeons (AAOS, www.aaos.org), Arthritis Foundation (www.arthritis.org), National Osteoporosis Foundation (NOF, www.nof.org), Orthopedic Research Society (ORS, www.ors. org), Society for Biomaterials (SFB, www.biomaterials.org), and Tissue Engineering and Regenerative Medicine International Society (TERMIS, www.termis.org).

2.15

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3 Replacement materials for facial reconstruction at the soft tissue– bone interface E. W E N T R U P-B Y R N E, Queensland University of Technology, Australia; L. G R Ø N DA H L and A. C H A N D L E R-T E M P L E, The University of Queensland, Australia

Abstract: The challenges faced by any tissue repair and regeneration process resulting from either trauma or disease are many and complex. Although it is of course impossible to identify any one anatomical region as being the most demanding in this respect, the craniofacial region surely qualifies. The judicious choice of available, well-defined and tested repair materials to be used in the reconstruction process by the multi-disciplinary team of reconstructive surgeons is critical. This chapter addresses one aspect of facial reconstruction that has been less well addressed in the literature; namely the materials used to repair and regenerate soft tissue both in terms of fillers and in terms of materials used at the hard–soft tissue interface. Key words: soft tissue–bone interface, expanded polytetrafluoroethylene, surface modification, biomineralisation, bioresorbable fillers.

3.1

Introduction

In spite of the beauty and complexity of the magnificently engineered structure that constitutes the human face, it appears that humankind has never been ‘satisfied’ with their faces and how they ‘look’. In his fascinating book, Landau states that documentation dating back to the Minoan Bronze Age around 3500 bc confirms that altering the human head and face occurred in several civilisations and he speculates that more recent evidence dates the practice even further back in the mists of time. The human face is intrinsically linked to a person’s identity. Our fascination with the human face ranges from the artist who continually paints self-portraits – none more famous than Rembrandt who painted over 90 (Osmond, 2000) – to one of the most recognised faces of all, ‘The Mona Lisa’, and the speculation about her identity (Rising, 2008). Scientists such as the renowned Erik Erikson spent a lifetime exploring the meaning of identity; in fact the term ‘identity crisis’ is attributed to him. One aspect of his research involved 51 © Woodhead Publishing Limited, 2010

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Frontal bone

Parietal bone

Sphenoid bone

Nasal bone

Temporal bone Occipital bone

Zygomatic bone Maxilla

Mandible

3.1 Lateral view of skull showing principal bones.

the ‘loss’ of identity in World War II soldiers as a result of facial injury (Landau, 1989). The psychological repercussions and effects on a person’s life of severe deformities resulting from either trauma or genetic malformations cannot be underestimated. As evidenced by the recent global publicity afforded to the first reports of facial or ‘near-total’ facial transplants, interest in restoring the aesthetic appearance is seen as just as important as restoration of function (BBC, 2009; Gonzalez, 2008; AFP, 2008). Figure 3.1 shows some of the principal bones of the craniofacial skeleton which forms the foundations for the aesthetic features of the human face. To quote from a presentation by Fialkov et al. (2000) at the Bone Engineering Workshop, held in Toronto in 1999, ‘The morphology of the entire facial skeleton, in particular the upper region, has significant cultural, sexual, and social implications. Concepts of beauty, youth and intelligence are associated with particular facial morphologies, and vary with ethnicity. Much of this morphology is based on the underlying bony structure of the facial skeleton.’ The main functions of the human craniofacial skeleton are threefold. First, the multilayered bone framework provides protection for vital


Replacement materials for facial reconstruction

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3.2 Anterior cut away view of the face, showing the soft tissue (left), and a radiograph showing the underlying bone structure (right).

structures such as the ocular and aural systems, the central nervous system and the upper aero-digestive tract. Second, it also forms the foundations for the aesthetic features of the human face. The third and final function is that, through providing the structural framework, lever and fulcrum, it renders mastication possible. Another important component of one’s facial features is the soft adipose (subcutaneous fat) tissue that forms the interface between the bones and the skin. Figure 3.2 shows an anterior cut away view of the face showing the soft tissue as well as a radiograph showing the underlying bone structure. The anatomical interrelationship between these ‘structures’ is a complex one and this, of course, means that in order to fulfil their functions, their repair and regeneration are also interdependent. Plastic surgery constitutes an enormous and sometimes much criticised industry, but it also includes the extremely important challenge of trauma repair. One example where facial surgery may be required and which is particularly relevant to this chapter is the case of human immunodeficiency virus (HIV)-related lipoatrophy (LA). As far as can be ascertained, Carr et al. (1998) were the first to describe this relatively newly discovered condition (HIV-LA). It involves loss of facial soft tissues, leading to serious changes (other areas of the body such as buttocks and feet can be also


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affected) associated with generalised lipoatrophy, apparently triggered by the antiretroviral therapy. Many patients perceive it as a highly stigmatising manifestation of their HIV infection since loss of facial fatty tissue leads to a gaunt appearance and may lead to issues in work, social, and personal relationships. Another serious issue is that it has been reported that patients reduce their adherence to their antiretroviral medication regime in order to avoid this highly visible wasting effect, thus jeopardising their treatment (Collins et al., 2000). Whether by choice or as a result of trauma, it is clear that the challenges faced by any repair and regeneration (RR) process are many and complex. Although it is, of course, impossible to identify any one anatomical region as being the most demanding in this respect, the craniofacial region must surely be a strong contender. The long healing and repair process starts with the triage team before finally reaching the reconstruction surgical team. Even before the RR process can properly begin, there is a long and difficult road for the trauma patient. For non-medical specialists, even a quick perusal of books such as ‘Head, Face and Neck Trauma’ edited by Stewart (2005) or ‘Evaluation and Treatment of Orbital Fractures: A Multidisciplinary Approach’, edited by Holck and Ng (2006) brings home the enormity of the task, not to mention the costs involved. Bone and soft tissue injuries are not always of immediate priority when other potentially vision- or life-threatening injuries are involved. However, once the immediate medical aspects of the trauma have been satisfied then the multi-disciplinary team of reconstructive surgeons become involved. A multidisciplinary approach and multiple skills are required in the ultimate effort to restore both function and appearance requiring access to a range of accredited, well-defined and tested repair materials to be used in the reconstruction process. In order to develop clinically usable products, the reconstructive surgeons and indirectly the materials scientists and engineers, must become conversant not only with their own fields of expertise as well as with the anatomy and specific functions of the repair sites involved, but also with the healing process. In addition to fractures of the facial and orbital bones, injuries to eyes, nose, mandibular, and lachrymal systems and soft tissues such as cartilage, muscles, tendons, ligaments, and nerves may all have to be considered. A comprehensive coverage of all of these is beyond the scope of this chapter. Hence we shall focus on one aspect of facial reconstruction that has been less well addressed in the literature; namely an overview of some of the materials used to repair and regenerate soft tissue both in terms of fillers and in terms of materials used at the hard–soft tissue interface. We will focus on the current range of commercially available repair materials, some relevant current research aimed at improving these materials, and finally the next generation of repair materials and new RR strategies.



3.2

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Facial reconstruction

There is a plethora of useful books describing the relevant tissues as well as the complexity of their interrelationships. As a starting point an almost historic, compact and very readable text is ‘Anatomy of the Head, Neck, Face and Jaws’ by Fried. It provides the anatomical language necessary to form a deep understanding of all the tissue structures from bone and cartilage to nerves and muscles (Fried, 1980). From a clinical perspective the craniofacial skeleton is divided into upper and lower regions. The main function of the lower facial skeleton is occlusal and includes the masticatory structures of the mandible and maxilla, whereas the upper facial skeleton serves as a protective device for housing vital organs. In their chapter on ‘Strategies for Bone Substitution in Craniofacial Surgery’, Fialkov et al. (2000), in addition to describing the interacting forces involved between the muscles and facial bones, also specify the importance of the overlying soft tissues. The morphology of the face depends not only on the underlying bony structure but also on the subcutaneous fat or adipose layer that interfaces between the bones and skin. Hence, in any major trauma or malformation, the repair and regeneration of more than one tissue will be required. Soft tissue repair requirements are to some extent dependent on bone repair. Some of the important and relevant conclusions these authors arrive at are: in order to withstand the overlying soft tissue compression, bone substitutes, resorbable scaffolds and osteogenic carriers all require a degree of rigidity in order to be effective in the reconstruction and augmentation of the upper facial skeleton and secondly, the requirements of bone substitutes in the upper facial skeleton are different from other parts of the body (Fialkov, 2000). As will be discussed in the following sections, the same can be said for the requirements of many soft tissue substitutes that form the interface with the hard tissue. One exception is in cases where the soft tissue is repaired using subcutaneous injections of fillers such as autologous adipose tissue. One significant problem encountered in facial repair is the fact that autologous bone resorption, as well as the short to medium term lifetimes of many soft tissue replacement materials, lead to volume changes that are highly visible in the face. Hence, where the repair and regeneration of both soft and hard facial tissues is concerned, the search for the ‘ideal’ implant or material that will maintain both volume and contour has lead to the widespread use of alloplastic (see Section 3.3.2) materials. From the available literature it is clear that there are still aspects of many craniofacial repair procedures and the materials used that would benefit from further in-depth fundamental research. Before embarking on soft tissue repair, a brief description of the tissues involved is pertinent.


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3.2.1 Tissues at the bone interface Bone is a vascularised tissue consisting of cells and a mineralised extracellular matrix scaffold. The principal mineral component of the scaffold is carbonated hydroxyapatite which, although closely related to hydroxyapatite [Ca10(PO4)6(OH)2], is unique. The most abundant of the collagens present is collagen Type I, which is also the second highest component. It provides flexibility and structure to the tissue. Bone also contains collagen XI, V and III, as well as proteoglycans such as chondroitin, versican and syndecan, serum proteins, phosphoproteins and gamma carboxy glutamate (Siebel et al., 2006). The craniofacial skeleton is made up of two distinct bone types: membranous bone and endochondral bone. In membranous bone formation, which is responsible for the majority of the bones, ossification through direct mineral deposition into the extracellular produced matrix is followed by transformation of the mesenchymal cells into osteoblasts. Ultimately, this leads to the formation of the frontal, parietal, nasal, zygoma, maxilla and mandibular bones (Fig. 3.1). Endochondral bone formation, on the other hand, is responsible for the occipital bone, nasal septum and cranial base. Here the initial cartilaginous template is mineralised. Osteoclasts invade and are replaced by osteoblasts which ultimately lead to bone formation. As a result of extensive remodelling, at maturity there is minimal original skeletal bone and this has important implications for the development of the materials used in RR strategies. Cartilage is an avascular tissue and one of four mineralised tissues found in the body. It consists of an extracellular matrix comprised mainly of a three-dimensional hydrated network of collagen fibres in which are embedded chondrocytes and proteoglycans. Depending on the type of cartilage, it may or may not be mineralised. For example, articular cartilage contains about 60% collagen Type II and 40% Types I, VI, IX, X and XI (Walsh, 2006). The extracellular matrix (ECM) contains 4–7% (wet weight) proteoglycans (PGs), with aggrecan the most abundant. From a material and chemical point of view, cartilage is a hydrogel due to its high water content and hence it is capable of resisting pressure. The PGs interact with water which swells them, thus stabilising the tissue and imparting compressive stiffness. This contributes to the viscoelastic properties of cartilage. Because it is anaerobic it also has low oxygen consumption. Bone and cartilage are very different materials and the interface between them is an indistinct ‘calcified zone’. Although collagen Type II is the most abundant cartilage collagen, calcifying cartilage is enriched with collagen Type X and contains extracellular matrix vesicles. Currently, there are as many as 28 different collagen types described in the literature. During the bone mineralisation process, which is regulated by growth factors, cytokines and hormones, apatite deposition occurs onto the collagen Type



57

I matrix (this being the most abundant bone collagen). According to one source, the distinction between bone and cartilage was already recognised by Aristotle, who separated fish into either cartilaginous or bony categories (Hall, 2005). However, it was not until the 1690s and the invention of the microscope that researchers such as Leewenhoek and Havers were able to investigate the intricate microstructures of bone and cartilage (Hall, 2005). Polarised light microscopy and scanning electron microscopy (SEM) studies made it possible to establish the architecture (de Visser et al., 2008). Adult articular cartilage has a zonal architecture which is determined by the alignment of its collagen fibres. The three zones of alignment are: the superficial zone (closest to the articular surface) where the fibres are aligned parallel to the articular surface, the radial zone (closest to the bone) where they are normal to the surface, and the transitional zone (between the superficial and radial zones) where a continuous variation in average fibre orientation is observed. In addition, the chondrocytes change morphology from flattened and more aligned in the calcified zone at the bone interface to rounder in the superficial zone. Muscles are organs of motion: folklore has it that although it takes 17 muscles to smile it takes 43 to frown. They are mainly composed of muscle cells, which in turn contain myofibrils. The hierarchical structure means that these contain sarcomeres which are composed of the proteins actin and myosin. In most muscles, all the fibres are oriented in the same direction, running in a line from the origin to the insertion. They usually connect two or more anatomical structures, one of which is capable of moving: bone and skin, two bones, two parts of skin, two organs, etc. The function of muscles is intrinsically linked to ligaments, nerves, tendons, aponeurosis and fascia: for example, these control the energy for the muscle contraction and its direction. All the muscles of facial expression are attached to skin; at least at the point of insertion. In addition, they are all superficial. Fried (1980) points out that they have the particular characteristic of coming in a wide range of sizes, shapes and strengths. As well as conveying emotions and displaying facial expressions, they fulfil many important functions such as closing the eyes and moving lips and mouth during mastication. Interestingly, the facial expression muscles work synergistically rather than independently. Hence, any repair is complicated by the fact that they are interdependent in many of their functions. Any material used to replace or repair bone, cartilage or skin will clearly have some relationship with the muscles connected to them, not to mention interfaces (whether lesser or greater) with the ligaments. Adipose tissue (AT) is a specialised, loose, non-fibrous connective tissue composed primarily of fat cells called adipocytes. Its most important role is in energy homeostasis. In mammals, fats, usually in the form of triglycerides, are stored either in white adipose tissue (WAT) or (to a lesser extent in


58


adults) brown adipose tissue (BAT). WAT contains a single lipid droplet, while BAT contains numerous smaller droplets and contains many more mitochondria. The latter are responsible for the brown colour. BAT contains more capillaries due to its greater need for oxygen and is the main fat present in newborns. It is also responsible for generating body heat. Later, this is mostly replaced with WAT, although some BAT remains in the neck and intrascapula regions. WAT acts as an insulator and a protective layer around vital organs (think of the kidneys). In the face it fulfils another vital function where it forms a cushioning layer with interfaces between bone and skin helping to shape and contour the facial features. The skin which forms the external covering of the body is its largest organ, constituting 15–20% of its total mass and a surface area of between 1.5 and 2 m2 for the average human. The term ‘integumentary system’ is used when derivatives such as mammary, sweat and sebaceous glands are included. Skin is broadly categorised as thick or thin; a reflection on its location since it can vary from 1 mm to over 5 mm such as on the palms of the hands or soles of the feet. However, although these terms, which refer only to the epidermis, are really only of interest from a histological point of view, they are mentioned to illustrate the fact that skin can differ anatomically. As well as the two main layers – the epidermis and the dermis – there is a third underlying hypodermis fatty (subcutaneous fat) layer (see Fig. 3.3). This layer is highly significant in facial repair and reconstruction (RR) and is relevant to this chapter. The epidermis is composed of a keratinised, stratified squamous epithelium derived from the ectoderm. It grows continuously but, through the process of desquamation, maintains its regular thickness. One of its primary functions is as a protective barrier from the environment. The dermis is a dense connective tissue whose functions include imparting mechanical support and strength as well as thickness to the skin. Two distinct layers are clearly identified using light microscopy: the papillary and the reticular. Just beneath the reticular layer are found the layers of the hypodermis: lobules which contain varying amounts of adipose tissue separated by connective tissue septa, smooth muscle and, in some sites, striated muscle. This is the layer that insulates inhabitants of cold climates. But it is also the layer that varies greatly from individual to individual, bestowing on each of us a large part of our ‘identity’ based on how thin or not ‘quite so thin’ we are. This is particularly true where the face is concerned.

3.2.2 Organs of special senses: eye, nose and ear Vision, hearing, equilibrium and smell are all intrinsically linked to the anatomy of the head. While recognising the importance of all of these there



Stratum corneum

59

Sebaceous gland

Skin

Subcutaneous fat (SF) Soft tissue Veins and arteries

(SF)

Muscle Soft tissue– bone interface

Bone

Bone

3.3 Tissue cross-section from skin to bone, including soft adipose tissue.

does seem to be some consensus that vision is the most important of all the senses and its protection is vital. Although the ‘eye’ itself can be considered as consisting of essential components such as the bulb and the accessory organs (muscles, eyelids and eyebrows, lachrymal system, etc.), where facial trauma is concerned the bony orbit is critically important because this is where the eye is housed and protected. Broadly speaking, it consists of the orbital rim which is relatively thick and the orbital walls which are much more fragile. A recent book ‘Biomaterials and Regenerative Medicine in Ophthalmology’ edited by Prof. Traian V. Chirila, Queensland Eye Institute, Australia (Chirila, 2010) covers many aspects of the repair of this particular system. The external nose is mainly composed of bone, cartilage, skin and mucosa. Nasal fractures are the most common facial trauma (∼40%) and for whole body trauma they rank third highest overall after wrist and clavicle (Dev, 2008). The history of medicine, and more particularly the materials used to repair the body, make fascinating reading, even if their historical accuracy is sometimes somewhat doubtful. One of the most interesting is the story


60


of Tycho Brahe’s artificial nose (1566) (Van Helden, 1995). After losing his nose (or part of it depending on which account one reads) in a duel, he had a gold/silver prosthesis fitted using an ‘adhesive balm’ to keep it in place. This incident apparently kindled his interest in medicine, although he is best remembered for his astronomical discoveries. Another well-documented case of nasal repair is found in the Edwin Smith Papyrus, which is the only surviving copy of an ancient Egyptian book on trauma surgery. It consists exclusively of cases beginning with the head and working down. According to one report, of the 48 surviving case-reports only one resorted to ‘magic’ (Wilkins, 1964). Another case described involved the ‘repositioning of the deviated nasal bones’. Treatment included ‘the use of internal splints, firm external splints and dressings made from linen and grease and honey’. Further fascinating reading can be found in Lascaratos et al. (2003) ‘From the Roots of Rhinology: The Reconstruction of Nasal Injuries by Hippocrates’. Returning to the present, Oeltjen and Hollier (2005) give an excellent overview of ‘Nasal and Naso–Orbital Ethmoid Fractures’ in ‘Head, Face and Neck Trauma’ edited by Stewart (2005). In addition to hearing, the ear fulfils two other functions: equilibrium and cosmesis. The external ear consists of two major parts; the auricula or pinna and the external auditory meatus (Fried, 1980). Even focusing on the pinna and cosmesis issues, its repair is by no means trivial. It is composed of soft tissues, with cartilage being dominant. The chapter by Chang on ‘Auricular Trauma’ gives a good overview of typical trauma, treatment options and strategies (Stewart, 2005). Typically, microsurgical techniques have been used since the 1980s to reattach partially avulsed pinna. In cases where the total external ear is missing from birth, surgery (four operations over 2 years), is possible and most successful when the patient is 6–8 years old. When the loss is due to radical cancer surgery, amputation, burns and/or congenital defects, then auricular prostheses are available (UT Southwestern, 2009).

3.3

Materials used in traditional interfacial repair

Since in this chapter we are focusing principally on repair of the tissues underlying the skin and interfaced between it and the facial bones, as shown in Fig. 3.3, we now need to consider the materials and current approaches to facial soft tissue repair. There is a host of products used for soft tissue augmentation or replacement. Table 3.1 lists a selection of currently available materials. Our focus is on the three main classes of materials used either as fillers or for other soft tissue replacement applications. Depending on the genre of journal (or research), the descriptive words used in the classification of materials can vary. For example, the terms temporary, semipermanent or permanent are often used, but in biomaterials science the



5% PAA injectable gel 2.5% PAA injectable gel

Beads suspended in bovine collagen gel Silicone Silicone Silicone

PMMA Silicone

PTFE PTFE-graphite PTFE-aluminium oxide Expanded PTFE Expanded PTFE tubing Expanded PTFE tubing Expanded PTFE with silver and chlorhexidine (antibacterial) Expanded PTFE dual porosity Expanded PTFE saline filled

Description/use

Polyacrylamide gels (PAAG)

Biostables PTFE

Material

ArteFill Silastic Silikon 1000 Adatosil 5000

Aquamid Eutrophill

Teflon Proplast I Proplast II Gore-Tex SoftForm UltraSoft MycroMesh Plus Advanta Fulfil

Trade name

Table 3.1 A selection of polymeric materials used as soft tissue fillers or in soft tissue repair

Artes Medical Dow Corning Alcon Labs Bausch & Lomb

Contura International Lab Procytech

Atrium Medical Products Evera Medical

Dow Chemical/Dupont Vitek Vitek WL Gore & Associates Tissue Technologies Tissue Technologies WL Gore & Associates

Manufacturer


Rooster-derived hyaluronic acid Rooster-derived hyaluronic acid Bacterial or non-animal stabilised Bacterial or non-animal stabilised Bacterial or non-animal stabilised Bacterial or non-animal stabilised

Gelfilm (absorbable gelatin film)

Hyaluronic acid

Gelatin

hyaluronic hyaluronic hyaluronic hyaluronic

Bovine collagen Bovine collagen Cross-linked bovine collagen Human collagen Human collagen Cross-linked human collagen Human harvested autologous cells Human cadaver allogeneic collagen Human harvested autologous cells Porcine collagen

Stimulatory filler

Bioresorbables Polylactide

Naturally derived Collagen

Description/use

Material

Table 3.1 Continued

acid acid acid acid

Gelfilm

Hylaform Hylaform plus Restylane Perlane Prevelle Juvederm

Zyderm I Zyderm II Zyplast Cosmo Derm I Cosmo Derm II Cosmo Plast Isolagen Dermalogen Autologen Evolence

Sculptra/NewFill

Trade name

Pharmacia & Upjohn Ophthalmic Division

Allergan, Inc. Allergan, Inc. Medicis Medicis Mentor Allergan, Inc.

Allergan, Inc. Allergan, Inc. Allergan, Inc. Allergan, Inc. Allergan, Inc. Allergan, Inc. Fibrocell Science, Inc. Collagenesis, Inc. Collagenesis, Inc. ColBar LifeScience Ltd

Sanofi Aventis

Manufacturer


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terms ‘bioresorbable’ (degrades in vivo), ‘non-biodegradable’ (permanent) and ‘naturally derived’ (sourced from living organisms) are preferred. Recently, it has been suggested by one surgeon that another approach could be to consider them ‘according to the reactions they induce within human tissue’ (Nicolau, 2008). Seeking to control the rate at which bioresorbable materials degrade (or erode) in vivo is one that attracts an enormous amount of research since the processes involved greatly influence vascularisation and/or new tissue growth, both of which in turn influence the extent of tissue regeneration. It should be pointed out here that, in general, naturally derived repair materials (e.g. adipose tissue) degrade very fast and often in an unpredictable manner. In contrast, the degradation rate of synthetic bioresorbable materials can often be tailored for a specific application. Klein and Elson (2000) listed the requirements of a soft-tissue augmentation material as being: • • • • • • • • •

potentially of high use cosmetically pleasing have minimum undesirable reactions have low ‘abuse’ potential or abuse not leading to significant morbidity rates non-teratogenic non-carcinogenic non-migratory capable of providing predictable, persistent, reproducible correction FDA [presumably also other regulatory bodies] approved if non-autologous

Because of the dominance of the ‘facial cosmetic surgery’ market, even a superficial search reveals that there is a plethora of injectable materials (or fillers) described in a huge variety of journals which are used to ‘improve’ facial appearances or rejuvenate ageing features. Comprehensive lists can be found in various reviews such as ‘Soft Tissue Augmentation: A review’ (Fernandez and Mackley, 2006) and ‘Long-lasting and Permanent Fillers: Biomaterial Influence over Host Tissue Response’ (Nicolau, 2008). This ‘market’ and the controversies surrounding its regulation in different countries are not of concern here. However, since many of the materials used also play an important role in the repair of facial defects, scars and trauma injuries, these are being included. As Klein and Elson (as well as many others) pointed out currently, there are no repair materials that fulfil all the desired criteria (Klein and Elson, 2000; Nicolau, 2007). So it is not surprising that it is also generally recognised that the ‘perfect’ filler has still to be found. In the words of Hirsch in her recent review on ‘Soft Tissue Augmentation’ (Hirsch and Cohen, 2006), ‘even a judicious selection of the


64


perfect product for a given indication in a particular patient does not guarantee a perfect outcome’. Time and time again in the literature there are warnings about using particular products without having the necessary knowledge and expertise. However, one thing any surgeon or material scientist will agree on is that this is an ambitious ‘wish list’ for any material to possess! We shall now discuss the most frequently used materials in more detail. Table 3.1 lists a selection of soft tissue augmentation and filling materials. This unprejudiced list is by no means complete but is an example of representative materials for each of the classes to be discussed. We have included autologous materials such as fat and collagen but not botulinum toxin.

3.3.1 Naturally derived materials Although two of the most frequently used naturally derived materials, ‘autologous fat tissue’ and ‘collagen-based materials’, lie outside the main focus of this chapter, we include a short overview of each. The oldest and most frequently used autograft is in fact ‘autologous fat tissue’. Neuber first used it in whole graft form in 1893 (Neuber, 1893). Since the 1920s it has been used in injectable form, but with the advent of liposuction its use has greatly increased. There is much controversy surrounding its efficacy, in particular its longevity, and this is reflected even in the titles of some recent reviews: ‘Autologous Fat Transfer for Facial Recontouring: Is there Science behind the Art?’ (Kaufman et al., 2007) and ‘Fat Grafting: Fact or Fiction?’ (Calabria and Hills, 2005). It is used as a non-permanent material to fill small defects and for scar repair. Because of the large degree of resorption (30–60%), substantial over-correction is necessary. The harvesting and tissue preparation techniques used are well documented. Despite the many disadvantages, its use has persisted because its autologous nature means biocompatibility is not an issue. Naturally derived materials come from a variety of sources including autologous (human) and allogenic sources such as bovine, porcine, avian, as well as bacterial. Collagen was one of the first naturally derived filler materials given FDA approval (Sarnoff et al., 2008). Bovine-derived collagen is one of the most popular injectable soft tissue fillers and is generally considered the gold standard against which all other materials are measured (Homicz and Watson, 2004; Klein and Elson, 2000). This is not really surprising, since collagen itself is a main structural element of both hard and soft tissues in mammals. Bovine collagen-based materials have been used for over 20 years and the main disadvantages associated with their use are well documented: a propensity for hypersensitivity reactions and limited efficacy lifetimes. The need for skin testing requires multiple visits and hence delayed treatment times. Abscess formation, tissue necrosis and



65

granulations as a result of foreign body reactions have all been reported (Hanke et al., 1991). Some of the many autologous and allogenic collagenbased materials that have been developed over the years are expensive, customised products. One such commercially autologous product is Isolagen® which contains the patient’s own living dermal fibroblast cells (Isolagen, 2007). More recently, in their review, Homicz and Watson attest that hyaluronic acid-based materials such as Restylane® and the Hylaform® range of products appear to be a ‘safe alternative’ to bovine collagen-based materials (Homicz and Watson, 2004). However, even though hyaluronic acid structures are believed to be identical across species (for example it is isolated from rooster combs and streptococcus), traces of hyaluronic acid-associated proteins can cause allergic reactions (reactions currently reported as COOH >> CONH2 ⬵ OH >> NH2 >> CH3 ⬵ 0 (Tanahashi and Matsuda, 1997). In addition, grafting phosphate-containing monomers onto polyethylene biomaterials resulted in doubling the amount of HAP growth on the modified material in SBF whereas a smaller effect was observed for the acrylic acid grafted material (Tretinnikov and Ikada, 1997). Subsequent in vivo studies on the phosphate-modified polyethylene biomaterial showed a significantly enhanced interface of the implant surface with the newly formed bone. This was attributed to the phosphate groups providing nucleation sites for HAP growth (Kamei et al., 1997). The initial studies on modifying the surface properties of ePTFE facial membranes involved the acrylate monomer monoacryloxyethyl phosphate (MAEP, Fig. 3.6a), using simultaneous grafting in water. Very low graft yields were observed. However, it was clear, from x-ray photoelectron spectroscopy (XPS) and in infra-red spectroscopy investigations, that the phosphate-containing graft copolymer had been successfully introduced (Grøndahl et al., 2002). Subsequent in vitro mineralisation studies using SBF of both phosphate- and carboxylate-modified ePTFE substrates revealed that not HAP but rather the more acidic calcium phosphate phase brushite or monetite formed predominantly, Fig. 3.6a (Grøndahl et al., 2003). The similarities in mineralisation outcomes for these two functionalities were later explained by the complex nature of the MAEP graft copolymer and homopolymer (Suzuki et al., 2006). Indeed it was found that the MAEP monomer is hydrolytically unstable and therefore forms a copolymer of MAEP and acrylic acid. Subsequent studies involved graft polymerisation of the methacrylate monomer methacryloxyethyl phosphate (MOEP, Fig. 3.6b) in various © Woodhead Publishing Limited, 2010

Replacement materials for facial reconstruction O O O O

77

O O

P

OH OH

P O

O

10 μm (a)

OH OH

20 μm (b)

3.6 (a) MAEP-grafted ePTFE and (b) MOEP-grafted ePTFE showing brushite or monetite (a) and HAP mineral phases (b) respectively.

solvents (e.g. methanol and methyl ethyl ketone (MEK)) (Wentrup-Byrne et al., 2005; Suzuki et al., 2005; Chandler-Temple et al., 2010a). It was found that the graft morphology varied with the solvent used (smooth morphology when grafted in methanol and globular morphology when grafted in MEK) and this was attributed to the solubility of the graft and homopolymers in the respective solvents (Wentrup-Byrne et al., 2005). Although, based on XPS examination, the chemistry of the two graft copolymers appeared to be very similar (i.e. they were both copolymers of MOEP and HEMA), the in vitro mineralisation outcomes were very different (Suzuki et al., 2005). From infra-red spectroscopic analysis of the mineralised layer it was found that only the surface grafted in methanol actually induced nucleation and growth of HAP (Fig. 3.6b). Other modifications resulted in the growth of a mixture of calcium phosphate phases. It was later shown that this difference was in fact due to the degree of crosslinking in the graft-copolymer (Suzuki et al., 2006). Depending on the grafting conditions, the nature of the graft-copolymer varied (Fig. 3.7a or b) and it was shown that a low degree of crosslinking (Fig. 3.7b) was required in order for HAP to nucleate (Fig. 3.6b). In addition, it was found that the crosslinking reactions of the graft copolymer were due to the presence of large amounts of a diene impurity in the monomers used in these studies. Subsequently, a method was devised to produce the non-branched graft copolymer (and homopolymer) without incorporation of the diene impurity (Grøndahl et al., 2008). As a very nice conclusion to this mineralisation study, it was found that the ePTFE membrane containing the non-branched MOEP copolymer induced rapid and pure HAP growth in SBF (Grøndahl et al., 2010). In addition to these promising in vitro mineralisation studies, it has also been shown that the grafted ePTFE membranes produced increased protein © Woodhead Publishing Limited, 2010

78

Biointegration of medical implant materials (a)

(b)

ePTFE

ePTFE

3.7 Grafted PTFE copolymer showing (a) no crosslinking and (b) some crosslinking.

adsorption and osteoblast attachment, which are both important for increasing the interfacial strength (Suzuki et al., 2005). Our recent work has demonstrated that in vitro macrophage response is affected by the types of proteins that adsorb from serum and can be minimised by judicious choice of monomer and solvent combinations during the grafting process (Chandler-Temple et al., 2010b). In conclusion, through selective modification of the material substrate surface, membranes can be improved (made more bioactive) and this results in stronger interfacing at the facial soft-tissue– bone interface.

3.5

Future trends

It would appear that the most desirable materials are those that are both biostable and capable of integrating at the soft–hard tissue interface. Since structurally graded PTFE materials are already commercially available, this concept is one which lends itself to further exploitation and development. Currently there are plenty of on-going fundamental materials research studies, including the chemical modification of surfaces of several welldefined polymers already used in facial repair and reconstruction. New developments should ideally combine both these aspects with the latest biomedical engineering technologies.

3.6

Acknowledgements

The authors (EW-B, LG and AC-T) would like to thank everyone including their students who, over the years (in particular Drs AC-T and Shuko Suzuki), worked on research projects the results of which are included in this chapter; Dr Richard Lewandowski for his generosity over many years in sharing his expertise in plastic surgery; Drs Bernard Môle and Pierre Nicolau (Paris), both of whom took the time to make a helpful contribution to our discussion on facial atrophy. We thank Sybil Curtis for her drawings in Figs 3.1 and 3.3, Dr Oliver Locos for Fig. 3.4 and finally Patrick J. Lynch,



79

medical illustrator and C. Carl Jaffe, MD, cardiologist for permission to use Fig. 3.2 (http://creativecommons.org/licenses/by/2.5/). Thanks are also due to Professor Graeme George whose helpful discussions in the beginning helped to initiate this chapter and who proofread it at the end.

3.7

References

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4 Corneal tissue engineering Y. - X. H UA N G, Ji Nan University, China

Abstract: This chapter reviews the current position in the field of corneal tissue engineering. First, the construction and biophysical properties of the human cornea and the special conditions required for corneal tissue regeneration are illustrated. Then the strategies and approaches used for engineering corneal tissue are introduced, including the traditional way of engineering corneal equivalents by seeding cells in biodegradable scaffolds in vitro; the method of using active cell-free artificial corneas to induce corneal tissue regeneration in vivo; and the methods based on stem cells and cell sheet technology. Key words: corneal tissue engineering, corneal equivalent, active artificial cornea, stem cell, cell sheet technology.

4.1

Introduction

There are several reasons why the cornea is an ideal candidate for tissue engineering research. First, corneal disease is the second most common cause of blindness and there are over 10 million patients worldwide who remain blind from corneal disease through a lack of donors for corneal transplantation. Therefore, there is a strong desire to produce artificial bioactive corneal tissue substitutes to overcome the problem of corneal tissue failure. Second, its construction is less complex than most tissues so the reconstruction of engineered substitutes should be simpler. Third, since it is an avascular tissue with immune privilege, there should be fewer problems regenerating its tissue in vivo. Therefore numerous attempts have been made to engineer corneal tissue using different approaches. Significant progress has been made and engineered corneal epitheliums are already in clinical use. This chapter will review the current position in the field of corneal tissue engineering, and introduce the strategies and approaches used. To enable readers to gain a better understanding of the tissue properties and the specific requirements of corneal tissue engineering, we will also describe the construction of the human cornea; its biophysical properties, especially its optical properties which account for its particular function as a transparent window; and the special conditions for tissue regeneration. Finally, we will discuss future challenges and prospects for corneal tissue engineering. 86 © Woodhead Publishing Limited, 2010

Corneal tissue engineering

4.2

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Characteristics of the human cornea and its regeneration

4.2.1 Construction of the human cornea The cornea is the outermost layer of the eye and represents 7% of a human eye’s surface area. It forms a firm shell with the sclera to protect the components of the eye from infection and injury and also serves as a transparent window for external images to enter the eye. The shape of a human cornea is a concave–convex lens with a radius of curvature of about 7.7 mm at the front surface, but with a value of about 6.8 mm at the back. It is ∼0.5 mm thick at the center, increasing to about 0.7 mm at the periphery. For most people, the corneal curvature on the front surface is different in the vertical and horizontal directions. The horizontal meridian is usually more curved in young corneas, but this reverses as corneas age. When viewed from the front, the cornea has a diameter of 11 mm in the vertical direction, which is slightly smaller than its diameter of 12 mm in the horizontal direction. With its curved shape, the cornea acts like a lens to focus images on the retina at the back of the eye, but causes a certain amount of astigmatism owing to its toricity. Essentially, the cornea is a connective tissue containing collagen in the form of fibrils and it is an avascular tissue. However, it can obtain its physiological requirements from the lacrimal fluid in front of it and the aqueous humor behind it. Structurally, the cornea contains five parallel layers (see Fig. 4.1). The outermost layer is the epithelium, whose role is to absorb nutrients and oxygen while protecting the eye. The epithelium is composed of about five stacks of epithelial cells and has a 4–7 μm thick superficial tear film. The outer epithelium can be divided into three regions: the central cornea, the limbus and the conjunctiva. The avascular cornea is connected to the blood system in the corneal limbus. The blood vessels of the limbus are not responsible for the nutrition of the cornea, but release leucocytes and proteins into the corneal stroma, which plays a role in the diseases and tissue regeneration of the cornea and conjunctiva (Maurice, 1969; Reim, 1982). Below the epithelium is the Bowman’s membrane, which has a thickness of 8–12 μm. The Bowman’s membrane consists of randomly arranged collagen fibrils, mainly Types I, III, V and VII. Beneath it is the stroma, a dense layer of connective tissue making up about 90% of the corneal thickness. The main constituents of the stroma are water (78% of the total weight at normal hydration); collagen (mostly Type I: 15%); glycosaminoglycans (GAGs) (1%); noncollagenous proteins (5%, some of which are bound to the GAGs); and salts (1%) (Fatt and Weissman, 1992). The stroma is composed of a stack of approximately 200 collagenous lamellae, consisting of


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Biointegration of medical implant materials Epithelium Basement membrane

Bowman’s layer

Stroma

Descemet’s membrane

Endothelium

4.1 Construction of the human cornea (adapted from Reichl, 2008).

interwoven collagen fibrils. It contains a population of keratocytes that are derived from neural crest cells, and these occupy 3 to 5% of the total volume (Fatt and Weissman, 1992). The keratocytes are flattened fibroblasts that normally lie quiescently but can be activated in response to injury. They are responsible for the secretion of the unique stromal extracellular matrix, and the synthesis of collagen. Beneath the stroma is a thin limiting lamina called Descemet’s membrane and beneath that, finally, is the endothelium. The endothelium itself is a monolayer of non-dividing cells. The cells are about 20 μm in diameter and 4 to 5 μm thick (Fatt and Weissman, 1992), and they are separated from each other by spaces about 20 nm wide. The intercellular spaces can act as slits to allow different compounds, such as water and oxygen, to diffuse through and thus control corneal hydration, which is crucial for the maintenance of corneal transparency.

4.2.2 Biophysical properties of the human cornea Acting as a firm shell to protect the components within the eye, the cornea has sufficient mechanical strength to maintain excess pressure, which varies between 10 and 21 mm Hg for normal humans (Pinsky et al., 2006), and withstand external knocks and the forces applied by the extraocular muscles during eye movement. The physical properties of the cornea depend on position and direction. Its Young’s modulus was reported (Howland et al.,



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1992; Smolek, 1994; Vito et al., 1989) to be typically about 1–10 MPa, ranging from 2.45 × 104 to 5.7 × 107 Pa, depending on the techniques used for the measurements. It has a high tensile strength along the direction of the collagen fibril axes. Therefore, circumferentially, the cornea is stiffest at the limbus, with a tensile modulus of ∼13 MPa, about 13 times greater than in the horizontal direction (Ruberti et al., 2007). Radially, as viewed from the front, the cornea is stiffest in the central region out to about 4 mm, with maximum strength in the inferior/superior and nasal/temporal directions. In addition, the cornea has very low strength from front to back (i.e. through its thickness), but has high tensile strength along the direction of the collagen fibril axes. As a soft tissue, the cornea is a viscoelastic material. Its strain depends on the time for which a stress acts. There is also a certain amount of creep if the stress is prolonged. Human corneas have a long-term creep component when put under prolonged stress in vitro. As the window of the eye, optically the cornea has several features. Firstly, it is a specialized avascular tissue that maintains a highly organized architecture, but contains no pigments or molecules that would absorb visible light; therefore it is extremely transparent. Its transmittivity is as high as 95% for visible light. At the nanoscopic level, its transparency is achieved from the excellent arrangement of the corneal collagen fibrils within the lamellae. The fibrils are quite uniform in diameter (30.8 nm) (Meek et al., 2003) and are positioned relative to each other with a high degree of lateral order and a relatively constant interfibrillar spacing of 60 nm, which is much smaller than the wavelength of visible light (Fatt and Weissman, 1992). Such an arrangement causes destructive interference of scattered light and constructive interference of directly transmitted light for all the visible wavelengths (Ijiri et al., 2006). Secondly, its refractive index is 1.376. This, together with a convex air– cornea interface, makes the cornea the major focusing component of the eye, acting like a lens to focus incident light on the retina. Over two thirds of the eye’s focusing occurs at its front surface. Thirdly, it is precisely curved and has a smooth surface. Its surface is coated by a tear film that is 4–7 μm thick. Optically, the tear film provides a very smooth surface over the cornea. The precise curvature of the cornea and the optical smoothness of the tear film together make it a highly efficient converging lens. The tear film also transports metabolic products to and from the cornea, preventing the cornea from drying out and thus helping to maintain its transparency and refracting power. Both the transparency and refractive index of the cornea are reported to be critically dependent on the hydration of the tissue (Meek et al., 2003). One more property which should be considered in tissue engineering is the cornea’s permeability. Since the cornea is the main barrier to substances


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entering the eye, the extent to which different compounds permeate the cornea is an important biopharmaceutical parameter. The cornea’s permeability not only represents its barrier property for the penetration profiles of substances, but also influences its hydration. As mentioned previously, the cornea consists of the epithelium, stroma and endothelium, which form the three primary layers through which substances can permeate. For transcorneal transport, the corneal epithelium usually presents a much greater barrier than the stroma. However, the penetration profiles of substances through the corneal tissues depend not only on the cornea’s barrier properties, but also on the chemical nature, size and conformation, lipid/water partition coefficient, and degree of ionization of the permeant molecules (Bijl et al., 1997, 1998, 2000, 2001). Ionic and relatively hydrophilic substances will diffuse slowly through the epithelium, while lipophilic agents will diffuse slowly through the underlying aqueous stroma. The permeability of the human cornea to different substances has been extensively investigated (Fatt and Hedbys, 1970; Weissman et al., 1983; Myung et al., 2006). However, we will mention only results for some of the substances that might be involved in corneal tissue engineering. The water flow conductivity is reported to be 19 × 10−13 cm4 dynes−1 sec−1 at 37 °C with physiological hydration (H = 3.2) (Fatt and Hedbys, 1970). The oxygen permeability is 29.51 × 10−11 ml O2 cm2/sec mL mmHg (SD = 5.62) (Weissman et al., 1983), and the diffusion coefficient of glucose is 3.0 ± 0.2 × 10−6 cm2/s (Myung et al., 2006). It should be noted that the cornea’s permeability strongly depends on its hydration, since a greater water content can decrease the volume fractions occupied by glycosaminoglycans (GAGs) and GAG-associated proteins. An increase in tissue hydration can significantly increase permeability; the accompanying increase in tissue thickness further decreases the tissue’s permeability. Hydration also influences other properties of the cornea. It affects the pressure-induced radial straining of the tissue and hence plays a part in determining corneal shape (Hjordtal, 1995). An increase in tissue hydration can also induce a loss of corneal transparency due to an increase in light scattering from the uptake of water by the stroma. It also induces a change in the corneal refractive index (Meek et al., 2003). Therefore, corneal hydration is a key factor in the biophysical behavior of this tissue.

4.3

Special conditions for wound healing and tissue regeneration of the cornea

In any method being used for corneal tissue engineering, there should be a process for wound healing and tissue regeneration. Therefore, knowledge



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of how the tissue is built, how it grows, how it is repaired, and what kinds of factors/substances are involved in the process, is very useful. We should bear in mind that the cornea is a specialized avascular tissue that must maintain a highly organized architecture in order to maintain its clarity. Therefore, its repair processes include mechanisms that deal with these specialized conditions. Initially, wound healing initiates the ‘reexpression’ of a number of genes; then, as a result, proteins and cellular products are synthesized and interact in a well-orchestrated process (Dayhaw-Barker, 1995a). The following phases usually occur in the process of corneal wound healing and tissue repair: inflammation, epithelialization, fibroplasia, extracellular matrix (ECM) deposition, and remodeling. The substances involved in the process are mesenchymal matrix proteins, proteoglycans, growth factors, proteolytic enzymes, inflammatory mediators and several cell types (Dayhaw-Barker, 1995b). The mesenchymal matrix proteins are responsible for structural, cellular and extracellular adhesion. The main matrix proteins found within the extracellular matrix are: (i) fibronectin (FN), which plays a role in cellular adhesion during migration; (ii) tenascin, a protein that is mainly expressed during either tissue repair or malignant processes; (iii) the integrins, cellular receptors for the extracellular compounds such as FN, laminin (LN), and collagens; (iv) LN, an integral part of most basement membranes, whose role is to promote the adherence of cells to the basement membrane, thus affecting cell movement, differentiation, and growth. Besides these proteins, the most notable fibrillar elements are the collagens, which form a group of at least twelve proteins (Montes et al., 1984; Kuhn, 1987; Burgeson, 1988; Marshall et al., 1993) and can be separated into striated and filamentous collagens. The striated ones are composed of Types I, II, III, and V. The filamentous collagens are primarily Types VI, VII, IX, and X. Each type plays a different role in the function of the cornea, such as increasing tensile strength, serving in some structural capacity, or helping to establish an anchoring site. The dermatan and keratin sulfate proteoglycans are the main proteoglycans or glycosaminoglycans (GAGs) found in corneal tissue. The GAGs not only play a role in maintaining tissue hydration (Katz et al., 1986), but also assist in the spacing and orientation of collagen fibers (Borcherding et al., 1975), which are critical for the tissue’s transparency (Tuft et al., 1993). The growth factors that affect the cornea are: the defensins, which originate from inflammatory cells and have an antimicrobial effect and the ability to stimulate epithelial cell growth (Murphy et al., 1993); the epidermal growth factor (EGF) family; the fibroblast growth factor, which mainly promotes cellular in vitro proliferation and differentiation with possible neurotropic and angiogenic properties; the insulin-like growth factors, such


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as IGFe l, which stimulate mitogenic activity and differentiation; the neuropeptides, which play a role in the stimulation of certain epithelial cell migration patterns (Nishida et al., 1992); the platelet-derived growth factors, which can stimulate the expression of IL-8 mRNA and the fibroblastic production of specific types of collagen and general matrix formation; the transforming growth factors (e.g. TGF-β), which regulate the matrix proteins and their integrin receptors; and lastly the retinoblastoma-derived growth factor, which also plays a role in the stimulation of corneal epithelial cell growth. The enzymes involved in the restructuring of the cornea are the serine proteinases, which include plasmin and its related compounds (Setten et al., 1989; Tervo et al., 1989), and the metal cofactor-requiring enzymes (Fini et al., 1992), including collagenase and the matrix metalloproteinases (MMP). The serine proteinases are responsible for the cleaving of matrix proteins and may influence the orderly progress of cellular migration. The cells involved in corneal tissue regeneration are the epithelial cells, stromal keratocytes and endothelial cells (Dayhaw-Barker, 1995a). The epithelial cells are highly involved in any epithelial insult. The limbus surrounding the cornea is a reservoir of corneal epithelial stem cells. Stromal keratocytes remodel collagen fibers in the implant, heal damaged collagen fibers, and create adhesions in the stromal tissue. The endothelial cells play a key role in transporting ions from the cornea, thus performing the function of a dehydrating pump (Sumide et al., 2006) to maintain the clarity of the cornea. Besides the above-mentioned compounds and cell types in the cornea, it should be noted that there are also some substances needed for wound healing and tissue regeneration in the tear layer, such as the tear proteins lysozyme and lactoferrin, both of which play a role in the defense mechanisms of the cornea and influence its susceptibility to infections (Reim et al., 1997). In addition, since the maintenance of the epithelial cover on the stroma is largely dependent on the wetting of its surface, a tear layer is needed for repairing the epithelial, otherwise dellen and persistent epithelial defects will develop. Various kinds of graft are usually involved in most cases of corneal tissue engineering; the process of tissue regeneration in the grafts is quite different from that of pure wound healing. Therefore, in addition to the information mentioned above, knowledge of the biological response to the grafts, the possibility of graft rejection and the process of re-establishment of normal corneal tissue in the grafts is required. When a graft is implanted, whether it is a corneal graft from a donor, a tissue-engineered corneal graft, or just a synthetic/composite scaffold, successful transplantation and tissue regeneration involve two processes: (i) repopulation of the graft by host cells, especially if the donor cells are



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depleted during the cryopreserve procedure; (ii) the maintenance or regeneration of the graft’s transparency. It has been proved that repopulation of the graft by host cells always happens, whether or not there are donor cells in the graft. The replacement of donor cells by host cells seems to be a normal and gradual process. In the long term, all cell types of the donor tend to become necrotic and be replaced by recipient cells. The epithelium and endothelium are replaced relatively early, within the first post-operative year. However, the replacement of stromal keratocytes takes much longer. It has been reported that a small proportion of stromal keratocytes can even survive for several years but they are all replaced eventually (Wollensak and Green, 1999; Henderson et al., 2001; Binder et al., 1982; Karakami et al., 1991). The repopulation and replacement by host keratocytes can occur without opacification of the graft. The factors that might contribute to a shortened cell cycle or early necrosis of donor cells were suggested to be (Wollensak and Green, 1999; Binder et al., 1982): preoperative cell death in cadaver donor tissue, mechanical trauma during the operation, attrition from aging, the toxic substance from the glycerin-preservation process remaining within the donor tissue, the absence of epithelial stem cells in the transplant, neuroparalytic mechanisms, disturbed metabolism because of the circular scar tissue and the sutures, chronic corneal edema, post-surgical inflammation, elevated intraocular pressure or pseudophakia, and the process of freezing for a cryopreserved graft. A piece of research on rabbits (Karakami et al., 1991) reported that keratocyte regeneration usually occurred within 21 days of implantation. The regenerating ‘keratoblasts’ were characterized by active mitochondria and an increase in the rough endoplasmic reticulum. Keratocytes tended to accumulate along the interface between the host and the corneal graft on day 16 and exhibited marked accumulation on day 45, indicating that keratocytes actively participated in the adhesion and healing of the host tissue and the graft. Following the proliferation and repopulation of keratocytes, collagen synthesis took place. The collagen synthesized by regenerating keratocytes was found to be incorporated more quickly into mature collagen forms. All these results suggest that the regenerated keratocytes play a role in remodeling collagen fibers in the graft, healing damaged collagen fibers in the host stroma, creating adhesions in the host tissue and the graft, and the recovery of corneal clarity. Therefore, to obtain transparency and a suitable refractive index, the donor graft has to be repopulated with the host’s keratocytes, there should be a reestablishment of a normal corneal extracellular matrix, and the stromal collagens should be re-aligned in a regular fashion (Binder et al., 1982, 1986; Zavala and Krumeich, 1987).


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According to the reports of some pieces of research (Ohno et al., 2002; Zhang et al., 2007), the corneal transparency was maintained for just one month after transplantation. In particular, the cryopreserved graft appeared somewhat cloudy throughout the post-operative week. The significant decrease in the transparency of the graft was usually due to a malfunction of any one of the three primary parts of the cornea, such as the loss of epithelium and the decomposition of collagen soon after the operation. In the cryopreserved corneas, most of the keratocytes were killed during the process of freezing, so very few normal-looking keratocytes were visible. On the 10th post-operative day, enlarged, activated keratocytes started to migrate into the periphery of the graft from the host stroma. The fiber diameter and interfibrillar spaces among the collagen fibers became larger than normal with a change in the Bowman’s membrane, which may be one of the reasons why the graft appeared cloudy. Even in the corneal graft which was preserved at just 4 °C before transplantation, the number of keratocytes was reduced to approximately one-third of the number seen in normal corneas that have not been operated on. After this, superficial neovascularization could be observed to grow at the peripheral regions of the graft about two weeks after transplantation and recede gradually from the fourth week. At the same time, the transparency of the graft improved because of the regeneration of the epithelium and the reconstruction of the corneal stroma. By the 16th week after grafting, the uniformly thin collagen fibrils were organized into lamellae, and no significant differences were found between the transparency and neovascularization of the graft and those of normal tissue. A series of remodeling processes will happen after grafting, such as inflammatory cell infiltration, stromal cell proliferation, epithelial mitosis, migration, collagen deposition, and nerve regeneration (Zhang et al., 2007). During the remodeling process, the cells from the corneal graft will become functional by promoting the ingrowth of host cells and the regeneration of damaged nerves from the surrounding tissue. In addition, many cytokines from the stromal cells can regulate the proliferation, motility, differentiation, and possibly other functions of the epithelial cells and eventually accelerate the improvement in corneal clarity (Wilson et al., 1999, 2003). The normal cornea is an immunologically privileged site (Barker and Billingham, 1973; Katami, 1991; Streilein, 1995). There are multiple mechanisms for maintaining its immune privileged status (Niederkorn, 1995, 2002; Niederkorn et al., 1989; Steinman and Nussenzweig, 1980; Tommila et al., 1987), including the lack of blood vessels, lack of lymphatics, blood–eye barrier, relative paucity of mature antigen-presenting cells (APCs) in the central cornea, presence of immunomodulatory factors in the aqueous humor, and the constitutive expression of CD 95 L (Fas ligand) within the



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eye (Panda et al., 2007; Osawa et al., 2004; Prendergast and Easty, 1991; Ray-Keil and Chandler, 1985, 1986; William et al., 1985). This privilege can be destroyed by inflammation and neovascularization. Therefore, steps have to be taken to reduce the risk of inflammation post grafting. Usually, an immunological reaction can be induced only up to six weeks after transplantation, corresponding to the time needed for the replacement of stromal cells. After this period, the donor stromal cells are either no longer present (Maumenee, 1951) or are present only in small amounts (Wollensak and Green, 1999). After an immunological reaction, corneal graft rejection may take place. Corneal graft rejection is defined as a complex immune-mediated process resulting in decompensation of the transplanted cornea. The process is initialized by recognition of the foreign histocompatibility antigens on the cells of the corneal grafts by the host immune system, which results in host sensitization. This is followed by a specific immune response against these antigens, which finally results in the decompensation of the graft tissue. We want to emphasize that corneal graft rejection is primarily a cell-mediated response; exogenetic cells would not only induce immunoreaction, but would also be obstacles to the repopulation of the host’s keratocytes and new tissue generation. Therefore, from this point of view, a graft without donor cells would have a lower risk of graft rejection.

4.4

Approaches to corneal tissue engineering

Recent approaches to corneal tissue engineering can be basically subdivided into four methods, according to their strategies or concepts. The first one, a traditional strategy, aims to seed and culture specific cell types in a biodegradable scaffold with a 3-D structure. The second strategy is based on the use of active scaffolds that closely mimic natural extracellular matrices, which are used in combination with signal biomolecules to induce corneal tissue regeneration in vivo. The third method is purely cell based, involving the dropping/injection of stem-cell suspensions or the transplantation of the cultured cells with supporting materials into a defect site or an injured corneal tissue. The fourth method focuses on the transplantation of autologous or allogeneic cell sheets without using scaffolds.

4.4.1 Traditional: cell-seeded biodegradable scaffold-based tissue engineering The strategy of the traditional method of corneal tissue engineering is to reconstruct a living corneal tissue in vitro by seeding and culturing cells in a biodegradable scaffold with biomolecules that promote the regeneration of tissues. The first step is the isolation and culture of cells. To recon-


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struct complete corneal tissue with all the three major cellular layers mentioned in Section 4.2.1, all the cell types in the corneal tissue (the epithelial cells, stromal keratocytes and endothelial cells) are needed. Since the stem cells located in the limbal portion surrounding the cornea exhibit extensive proliferation potential (self-renewal capacity) and appropriate differentiation abilities, they are usually taken as the best cell source for epithelial cells. The epithelial cells can be isolated from the limbus by dispase digestion using the method described by Green (Rheinwald and Green, 1975) that promotes epithelial cell proliferation. Human corneal keratocytes, on the other hand, can be obtained from the keratocyte contamination in epithelial cell cultures by changing the culture medium for a keratocyte culture medium. Since the medium is not adequate for the proliferation of epithelial cells, only keratocytes will be left in the subcultures. Of course, corneal keratocyte cultures can also be established from the corneal stroma (Germain et al., 2000). Human corneal endothelial cells are usually isolated by the enzymatic treatment of excised corneas (Engelmann et al., 1988). After obtaining the cells, the second step is the production of the stroma tissue. Different kinds of material have been used as the scaffold for reconstructing the stroma tissue, such as human Type I and III collagen (Germain et al., 1999), bovine Type I collagen (Minami et al., 1993), polyglycolic acid (Hu et al., 2004) and a collagen–chondroitin sulfate substrate (Griffith et al., 1999). The reconstructed stromal tissue is produced by mixing corneal keratocytes with the materials, then pouring the mixture into a Petri dish containing an anchorage ring and incubating for gelation. Fibroblasts reorganize the extracellular matrix on the culture into a better substrate for epithelial cells. The stroma reconstructed in this way resembles the corneal stroma but the Descemet’s and Bowman’s membranes are absent (Germain et al., 1999). The next step is to reconstruct the layers of the epithelium and endothelium. Epithelial and endothelial cells are seeded and cultured respectively on the top and the bottom of the reconstructed stroma under either AIC (air-interfaced culture) or LCC (liquid-covered culture) conditions for several more days (Germain et al., 1999; Schneider et al., 1999). The resulting corneas have been found to be histologically similar to a native cornea and express components of the epithelial basement membrane at the epithelium–stroma junction. Although some of the engineered corneal tissues constructed in this way have shown morphology similar to a native cornea, including well-defined epithelial, stromal, and endothelial cell layers, they have been used mainly as models for physiological, toxicological, or pharmacological investigations, as they are not yet ready to be used for transplantation. The reason is that most of them do not have enough tensile strength for surgical manipulation and



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fixation. A detailed study of these kinds of corneal equivalent is still needed to further improve their tensile strength for surgical applications.

4.4.2 Active scaffold which induces corneal tissue regeneration in vivo As mentioned in the last section, corneal equivalents constructed using the traditional method of tissue engineering do not have enough tensile strength to form a surface curvature or to allow surgical manipulation and fixation for transplantation. In addition, in order to construct corneal equivalents, several cell lines (epithelial cells, stromal keratocytes and endothelial cells), at least, have to be prepared and cultured for the regeneration of the different layers of the cornea. The processes of culturing and proliferating each cell type on substrates are complicated and time consuming. Furthermore, even if such a corneal equivalent can be built and has enough tensile strength to be transplanted into a recipient’s eye, we have learned from Section 4.3 that to obtain transparency and a suitable refractive index, the donor graft has to be repopulated with the host’s cells. The previously seeded allogeneic cells in the graft tend to be necrotic and are replaced by recipient cells just after transplantation. They may be acting as a biological medication and stimulate proliferation of the recipient’s residual cells in some way, such as producing extracellular matrix proteins as well as cytokines and growth factors. However, while they induce immunoreaction, the dead donor cells will also be obstacles to the repopulation of the host’s cells. Moreover, in most cases, cell-seeded grafts have to be cryopreserved before transplantation. It has been proved that most of the keratocytes will be killed during the process of freezing; even in a corneal graft which was preserved at just 4 °C before transplantation, the number of keratocytes was reduced to approximately one-third of that seen in normal corneas that have not been operated on (Ohno et al., 2002). Therefore, the cultured donor cells are not actually necessary, since the positive role they may play, if any, comes mainly from what they produce during culturing, such as extracellular matrix proteins, cytokines and growth factors. On the other hand, some reports about the continuous maturation and differentiation of immature or incomplete tissues after transplantation (Compton et al., 1989; Germain et al., 1995) suggest that there is a sophisticated regulating system in the human body that guides and controls the growth of different cells, with each kind of cell being guided to grow in their proper location to form different tissues. Therefore, it would be better to let the human body do the work as a bioreactor to culture cells and generate new tissues in vivo. Based on these considerations, in 2002, Huang et al. proposed a novel strategy for corneal tissue engineering (Huang and Li, 2002a, b, c, 2003,


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2007). They suggested using a cell-free active artificial cornea to induce new corneal tissue formation in vivo. The active artificial cornea they produced had not only a similar composition (such as collagen and extracellular matrix) to the human cornea and the required optical and biomechanical properties, but also some signal molecules such as growth factor. They believed that an acellular scaffold with such properties could induce the recipient’s cells to grow in it then repopulate it and generate new corneal tissue, as happens in an implanted corneal graft. The in vivo ‘bioreactor’ can provide not only the most suitable environmental conditions, but also all the substances needed for tissue regeneration as mentioned in Section 4.3. Since there were no dead donor’s cells that needed to be removed to let the host cells migrate in, the tissue regeneration of the acellular graft was expected to occur in an even shorter period than for a corneal allograft (Huang and Li, 2007). The active artificial cornea was designed to mimic the human cornea and was composed of collagen, chitosan and glycosaminoglycans (GAGs). Collagen is the major structural protein in the human corneal stroma and has many desirable features. It has low immunoreaction and the ability to promote normal tissue regeneration. Chitosan has excellent biocompatibility with the human body. Its interaction with collagen can inhibit collagenase and protease to reduce the biodegradation of the collagen. It also has a high tensile strength to support the formation of collagen into different shapes. The third material, glycosaminoglycan (GAG), was cross-linked to form the extracellular matrix for cell adhesion, migration, proliferation and tissue remodeling. It also gave the new artificial cornea a higher elastic modulus and a more porous structure. Type I collagen powder and chitosan were dissolved in water of pH < 4.5, and the concentration of collagen was adjusted to 1% by acetic acid aqueous solution. The solution was stirred by a homogenizer for 30 min, and then the GAG was slowly added and homogenized for 45 min. The GAG/ collagen–chitosan homogenous solution was spread on a glass plate and heated at 35 °C for 48 h to obtain a completely dry membrane. The dry membrane was then immersed into deionized water for l h to be modified. The artificial cornea displayed similar optical and mechanical properties to a human cornea. It had an optical transmittance of ≥95%; a refractive index of 1.368∼1.382; a tensile strength of 167 kg/cm2 (dry state); an ultimate elongation of 300%; and its degree of swelling in water was 65–75%. With such physical properties, the artificial cornea substrate can easily be constructed into different shapes and curvatures for transplantation (see Fig. 4.2), and will be well tolerated by the recipient. Animal (rabbit) artificial corneal implantation was performed to prove the efficacy of this approach. The implants were prepared in the shape of



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4.2 Artificial cornea lenticules with the function of inducing corneal tissue regeneration in vivo. The diameter of the lenticules is 8.05 mm, the radius of their surface curvature is 6.76 mm.

4.3 Specimen taken on the 60th day. Some host’s keratocytes accumulated along the interface between the host and the graft, and grew into the superficial lamellae of the implant; no inflammatory cell was observed.

discs with a diameter of 8 mm and thickness of 0.3 mm, then rinsed and aseptized in a sterile PBS buffer. In the implantations, the rabbit corneas were excised down to the middle of the stromal layer at the center of the cornea. The active artificial cornea discs, after being administered with a growth factor (bFGF), were inserted into the recipient corneal beds, and two interrupted nylon sutures were used to close the incision. Although the animals were given the injection against anti-immunologic rejection only on the 15th day, the corneas of the animals were observed to maintain a smooth surface and clear stroma as well as transparency. The eyes were clear and had only a little blood-vessel hyperplasia on the 10th day, and no inflammation and hyperemia on the 30th day. A specimen taken on the 60th day, shown in Fig. 4.3, indicated that the implant and the native corneal tissue had good compatibility and connection. Some keratocytes could be seen to accumulate along the interface between the host and the graft and these grew into the superficial lamellae of the implant, while no inflammatory cells were observed. The biodegradability measurement showed


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4.4 Specimen taken on the 120th day post-operatively. Only a small portion of the implant was still degrading, accompanied by metrocytes growing at 4 months. New corneal tissue formation can be observed in the region taken by the implant originally (inside the dashed frame). The newly generated tissue and the host’s deep stroma have the same tinctorial properties; its collagens have good regularity and integration with the native ones.

that there was 30% biodegradation between the 45th and 60th day post-operatively. Figure 4.4 shows a specimen taken on the 120th day post-operatively. It can clearly be seen that the implant was almost totally degraded. In the region occupied by the implant originally (the region in the dashed frame), new corneal tissue had been generated by the repopulating keratocytes, and there were plenty of corneal metrocytes in the zone where the artificial cornea was degrading. The newly generated tissue and the host’s stroma had the same tinctorial properties; its collagens had good regularity and integration with the native ones. The follow-up observation showed that the new corneal tissue generation and implant degradation were completed between the 120th and 150th day. Complete tissue regeneration in the control group using corneal allografts as implants required an extra month. The new generation of the corneal tissue was proved by the specimen taken 6 months after the transplantation. It was about 0.3 mm thicker than the normal (control) cornea. This indicates that it really did contain newly generated cornea tissue. Slit-lamp microscope imaging on the eye of the animal just before sacrifice 5 months post-operatively showed that the



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cornea was clear, smooth and homogeneous in thickness. Animal experiments of lamellar keratoplasty using the novel artificial cornea also showed similar results, and complete re-epithelialization on the surface of the cornea with the implant was found on the 5th day after transplantation. In later experiments performed in a similar way by other authors, using materials such as hydrated collagen and N-isopropylacrylamide copolymerbased ECMs to fabricate the implants, tissue regeneration was also observed in the acellular implants in rabbits and mini-pigs (Griffith et al., 2008; Liu et al., 2006). The efficacy of this approach was further proved. In summary, the strategy of using an active acellular scaffold to induce host cell repopulation and new cornea tissue generation in vivo opens up a new approach to corneal tissue engineering. Since there was no exogenetic cell, the active scaffold had only a small rejection reaction, mainly induced by the suture after implantation. On the other hand, its GAG and growth factor components had an anti-inflammatory effect and gave the autologous cells the signal to grow, so new tissue was generated faster in the active acellular scaffold than in the one using homologuous corneal lamina for the implants. As we mentioned previously, corneal transplantation has a special requirement that other tissue transplantation does not: to obtain transparency and a suitable refractive index, the donor implant has to be repopulated with the host’s keratocytes, and the stromal collagens should be re-aligned in regularity (Binder et al., 1982, 1986; Zavala and Krumeich, 1987). So no matter whether it is a homologuous cornea or a cornea equivalent produced in a typical tissue engineering process, they both have to undergo such a process. They have no advantage over an artificial cornea without cell culturing before transplantation. On the contrary, their exogenetic cells may induce immunoreaction and would be obstacles to the repopulation of the host’s keratocytes and new tissue generation. Therefore, a cell-free artificial cornea that induces new corneal tissue generation in vivo would be a better method for corneal transplantation if the host can still generate keratocytes. Its advantage is fairly obvious: not only can it be used universally for different implantation surgeries for various people, but it is also suitable for long-term storage and longdistance transportation. In addition, by having similar optical and mechanical properties to the human cornea, the novel artificial cornea is well tolerated by the patient and is easy to prepare and process into different shapes and sizes on a large scale. It can also be processed into lenticules with different optical powers for epikeratophakia. Its immediate applications include increasing the thickness of a cornea for patients with a thinner cornea or critical corneal thickness, resulting from radial keratoplasty or other causes such as wearing contact lenses. It can also be used for refractive keratophakia.


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4.4.3 Stem-cell tissue engineering It is well known that stem cells have the unique capability to self-renew and to generate committed progenitors. They can differentiate into the cell lineages of the tissue of origin after a limited number of cell divisions, so stem cells have been widely used for engineering different tissues, including corneal tissue. For corneas, the self-renewal of corneal epithelium can usually be achieved by the stem cells located in the limbal region of the cornea. However, some pathological ocular conditions, such as severe chemical or thermal burns, genetic disorders, microbial keratitis, collagen vascular diseases and immunological disorders, can cause permanent damage of these stem cells. When limbal stem cells are destroyed, the conjunctival epithelial cells may migrate to cover the denuded corneal surface. Such a resurfacing would result in an irregular opaque ocular surface as well as stromal neovascularization, chronic inflammation, and persistent epithelial defects. Conjunctivalization of the cornea may lead to irreversible corneal scarring and impaired vision. The traditional way to repair the corneal epithelium and restore vision in these patients was to perform limbal tissue transplantation or combined limbal corneal transplantation, but this was restricted by a shortage of donor corneas. Therefore, a method of engineering corneal tissue from allogeneic or autogenetic (in case of unilateral limbal stem-cell deficiency) stem cells was developed to overcome this difficulty (Lehrer et al., 1998; Sangwan et al., 2003; Schwab et al., 2000b; Tsai et al., 2000; Sun and Green, 1977). Compared with whole organ transplantation, this method has several advantages: it can provide a potentially unlimited resource for corneal epithelial stem-cell transplantation; it is an easy, safe and less painful procedure; and it can offer the potential for in vitro manipulation of the graft prior to implantation to eliminate unwanted features such as the depletion of antigen-presenting cells from cultivated cell populations. The clinical use of cultured limbal stem cells was first reported by Pellegrini et al. in 1997 (Pellegrini and Traverso, 1997). Since then, many reports of the clinical use of this technology have been published, and in several clinical centers it has become a routine stem-cell therapy for the permanent regeneration of a corneal epithelium in patients with limbal stem-cell deficiency. According to Pellegrini, epithelial cell cultures (Pellegrini and Traverso, 1997) were established from 1–2 mm2 full-thickness biopsy samples taken from the ocular surface (bulbar conjunctiva, cornea, and limbus) of donors. The biopsy samples were minced and treated with trypsin at 37 °C for 3 h. The cells were then plated on a feeder layer of lethally irradiated 3T3-J2 cells (2.4 × 104/cm2) and cultured in 5% carbon dioxide in Dulbecco-Vogt



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Eagle’s and Ham’s F12 media (3:1 mixture). Before transplanting the cultured limbal epithelial cells, the conjunctival epithelium covering the cornea and limbus of the injured eye had to be removed. After placing the cultured epithelial graft on the prepared eye, a soft therapeutic hydrophilic contact lens was placed over the graft to provide an efficient protective bandage to shield the graft and provide a structural framework upon which the regenerating epithelium could grow. The transplantation of limbal epithelial cells cultured from an organ donor has to be performed in combination with immunosuppressive therapy, and the approach has a very low success rate in the long term, as compared with autologous cells. Interestingly, it was found that the majority of successful cases eventually showed only recipient DNA in the regenerated epithelium (Henderson et al., 2001). This is not so surprising, since we have mentioned in Section 4.3 that over the long-term in tissue regeneration, all cell types of the donor tend to become necrotic and be replaced by recipient cells, though this time, the recipient’s limbal stem cells were believed to be destroyed, and some of the cases even had total limbal stem-cell deficiency. Therefore, it was suggested that allogeneic cultures were acting as a biological medication and actually stimulating the proliferation of residual stem cells of the patient (Pellegrini et al., 2007). Owing to the fragility of the epithelium, the clinical application of cultured human limbal cells in this way could not be reproduced on a large scale (Pellegrini and Traverso, 1997). Therefore, some supporting materials have been used for cell culture, transportation and transplantation onto patients, such as fibrin glue (Rama, 2001), amniotic membrane (Koizumi, 2001; Sangwan et al., 2003; Schwab et al., 2000; Tsai et al., 2000), collagen sponges or strips, and devitalized membranes or polymers (Nishida et al., 2004b). Although providing the cultured epithelium with a basement membrane is likely to improve graft ‘take’, and may even promote the survival of cultured stem cells (Daniels et al., 2001b), alternative culture methods have, nevertheless, been tried. Many efforts have been focused on the removal of the feeder layer and/or fetal calf serum, on the assumption that there are potential risks from the use of these supporting materials. Another idea being considered is that using new culture media or cultivating limbal cells onto different carriers, even with cells in suspension, may help the transplantation to be performed by the dropping or injection of stem-cell suspensions into a defect site or an injured corneal tissue. In one of the approaches not using a feeder-cell layer (Qian et al., 2001), the cells were cultured in the following way: freshly excised limbal tissue was immersed in Dispase II (1.2 unit/mL) for 4 hours at 37 °C under 5% CO2. Epithelial stem cells were collected by scraping them off the limbal tissue. The cell suspension was then centrifuged at 1500 rpm for 10 min and


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washed in MEM once. The cell pellet was re-suspended in the culture medium and seeded in a 36 mm culture dish. Cells were cultured at 37 °C under 5% CO2 for 2 days with a culture medium. The medium was then changed to a serum-free keratinocyte culture medium for another 10 to 14 days. When the growth of epithelial cells reached a confluent layer (1–3 × 105) in 10 to 15 days from one corneal strip, cells were released by trypsin and transplanted on the 14th day after limbectomy. An average of 3 × 105 cells in 0.2 mL of limbal cell culture medium were transplanted in the form of eye drops. Cells were transferred to both the conjunctival sac and a soft therapeutic hydrophilic contact lens, which was then placed to cover the ocular surface. No significant graft rejection was found with allogeneic stem-cell transplantation, even in the absence of immunosuppressive therapy. The method seems promising, since corneal epithelium was regenerated in 30 out of 36 (83.3%) allograft and 35 out of 41 (85.3%) autograft eyes. However, it has been used only on rabbits and its clinical performance on humans has not yet been proved. Until now, the main aims of stem-cell therapy have been to promote the re-epithelialization of the cornea, provide a stable epithelium, prevent regression of new vessels, and restore epithelial clarity. Can similar methods be applied to the stroma and endothelium of the cornea? Does the presence of a graft acting as a biological dressing reduce the hostility of the local environment such that any remaining host stem cells can survive and function? Why does a transplanted cell migrate and differentiate appropriately or not, and how does the environment stimulate such parameters? These are questions which require further investigation. Human embryonic stem cells (hESCs) could be a potential source of corneal cells for generating corneal tissues for transplantation. The advantages of using hESCs are obvious, such as the possibility of studying the pathways involved in cell-lineage differentiation and having an unlimited source of corneal epithelial cells for transplantation. Committed stem cells have been obtained from mouse embryonic stem cells for corneal reconstruction in mice (Ahmad et al., 2007) and corneal epithelial cells have also been produced from hESCs (Homma et al., 2004). It has even been proposed that stem cells could be transported to diverse tissue compartments via the blood stream, and that on arrival the cells would generate the appropriate cell types in response to local signals (Daniels et al., 2001a).

4.4.4 Corneal tissue engineering based on cell sheet technology Although the potential for stem-cell tissue engineering is significant, there are still some critical problems. For instance, with the dropping/injection of single-cell suspensions, in many cases the implanted cells cannot be retained



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around the target tissue, so it is difficult to control their size, shape and location after the transplantation. The use of supporting materials makes handling the cell culture easier, but their presence can potentially influence the post-operative clinical outcomes. For example, an amniotic membrane would persist between the corneal stroma and the expanded epithelial cells after transplantation (Kinoshita et al., 2004), and could even affect the optical transparency of the fabricated constructs. Although fibrin gel degrades rapidly after transplantation into the corneal stroma, its biodegradation can induce inflammation and thus might possibly result in microtrauma to the corneal stroma. Moreover, the possibility of infection from the use of biological carriers cannot be completely excluded (Yang et al., 2006). To solve these problems, an alternative method called cell sheet technology has been developed, which involves the creation of carrier-free constructs that can still be easily manipulated during surgical operations (Yamada et al., 1990; Okano et al., 1993, 1995). Cell sheet technology mainly employs a ‘thermo-responsive culture dish’ to enable reversible cell adhesion to and detachment from the dish surface by the controllable hydrophobicity of the surface (Yamada et al., 1990; Okano et al., 1993, 1995). The technology enables a non-invasive harvest of cultured cells to be made as an intact monolayer cell sheet, including the deposited extra-cellular matrix (ECM). The monolayer cell sheet can be collected simply by reducing the culture temperature to below 32 °C for less than 30 minutes, without using any enzymes or chelating agents. The technology also enables the direct transplantation of cell sheets into host tissues without scaffolds, fixation, or sutures. The main material used for the thermo-responsive graft-cell culture dish is poly(N-isopropylacrylamide) (PIPAAm). The temperature-responsive polymer exhibits thermo-responsive hydrophobicity changes in aqueous solutions (Heskins et al., 1968). By means of electron beam irradiation of an IPAAm monomer on polystyrene dishes, the monomer is polymerized and covalently bonded with the dish surface. At temperatures below 32 °C, PIPAAm molecules are highly hydrated, so the PIPAAm grafted surfaces are hydrophilic. However, when the temperature rises above 32 °C, the surfaces suddenly change to hydrophobic due to the extensive dehydration of the PIPAAm molecules. When cells attach to a material surface, they physically adhere onto hydrophobic surfaces better than onto hydrophilic surfaces. After a period of passive adhesion, cells change their morphology to spread due to metabolic processes using ATP. On the hydrophobic surface of a thermoresponsive culture dish, adhered cells can proliferate normally to confluency and express normal phenotypic markers for each cell type. Spread cells rarely change their shape to a round one in order to detach from the hydrophobic surface without the release of surface-engaged receptors. When the


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Suspended cells

Cell adhesive at 37 °C

Corneal epithelial cell sheet

Donut-shaped PVDF membrane

Non-adhesive at 20 °C

Temperature reduction

Transplanted to ablated stromal bed

After removal of donut-shaped supporter

4.5 The procedure of epithelial cell sheet preparation and transplantation (Hayashida et al., 2006). PVDF, polyvinylidene difluoride.

temperature is below 32 °C (typically 20 °C for 30 minutes), the surface becomes hydrophilic, and spread and adhered cells can be spontaneously detached from the thermo-responsive culture dish surface due to hydration of the surface. This detachment process does not damage the cells because the temperature change is a physically mild treatment, and no enzymes such as trypsin or a chelator such as EDTA are required. The monolayers of cells have even been proved to maintain basal surface ECM proteins after detachment (Kushida et al., 1999). It has been proved that various kinds of cell can be harvested on the PIPAAm grafted culture dish under general incubation conditions of 5% CO2 and 37 °C (Matsuda et al., 2007). Figure 4.5 shows the procedure for epithelial cell sheet preparation and transplantation (Hayashida et al., 2006). By using cell sheet technology, one can reconstruct tissues in three general ways. For tissues such as the corneal epithelium, single cell sheets can be directly transplanted into host tissues (Nishida et al., 2004a). To create 3-D structures, one can perform homotypic layering of cell sheets (Shimizu et al., 2002, 2003). For more complex laminar structures, the method of stratifying different cell sheets should be considered (Harimoto et al., 2002).



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Cell sheet technology has been successfully applied to corneal tissue engineering (Nishida et al., 2004a). In the reconstruction of the corneal epithelium, limbal epithelial stem cells were isolated and cultured on the temperature-responsive culture surfaces at 37 °C. After the cells reached confluency, the culture temperature was decreased to 20 °C for 30 min, when the cell sheets, together with their deposited ECM, could be easily harvested and adhered to the host corneal stroma without the need for sutures (Nishida et al., 2004a). Since the corneal epithelial cell sheets were harvested without using dispase, they were less fragile and contained both cell-to-cell junctions and ECM proteins that can be damaged by dispase. It was demonstrated that a well-formed epithelial sheet can be transplanted without the need for any carrier substrate, such as an amniotic membrane or fibrin gel, and the transplanted cell sheets can attach to the corneal stroma in 5 minutes without the need for sutures. It was found that in patients receiving corneal epithelial cell sheet transplantation, the corneal surface remained clear, with significantly improved visual acuity, more than one year after surgery. In cases of severe disease or bilateral limbal stem-cell deficiency, damage to both eyes prevents the use of autologous corneal epithelial stem cells. In these circumstances, Nishida et al. demonstrated that autologous oral mucosal epithelial cell sheets could be used as an alternative to corneal epithelial cell sheets (Nishida et al., 2004b, Hayashida et al., 2005). They found that oral mucosal epithelial cell sheets fabricated on temperatureresponsive dishes could be harvested and transplanted in the same manner as corneal epithelial sheets. The oral mucosal epithelial sheets fabricated in this way resembled the native corneal epithelium even more closely than the native oral mucosal epithelium. Results from all the human trials they performed demonstrated remarkably improved visual acuity, with all corneas maintaining a clear and smooth surface. They also found that in a rabbit model, after transplantation, oral mucosal epithelial cell sheets underwent changes in their keratin expression profiles towards a corneal phenotype (Hayashida et al., 2005). They believed that the results indicated that the direct interaction between the transplanted cell sheets and the underlying corneal stroma might have had an effect on the phenotypic modulation of the oral mucosal epithelial cells. These findings are encouraging, since using autologous oral mucosal epithelial cell sheets to replace allogeneic corneal epithelial cell sheets can prevent many of the problems induced by allogeneic immunorejection and immunosuppression. They also allow patients with bilateral limbal stem-cell deficiency to be treated easily and provide a sufficient source of epithelial cell sheets for transplantation. However, the interpretation of the findings remains somewhat unclear, and there are questions that need answering, requiring further investigations.


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It was also demonstrated that corneal endothelial cell sheets (Nakamura and Kinoshita, 2003; Nishida et al., 2004b; Sumide et al., 2006) can be created in a similar way. The results of transplanting human corneal endothelial cell (HCEC) sheets into a rabbit model showed that the implanted HCEC sheets were able to maintain a viable monolayer structure beneath the corneal stroma that closely resembled the native endothelium. In summary, cell sheet technology has the following advantages. As the cultured cells are harvested only by a mild, low-temperature treatment without the use of proteolytic enzymes, cell sheet technology allows the non-invasive harvest of cultured cells as an intact monolayer cell sheet including deposited ECM (Kushida et al., 1999). By means of their deposited ECM, cell sheets can be directly attached to host tissues and even wound sites without scaffolds, fixation, or sutures, with minimal cell loss and fewer inflammatory responses.

4.5

Future trends

Over a development period of almost 20 years, although corneal tissue engineering has made significant progress and shown great promise for supplying engineered bioactive organ and tissue substitutes, it has not advanced as fast as people expected. As yet, it is still a great challenge to fabricate a complete substitute for the whole corneal tissue with all the three primary layers, and mimic its optical and mechanical properties. The problems and obstacles are numerous. For instance, although the cornea is a less complex tissue than most of the tissues in the human body, at the same time it has special properties, since it requires high transparency and a specific refractive index, besides possessing specific biomechanical properties. Even though it consists of only five layers, it is composed of many substances. In the extracellular matrix of its stoma alone, there are eight types of collagens and some other substances such as GAGs. Each of them plays a certain role in the biophysical properties of the cornea. They are organized in regularity to ensure transparency, tensile strength, and hydration, as mentioned in Section 4.3. However, in constructing a corneal substitute using the present techniques of tissue engineering, it is very difficult to fabricate one with all the substances that appear in the human corneal stroma in similar proportions, although the techniques for building 3-D corneal constructs composed of all the three cell types of human cells already exist. More difficult is the establishment of their particular organization, which accounts for the transparency, tensile strength and hydration. In addition, unlike the in vivo ‘bioreactor’, the in vitro bioreactor at present cannot provide all the environmental conditions and substances needed for corneal tissue regeneration to culture a corneal substitute with the structure and properties that mimic the human cornea. Furthermore, after a graft is



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implanted, the repopulation of host cells and the remodeling of collagens always occur in order to maintain or re-establish the transparency of the implanted graft. Therefore, an engineered cell-seeded graft, especially the allogeneic cell-seeded ones, will face more problems of immunological reaction and inflammatory responses, because corneal graft rejection is primarily a cell-mediated response. The greater problem is whether its integration could be supported by the host tissue. Therefore, the strategy and approaches for further development of corneal tissue engineering should be reconsidered. Do we need to construct a complete substitute for the whole corneal tissue or just replace its damaged portions? In most cases, the cornea is only partly damaged. So on the whole, only substitutes for the damage portions are required; a complete corneal substitute with all the three primary layers is seldom needed. For this particular circumstance, at present, cell sheet technology is probably the right way to supply engineered epithelium and endothelium. The technique of inducing tissue regeneration in vivo by an active cell-free artificial cornea, on the other hand, is the most suitable means of achieving regeneration of the stroma. A combination of the two techniques, therefore, is expected to be able to regenerate the whole cornea. Certainly stem-cell technology, especially embryonic stem-cell technology, may provide a more efficient means of achieving regeneration of the whole cornea in the future. As embryonic stem-cell culture and differentiation technologies improve, it is hoped that committed stem cells will be obtained to reconstruct the whole cornea in vivo/vitro, or stem cells will be transported to diverse tissue compartments of the cornea via the bloodstream or other means, and that on arrival, the cells can generate the appropriate cell types and tissues in response to local signals. Besides these issues, there are two more puzzles: should we use just autologous cells (including autologous oral mucosal epithelial cells) to culture engineered corneal tissue or also use allogeneic cells? Should the tissue be cultured in vitro or in the body? We hope that this chapter will help readers to discover the right answers.

4.6

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5 Tissue engineering for small-diameter vascular grafts J. I. R O T M A N S, Leiden University Medical Centre, The Netherlands; and J. H. C A M P B E L L, University of Queensland, Australia

Abstract: This chapter discusses the current status of research in the field of vascular tissue engineering. The chapter first reviews the clinical need for tissue engineered blood vessels and the required characteristics of engineered vascular constructs. Subsequently, a variety of approaches for vascular tissue engineering are discussed that have been utilized in an effort to design the optimal arterial substitute. Finally, the major hurdles on the way to widespread clinical use are reviewed. Key words: vascular tissue engineering, prosthetic vascular graft, scaffold, stem cells.

5.1

Introduction

5.1.1 The global burden of vascular disease Atherosclerotic vascular diseases, in the form of coronary artery and peripheral vascular disease, are the leading cause of mortality in the western world (Lopez et al., 2006). Although the incidence of cardiovascular events seems to have decreased slightly in the last ten years, a renewed increase is expected in the coming decades due to the global epidemic of obesity and diabetes mellitus (King et al., 1998). Many patients require surgical procedures for replacement of diseased blood vessels in case of critical ischemia. In the United States, more than 1.4 million arterial bypass operations are performed each year (DeFrances and Podgornik, 2006). In addition, 1.6 million patients on chronic hemodialysis worldwide require frequent surgical interventions to create and maintain adequate vascular access (Schwab et al., 1999). Indeed, surgical interventions relating to hemodialysis access are becoming one of the most common procedures performed by a typical vascular surgeon. Unfortunately, adequate tissue for vascular grafting is lacking in many patients due to previous surgical interventions and preexisting vascular disease. Between 10 and 40% of patients do not have veins suitable for grafting owing to pre-existing vascular disease, vein stripping, or vein harvesting for prior vascular procedures (Piccone, 1987). 116 © Woodhead Publishing Limited, 2010

Tissue engineering for small-diameter vascular grafts

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5.1.2 Limitations of prosthetic grafts In the absence of suitable autologous tissues, non-absorbable prosthetic grafts are frequently used. Today, the two most common prosthetic graft materials are polyethylene terephthalate, or Dacron and expanded polytetrafluorethylene (ePTFE). These prosthetic grafts do not match the efficacy of native vessels, particularly in small-diameter applications. Dacron and ePTFE grafts are much stiffer than the elastic arteries to which they are attached, and lack an anti-thrombogenic endothelial layer at the luminal side (Sarkar et al., 2006). When used as bypass arteries that are less than 6 mm in diameter, thrombosis rates are greater than 40% after six months (Sayers et al., 1998). Similar failure rates are observed when prosthetic grafts are used as vascular access for hemodialysis (Schwab et al., 1999). The limited durability of prosthetic grafts relates to (i) acute thrombosis caused by the absence of functional endothelium, (ii) formation of stenotic lesions due to proliferation of vascular smooth muscle cells (i.e. intimal hyperplasia) and (iii) graft infection (Rotmans et al., 2005b).

5.1.3 Clinical perspective of vascular tissue engineering In view of the above limitations, there is a considerable clinical need for alternatives to autologous veins and prosthetic grafts used in surgical procedures such as lower limb arterial bypass grafting, arteriovenous grafts for hemodialysis and coronary artery bypass grafting. Tissue engineering, defined as the combination of cells, scaffolding, and signaling to form biologically active tissues (Riha et al., 2005), offers potential alternatives to fulfil this urgent clinical need. In fact, tissue engineered blood vessels (TEBV) could ultimately be superior to venous transplants since they would provide diameter-matched conduits without existing disease, valves and bifurcations. Furthermore, TEBVs may have the potential to grow and adapt to changing hemodynamic circumstances, when implanted in pediatric patients. As a consequence of the enormous burden of cardiovascular disease, the development of highly sophisticated biomedical technologies, as well as the increasing knowledge about (stem) cell biology, vascular tissue engineering is a rapidly emerging field in medical research. Since the first attempt by Weinberg and Bell (1986) to engineer new blood vessels, techniques to construct adequate TEBVs have witnessed substantial improvements. However, none of the currently available TEBVs fulfil the high-demanding requirements for arterial substitutes. In the current chapter, we describe several methods that have been utilized in an effort to produce artificial arteries. Furthermore, we will address the main limitations of the current TEBVs and attempt to sketch the main hurdles that need to be taken before successful clinical application becomes reality.


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5.2

Required characteristics of tissue engineered blood vessels

5.2.1 Anatomy and function of native arteries In the field of tissue engineering, it is commonly thought that the best way to successfully mimic the mechanical behavior of native arteries is to engineer tissue with similar composition and structure. Therefore, a successful approach in vascular tissue engineering starts with sufficient knowledge about anatomy and functions of these tissues. Native blood vessels have a concentric layered structure, with each layer being distinct in its cell and protein composition (Fig. 5.1). The tunica intima is the inner layer of the vessel wall that consists of a monolayer of specialized endothelial cells (ECs) together with a subendothelial layer which is composed of connective tissue. This layer possesses a variety of functions, including prevention of clot formation and inflammation of the underlying tissue, as well as signaling functions to the muscular component of the vessel wall (Cines et al., 1998). The internal elastic lamina (IEL) consists of crosslinked elastic fibers and forms the partition between the tunica intima and the tunica media. Elastin is the dominant extracellular matrix protein deposited in the arterial wall and can contribute up to 50% of its dry

V

A

M I

E

EC

Layer I Tunica intima Endothelial cells Glycocalyx Subendothelial extracellular matrix M Tunica media Smooth muscle cells Collagen fibers Elastic fibers Glycoproteins

Function Prevents coagulation Prevents inflammation Signaling to muscle layer

Vasomotion Tensile stiffness Tensile elasticity

A Tunica adventitia Fibroblasts Collagen fibers Vaso vasorum

Nutritional supply Structural support Role in vascular repair

E Elastic laminae Elastic fibers

Tensile elasticity

5.1 Schematic representation of the arterial wall and its function. The left panel is a combined figure of an elastin van Gieson-stained section (left half) and a schematic cartoon of the arterial wall (right half). V, vaso vasorum; EC, endothelial cell.



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weight (Karnik et al., 2003). The tunica media is the muscular layer of the artery which is formed by contractile vascular smooth muscle cells (VSMCs), elastin, and collagens Type I and III, as well as proteoglycans. The mechanical properties of native arteries rely on all these components of the tunica media (Bank et al., 1996). The collagen fibers are aligned circumferentially along the axis of the vessel and provide tensile stiffness (Gelse et al., 2003), required for resistance against rupture. Proteoglycans such as heparan sulfate, chondroitin sulfate, and dermatan sulfate contribute to the compressibility of the arterial wall (Gandley et al., 1997). Elastin is arranged as extracellular concentric cylinders in the native artery, confers elasticity of the vascular wall, and acts as a potent autocrine regulator of VSMC proliferation (Patel et al., 2006). In response to stimuli from the ECs or cytokines from the blood, VSMCs contract or dilate in a coordinated fashion which, in turn, leads to constriction or dilation of the artery. In addition, innervation of the medial layer of larger arteries procures direct communication between the nervous system and the vasculature. Peripheral to the tunica media lies another layer of elastic fibers called the external elastic lamina (EEL) which separates the medial layer from the tunica adventitia. The latter layer consists of a loose collagen matrix with embedded fibroblasts. Small blood vessels called vaso vasorum are also present in this layer, which provides nutritional supply for the vessel wall. Besides provision of structural support, the adventitia is thought to play a role in vascular repair after injury (Li et al., 2000).

5.2.2 Required qualities of tissue engineered blood vessels Vascular tissue engineering aims to obviate the twin obstacles of thrombogenicity and immunogenicity by using autologous cells and a suitable matrix with ideal mechanical properties. Proposed criteria for adequate mechanical properties of TEBV include a burst pressure of >1700 mmHg (L’Heureux et al., 2007a) and suture holding strength of >50 grams-force (Konig et al., 2008). Furthermore, TEBV should retain axial and radial compliance, thereby preserving resistance against aneurysm formation, even up to several years after implantation. In addition, TEBVs should be nontoxic, noncarcinogenic, leak-proof and, in the case of hemodialysis access grafts, resistant to repetitive punctioning. Finally, they should not be susceptible to infection and allow manufacturing in a relatively short space of time with differing specifications in terms of diameter and length. Unarguably, these extensive requirements need to be assessed in long-term follow-up studies in relevant animal models before clinical application in humans can be considered. In order to reach this ultimate goal, close collaboration of vascular biologists with specific expertise in cell-matrix interactions, (chemical) engineers and clinicians is of utmost importance.


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5.3

Approaches to vascular tissue engineering

Tissue engineering has been described as ‘an interdisciplinary field that applies the principles and methods of engineering and the life sciences toward the development of biological substitutes that restore, maintain and improve tissue function’ (Fuchs et al., 2001). In the field of vascular tissue engineering, a variety of methods have been utilized to synthesize new vascular conduits including (i) cell seeding of conventional non-absorbable prosthetic grafts, (ii) in vitro engineering of TEBVs using scaffolds based on synthetic polymers or natural macromolecules, (iii) use of decellularized allografts or xenografts as scaffolds and (iv) the use of self-assembled scaffolds (Table 5.1). In scaffold-based approaches, the scaffold functions to bring the various cells of the vasculature in close proximity to one another to enable intercellular signaling, which in turn facilitates cell adhesion, migration and differentiation.

5.3.1 Seeding of non-absorbable prosthetic grafts Prosthetic grafts have been used in vascular surgery since 1952. Shortly after their introduction, researchers concluded that the patency of smalldiameter prosthetic grafts was limited due to thrombosis, intimal hyperplasia and infection (Kapadia et al., 2008). Since then, several attempts have been made to modify graft surface in an effort to reduce thrombogenicity, to decrease the development of intimal hyperplasia and to improve host incorporation and healing.

Cell seeding of prosthetic grafts The absence of endothelial layer and contact of blood components to the prosthetic material are thought to contribute to graft failure. In humans, prosthetic graft endothelialization is very slow and usually does not extend beyond 1–2 cm of the graft edges (Rahlf et al., 1986). Consequently, ECseeding at the luminal surface of prosthetic vascular grafts is a valuable strategy to improve graft patency. In an effort to restore a physiological endothelial barrier, Herring et al. (1978) were the first to show the benefits of seeding ECs on polyethylene prosthesis in a canine model. Subsequent clinical trials in humans using ePTFE grafts seeded with venous ECs for infra-inguinal bypasses, reported increased patency rates for up to nine years of follow-up (65% for the endothelialized group versus 16% patency of non-seeded grafts, p = 0.02) (Deutsch et al., 1999). Microscopic analysis



Scaffold

ePTFE

ePTFE

ePTFE

ePTFE Polyurethane

ePTFE

ePTFE

ePTFE Collagen

Fibrin

Elastin

First author (year)

Deutsch (1999)

Bhattacharya (2000)

Griese (2003)

Rotmans (2005) Batchelor (2003)

Dorrucci (2008)

Cagiannos (2005)

Lee (2007) Weiberg (1986)

Swartz (2005)

Leach (2005)

VSMCs

Paclitaxel ECs/VSMCs/ fibroblasts ECs/VSMCs

Sirolimus

Heparin

Circulating EPCs NO-donor

Circulating EPCs

BM-EPCs

ECs

Cells/compound

N/A Burst strength, 323 mmHg 7% of arterial break tension Elastic modulus, 900 mmHg

N/A

N/A

N/A N/A

N/A

N/A

N/A

Mechanical properties

–

Sheep

Pig –

Pig

Human

Pig Sheep

Rabbit

Dog

Human

Animal/human

All grafts (n = 2) patent at 15 weeks Not tested in vivo

65% patency at 9 years of follow up Accelerated endothelialization Accelerated endothelialization 3-fold increase in IH Increased thrombus-free surface 2-year primary patency rate of 85% Decreased stenosis in the outflow graft 100% patency at 12 weeks Not tested in vivo

Results from in vivo studies

Table 5.1 Summary of evaluated strategies for vascular tissue engineering. Abbreviations: ePTFE, expanded polytetrafluorethylene; ECs, endothelial cells; BM-EPCs, bone marrow endothelial progenitor cells; VSMCs, vascular smooth muscle cells; N/A, not available; IH, intimal hyperplasia; PLGA, poly[lactic-co-(glycolic acid)]; PLA, polylactic acid; PCL, polyεcaprolactone; Decell., decellularized


Hyaluronan

Polyglycolic acid

PLGA

PCL–PLA

PCL–PLA

Decel. porcine aorta Decel. porcine vein Decel. bovine ureters Decel. intestinal mucosa Self-assembled in vitro Self-assembled in vivo Self-assembled in vivo

Arrigoni (2006)

Nklason (1999)

Iwai (2004)

Shin’oka (2001)

Matsumura (2003)

Amiel (2006)

Cheu (2004)

Sparks (1973)

L’Heureux (2006)

Shell (2005)

Chemla (2008)

Katzman (2005)

Scaffold

First author (year)

Table 5.1 Continued

Tissue capsule

Tissue capsule

ECs/fibroblasts

None

None

None

ECs/VSMCs/ fibroblasts Bone marrow cells ECs

ECs/VSMCs

ECs/VSMCs

VSMCs

Cells/compound

Burst strength >2500 mmHg

Burst strength >2000 mmHg N/A

N/A

N/A

Burst strength >1000 mmHg N/A

N/A

33% of arterial tensile strength 24% of arterial burst strength 900% of arterial tensile trength N/A


Dog

Human

Human

Pig

Human

Human

–

Human

Human

Dog

Pig

–

Animal/human

73% patency at 6.5 months

80% patency at 5 months (ongoing study) 20% patency at 6 months

Similar patency as ePTFE grafts Similar patency as ePTFE grafts No significant reduction in IH

Patent at 3 years in pulmonary circulation No complications in all 17 patients Not tested in vivo

100% patency at 6 months

100% patency at 24 days

Not tested in vivo

Results from in vivo studies


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of nine failed seeded-grafts at a mean follow up of 39 months confirmed the presence of an endothelium on all specimens (Deutsch et al., 2008). In these clinical studies, a two-stage procedure was utilized in which harvested ECs were expanded in vitro for four weeks before the seeded graft was implanted in the same patient. As a consequence, this time- and costintensive method is not suitable for emergency surgery. To circumvent these limitations, one-stage procedures have been developed in which cell harvest, seeding and graft implantation can be completed during one surgical procedure. Unfortunately, none of the clinical studies using a one-stage procedure showed favorable results over non-seeded prosthetic grafts (Bordenave et al., 2005). As an alternative to mature ECs, recent studies have focused on the use of endothelial progenitor cells (EPCs) for cell seeding procedures. EPCs are a subset of CD34+ cells with the potential to proliferate and differentiate into mature ECs (Asahara et al., 1997). Several studies have emphasized that circulating EPCs have the capacity to home to sites of vascular injury, thus promoting the process of re-endothelialization (Walter et al., 2002; Werner et al., 2003). Initial seeding studies were performed by Bhattacharya et al. (2000), who harvested CD34+ EPC from the bone-marrow of dogs. Subsequently, these cells were seeded on polyester vascular grafts using a two-stage procedure. At four weeks after graft implantation in the thoracic aorta, the authors observed a significant increase in graft endothelialization (92% versus 26.6% in non-seeded grafts, p < 0.01). The feasibility of EPC-seeding in a two-stage procedure was confirmed by Griese et al. (2003), who harvested CD34+ EPC from peripheral blood of rabbits. Again, accelerated endothelialization was observed at four weeks after grafting. Unfortunately, the authors did not report on the effect of EPC-seeding on the degree of neointimal formation in the anastomotic area. In an effort to ‘auto-endothelialize’ prosthetic grafts, we evaluated the efficacy of anti-CD34 coated ePTFE grafts (Rotmans et al., 2005a). Their ability to capture EPCs in vivo was assessed in a porcine model of arteriovenous graft failure in which prosthetic grafts were implanted between the carotid artery and the jugular vein (Rotmans et al., 2003). In spite of cellular coverage of the luminal surface, we observed a three-fold increase in neo-intima formation in anti-CD34 coated grafts after 28 days of follow-up. Obviously, the attracted cells enhanced proliferation of VSMCs in the anastomotic area or transdifferentiated into VSMCs themselves. Additional studies will reveal whether and to what extent these captured cells can be stimulated to regain their endothelial and vasculoprotective function. In recent years, several other cell populations, including CD14+ monocytes (Harraz et al., 2001), mesenchymal stem cells (Oswald et al., 2004) and VSMCs (Wang et al., 2006) have been shown to possess the ability to differentiate into functional


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ECs. Future studies will show if these cells are a suitable source for cell seeding procedures. Graft seeding with small molecules In view of the important role of nitric oxide (NO) in maintaining vascular homeostasis, NO-releasing systems have been developed for stent and graft coatings. NO is a potent vasodilator, with protective effects on ECs. Furthermore, NO inhibits thrombus formation and proliferation of VSMCs (Vallance and Collier, 1994). Since NO has a half-life of only a few seconds, several researchers have focused on the immobilization of NO-donor compounds within biomaterials. Although different studies using NO-releasing grafts showed favorable effects on thrombus formation and intimal hyperplasia (Smith et al., 1996; Batchelor et al., 2003), concerns have been raised regarding the formation of toxic metabolites and the finite reservoir of NO that exists within prosthetic grafts (Reynolds et al., 2004; Kapadia et al., 2008). Currently, clinical data on the use of NO-donor coated grafts are lacking. Protein coating or binding proteins to grafts is another approach for improving graft patency. Recently, several commercially available and FDA-approved grafts have used heparin-coating technologies to improve graft patency (Kapadia et al., 2008). Although no prospective, randomizedcontrolled clinical trials have been conducted, the two-year primary patency rate of 85% of heparin-bonded ePTFE grafts used for infragenicular arterial bypass surgery suggests superior patency rates, when compared to historical controls (Dorrucci et al., 2008). Although these studies are promising, the risk of heparin-induced thrombocytopenia, which can be lethal, remains a drawback (Kapadia et al., 2008). Alternative compounds for graft seeding include sirolimus and paclitaxel. Indeed, sirolimus- and paclitaxel-eluting stents have been used successfully in occlusive coronary artery disease, and have been shown to reduce in-stent restenosis (Eisenberg and Konnyu, 2006). Both these agents possess strong anti-proliferative effects on VSMCs. Lee et al. (2007) assessed the potential of paclitaxel-eluting prosthetic grafts, which were evaluated in a porcine arteriovenous graft model. When compared to conventional prosthetic grafts, paclitaxel-eluting grafts showed better survival than uncoated grafts at 12 weeks after implantation (100% versus 25%, respectively, p = 0.01). Cagiannos et al. (2005) used sirolimus-eluting grafts as iliac artery bypass grafts in pigs and showed a significant reduction in per cent cross-sectional narrowing as well. Future clinical trials will reveal if these encouraging results translate into increased patency rates of prosthetic grafts in patients who require an arterial bypass or hemodialysis access.



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5.3.2 In vitro tissue engineering using scaffolds of natural macromolecules The archetypal tissue engineering approach is to use a scaffold that is seeded with vascular cells in vitro. The properties of an ideal scaffold have been identified as (i) sufficient porosity to allow cell/tissue growth as well as transport of nutrients, (ii) suitable surface chemistry for cell attachment and proliferation and (iii) pseudo-physiological mechanical properties (Hutmacher, 2001). Natural macromolecules such as collagen as scaffold components have the advantage of facilitating cell binding. Collagen scaffolds Collagen Type I is often used in this context because it is abundant in many tissues and can be isolated, solubilized and subsequently poured on a mold of the desired (tubular) shape. Weinberg and Bell (1986) published the first study in which a cell-seeded collagen scaffold was used to engineer a TEBV in vitro. In their study, a multilayered tube was constructed that included bovine ECs, VSMCs, as well as fibroblasts, mimicking the structure of native arteries. Their initial constructs were so highly distensible that they ruptured at very low pressures (2000 mmHg. In subsequent studies, solely fibroblasts and ECs were used (L’Heureux et al., 2006). When implanted as interposition grafts in the abdominal aorta of nude rats, 86% of these TEBVs remained patent up to 32 weeks of follow-up. Furthermore, no signs of luminal narrowing or aneurysm formation were observed. Subsequently, autologous TEBVs were implanted as arteriovenous access grafts in six patients on chronic hemodialysis (L’Heureux et al., 2007b). Total production time for the graft ranged between 6 and 9 months and graft length varied from 14 to 30 cm. In the first patient, the TEBV was used for more than 13 months, until the patient underwent successful kidney transplantation. One patient died of unrelated causes while thrombotic failure was observed in one patient at 12 weeks, which was attributed to low flow rate. In the remaining three patients, the vessel was still patent at 5 months after implantation. Furthermore, time to hemostasis after dialysis access was lower for the TEBVs than for ePTFE grafts (L’Heureux et al., 2007a). Recently, the authors reported 60% primary patency rate at 6 months after implantation (McAllister et al., 2009). These preliminary results are encouraging, especially since arteriovenous hemodialysis accesses represent the most challenging mechanical environment in the field of vascular surgery. Besides the repetitive punctioning of the conduit for dialysis, the latter relates to a turbulent flow pattern at the anastomotic areas and the high flow rate which usually exceed 800 mL/min. The major limitations of this in vitro approach are the laborious and time-consuming procedures (28 weeks) (L’Heureux et al., 2006) that are required to grow vital and sterile blood vessels.

In vivo engineering of self-assembled vascular grafts The ultimate goal of these in vivo approaches is that patients grow their own arteries within their own body. The rationale for these methods stems from the observation that implantation of prosthetic materials in the human body initiates an inflammatory response that culminates in the formation of an autologous fibrocellular capsule. Proponents of these methods,



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including ourselves, hypothesized that under the guidance of mechanical stimuli and the new microenvironment, this tissue capsule will transform into a fully functional artery once implanted into the vasculature (Efendy et al., 2000). The foreign body material (called mandril) serves as a mold, which is removed before vascular grafting. Such a method takes advantage of the plasticity of tissues since it appears that it is the cell and matrix environment rather than their source that determines eventual function (Edelman, 1999). In addition, major limitations of in vitro tissue engineering approaches are circumvented. Indeed, no time-consuming cell culture steps are required. Furthermore, a major difficulty with any in vitro vascular engineering approach is that cells in culture alter their phenotype as well as their immunogenic properties (Thomas and Campbell, 2001). For example, VSMCs in culture modulate from a contractile into a synthetic phenotype (Chamley-Campbell et al., 1979) whereas healthy arteries are composed of contractile VSMCs. Sparks’ mandril Many years ago, Sparks (1973) published a revolutionary in vivo method for in vivo tissue engineering of vascular grafts. His method consisted of preparing a silicone rod of desired diameter and length with coverings of Dacron mesh tubes. This mandril was implanted at the location of the contemplated arterial grafting procedure. After a waiting period of 5 to 8 weeks, during which an autologous tube grew around the silicone mandril and encapsulated the knitted Dacron tubes within its wall, the mandril was withdrawn and the new Dacron reinforced autologous tube was implanted as arterial bypass graft. In vivo studies in dogs showed promising results but clinical studies using these TEBVs as femoropopliteal bypass grafts failed, mainly due to unacceptable high rates of thrombosis and aneurysm formation (Roberts and Hopkinson, 1977). Most likely, the latter failure relates to the absence of an endothelial, anti-thrombogenic layer at the luminal site of the graft and the lack of elasticity (Opitz et al., 2004). Peritoneal cavity as an in vivo bioreactor Ten years ago, we proposed the concept of using the peritoneal cavity as a bioreactor in which to grow an ‘artery’ (Campbell et al., 1999). This site not only serves as a humidified bioreactor in vivo, but also supplies the source of cells, nutrients, growth factors, and other necessary components to create new arteries. Initial studies in rats and rabbits showed that the implanted foreign bodies in the peritoneal cavity became encapsulated with cells of hemopoietic origin (Campbell et al., 2000). Subsequent studies in dogs revealed optimal tissue capsule formation when polyethylene


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Biointegration of medical implant materials (b)

L

(c)

L

5.2 Morphology of TEBVs obtained from the peritoneal cavity. (a) Central region of a TEBV, 6.5 months after it had been grafted into the femoral artery in a dog. Photomicrograph of immunostaining for (b) smooth muscle myosin and (c) Ulex lectin of a TEBV, 4 months after grafting into the femoral artery of a dog. In (b) and (c), the reactive cells (which express the antigen) are stained white. Smooth muscle myosin and Ulex lectin are markers for VSMCs and ECs, respectively. In (b) and (c), the luminal side of the tissue is depicted with the letter L.

tubing encapsulated by Dexon mesh were inserted (Chue et al., 2004). Within 2–3 weeks after insertion of the tubing, the encapsulating cells gradually trans-differentiated into myofibroblasts that subsequently synthesized a collagen matrix, forming a capsule of living tissue covered by a single layer of mesothelium on the outer surface. Using a transgenic mouse model, a PhD student in our laboratory, Jane Mooney, has recently shown the origin of these cells to be monocyte/macrophages (Mooney et al., 2010). After removal of the polyethylene tubing, the tissue capsules were transplanted as interposition grafts in the femoral artery of the same animal in which they were grown. By doing so, the capsules transformed to arteries that consist of ECs, VSMCs, fibroblasts and extracellular matrix proteins, including collagen and elastin (Fig. 5.2). At time of harvest, between 3 and 6.5 months, 73% of TEBVs remained patent. The presence of elastin in this engineered vessel after grafting is particularly encouraging, because the lack of elastin in engineered grafts is believed to cause late dilatation of engineered conduits in high-pressure arterial circuits (Opitz et al., 2004) and thus may contribute to aneurysm formation. Similar results were obtained when the TEBVs were grafted in the aorta of rats (Campbell et al., 1999) or in the carotid artery of rabbits (Campbell et al., 2000). Functionally, these grafts revealed remarkable biocompatibility, together with burst strengths of >2500 mmHg, elasticity and vascular reactivity (Chue et al., 2004). By 3–4 months, the artificial arteries in the high-pressure sites have doubled in thickness, and an ‘adventitia’ containing vasa vasorum has developed on their outer surface (Thomas et al., 2003). Currently, we are evaluating the efficacy of these artificial arteries in improving patency of arteriovenous grafts in pigs. These studies will serve



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as the litmus test before human trials can be initiated. Subsequently, it remains to be tested whether similar tubes of living tissue can be grown in the peritoneal cavity of humans, particularly those with conditions such as diabetes mellitus or chronic renal failure. Indeed, recent studies showed impaired progenitor cell function in patients with these diseases (Loomans et al., 2004; Westerweel et al., 2007).

5.3.6 Use of stem cells for vascular tissue engineering The rapid expansion of knowledge about stem cell biology provided an enormous boost for new studies in the field of (vascular) tissue engineering. Stem or progenitor cells are characterized by their capacity to differentiate into specialized cell types and they can be found in the embryo or adult. Human embryonic stem cells have ethical constraints and the desire for autologous grafts make adult progenitor cells more suitable for tissue engineering. Adult stem cells have been isolated from a variety of sources, such as blood, bone marrow and fat (Sales et al., 2005). In view of their potential to differentiate into vascular cells, endothelial progenitor cells (EPCs) (Asahara et al., 1997) and smooth muscle progenitor cells (SPCs) (Saiura et al., 2001) have been utilized for seeding of prosthetic grafts (see Section 5.3 ‘cell seeding of prosthetic grafts’) and xenogeneic decellularized scaffolds. Kaushal et al. (2001) showed encouraging results of decellularized porcine vascular scaffolds that were seeded with circulating EPCs, and subsequently implanted as a xenogeneic interposition graft in sheep. While all (n = 4) non-seeded conduits thrombosed, all (n = 7) EPC-seeded were patent after 130 days of follow-up. Despite their successful studies in sheep, no reports have been published on EPC-seeded xenogeneic scaffolds in humans thus far. Cho et al. (2005) harvested canine bone marrow-derived progenitor cells which were differentiated into VSMCs and ECs in vitro. Subsequently, these cells were seeded onto decellularized canine carotid arteries. Eight weeks after implanting these TEBVs as interposition grafts in carotid arteries in dogs, all grafts remained patent and showed remarkable regeneration of the three-layered structure of native arteries. Although these animal studies are very encouraging, the authors have not yet attempted to replicate this success in human studies. In addition, none of these studies has shown that stem cells may enhance elastogenesis, which is a prerequisite for successful application of TEBV in the systemic arterial circulation. In contrast, stem-cell seeded scaffolds have already been implanted into the pulmonary circulation in humans. As already mentioned, the low blood pressure in the pulmonary circulation may favor successful application of TEBVs since the biomechanical properties of TEBV are not as high as are required for the systemic arterial circulation. In 2001, researchers at The


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Heart Institute of Japan started to use bone-marrow-cell-seeded PCLbased scaffolds for pulmonary artery reconstruction in pediatric patients (Matsumura et al., 2003). The authors reported on successful implantation of TEBVs in 17 patients. No complications such as thrombosis or stenosis were observed. Long-term follow-up studies should reveal the ultimate clinical value of this technique. For future studies, mesenchymal stem cells (MSCs) could serve as an alternative source for vascular tissue engineering since differentiation of these cells into VSMCs (Narita et al., 2008) and ECs (Yue et al., 2008) has recently been demonstrated.

5.4

Future trends

5.4.1 Validation in relevant large animal models In medical research, there is a tendency to suggest a potential breakthrough once short-term results of new therapies are obtained. The same is true for tissue engineering approaches. However, long-term follow-up studies in relevant large animal models are a prerequisite before clinical trials are initiated. For cardiovascular studies, pigs are frequently used as models because of their analogous vascular anatomy, size, and physiology (Ferrell et al., 1992). In our opinion, the arteriovenous graft model is the optimal application for preclinical and clinical evaluation because (i) there is an urgent need for better vascular accesses, (ii) graft failure is unlikely to be life threatening and (iii) it is the most challenging model with respect to the mechanical environment. The latter relates to the turbulent flow pattern and blood flow rates usually exceeding 800 mL/min, the gradual decline in blood pressure along the course of the graft and the repetitive punctioning of the graft during hemodialysis sessions (Rotmans et al., 2005b). With regard to these animal studies, it is important to notice that elasticity of arteries does vary between species. Porcine coronary and internal mammary artery walls were found to be three times more elastic than human arteries (van Andel et al., 2003). The latter may result in underestimation of wall stress and the risk of rupture when TEBVs are evaluated in pigs.

5.4.2 Hurdles on the way to widespread clinical use Despite the significant improvements that have been achieved in the last two decades, the ideal substitute for small-diameter arteries still remains to be constructed. For future studies, elastic compliance is one of the most crucial characteristics of TEBVs that has to be ameliorated. Not surprisingly, most TEBVs that were generated under static conditions lack tensile strength and elasticity. Indeed, several studies showed that shear stress is an important stimulator of collagen (Jin et al., 2001) and elastin (Isenberg



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and Tranquillo, 2003) synthesis and induces the circumferential orientation of VSMCs (Seliktar et al., 2000). In subsequent efforts to design an ideal tissue engineered blood vessel, many investigators examined different types of mechanical stimulation techniques using bioreactors in order to mimic the exposure to shear stress in vitro. Although most of them succeeded in improving burst strength as a result of collagen synthesis, most failed to induce elastin synthesis (Niklason et al., 1999; Hoerstrup et al., 2002). Biosynthesis and subsequent crosslinking of elastin appears to be one of the most complex and tightly regulated processes during the maturation of blood vessels (Bunda et al., 2005). Incorporation of this knowledge about elastin metabolism in new strategies for vascular tissue engineering is a prerequisite to improve tissue elasticity. Another important issue that has to be addressed in future studies is the formation of vaso vasorum in TEBVs to ensure sufficient nutritional supply of the thickening vascular construct. Indeed, if oxygen concentrations are inadequate, cell proliferation ceases and cell viability begins to break down (Kellner et al., 2002). In the past, the adventitial segment of the vessel wall has received limited attention compared with the endothelium. However, the adventitia (in which the vaso vasorum is located) has emerged as an active participant in vascular homeostasis and remodeling (Sartore et al., 2001). Therefore, adventitial delivery of progenitor cells, vascular growth factors or gene therapeutic constructs may enhance the functional capacity of this important layer of TEBVs. Besides the above-mentioned anatomical and functional requirements, a limited time frame required for the production of TEBVs is essential for future large-scale clinical application. Therefore, time consuming in vitro cell culture procedures need to be reduced as much as possible. In this respect, it remains to be determined if EC seeding of engineered constructs prior to implantation is a prerequisite for successful application. Although prior cell seeding has obvious advantages for immediate graft function, cellular retention after implantation is usually low (Bhat et al., 1998). Omitting this step could therefore save a substantial amount of valuable time. Evidence for potential success of such an approach comes from human (Gravanis and Roubin, 1989) and animal studies (Reidy and Schwartz, 1981) that revealed spontaneous reendothelialization of denudated arterial segments, for instance after balloon angioplasty. Furthermore, some decellularized xenogeneic scaffolds (Bergmeister et al., 2008) have shown to endothelialize spontaneously after implantation into the arterial circulation. Additional modifications that increase mobilization and homing of EPCs may further enhance the formation of an endothelial layer on the luminal surface of TEBVs in vivo. For instance, incorporation of growth factors such as transforming growth factor-β and vascular endothelial growth factor into engineered constructs may improve endothelialization,


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since these growth factors have emerged as pivotal regulators of EPC function (Hristov et al., 2003). Such efforts to facilitate in vivo endothelialization may kill two birds with one stone since laborious in vitro seeding procedures are circumvented while the full regenerative potential of vascular progenitor cells can be exploited. As a result of recent advances of the field of regenerative medicine, it has been hypothesized that the ultimate configuration of an autologous cell-based vascular graft need not be determined at outset by the cells that comprise the device, but rather by the dynamic environment, wherein the body modifies the tissue-engineered construct to meet local flow, metabolic and inflammatory requirements (Edelman, 1999).

5.5

Conclusion

Vascular tissue engineering is a rapidly expanding field and represents the newest concept in order to alleviate the limitations of small-diameter prosthetic grafts. Although the use of human TEBVs in the pulmonary circulation has shown inspiring clinical success in pediatric patients, the promise of an ‘off-the-shelf’ tissue-engineering graft for adult revascularization remains unrealized. When compared to other organs, such as the kidney, the structure of blood vessels appears relatively simple. However, despite more than twenty years of research, the ‘Holy Grail’ is still to be discovered. One hundred and fifty years after the publication of ‘On the Origin of Species’ by Charles Darwin, that work should not only result in enormous admiration for the evolution of mankind, but should also encourage the scientific community to continue its efforts to create autologous vascular substitutes. In this respect, the broadly supported focus of stem cell biology, regenerative medicine and tissue engineering in medical research is encouraging. In the past, progress in tissue engineering may have been stunted by the existence of separate worlds in which chemical engineers and biomedical scientists have been working. To reach the ultimate goal of manufacturing adequate tissue engineered blood vessels within the foreseeable future, strong collaboration of engineers, biologists and clinicians is of vital importance.

5.6

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6 Stem cells for organ regeneration K. D. D E B, Dayananda Sagar Institutions, India

Abstract: Tissue engineering aims at regenerating and restoring organ function by exploiting the self-healing properties of body tissues. To achieve this, cells from different sources, growth inducing molecules, and a biomaterial-support called a scaffold are used in isolation or in combination to reconstruct organs. However, regeneration of a functional tissue or organ has so far remained a challenge. With the recent advances in adult- and embryonic-stem cells technologies, the dream of reconstructing an entire organ or at least parts of it, seems likely to be realized soon. Moreover stem cells provide tissue engineers the freedom to choose cells at a more defined or intermediate stage of differentiation. Such advances will lead to better integration of the tissue engineered grafts in the host. Organ re-construction is performed either in vitro in laboratories and may be made available on demand; or the regenerating factors or microenvironment can be engineered in vivo for organs to grow inside the body. Integration of stem cells and tissue engineering has resulted in tremendous progress, poised to bring a paradigm shift in regenerative medicine. This chapter presents an overview of the basics and the major advances made in this area so far. Key words: stem cell, organ regeneration, tissue engineering, scaffolds, biomaterials.

6.1

Introduction

Advances in modern medicine have evolved from multidisciplinary research across a wide range of scientific and engineering disciplines. Tissue engineering is defined as an ‘interdisciplinary field that applies the principles of engineering and life sciences toward the development of biological substitutes that restore, maintain or improve tissue function’ (Langer and Vacanti, 1993; Nerem, 1991). It is based on principles to exploit the self-healing potential of an organ to regenerate itself. Increase in average life span of human beings over the last couple of decades has resulted in a vast proportion of patients suffering from agerelated degenerative diseases, organ failures, congenital malformation or other debilitating conditions. Until recently, treatment options for such patients were limited to either prosthetic implants or organ transplantations. In the first approach, reconstructive surgery is performed by inserting 147 © Woodhead Publishing Limited, 2010

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an artificial organ substitute, also called a biomedical device. Examples include artificial pacemakers, heart valves, gold teeth, fracture fixation plates, kidney hemodialyzers and prosthetic joints, to substitute for the damaged organs. However, in most of these cases the artificial organs cannot perform all the functions of the organ, and lead to progressive deterioration of the health of the patient (Tabata, 2000). Moreover, many of these devices lack lifelong durability. In an alternative approach, an organ obtained from a live donor or cadaver is transplanted into the patient. Though mostly successful, such transplantations are limited by the number of matching donors available. Both implanted devices and transplanted organs have to face immune rejection from the patient. The advent of immunosuppressive drugs such as glucocorticoids, cyclosporine, and later azathioprine and monoclonal antibodies, did revolutionize the practice of transplantation medicine to a certain extent (Italia et al., 2006), and enabled organ substitutes to overcome foreign body rejections. These drugs, however, have severe adverse effects for long-term therapy and predispose patients to recurrent infections. Hence, the clinical solutions for organ failure so far remain imperfect. Organogenesis is induced to occur in laboratories and transplanted into the individuals, or the necessary factors are delivered in the injury site in vivo. This is achieved by engineering in tandem one or more of the three elements of tissue engineering – cells, tissue-inducing factors and materials – to support and guide tissue growth (Langer and Vacanti, 1993). It can be implied that a typical tissue engineer has to work at the interface of biological, chemical, physical, engineering, medicine, genetics, pharmaceutical, and materials sciences to achieve the desired goals. Tissue engineering is therefore a relatively modern branch of science, as compared to the allied practices of regenerative medicine and cell-based therapies. Regeneration and stem cells have been widely studied subjects, particularly in insect and other organisms. The ability of lizards to regenerate lost tails is appreciated with amazement by most of us during our childhood. According to the Book of Genesis, 2:21, a rib was harvested from Adam, the first donor, and was used to fashion Eve. In Greek mythology a story goes that Prometheus was to be punished for giving fire to mankind, and an eagle was sent to devour his liver; however, he survived because he could quickly regenerate his liver every day (Rosenthal, 2003). In Ancient India, Ayurvedic texts dating back to 600 bc contain a detailed description of protocols to carry out rhinoplasty, a procedure intended to resurrect an injured nose with skin obtained from other parts of the body. Shusruta, the surgeon who described such a procedure, is regarded as a pioneer of skin grafting and plastic surgery in India (Rana and Arora, 2002). In modern times, routine graftings with cornea were performed in Vienna during the


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early years of the 19th century (Gardner, 2007). In the 1930s, a Swiss physician, Dr Paul Niehans, also known as the father of cell therapy, injected a suspension of steer thyroid cells into a critical patient suffering from thyroid deficiency. The patient went on to live for the next 30 years and this heralded the world of modern cell therapy (www.healthatoz.com/celltherapy). The clinical success of the first ‘artificial kidney dialyzers’ in 1944 (Kolff et al., 1944) ushered in the era of engineering sciences’ entry into healthcare research. The next big breakthrough in regenerative medicine was the first successful heart transplantation by Dr Christiaan Barnard in the mid-1960s (Barnard, 1967). Thereafter, in the 1970s artificial pacemakers were successfully introduced in clinical applications (Niklason and Langer, 2001). This was followed by approval of cyclosporine for clinical use as an immunosuppressant in 1983. These events together brought great strides in the management of patients affected with some forms of organ failure. However, concerns over the long-term use of artificial organs and immunosuppressive drugs surfaced soon, and unavailability of matching donors made the situation no better. In modern times, the earliest applications of tissue engineering in clinical practice started when Drs Bruke, Yannas and Bell prepared collagen-based matrices, seeded with dermal fibroblasts, to treat burn victims, in the early 1980s. The first in vitro skin grafting was successfully tested clinically at around the same time (Yannas et al., 1982). Since then, tissue engineering has made a formal entry in the scientific literature, and became more popular during the late 1980s and early 1990s.

6.2

Basic components of tissue engineering

Tissue engineering for organ regeneration mainly utilizes various isolated cell types and aims at creating a structurally and functionally active organ system that can substitute for the damaged tissue in vivo. As mentioned earlier, tissue engineers try to use cells, tissue regenerating factors and scaffolding materials to create new tissues. These are the three primary elements of tissue engineering. A tissue engineer has to select a right mix and a formula comprising these components and apply them in conjunction to fulfill the objectives. With the advent of stem cell technologies and isolation procedures, tissue engineers are now trying to harness the ability of stem cells to self-renew and differentiate to various lineages.

6.2.1 Cell sources and types Cells available for tissue engineering are classified in various ways. Depending upon the donor of the cells, there are three types of tissue engineering – autologous, allologous and xenologous. In Dr Niehans’ experiment


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described in Section 6.1, he obtained cells from a different species, the steer, and used them to restore a patient’s dysfunctional organ. The procedure of introducing cells obtained from an individual of a different species is called xeno-grafting. Xeno-grafts are susceptible to immune reactions and may produce the challenge of cross infections (Appel et al., 2000; Platt, 2000). However, as some studies have shown, cross infectivity is always a possibility and not a necessity (Levy et al., 2000). Apart from this, the host will identify the xeno-graft as a foreign body and will try to reject these cells, resulting in inflammation and so called graft-versus-host reactions. Thus, xeno-grafting is accompanied by supplements of immunosuppressive drugs such as cyclosporine. If tissue engineering is performed utilizing cells of an individual of the same species as the patient, it is called allologous grafting. It is desired that the donor and the recipient share the same human leukocyte antigen for graft survival and cell acceptance. However, it is often very difficult to obtain matching donors, and immuno-modulators may be required in these procedures. This brings us to the most desirable situation for tissue engineering, at least from an immunological point of view. In this instance, cells harvested from a patient himself are induced to grow and substitute the distorted organ. This can be accomplished either in vitro or in vivo. Such a procedure is known as auto-grafting or autologous tissue engineering. An alternative classification of cells available for tissue engineering is based upon their morphological and physiological characteristics. Thus, tissue engineers have the choice of using: (i) tissue-specific differentiated cells, (ii) adult progenitor or adult stem cells, (iii) embryonic stem cells, and the more recently created (iv) induced pluripotent (iPS) stem cells. The following sections cover each of these cell types. Somatic cells In vitro expansion of cells to significant quantities has been a limitation when developing cell-based regenerative medicine techniques for organ replacement. Though some organs, such as the liver, have a high regenerative capacity, hepatocyte growth and expansion in vitro can be difficult. By studying the privileged sites for committed precursor cells in specific organs, as well as exploring the conditions that promote differentiation and/or selfrenewal of these cells, it may be possible to overcome the obstacles that limit cell expansion in vitro. Organ regeneration and tissue engineering attempts with most adult cell types have met with varying degrees of success. Adult autologous tissue-specific cells are harvested from the patient, expanded ex vivo and reintroduced into the body. A limitation of using mature cells for tissue engineering is, with the exception of some cell types such as the dermal fibroblasts, keratinocytes, and some epithelium cells such



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as corneal epithelia, cartilage and adipocytes, harvesting adult tissuespecific cells is difficult and requires invasive surgeries, if not biopsies. Such surgeries may result in donor site cell morbidity and scarring. The two most developed autologous mature cell therapies that have advanced from the laboratory to the clinic involve the repair of cartilage using autologous chondrocytes, and the treatment of burns with autologous cultured keratinocytes (Fodor, 2003). A few tissue engineered products based on allologous adult mature cells such as Apligraft (composed of neonatal foreskin keratinocytes and dermal fibroblasts), are available in the market and many others are actively being investigated (Chu et al., 1995; Eaglstein and Falanga, 1998; Kondziolka et al., 2000). Xenologous adult cells are likely to pose significant challenges. Clinical trials with transgenically engineered pig livers to detoxify the blood of patients suffering from fulminant hepatic failure (FHF) via extracoporial perfusion has been conducted (Levy et al., 2000). Tissue engineering with allologous and xenologous cells need to apply technologies such as cell encapsulation for immune isolation, modified tissue processing methods, tolerance induction in patients and production of transgenic organisms with humanized antigenic properties to overcome their inherent disadvantages. Galα1,3 gal transferase null transgenic pigs have already been produced and they represent a significant development towards eliminating both hyperacute and acute vascular rejection and extended survival of xenologous cells and organs (Phelps et al., 2003). Despite the advent of stem cells, adult cells to some extent continue to excite tissue engineers in their pursuit to create new tissues (Kucia et al., 2007). Somatic stem cells Adult stem cells have been known for more than two decades now; however, it is only the recent advances in our understanding of stem cells that offer the promise to convert all the expectations of tissue engineering into reality, and revolutionize the practice of regenerative medicine. The reasons why stem cells appear to be such ideal candidates for regenerative medicine and tissue engineering are attributed to their characteristic ability to self-renew and undergo multiphenotypic differentiation. There are three types of stem cells – embryonic, germinal and somatic/adult/foetal stem cells. The potential of stem cells to give rise to a number of different cell types determines their therapeutic utility. Thus stem cells are classified as pluripotent (embryonic stem cells or ES cells), multipotent (adult stem cells), and unipotent (epidermal stem cells) (Serakinci and Keith, 2006). Thus, discounting other factors, ES, which can get differentiated to a wider number of cell phenotypes compared to adult stem cells, should be better candidates for tissue engineering. However, adult stem cells (except neural


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stem cells) are easier to harvest. Almost all organs of the adult human body contain a specific mass of cells possessing the capacity to self-renew and differentiate. These cells may be quiescent in some tissue beds while active in others. The bone marrow is one of the most abundant sources of selfreplicating mesenchymal and hematopoetic stem cells, both of which are multipotent in character. Human MSC can give rise to a variety of tissues such as adipocytes, bone, cartilage and muscles, and can be used for tissue engineering purposes. In addition, neuronal stem cells isolated from the brain or the spinal cord are relatively easy to expand in culture and can be used for various clinical ailments such as Parkinson’s, Alzheimer’s and other brain trauma. However, adult stem cells lack the immortality of ESCs – a desirable feature for tissue engineering. This feature can be introduced by genetic transductions of cells, e.g. ectopic expression of the human telomerase reverse transciptase gene may render them capable of infinite proliferation (Natesan, 2005). Recently it has been shown that, under certain conditions, mature adult cells can be induced to form a more primitive phenotype by a process termed ‘dedifferentiation’. This technology is discussed in more detail in the section for induced pluripotent cells (iPS), which follows later. Adult stem cells isolated from a particular tissue can also form cells of another phenotype, this phenomenon being known as ‘trans-differentiation’. For example, co-culturing neural stem cells with muscle progenitor cells has been shown to form muscle without the need for any other factor (Galli et al., 2000). Both dedifferentiation and trans-differentiation, when established completely, will greatly increase the clinical applicability of adult stem cell therapy (Oliveri, 2007; Thowfeequ et al., 2007). Adult stem cells derived from adult bone marrow and muscles are currently being used by several workers for tissue engineering experiments (Wei et al., 2007). Besides this, it is also believed that adult tissues harvest a small quiescent population of pluripotent stem cells known as the embryonic-like stem cells (ELSCs) (Young and Black, 2004). It is thought that, during development, a small number of cells escape the development continuum, migrate and home-in at various parts/tissues of the adult. We have seen that the umbilical cord Wharton’s jelly is a rich source of such population of stem cells (unpublished data). Also, the placental and amniotic fluid stem cells have been shown to be very rich sources of adult multipotent stem cells. The following section presents a brief description of these cells. Placental and amniotic fluid stem cells Both amniotic fluid and placenta are known to contain multiple partially differentiated cell types derived from the developing fetus. These cells can be obtained from amniocentesis, or from chorionic villous sampling in the



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developing fetus, or from the placenta at the time of birth. Stem cell populations from these sources, called amniotic fluid and placental stem cells (AFPSCs), that express embryonic and adult stem cell markers have been successfully isolated (DeCoppi et al., 2007). The undifferentiated stem cells expand extensively without the need for a feeder cell layer and double every 36 hours. Unlike human embryonic stem cells, the AFPSCs do not form tumors in vivo. Lines maintained for over 250 population doublings retained long telomeres and a normal karyotype. These cells can be differentiated into neuronal lineage secreting the neurotransmitter Lglutamate or expressing G-protein-gated inwardly rectifying potassium (GIRK) channels, hepatic lineage cells producing urea, and osteogenic lineage cells forming tissue engineered bone. Embryonic stem (ES) cells and derivatives In 1981, pluripotent cells were found in the inner cell mass of the human embryo, and the term ‘human embryonic stem cell’ was coined (Martin, 1981). These cells are known to differentiate into all cells of the human body, excluding placental cells (only cells derived from the morula are totipotent; that is, able to develop into all cells of the human body). These cells have great therapeutic potential, but their use is currently limited by several factors, both biological and ethical. The political controversy surrounding ES cells started in 1998 with the creation of human ES cells from discarded, non-transferred human embryos. Human ES cells were isolated from the inner cell mass of a blastocyst (an embryo five days postfertilization) using an immunosurgical technique where complement proteins and antibodies lyse the trophectoderm so that only the inner cell mass survives. Given that some cells cannot be expanded ex vivo, ES cells could be the ideal resource for tissue engineering because of their fundamental properties: the ability to self-renew indefinitely and the ability to differentiate into cells from all three embryonic germ layers (Deb et al., 2008). These cells have demonstrated longevity in culture and can maintain their undifferentiated state for at least 80 passages when grown using current published protocols. During the past few years, ES cells have evolved as a strong contender for cell therapies and tissue engineering (Deb and Sarda, 2008). These cells offer unique advantages over other cell types for therapy, and some of these therapies based on ES cell progenitors are now close to approval. While protocols for derivation of almost every type of cell from mouse ES cells have been published, derivation of many different cell phenotypes from human ES cells are also being worked out. Human ES cells can give rise to neurons, cardiomyocytes, pancreatic islet cells, hematopoeitic cells, RBCs, etc., under controlled culture conditions. Guided differentiation of human


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ES cells towards dopaminergic and motor neurons in the laboratory is made possible (Wichterle et al., 2002). ES cells allow tissue engineers to define a specific stage of differentiation for transplantation. ES cells also exhibit features that bestow upon them unlimited potential for self-renewal, thus making them prime candidates for cell banking (Mitjavila-Garcia et al., 2005). Banked cells can then be used for therapy when needed. A focus of current research on ES cells is to provide karyotypic stability over repeated passages (Buzzard et al., 2004). Human ES cells cultured on mouse fibroblast feeder cells pose a threat for xenogenic contamination and are therefore unsafe for clinical applications. Through modern tissue engineering, considerable efforts are being directed towards eliminating feeder layers so as to make them more useful for therapeutic purposes (Amit et al., 2004). However, clinical application of ES cells is limited because they represent an allogenic resource and thus have the potential to evoke an immune response. Beside this, the ES cells, when injected, give rise to teratocarcinomas, and this has so far remained a major hurdle. Better differentiation protocols with higher efficiency are required to obtain terminally differentiated derivatives before application in regenerative medicine. Also, there is a need to develop technologies to sort out undifferentiated populations of ES cells following a terminal differentiation protocol (Deb and Sarda, 2008). Alternative and new stem cell technologies (such as somatic cell nuclear transfer and reprogramming) hold promise to overcome the current limitations.

6.2.2 Therapeutic cloning: somatic cell nuclear transfer Somatic cell nuclear transfer (SCNT) is a technique used to derive patientspecific pluripotent stem cells. SCNT basically means the removal of an oocyte nucleus in vitro, followed by its replacement with a nucleus derived from a somatic cell obtained from a patient. Activation of this oocyte to stimulate cell divisions up to the blastocyst stage, is then attained either with chemicals or electricity. At this time, the inner cell mass is isolated and cultured, resulting in ES cells that are genetically identical to the patient. It has been shown that nuclear transferred ES cells derived from fibroblasts, lymphocytes, and olfactory neurons are pluripotent and generate live pups after tetraploid blastocyst complementation, showing the same developmental potential as fertilized blastocysts (Byrne et al., 2007; Meissner et al., 2007; Mitalipov, 2007; Wernig et al., 2007). Thus, the resulting ES cells are a perfect match to the patient’s immune system and would prevent rejection. Although SCNT-derived ES cells contain the nuclear genome of the donor cells, the mitochondrial DNA (mtDNA) contained in the oocyte could lead to immunogenicity after transplantation. Lanza et al. carried out an elegant experiment to assess the histocompatibilty of nuclear



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transfer-generated tissues. They microinjected the nucleus of a bovine skin fibroblast into an enucleated oocyte, and blastocysts were generated (Lanza et al., 2002). The blastocysts were implanted (reproductive cloning), with a purpose of harvesting renal, cardiac and skeletal muscle cells, which were then expanded in vitro, and seeded onto biodegradable scaffolds. These scaffolds were then implanted into the donor from whom the cells were cloned, to determine if the cells were histocompatible. The experiment revealed that cloned renal cells showed no evidence of T-cell response, suggesting that rejection would not necessarily occur in the presence of oocytederived mtDNA. These findings demonstrate the power of SCNT in overcoming the histocompatibility problem of stem cell therapy. Although promising, SCNT has certain technical and ethical limitations that require further improvement before its clinical application. The ethical concern is pertaining to the potential of the embryos resulting from SCNT to develop into cloned embryos if implanted into a uterus. Many animal studies have shown that blastocysts generated from SCNT can give rise to a liveborn infant that is a clone of the donor. In 1997, for example, a sheep named Dolly was derived from an adult somatic cell using nuclear transfer. This process is known as reproductive cloning, which is banned in most countries for human applications. In contrast, therapeutic cloning is used to generate only ES cell lines whose genetic material is identical to that of their source, and the generated blastocysts are never implanted into a uterus. In this case, blastocysts are only allowed to grow in culture until they reach a 100-cell stage, from which ES cells can be obtained. Recently, nonhuman primate ES cell lines were generated by SCNT of nuclei from adult skin fibroblasts (Byrne et al., 2007; Mitalipov, 2007). A total of 304 oocytes yielded 35 blastocysts, from which two ES cell lines were derived. Both lines demonstrated typical ES cell morphology. They also demonstrated selfrenewal and expressed the stem cell markers OCT4, SSEA4, LEFTYA, TDGF, TRA1-60 and TRA1-80. To test their differentiation potential, the cells were exposed to cardiac and neural differentiation conditions, and these experiments resulted in cells that expressed markers of the specified lineages. When injected into SCID (severe combined immunodeficiency) mice, the SCNT-derived non-human primate ES cells induced teratomas that contained differentiated cell types from all three embryonic germ layers. A careful assessment of the quality of the lines must be done before SCNT-derived ES cells can be used for therapy. Some cell lines generated by SCNT have been shown to contain chromosomal translocations and it is not known whether these abnormalities originated from aneuploid embryos or if they occurred during ES cell isolation and culture. In addition, the low efficiency of SNCT (0.7%) and the inadequate supply of human oocytes further limits the therapeutic applications of this technique.


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6.2.3 Reprogrammed somatic induced pluripotent stem (iPS) cells Very recently, many exciting reports of the successful transformation of adult cells into pluripotent stem cells through various kinds of genetic ‘reprogramming’ have been published. Reprogramming is a technique used to de-differentiate adult somatic cells to produce patient-specific pluripotent stem cells. Yamanaka was the first to discover that mouse embryonic fibroblasts (MEFs) and adult mouse fibroblasts could be genetically modified and reprogrammed into an ‘induced pluripotent state’ (iPS) (Takahashi and Yamanaka, 2006). Cells generated by reprogramming would be genetically identical to the somatic cells (and thus, the patient who donated these cells) and therefore would not be rejected. Yamanaka’s group used MEFs engineered to express a neomycin resistance gene from the Fbx15 locus, a gene specifically expressed only in ES cells. They examined twenty-four genes that were thought to be important for embryonic stem cells and identified four key genes that, when introduced into the reporter fibroblasts, resulted in drug-resistant cells. These were Oct3/4, Sox2, c-Myc, and Klf4. This experiment indicated that expression of the four genes in these transgenic MEFs led to expression of a gene specific for ES cells. The resultant iPS cells possessed the immortal growth characteristics of self-renewing ES cells, expressed genes specific for ES cells, and generated embryoid bodies in vitro and teratomas in vivo. When iPS cells were injected into mouse blastocysts, they contributed to a variety of cell types. However, although iPS cells selected in this way were pluripotent, they were not identical to ES cells. Unlike ES cells, chimeras made from iPS cells did not result in full-term pregnancies. Gene expression profiles of the iPS cells showed that they possessed a distinct gene expression signature that was different from that of ES cells. In addition, the epigenetic state of the iPS cells was somewhere between that found in somatic cells and that found in ES cells, suggesting that the reprogramming was incomplete. These results were improved significantly by Wernig et al. (2007). Fibroblasts were infected with retroviral vectors and selected for the activation of endogenous Oct4 or Nanog genes. Results from this study showed that DNA methylation, gene expression profiles, and the chromatin state of the reprogrammed cells were similar to those of ES cells. Teratomas induced by these cells contained differentiated cell types representing all three embryonic germ layers. Most importantly, the reprogrammed cells from this experiment were able to form viable chimeras and contribute to the germ line like ES cells, suggesting that these iPS cells were completely reprogrammed. Wernig et al. observed that the number of reprogrammed colonies increased when drug selection was initiated later (day 20 rather than day 3 post-transduction). This suggests that reprogramming is a slow and



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gradual process and may explain why previous attempts resulted in incomplete reprogramming. It was recently shown that reprogramming of human cells is possible (Takahashi et al., 2007; Yu et al., 2007). Yamanaka’s group showed that retrovirus-mediated transfection of Oct3/4, Sox2, Klf4, and c-Myc generates human iPS cells that are similar to hES cells in terms of morphology, proliferation, gene expression, surface markers, and teratoma formation. Thompson’s group showed that retroviral transduction of OCT4, SOX2, NANOG, and LIN28 could generate pluripotent stem cells without introducing any oncogenes (c-Myc). Both studies showed that human iPS were similar but not identical to hES cells (Takahashi et al., 2007; Yu et al., 2007). Another concern is that these iPS cells contain three to six retroviral integrations (one for each factor), which may increase the risk of tumorigenesis. It was found that tumor formation occurred in chimeric mice generated from Nanog-iPS cells and also that 20% of the offspring developed tumors due to the retroviral expression of c-Myc. An alternative approach using a transient expression method, such as a non-integrating adenovirusmediated system, was successfully used to produce iPS cells recently. Both Jaenisch and Yamanaka showed strong silencing of the viral-controlled transcripts in iPS cells (Okita et al., 2007; Meissner et al., 2007). This indicates that these viral genes are required only for the induction, not the maintenance, of pluripotency. Another concern about the findings was the use of transgenic donor cells for reprogrammed cells in the mouse studies. Both studies used donor cells from transgenic mice harboring a drug resistance gene driven by Fbx15, Oct3/4, or Nanog promoters so that if these ES cell-specific genes were activated, the resulting cells could be easily selected using neomycin. It was important to assess whether iPS cells could be derived from nontransgenic donor cells, for applications in therapy. Wild type MEF and adult skin cells were therefore retrovirally transduced with Oct3/4, Sox2, c-Myc, and Klf4 and ES-like colonies were isolated by morphology alone, without the use of drug selection for Oct4 or Nanog (Meissner et al., 2007). These iPS cells from wild type donor cells formed teratomas and generated live chimeras. This study suggests that transgenic donor cells are not necessary to generate IPS cells. Although very exciting, there are still many questions about the reprogramming that need to be addressed before these cells become clinically eligible. Use of retroviral and adenoviral constructs for genetically modifying cells may face problems in getting approvals from the regulatory agencies in many countries such as India. Trials are therefore being carried out in our laboratories to reprogram cells by using chemical entities or natural compounds. We are also trying to reprogram cells by changing their niche, and the microenvironment. Micro-encapsulation of the cells using different


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biomaterials is already being carried out in our laboratories to achieve such de-differentiation from somatic cells. Tissue engineering may therefore have a major role to play in this area.

6.2.4 Growth factors For successful application of cells in tissue engineering, it is essential to understand and modulate the key signaling mechanisms involved in cellular differentiation and proliferation. The processes of differentiation and proliferation are controlled by both intrinsic regulators and the extracellular environment. The intrinsic environment consists of a cocktail of growth factors, signaling molecules, and hormones. The signaling pathways are triggered when an appropriate receptor is activated by a ligand. This ligand can be a protein, mechanical stimulus, molecules like retinoic acid (Wichterle et al., 2002) or a synthetic small molecule. Mechanisms of action of most drug molecules in clinical practice today are based on such ligand–receptor interaction. FGF, Wnt, Hedgehog, TGF/BMP, and Notch are some of the signaling pathways responsible for mammalian organogenesis. These pathways are also supposed to play a significant role in tissue repair and regeneration (Wu and Sheng, 2006). Kinase proteins on cell surfaces are postulated to be some of the key receptors for regulation of cell differentiation (Rzucidlo et al., 2007; Sato et al., 2004). Interestingly, protein kinases are favored by medicinal chemists also as potential targets for many disorders (Thaimattam et al., 2007). Thus a few investigators, aided by advances in proteomics and bioinformatics tools, have already started designing and screening libraries of synthetic small molecules that selectively bind to these receptors and bring about morphogenesis of cells into a specific lineage (Sheng et al., 2003). The quintessential role of telomerase in regulating cell proliferation and division is well appreciated (Klinger et al., 2006; Peterson and Niklason, 2007). In the years to come, synthetic molecules which can bind selectively to the signaling receptors bcl-2 and telomerase will become an indispensable tool for the tissue engineer in regenerative therapy. Some of the growth factors used for tissue engineering purpose are fibroblasts growth factor, bone morphogenic protein, insulin-like growth factor, and vascular endothelial growth factor, amongst others.

6.2.5 General scaffolds and materials Neither isolated cells nor growth factors can be relied upon to produce tissues of all practical sizes and shapes. Here, tissue engineering enters the domain of material science. Development of biomaterials for tissue engineering is as exciting a field of research as stem cells and molecular biology. The need of biomaterials can be understood looking at the role of extra



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cellular matrix (ECM) in native tissue environment. This matrix, composed of various biopolymers conjugated with glycol-saccharides and proteins, is responsible for providing mechanical strength and physical support to the growing cells, and directs their proliferation, differentiation, and morphogenesis (Zagris, 2001). It allows cells to assume a particular shape and size. During any tissue damage, a significant amount of ECM is also lost. Scaffolds are those structures that are fabricated to replace ECM in vivo or provide an artificial ECM for in vitro regeneration. It is sometimes possible to regenerate tissues by providing just the scaffold at the site of injury. Thus, cells get the necessary support to grow. As the growing cells can secrete their own ECM, an ideal scaffold should degrade with time, without leaving any scar. The degradation products should be non-toxic. In tissues, where the cells have little inherent regeneration potential, the scaffold itself may not be sufficient. In these cases, scaffolds are seeded with replicative cells and suitable growth factors to form the desired tissue. The characteristics of a good scaffold material are high porosity and proper pore size, high surface-area-to-volume ratio – this allows maximum cell–cell interaction and cell–scaffold interaction, biodegradability, high mechanical strength, and high degree of biocompatibility (Ma, 2004). Below we present some of the materials used in tissue engineering. Metals The impact of metallic implants prevails in orthopedics, dentistry, craniofacial surgeries, etc. Stainless steel, cobalt and titanium based alloys are the metals mostly used in medicine (Brunski, 1996). Inertness coupled with mechanical integrity has made metals suitable for implantation. Though more attention is paid to other types of biomaterials for tissue engineering, metals such as titanium do find continued use as an adjunct material (Liu et al., 2006). Ceramics Ceramics and bio-glasses comprise a range of inorganic materials with established use in dentistry and some medical application such as eye glasses. These materials include alumina, silicon dioxide, TiO2, CaO and P2O5. Some porous ceramics isolated from corals appear to be suitable, for they can effectively provide surface for bone remodeling, especially when load bearing is not the primary criterion (Schors and Holmes, 1993). Amongst all ceramic materials, calcium phosphate has gained widespread acceptability for tissue engineering. Calcium phosphate and hydroxyapetite impregnated with growth factors and antibiotics appear attractive propositions for scaffold fabrication (Nie and Wang, 2007).


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Natural polymers Natural polymers mimic the native ECM down to the molecular level and are better materials for the biological system to recognize. Also, many of them possess the property of self-assembly and hence it is simple to build the ECM in vitro from such polymers. They are degraded into noninflammatory products (a significant advantage over their synthetic counterparts), but these polymers may be immunogenic in nature. Since they are obtained from natural sources, they carry a risk of microbiological contamination, and their exact composition is often difficult to ascertain and reproduce. Many of the components are proteinacious in nature so they are less amenable to the rigors of processing conditions. Collagen, dextran, fibrin, silk, gelatin, sodium alginate, chitosan, and hyaluronic acid are some of such materials that are extensively used for tissue engineering (Yannas, 1993). In fact, collagen was one of the first materials to be applied successfully for clinical tissue engineering (Burke et al., 1981).

Synthetic polymers As a tissue engineer, biomaterial needs to play the role of a temporary scaffold, wherein it is supposed to degrade at a rate corresponding to the kinetics of neo-tissue formation; novel biodegradable materials have gained preference over erstwhile popular synthetic polymers such as polymethylmethacrylate (PMMA) and polyethylene. Linear aliphatic polyesters are amongst the most popular materials for tissue fabrication. Some of their fascinating properties are controlled degradation, and the ability to entrap various drug molecules and yield materials of different strengths. For example, polyglycolic acid (PGA) is a hydrophilic polymer with very fast degradation rate. Polylactic acid, on the other hand, is hydrophobic and persists for over a year in vivo. By blending these polymers in different ratios, materials with customized biological properties can be synthesized. This approach has already evinced much interest in drug delivery, and tissue engineering is fast picking up. Another important advantage is the vast amount of chemical modification that can be applied to these polymers to get desired biocompatibility and cell–material interactions. Polycaprolactone, polyphospharazenes and polyanhydrides are other materials of this class with potential use in tissue engineering (Katti et al., 2002; SokolskyPapkov et al., 2007). Biodegradable polymer scaffolds with interconnected pores allow formation of complex D tissues during differentiation. The scaffold, designed to withstand the compressive strength of cells, provides physical cues for cell orientation and spreading, and promotes differentiation, proliferation and organization of cells. The pores provide space for remodeling of tissue structures (Richardson et al., 2001). Variation of growth



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factor delivery from these scaffolds can lead to formation of desired tissuelike cartilage, and neural and blood vessels. Such engineered tissue-like structures can then be directly used for transplantation (Vacanti and Langer, 1999). From the elements of tissue engineering it can be implied that this multidisciplinary field involves interplay between some of the most exciting fields of modern science, namely, stem cell biology, molecular genetics, protein drug delivery and biomaterials sciences, along with the essential tenets of physics, chemistry, medicine and engineering. Advances in the regeneration of specific tissues and organs using tissue engineering approaches are discussed in the next section.

6.3

Tissue engineering and stem cells in organ regeneration

Currently, the tools to create scaffolds and structures for organ regeneration are limited, and to have precise control over the mechanical and microenvironmental variables is difficult. The other limitation is that there are only a few molecular, cellular, and tissue levels tests to evaluate what is actually happening within the constructs. However, tissue engineers are doing their best in trying to mimic the environment of the tissue. This section presents a brief account of the developments in several major areas of organ regeneration.

6.3.1 Prospects for liver regeneration Liver is one of the most perfused organs and is usually one of the first organs to be exposed to xenobiotics and pathogens. Liver cirrhosis, acute liver failure, and viral hepatitis are known to affect a significantly large number of patients. Many of them are hospitalized at critical stages of organ failure and need immediate organ substitution. Bioartificial liver support systems or liver bioreactors have thus occupied biomedical engineers’ efforts for a long time (Sussman and Kelly, 1993). Some of them are widely used clinically, while modifications on others continue to be made. They are, however, not permanent solutions. For such patients, liver tissue engineering is an exciting prospect. Fortunately, compared to functional cells of other vital organs, liver cells (hepatocytes) can regulate their growth and size in response to damage. This phenomenon is referred to as compensatory hyperplasia (Kay and Fausto, 1997). Thus, obtaining part of liver cells for allogenic or autologous engineering does not cause severe damage to the donor organ. However, obtaining donors for all such patients is a daunting task, as indicated by the American Liver Foundation – only 3000 donors for about 30 000 patients affected with one form or another of liver failure


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(Langer and Vacanti, 1993). In such cases the patients may be assisted on an artificial bioreactor, and meanwhile a fully functional organ can be grown from hepatocytes of the donor or the patient himself. Functional tissue engineered liver can then be transplanted. It has been possible to derive functional liver cells from human embryonic stem cells. Also hepatocyte precursor cells have been isolated successfully from liver tissue. With such advances it would be possible to grow the hepatocytes in vitro on supporting three-dimensional scaffolds, generating artificial livers ready for transplantation. A second avenue for liver tissue engineering lies in treatment of metabolic disorders by gene therapy (Kay and Woo, 1994). Liver is a major biosynthetic factory of the body and is responsible for secretion (and removal) of many essential biomolecules in (from) circulation. Tissue engineering promises cures for various hepato-deficiency disorders that manifest in metabolic abnormalities. The methodology adopted involves isolation of the hepatocytes incapable of synthesizing a particular molecule due to genetic abnormality, followed by genetic engineering of these cells, and their re-introduction in patients. Such attempts have been tried in patients suffering from familial hypercholesterolamia (Grossman et al., 1994).

6.3.2 Myocardial tissue engineering Worldwide, thousands of patients are afflicted with myocardial infarction (MI), which deteriorates to advanced heart failure despite optimal medical therapy. MI results in more hospitalization than all forms of cancer and is associated with a significant number of morbidities (Capi and Gepstein, 2006; Jessup and Brozena, 2003). Limitations of current therapies and transplantations have pushed great interest in tissue engineering to repair the failing heart. Human heart comprises terminally differentiated cardiomyocytes, which have very limited potential for self-renewal. Today, unlike heart valves or blood vessels, heart muscle has no replacement alternatives and engineering a heart muscle remains the most critical factor for myocardial tissue engineering. Towards this end, various cell sources have been investigated for their potential to yield functional cardiac muscle cells. Fetal cardiomyocytes, embryonic stem cells, skeletal myoblasts, crude bonemarrow cells, mesenchymal stem cells, hematopoietic stem cells, fibroblasts smooth muscle cells, etc., have been proposed by various authors. Recent reports indicate that mesenchymal stem cells collected from the menstrual blood are the best sources for cardiomyocyte differentiation. Extensive work in the isolation of correct cell type is still required, as the ability of skeletal cells to reproducibly differentiate into contractile and electrophysiologically functional cardiac cells is debated, and concerns over their safetyrelated issues (arrhythmias for skeletal progenitors and calcification with



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bone marrow cells) remain unaddressed (Leor et al., 2005). For in vitro cardiac tissue engineering, bioreactors are designed to simulate adjustable pulsatile fluid flow and varying levels of pressure, as encountered by cardiac cells in real time situations. Investigations by one group have identified deciding factors which lead to strongly contracting (up to 3 mN/mm2) and morphologically highly differentiated cardiac muscle constructs named ‘engineered heart tissue’ (EHT). They believe that (i) addition of Matrigel to the reconstitution mixture (only in rat EHT), (ii) EHT culture under cyclic exposure to mechanical loads, (iii) circular shape of the bioreactor, in contrast to EHT patches, and (iv) utilization of cell mixtures rather than purified cardiac myocyte populations are key factors for the in vitro construction of EHT (Zimmermann et al., 2004). The achievements in the field of myocardial tissue engineering are applied in parallel to the repair of congenital cardiac defects. For example, a study has demonstrated full replacement of the ventricular free wall by seeding mesenchymal stem cells in a scaffold made of polytetrafluoroethylene, polylactide mesh, and Type-I and -IV collagen hydrogel in an animal model is possible (Krupnick et al., 2002).

6.3.3 Corneal regeneration and tissue engineering of the eye Of various ocular problems, corneal and conjunctival epithelial cell injury, degenerations, and abnormalities are relatively common. Persistent epithelial defects caused by microbes, chemicals, iatrogenic, physical, and congenital insults, along with retinal deficiencies pose strong threats to vision. Surface diseases such as Stevens–Johnson’s syndrome, chemical and thermal burns, recurrent pterygia, ocular tumors, immunologic conditions, radiation injury, inherited and congenital syndromes, aniridia, and ocular pemphigoid, severely compromise the ocular surface and cause catastrophic visual loss in otherwise potentially healthy eyes. Treatment is expensive, frustrating, time-consuming, and often unsuccessful (Schwab, 1999). Tissue engineering of autologous limbal epithelial cells cultured on amniotic membrane and 3T3 fibroblasts is found to be a simple and effective method of reconstructing the corneal surface and restoring useful vision in patients with unilateral deficiency of limbal epithelial cells (Pellegrini et al., 1997; Tsai et al., 2000). This is also discussed in more detail in an earlier chapter on this topic in this book.

6.3.4 Cartilage tissue engineering Damage to the articular cartilage by trauma or degenerative joint diseases, such as primary osteoarthritis, causes disabling joint pains. Articular


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cartilage facilitates movement by minimizing friction between joints (as lubricants) and allows load-bearing through distribution of mechanical stress and acts as a shock absorber against compressive forces. Once damaged, cartilage tissue possesses very limited potential for healing. Current treatment methods for restoration of function of articular cartilage, other than total joint arthroplasty, are autografting, allografting, periosteal and perichondrial grafting, stimulation of intrinsic regeneration by intentionally drilling full-thickness defects, pharmacological intervention, and, autologous cell transplantation such as the periosteal flap technique marketed by Genzyme Corp. (Cambridge, MA, USA). Despite these, cartilage damage often cannot be repaired to a fully functional normal state (Brittberg et al., 1994; Tuan et al., 2003). Scaffolds such as alginate and collagen hydrogels which can maintain the rounded morphology of chondrocytes are some of the favorable cell carriers for cartilage tissue engineering (Kuo and Tuan, 2003). In addition, genetic elements responsible for chondrogenic differentiation have been identified and will aid articular engineering (Bursell et al., 2007).

6.3.5 Tissue engineering of the bone As life expectancy increases, so does the need to treat large-sized bone defects. New biomaterials combined with osteogenic cells are now being developed as an alternative to bone grafts. For example, bioceramics such as hydroxyapatite and tricalcium phosphate can be manipulated to match the mechanical properties of bone and are generally considered osteoinductive (inducing new bone formation), and with highly porous architecture, they are considered osteoconductive (favoring new bone tissue ingrowth). Porous ceramics of hydroxyapatite and β-tricalcium phosphate loaded with MSCs were shown to be capable of healing critical-sized segmental bone defects not capable of being healed by implantation of scaffold alone. MSCs seeded in bioceramics have been shown to regenerate bone in a variety of studies (Bruder et al., 1998). Bone tissue formation after in vivo transplantation of autologous bone marrow-derived cells in the bone defect of a geriatric patient has been achieved in the clinic. Bone marrow cells were expanded in vitro during which time they were exposed to novel recombinant human transforming growth factor 1, fusion protein bearing a collagen-binding domain, dexamethasone and β-glycerophosphate, followed by loading them into porous ceramic scaffolds (Becerra et al., 2006). Current research is focused on evaluation of biomaterials for their ability to promote osteogenesis (Hee and Nicoll, 2006; Jager et al., 2005).

6.3.6 Vascular and valvular tissue engineering Peripheral vascular diseases and diseases of coronary arteries, e.g. atherosclerosis, have immense implications on healthcare. Coronary artery bypass



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and balloon angioplasty, along with use of pharmacotherapy are current treatment modalities. While balloon angioplasty is considered a simple and effective tool in the armory of interventional cardiologists, recurrent restenosis largely limits its universal application. Working out a solution to recurrent restenosis after balloon angioplasty of arteries is an exciting area of research in itself and has led to the development of drug-eluting stents (King, 2007). Autologous coronary bypass surgery, on the other hand, is difficult in a significant population of patients due to lack of arteries suitable for the purpose (Yow et al., 2006). Existing cardiac disease further complicates the condition of patients by retarding the natural healing process of both small capillaries (angiogenesis) and large vessels (arteriogenesis) (Hill et al., 2003). Thus, the great clinical need to restore vascular functions has driven the medical and bioengineering community towards fabrication of blood vessel substitutes, called ‘designer vessels’, tissueengineered vascular conduits, arterial grafts, etc. Such conduits must be nonthrombogenic, nonimmunogenic, and suturable, and must possess adequate mechanical strengths and appropriate functional and healing responses. Conduits made of Dacron or expanded polytetrafluoroethylene (ePTFE) are known to perform adequately in high-flow, large-diameter environments and were primitive strides in this area. However, these materials do not show proper utility for small-diameter vessel ( Cu2+ > Cd2+ > Ba2+ > Sr2+ > Ca2+ > Co2+ = Ni2+ = Zn2+ > Mn2+. The association between the alginate polymer chain and divalent cations has been popularly described using the ‘egg-box’ model (Fig. 9.2). However, the details of this model have been questioned by many researchers (Braccini and Pérez, 2001; Kim and Han, 2000; Grant et al., 1973;


240

Biointegration of medical implant materials Polymer strand 1 O COO HO O

HO

HO

HO

O

O COO− HO O

H

−

HO

O

H

H

H

O O

O H

HO

O

Ca2+

O

HO

OH

O

O O

OH

O

O OH O−OC H

O

O

H

OH H

Ca2+ O− OH O OH − O OC O

O

O

O OH

O

O OH

Polymer strand 2

9.2 Schematic representation of the interstrand crosslinking of alginate by calcium ions in the so-called ‘egg-box’ model.

Sikorski et al., 2007). The binding involves coordination of one divalent cation with four oxygen atoms. Braccini and Pérez (2001) proposed that the parallel and antiparallel arrangement of 21 helical chains provides a favorable association, also providing a compact cavity for a highly cooperative binding with the divalent cation. More recently, Li et al. proposed 31 helical conformation based on x-ray diffraction of Ca-alginate gels that were formed by a slow gelation method (Li et al., 2007). Divalent cations such as Ca2+ and Ba2+ bind preferentially to the G blocks and thus alginate gels with lower G content (higher M content) have lower strength and stability (Mørch et al., 2008; Jørgensen et al., 2007). Grant et al. (1973) demonstrated that although the interactions of alginate with cations are dominated by G blocks, once the G binding sites are saturated the threshold is passed over to the MG blocks. Thus, a more robust hydrogel can be engineered by using a G-rich alginate or by increasing the amount of G blocks by enzymatic epimerization of the polymer (Mørch et al., 2007). Ionically crosslinked alginate microcapsules can be prepared by various methods (Fig. 9.3) including suspending the extruded alginate beads in a crosslinker solution (Skaugrud et al., 1999; Aslani and Kennedy, 1996) or emulsification followed by addition of a crosslinking solution (Fundueanu et al., 1999). Modifications of the emulsification technique to achieve internal or external gelation have also been reported (Poncelet et al., 1992; Chan et al., 2002). A combination of two methods can also be used to produce desirable gels or to achieve better control of the gelation process (Tan and Takeuchi, 2007). Ionic crosslinking of alginate membranes is commonly


Alginate-based drug delivery devices

241

Sodium alginate solution

Immersion Incorporated calcium carbonate

Calcium chloride solution

Acidic solution

Injecting barium chloride crystals

Barium chloride solution

9.3 Illustration of the common techniques used to produce ionically crosslinked alginate capsules.

achieved by the immersion technique, where the xerogel is placed in a crosslinker solution allowing the crosslinking agent to diffuse into the matrix (Corkhill et al., 1990; Pavlath et al., 1999; al Musa et al., 1999; Jejurikar et al., 2008). Recently a new method of pressure-assisted diffusion was developed to achieve more uniform crosslinking (Jejurikar et al., 2008). The internal structure of alginate hydrogels can be investigated by cryogenic SEM, which allows evaluation of pore size and distribution of the hydrated gel (Fig. 9.4). These ionically crosslinked hydrogels swell significantly in an aqueous media. Ca-alginate capsules are known to swell more than 90% over long-term immersion in physiological media under optimal pH and temperature conditions. The extent of swelling of a crosslinked alginate gel is dependent on temperature, pH, and ion concentration of the swelling media, and also on the crosslinking gradient through the gel matrix (Moe et al., 1993). In a recent study, Qin reported that Ca-alginate gels with higher M content swell more than those with higher G content, which reflects the preferential binding of the cations to the G sub-units (Qin, 2008). The mechanical properties of alginate hydrogels change over an initial period of immersion in a physiologically relevant solution (Le Roux et al., 1999; Kuo and Ma, 2008) and are dependent on the technique and material used to prepare the hydrogels, as illustrated in Table 9.1. The stress–strain response of the hydrogel is dependent upon the extent of crosslinking, the porous structure of the matrix and the water flow within.


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9.4 Cryogenic scanning electron micrograph of a hydrated calcium crosslinked alginate matrix showing its porous nature.

Table 9.1 Mechanical properties of calcium crosslinked alginate hydrogels Alginate sample and method for preparation Capsules prepared by extrusion method (Nguyen et al., 2009) Cylinders prepared by internal gelation method (le Roux et al., 1999) Discs prepared by emulsion method (Webber and Shull, 2004) Discs prepared by internal gelation method (Drury et al., 2004) Films prepared by immersion (Jejurikar et al., 2008) Films prepared by pressure assisted diffusion (Jejurikar et al., 2008)


Maximum modulus (kPa)

Compression

490

Compression

10

Compression

100

Tensile

50

Tensile

5 400

Tensile

16 500

Degradation of ionically crosslinked alginate gels is initiated by loss of crosslinking cation, followed by loss of high and low molecular weight polymer strands. This affects the gel strength, stability, pore size and pore distribution (Shoichet et al., 1995). The rate of degradation is affected by the alginate properties (in particular the M/G ratio) and also by the solution



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in which the alginate gel is suspended (Tan et al., 2009). The effect of the calcium concentration in the medium was investigated in detail by Kuo and Ma (2008) and it was found that a high concentration (5 mM) resulted in greater retention of the crosslinking density than a lower concentration (e.g. 2–4 mM). It is of course important to realize that the mechanism and rate of degradation in vivo is also dependent on the site of implantation and the local cellular environment.

9.2.4 Chemically modified alginate Chemical modification of the alginate biopolymer can change the reactivity of alginate either by introducing highly reactive functional groups (e.g. aldehyde groups) or by introducing chemical (e.g. phosphate) or biochemical (e.g. amino acids) groups that can increase the biointegration of alginate-based materials. Since the alginate biopolymer itself contains both hydroxyl and carboxylic acid functional groups, it offers great versatility for chemical modification, as will be illustrated below. It has recently been demonstrated that phosphorylation of alginate can be achieved using the so-called urea phosphate method (Coleman, 2007). This entails reacting a dispersion of the biopolymer with phosphoric acid in DMF in the presence of urea (Mucalo et al., 1995). In this manner, phosphate groups are introduced at C1, C2 and C3 through reactions with the hydroxyl groups. Characterization by NMR (1H, 13C and 31P) established that the dominant site of phosphorylation on d-mannuronic acid residues was M-3, which indeed is the equatorial group and therefore more reactive than its axial counterparts due to reduced steric hindrance. Because of the high acidity of the solution during the phosphorylation reaction, some degradation of the biopolymer occurs concurrently with the phosphorylation reaction. It was thus possible to achieve up to 26% phosphorylation (i.e. 26% of all alginate subunits carry a phosphate group) and this is paralleled with a decrease in molecular weight from approximately 140 to 40 kDa. Coupling of single amino acids (e.g. cysteine) or small peptides (e.g. containing the RGD sequence) to alginates has been studied in detail. In all cases the amine terminal of the amino acid/peptide is reacted with the carboxylic acid group of the alginate using carbodiimide chemistry (i.e. activation of the carboxylic acid group). More specifically, cysteine, has been coupled to alginate with 7% substitution (of alginate subunits, 400 μmol/g) when using stoichiometric amounts of cysteine, as determined photometrically (Greimel et al., 2007). The aim was to improve the mucoadhesive properties of ionically crosslinked capsules used for oral delivery (Section 9.3.2). Other amino acids including lysine, arginine, aspartic acid and phenylalanine have likewise been coupled to alginate (Zhu, 2002). Successful reaction was verified from FTIR and 1H NMR; and the amino acid content


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was found to be 3% (alginate subunits) based on analysis by UV spectrometry. The RGD peptide sequence is used extensively in biomaterials science due to its cell adhesion properties. A number of RGD-containing peptides have been coupled to alginate using carbodiimide chemistry; including GRGDY (Rowley et al., 1999), GGGGRGDSP (Drury et al., 2005), and GGGGRGDSY (Connelly et al., 2007). The coupling efficiency was determined using 125I labeled peptide and found to be 25% (alginate subunits) without further improvement with increased carboxylic acid activation (Rowley et al., 1999). These peptide-modified alginate biopolymers have found many applications, some of which will be discussed in the following sections of this chapter. Amphiphilic alginate derivatives have been synthesized by reacting the carboxylate groups of alginate, transformed into a tertrabutyl ammonium salt, in the organic solvent dimethylsulfoxide with alkyl halides (e.g. dodecyl bromide) (Pelletier et al., 2000). In this manner, the long alkyl chains were linked to the alginate biopolymer via ester linkages. An 8% (C12 alkyl chain length) and 1.3% (C18 alkyl chain length) substitution ratio could be achieved in this manner, as determined through gas chromatography of the alkyl alcohol obtained from alkaline hydrolysis of the amphiphilic polymer. An alternative method of reacting the carboxylic acid group of alginate, pre-activated using chloro-1-methylpyridinium iodide (CMPI), with an alkylamine (e.g. dodecylamine) in dimethyl formamide yielded substitution ratios in the range of 2–17% and this could be easily controlled by the amount of CMPI used (Vallee et al., 2009). Such amphiphilic alginate biopolymers can form physical gels due to hydrophobic interactions, and these gels can be further strengthened by calcium crosslinking (de Boisseson et al., 2004). Oxidation of alginate using periodate was first described by Malaprade (1928). It involves cleavage of the C2–C3 bond, transforming the uronic acid subunits into an open chain adduct containing a dialdehyde known as alginate dialdehyde (ADA) (Fig. 9.5). These aldehyde groups can react spontaneously with hydroxyl groups present on the adjacent uronic acid subunits in the polymer chain to form a hemiacetal (Balakrishnan et al., 2005). Its formation can prevent complete oxidation of the biopolymer and consumption of sodium periodate. The open chain adduct created by the periodate oxidation of alginate is much more susceptible to hydrolytic scission than unmodified alginate (Bouhadir et al., 2001). Oxidation of alginate is usually carried out in an aqueous solution at room temperature (Balakrishnan et al., 2005; Gomez et al., 2007). The degree of oxidation can be controlled, and up to 87% oxidation has been demonstrated (Balakrishnan et al., 2005). The degree of oxidation can be determined via iodometry by titrating any periodate remaining after reaction (Whistler and Wolfrom, 1962). Characterization is carried out using FTIR (formation of the alde-


Alginate-based drug delivery devices –

–

OOC

OH

OHO

OH –

OOC O

O

OH

OOC

245

OH O

O HO

Sodium periodate (aqueous) –

–

OOC

OHO

OH

–

OOC O

CH O

O HC

OOC

O HO

OH O

:O:

9.5 Reaction scheme of the oxidation reaction of alginate forming dialdehyde alginate (ADA).

hyde and acetal bands), 1H NMR (aldehyde content) (Gomez et al., 2007), and 13C NMR (to confirm the polymer scissioning that is accompanied by oxidation). Extensive scissioning of the polymer chain is reflected in the decrease in molecular weight with an increase in the amount of oxidation. Vold et al. (2006) and Smidsrød and Painter (1973) demonstrated changes in the intrinsic viscosity and chain stiffness in the oxidized alginate samples and concluded that the stiffness and the length of chains were decreased with an increase in the oxidation of alginate samples. The enhanced flexibility of the polymer chains was linked to the open-chain adducts. The gelation ability of ADA in the presence of Ca2+ ions is maintained at low oxidation levels (e.g. 10%); however, the Ca-ADA hydrogels show reduced mechanical properties. This is related to the reduced molar mass and number of GG blocks participating in binding with divalent cations. ADAs are generally biocompatible; however, alginates oxidized to a very high degree can induce an immunological response as a consequence of the presence of free radicals. It has been reported that alginates are resistant to biodegradation and are not broken down in mammals (Al-Shamkhani and Duncan, 1995). However, a controlled degradation of alginate can be achieved by partially oxidizing the alginate polymer, which opens avenues for use of ADA in tissue engineering applications (Bouhadir et al., 2001).

9.2.5 Covalently crosslinked matrices Covalently crosslinked hydrogels, possessing increased stability compared to their ionically crosslinked counterparts, allow for a greater control over mechanical and swelling properties of the gels, and can be engineered to suit various applications such as cartilage repair, wound dressing and drug


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delivery. As discussed above, the hydroxyl and carboxylic acid functional groups of alginate offer great versatility for chemical modification, including crosslinking reactions. Formation of covalently crosslinked hydrogels can thus be achieved by reaction of these functional groups with complementary reactive groups (Augst et al., 2006). Crosslinking molecules such as dialdehydes and diamines have been used with suitable catalysts to create hydrogels. A crosslinking reaction between the hydroxyl groups of alginate and the aldehyde groups of typically gluteraldehyde (1,5-pentanedial) is one approach that has been studied extensively (Chan et al., 2008; Yeom and Lee, 1998). The chemical reaction leads to the formation of acetal groups in the crosslinked network. Kim et al. demonstrated that gluteraldehydecrosslinked alginate hydrogel fibers exhibit high absorbency (Kim et al., 2000), which could be increased by decreasing the crosslinker concentration to just above the critical concentration required for crosslinking of the polymer in agreement with general hydrogel theory. These hydrogels have potential application in products such as additives for soil in agriculture, water-blocking tapes, sanitary napkins, disposable diapers or drug delivery systems, where water retention is important, or as membrane material for pervaporation processes. However, gluteraldehyde is toxic to cells even in low concentration and hence these hydrogels are not ideal candidates for biomedical applications. Alternatively, alginate dialdehyde (ADA) obtained by periodate oxidation of alginate can be used as the aldehyde crosslinker to react both intra-molecularily with other parts of the ADA and intermolecularily with unmodified alginate (Jejurikar et al., 2009). These gels displayed very high swelling characteristics and good mechanical strength which could be controlled by the degree of oxidation of ADA, as well as the ratio of ADA to alginate. ADA has also been crosslinked with adipic acid hydrazide. The swelling and degradation properties of these hydrogels were controlled by varying the amount of adipic acid hydrazide in the reaction mixture (Lee et al., 2004). Direct crosslinking of hydroxyl and carboxylic acid groups of alginate using the non-toxic carbodiimide (EDC) catalyst has been demonstrated (Rowley et al., 1999; Xu et al., 2003a). EDC facilitates formation of ester linkages through an intermediate and does not itself remain a part of the structure. Xu et al. (2003b) made blends of alginate and carrageenan and used EDC chemistry to create more durable hydrogels. These hydrogels predominantly contained alginate, thus preserving their water absorbing properties. Xu et al. demonstrated that these hydrogels paralleled the gluteraldehyde-crosslinked hydrogels as good candidates for membrane materials for the pervaporation process (Xu et al., 2003a). EDC chemistry, when coupled with the co-reactant N-hydroxyl succinimide (NHS), has been utilized to facilitate amide bond formation between the carboxylic acid groups



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of alginate and the amine groups of diamines. This strategy has been utilized with methyl ester l-lysine, ethylene diamine, and poly(ethylene glycol) (PEG) diamines (Lee et al., 2000a; Hosoya et al., 2004; Eiselt et al., 1998, 1999). Eiselt et al. developed alginate hydrogels with demonstrated improved mechanical properties, compared to ionically crosslinked alginate hydrogels, by crosslinking alginate and PEG-diamines (Eiselt et al., 1998, 1999). The mechanical properties were controlled by the molecular weight of the PEG and the amount of PEG-diamines introduced into the crosslinked network. Alginate crosslinked with ethylenediamine and/or an amine silane have been reported as useful for making bone-like apatite, artificial skin, and nerves (Hosoya et al., 2004). A number of other methods have also been explored for the covalent crosslinking of alginate, including irradiation and enzymatic techniques. Alginate has been modified using methacrylic anhydride (Smeds and Grinstaff, 2001), leading to an ester linkage between alginate and the vinyl-group containing moiety. These modified polymers in the presence of a photoinitiator and UV irradiation produced covalently crosslinked hydrogels exhibiting desirable swelling and mechanical properties. These gels are of great interest for sutureless closure of wounds and tissue reconstruction/ artificial organs. A recent example of enzymatically crosslinked alginate utilizes tyramine-modified alginate (coupled via amide linkages using EDC/ NHS chemistry) (Sakai and Kawakami, 2008). Horseradish peroxidase was applied to catalyze the oxidation reaction, coupling the phenols, thereby creating covalent crosslinks. This was done on calcium crosslinked gels and it was found that the swelling was reduced and could be controlled by the reaction conditions.

9.3

Drug delivery using alginate matrices

The transport of pharmaceuticals to the desired target site and their release in a controlled and sustained way represents a massive research field of drug delivery. To simply focus on alginate-based systems and claim that these represent the state of the art in this field would be naïve. However, a number of strategies in drug delivery can be illustrated through important examples in polysaccharide-based systems. Drug delivery capsules are designed with two key objectives in mind: the safe transport of the active molecule to the active target site and minimal impact on the host biological system. Typically, engineering of the structure of the assembly is either directed towards improved permeability, stability and retention (passive targeting) or the modification of the functionality of the capsule to improve localized delivery and/or biointegration (active targeting). A number of these strategies are illustrated in Fig. 9.6 and will be discussed further below (Sections 9.3.1 and 9.3.3).


248

Biointegration of medical implant materials Key Alginate

Modified alginate Calcium ion

Cationic polyelectrolyte Anionic polyelectrolyte Covalently crosslinked polyelectrolyte complex

(e) LbL coating

(a) (d)

(f)

(b)

(c)

9.6 Schematic illustration of various approaches used to control permeability and biointegration. (a) Simple crosslinked alginate matrix; (b) crosslinked matrix produced from chemically modified alginate; (c) post modification of alginate matrix with chemical or biological moieties; (d) LbL assembly around alginate matrix; (e) covalent crosslinking of LbL assembly; (f) post modification of LbL assembly.

The strategies that are implemented depend on the nature of the drug to be delivered and the target environment (Table 9.2). Oral delivery of drugs represents a particular challenge due to the complex gastrointestinal environment, and there are several examples of research directed towards acquiring the desired characteristics for efficacy via this delivery route. These strategies include: modification of alginate to introduce binding through disulfide bonds with mucous glycoproteins (Greimel et al., 2007); novel crosslinking strategies to strengthen capsules (Anal and Stevens, 2005); application of a covalently crosslinked chitosan coat (Taqieddin and Amiji, 2004), and introduction of buoyancy to evade gastric emptying (Shishu et al., 2007). The layer-by-layer (LbL) assembly approach has been utilized for a number of different applications with the aim of modulating the release profile of the encapsulated drug through the formation of a diffusional barrier around the alginate capsule. In the work by Matsusaki



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Table 9.2 Common drug delivery environments Delivery site

Desirable properties

Gastrointestinal

Stability in low pH environment. Non-covalent associations with mucous gel layer. Prolonged gastric residence time. Enhanced lifetime within circulatory system. Small size to pass through capillaries. High surface area to volume ratio to maximize contact. Efficient transfer to blood or lymphatic system. High volume and sustained release. Bioresorbable or biomimetic.

Intravenous Peritoneal cavity Subcutaneous Hard tissue fracture

et al. (2007) encapsulating vascular endothelial growth factor (VEGF) in alginate capsules, it was found that varying the number of layers in the LbL assembly had a significant effect on the rate of VEGF release. Modification of microcapsules to reduce the local inflammatory or tissue response in the host represents an emerging area of interest in biointegration (Section 9.3.3). Strategies include the modification of alginate to incorporate antibodies on the outer surface of the microcapsule, e.g. anti-TNF-α (Leung et al., 2008) and heparin as the outer layer of a multilayered capsule assembled by physical adsorption (Bünger et al., 2003).

9.3.1 Use of chemically modified alginate The uses of chemically modified alginate in drug delivery applications have been explored for different aspects of biointegration. The nature of the physicochemical relationship between drug molecule and carrier matrix impacts on the drug release profile and so the design of capsules represents a balance between the matrix/environment and drug/matrix associations. Examples considered here include the use of thiolated alginate for mucoadhesion, the use of amphiphilic alginate derivatives for enhancing protein retention, and the use of oxidized alginate for both formation of a covalently crosslinked matrix and covalent attachment of a drug molecule. While the first example is mainly driven by the interactions of the delivery system with the in vivo environment, the last two examples explore the drug/matrix interactions. Thiomers have been proposed to be the next generation of mucoadhesive polymers to be used in oral drug delivery applications (Bernkop-Schnürch and Greimel, 2005). They have been proposed to act by thiol/disulfide bond exchange reactions and oxidation reactions with cysteine-rich subdomains of mucus glycoproteins, thus forming covalent linkages to mucin via


250


disulfide bonds. In their study on producing mucoadhesive alginate delivery capsules, Greimel et al. (2007) used mixtures of thiolated alginate (modified by cysteine attachment, Section 9.2.4) and poly acrylic acid (PAA) to encapsulate insulin. They found that the introduction of the thiol group offered additional benefit to the delivery system. It increased the encapsulation efficiency from 15% (unmodified alginate) to 65% (thiolated alginate/ PAA) and decreased the release rate. In addition, capsule stability in simulated intestinal fluid was significantly enhanced and this was attributed to the formation of internal disulfide bonds during capsule formation (i.e. 82–85% of free thiol groups were oxidized). These particles thus show some promising characteristics for use in oral drug delivery. Hydrophobically modified alginates have been trialled for encapsulation of proteins through the incorporation long-chain alkyl groups, e.g. C12 via ester bonds (Leonard et al., 2004). These amphiphilic alginate derivatives were used to encapsulate a series of model proteins with very high encapsulation efficiencies (70–100%), in which both hydrophobic interactions and calcium ion crosslinking were contributing to the bead formation process. In addition, the strong interactions between the proteins and the amphiphilic alginate derivatives were illustrated by retention of the proteins in simple buffers. This was in contrast to the rapid release of these proteins from unmodified alginate capsules. Protein release from the amphiphilic capsules was observed only in the presence of the esterase lipase, which hydrolyzes the ester bond, or in the presence of surfactants that disrupt the hydrophobic interactions in the capsule. An elegant use of oxidized alginate (alginate dialdehyde, ADA) in drug delivery has been demonstrated by Balakrishnan and Jayakrishnan (2005). They produced an injectable scaffold from ADA and gelatin in the presence of small amounts of borax. It was observed that the gelling time decreased with increasing concentration of all reagents, while it increased with the degree of oxidation of ADA. Using primaquine as a model drug, they demonstrated slower drug release rates when using ADA of a high degree of oxidation and attributed this to the ability of the drug to undergo a Schiff’s reaction with the aldehyde groups, requiring subsequent degradation (rather than simple diffusion) processes to take place in order to release the drug.

9.3.2 Composite matrices for targeted delivery Biocomposites are widely used in bone repair and regeneration applications (Grøndahl and Jack, 2009). Due to their osteoconductive properties, hydroxyapatite (HAP) and other calcium phoshate phases have been widely researched for such biocomposite fabrication. A number of studies have shown promising in vitro bioactivity of HAP/alginate composite matri-



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ces for bone tissue engineering applications (Lin and Yeh, 2004; Turco et al., 2009). Maruyama (1995) investigated the in vivo performance of HAP/ alginate composite capsules and found them to show excellent osteoconductivity properties, bridging the gap between the implants and the cortical bone without adverse effects. In addition, they proposed that uniformly packed spherical particles with monodisperse porosity supported the rate of bone growth. Sivakumar and Rao (2003) applied this knowledge to the development of a delivery system for the antibiotic gentamicin, using coralline hydroxyapatite microspheres to encapsulate the drug. With the same general aim of drug delivery to bone fractures, the antimicrobial drug Biocide 1 was encapsulated in HAP/alginate granules (Krylova et al., 2002); the antibiotic gentamycin was encapsulated in HAP/alginate particles (Paul and Sharma, 1997); the therapeutic enzyme glucocerebrosidase was encapsulated in HAP/alginate and calcium titanium phosphate/alginate microspheres (Ribeiro et al., 2004); and a glucosaminoglycan was encapsulated in HAP/alginate capsules (Tan et al., 2009). Among the studies reporting drug delivery from HAP/alginate systems, only two include a comparison to the release from a pure alginate system (Ribeiro et al., 2004; Tan et al., 2009). In the work by Ribeiro et al. (2004), it was reported that the protein drug was released more rapidly from the pure alginate capsules than from composite ones when the protein was pre-adsorbed to the inorganic particles. This was attributed to the high affinity of HAP for the protein. A similar effect was observed in the study encapsulating heparin, where some retention of the drug was observed for the composite capsules (Tan et al., 2009). However, the systems in this study were very complex and the release rate was affected by the different types of intermolecular interactions between the glucosaminoglycan molecules and the alginate and HAP matrix components.

9.3.3 Microencapsulation for transplantation The transplantation of cells is a viable strategy in drug delivery, where therapeutic cells can be derived either from the same species (allograft) or, more commonly, a different species (xenograft). The relatively high biocompatibility and low cytotoxicity of purified alginate have made it a popular biomaterial for the microencapsulation, xenotransplantation and immunoisolation of these cells. The provenance of this field can be traced to the research of Lim and Sun who, in 1980, reported the assembly of a calcium alginate–polylysine–alginate (APA) microcapsule around pancreatic islets (Lim and Sun, 1980). A large body of the subsequent literature has been directed towards attempts to create the bioartificial pancreas through the microencapsulation and immunoisolation of pancreatic islets; there are multiple reviews relating to this specific application (Wilson and Chaikof,


252


2008; Lacik, 2006; de Vos et al., 2006; Narang and Mahato, 2006). One of the hurdles to progression in this field has been the diversity of methodologies and materials adopted for microencapsulation, resulting in a consensus that controlled systematic studies are required. In recent years, researchers have collaborated towards this goal, catalyzed by the EU-COST project (http:// cost865.bioencapsulation.net/). Despite the variability in methodologies, the substantial number of studies published prior to 2006 has enabled refinement of a suite of desirable and undesirable properties in alginatebased microcapsules which contribute to their potential success in their application (de Vos et al., 2006; Lacik, 2006; Zimmermann et al., 2005). These characteristics have been summarized in Table 9.3. Despite the demand for the standardization of methodologies surrounding the microencapsulation of pancreatic islets, it is evident that the microcapsule matrix must be individually tailored for each cellular application (Fig. 9.7). One of the challenges in optimizing the microcapsule for a desired application relates to the strength and permeability of the alginate-based matrix, with undesirable outcomes including rupture or the escape of cells during proliferation and differentiation (Li et al., 2008; Lee et al., 2009). The porosity and strength of the matrix can be manipulated through the selection of the type of alginate, with the ratio of M/G units dictating the extent of crosslinking (refer to Section 9.2.3). There are multiple approaches that have been adopted for a wide range of cellular systems and target environments, a summary of the most recent studies applying microencapsulation in alginate-based matrices being provided in Table 9.4 demonstrating the diversity and expansion of this therapeutic approach. While there is advancement in the optimization of the matrices, these studies continue to raise challenges such as promotion of the process of vascularization to optimize oxygen exchange, which may be overshadowed by the detrimental outcome of fibrotic overgrowth (Wilson and Chaikof, 2008). A survey of the strategies summarized in Table 9.4 reveals that there are several levels of complexity in the structural matrix of the microcapsules being trialled. These include: simple crosslinked gels (Fig. 9.6a); multilayered self-assembled polyelectrolyte complexes (Fig. 9.6d); and post modification to introduce desirable attributes (Fig. 9.6e and f). Each of these levels is described in detail below. 1

Alginate hydrogels ionically crosslinked by divalent cations are the simplest systems that have been adopted for cell microencapsulation and, of these, the most widely adopted continues to be Ca-alginate hydrogels. The process of forming hydrogels around the cells by this approach is very gentle, with low shear stresses and minimal exposure to reagents. The characteristic properties of the type of alginate selected for this purpose have a significant impact on the subsequent viability


Table 9.3 Desirable attributes in the assembly of alginate-based microcapsules Desirable

Undesirable

Alginate of sufficient purity so that it does not induce a cytotoxic response in the hosted cells. Alginates with an endotoxin content >100 EU/g are not suitable for in vivo studies. A robust structure which can withstand localized compression and shear stresses imparted by the target environment. Longevity of the matrix resistant to biodegradation. Tailoring of the capsule through a blend of a high G alginate imparting mechanical strength and a high M alginate promoting elasticity when required. A semi-permeable membrane which can retain the cells that it surrounds and simultaneously permits diffusion of both the nutrients and the active biomolecules. Oxygen transport to the encapsulated cells is a critical process. The pore dimensions and their interconnectivity in the polymer network control diffusion.

Toxins inherent in unpurified alginates include pyrogens, mitogens, polyphenols, and peptides. These substances invoke recognition by macrophages and induce fibrotic overgrowth of the microcapsules. A low degree of cross-linking results in a structure which swells and degrades in the presence of biological fluids. This also relates to the M/G ratio of the alginate and the viscosity. In contrast, a membrane that is too rigid (high degree of cross-linking) ruptures easily.

The outer surface of the microcapsule should be immunosilent. In APA capsules, high M alginates are preferred as they mask polycations more efficiently through stronger electrostatic interactions.

A surface-to-volume ratio that optimizes implantation efficiency and diffusional transfer processes.

The ability to tailor the capsule towards the host environment. Minimal adhesion or fibrotic overgrowth is tolerable for microcapsules implanted in the peritoneal cavity; however, vascularization is required for subcutaneous transplantation.

A semi-permeable membrane in which the porosity is either too small and prevents diffusion of essential nutrients or oxygen (hypoxia) to the encapsulated cells, or is too large, enabling cells to escape, inducing an immunogenic response, transfection or for cytotoxic species (e.g. inflammatory cytokines such as tumor necrosis factor TNF-α) to enter the microcapsule. Exposure of functionalities that cause a host immune response in the form of attack by humoral and T-cell mediated processes or the formation of a collagenous layer to isolate the ‘foreign’ body. In addition, a rough outer surface of the capsule encourages fibrotic overgrowth, which inhibits diffusion of nutrients and oxygen to the encapsulated cells. Large microcapsules which restrict diffusion of nutrients and oxygen to the encapsulated cells. The surfaceto-volume ratio results in inefficient bioactive molecule release. Physicochemical properties of the outer microcapsule surface that are detrimental to biointegration.


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Biointegration of medical implant materials Desirable: • oxygen and nutrients • signalling biomolecules Undesirable: • stimulus of macrophages, monocytes or fibroblasts which produce antibodies, cytokines or free radicals

Key Encapsulated cell Alginate Calcium ion Cationic polyelectrolyte Anionic polyelectrolyte

Desirable: • cellular waste products • secreted active biomolecules Undesirable: • escaping cells • immunogenic proteins

9.7 Schematic representation of the most common example of a multilayered assembly applied to microencapsulate cells. A summary of the desirable and undesirable processes that can impact on the viability of the cells is shown.

and function of the encapsulated cells. The composition can be altered to vary the mechanical properties and density of the Ca-alginate matrix which can impact on cell proliferation or differentiation, e.g. in the maturation of encapsulated follicles (West et al., 2007). However, the instability of Ca-alginate capsules in vivo due to exchange of crosslinking ions with native ions has led to the increasing use of Ba2+ as the preferred ionic crosslinker. Microcapsules prepared from Ba-alginate hydrogels have been demonstrated to possess improved mechanical properties and an increased capsule lifetime by a factor of two for high G alginates (Mørch et al., 2006), although the ratio of G content needs to be greater than 60%. These microcapsules do not completely prevent the undesirable diffusion of IgG and several cytokines to the cells within the interior of the microcapsule; however, the evidence indicates that the encapsulated islets have sufficient protection from the local environment (Omer et al., 2005). The diffusion of displaced Ca2+ ions within the hydrogel has been reported as having a detrimental impact on the viability of adjacent encapsulated Sertoli cells and the application of



Glial cell line-derived neurotrophic factor, GDNF Osteocalcin

Schwann

Bone marrow cells (BMC)

Foetal myoblasts

Fischer rat 3T3 fibroblasts

C2C12 myoblasts

HEK293

Crandall–Reece feline kidney (CRFK)

Glial cell line-derived neurotrophic factor, GDNF Therapeutic gene products Total proteins

Glial cell line-derived neurotrophic factor, GDNF Erythropoietin, EPO

Steroids

Follicles

Preosteoblastic (MC3T3-E1)

Secreted biomolecule

Cell system

Purified, intermediate G

Not specified

High G content Low viscosity A range of purity alginates and viscosities Not specified

Purified, high G and both high and low mw RGD-modified alginate† Tyramine-modified alginate for covalent cross-links Non-purified

Purified; G content 55–65%, oxidized and irradiated Filtered High M content

Alginate purity and viscosity

(Ca-Alg)-PLL-Alg

In vitro

In vitro

Striatum

(Ca-Alg)-PLL-Alg (Ca-Alg)-PLL-Alg

Subcutaneously and intraperitoneally

In vitro

Ca-Alg/collagen composite (Ca-Alg)-PLL-Alg

In vitro

(Tyr-Alg) + (Ca-Alg)

In vitro

(Ba-Alg) In vitro

In vitro

(Ca-Alg)

(RGD-Ca-Alg)

In vitro or in vivo

Microcapsule assemblies

Table 9.4 Categories of cells other than pancreatic islets that have been immunoisolated in alginate microcapsules

Abbah et al., 2008

Li et al., 2008

Grandoso et al., 2007

Orive et al., 2005 Ponce et al., 2006

Sakai and Kawakami, 2007 Lee et al., 2009

Evangelista et al., 2007

de Guzman et al., 2008

West et al., 2007

References


A-aromatase and IGF-1 Secreted alkaline phosphatase (SEAP) reporter enzyme Metabolizing molecules diffusing into capsule Glycosaminoglycan and collagen *

Sertoli cells

Endostatin

Recombinant CHO

Filtered

Purified low viscosity, M rich

Filtered G content not specified G content 69%, 130 K mw Ultrapure, G content 67%, low viscosity

Ultrapure low viscosity G alginate

High purity

Alginate purity and viscosity

Zhang et al., 2008 Zhang et al., 2007

Peritoneal

Marsich et al., 2008 Dusseault et al., 2008

Lin et al., 2008

Luca et al., 2007 Wikström et al., 2008

References

Peritoneal

Peritoneal

In vitro

Ca-Alg+Chitlac‡ (Ca-Alg)-(PLLANB-NOS)-Alg PLL modified with photo-crosslinker to introduce covalent bonds (Ca-Alg)-Chi-Alg (Ca-Alg)-PLL-Alg (Ca removed by chelation) (Ca-Alg)-PLL-Alg (Ca removed by chelation)

Oral administration to stomach

In vitro

Peritoneal

In vitro or in vivo

(Ca-Alg)-Chi-Alg (Ca-Alg)-PLL-Alg

(Ca-Alg)-PLL-Alg Ba-Alg (Ca-Alg)-PLL-Alg (Ba-Alg)-PLL-Alg (Sr-Alg)-PLL-Alg

Microcapsule assemblies

* Study demonstrated retention of immortal cells in microcapsule as proof of concept for similar cell lines that possess risk of malignant transformation. † Modification of alginate to enhance adhesion, proliferation and differentiation. ‡ Chitlac, lactose modified chitosan.

Bone morphologenic protein (BMP)

Embryonic fibroblasts (C3H10T1/2 line)

EL-4 thymoma

Chrondrocytes

E Coli DH5

Human retinal pigment epithelial cell line (ARPE-19)

Secreted biomolecule

Cell system

Table 9.4 Continued


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Ba-alginate microcapsules was reported a successful alternative strategy (Luca et al., 2007). 2 A multilayered self-assembled polyelectrolyte complex (e.g. the APA system) is typically produced by the following sequence: a Ca-alginate hydrogel capsule formed around the cells and then multiple polyelectrolyte layers of opposing charge (e.g. PLL and alginate) are selfassembled through the combination of electrostatic intermolecular forces, hydrogen bonding and polymer flocculation to strengthen the capsule. This assembly remains the basis of the majority of cellular encapsulation matrices. The structure of these capsules is often naively represented as discrete layers; however, evidence has been provided for the formation of polyelectrolyte complexes between alginate and PLL where PLL penetrates the alginate core up to a depth of 30 μm (de Vos et al., 2006). The most significant challenge is to achieve full screening of the exposed charges on the cationic polyelectrolyte, otherwise these will invoke an immune response. PLL inefficiently complexed with an outer alginate layer can potentially remain exposed at the outermost surface of the capsule (Tam et al., 2005). The APA microcapsule offers greater mechanical strength and protection of the encapsulated cells than the Ca-alginate hydrogel capsule; however it must be considered that the longevity of the process, which includes multiple washing steps, can itself be detrimental to the cells. 3 Multilayered assemblies, where the analogues of the APA matrix components are modified to introduce desirable attributes such as substitution of divalent ions or introduction of covalent bonds between biopolymers. The standard APA capsule adopts Ca2+ ions as the crosslinker for the inner capsule. However, the influence of the size of the crosslinking ion (Ca2+, Ba2+, Sr2+) on the structure of the resulting matrix and encapsulated cell viability has been explored for ARPE-19 cells in vitro (Wikström et al., 2008): Sr2+ was found to be unsuitable. The introduction of covalent linkages increased the mechanical strength and reduced the swelling of the capsule matrix in the host environment. The most common modification to the self-assembled LbL matrix is the introduction of covalent bonds between alginate and PLL; for example, those achieved by the introduction of a photocrosslinkable moiety onto the PLL to strengthen the microcapsule (e.g. Dusseault et al., 2008) or the introduction of enzymatically crosslinkable groups (Sakai and Kawakami, 2008). The mechanical properties of the capsule become important in the encapsulation of cells that are required to differentiate and proliferate, and an example of a post-modification strategy is the removal of calcium from the internal hydrogel, to encourage cell growth and active biomolecule


258


production (Zhang et al., 2008). The modification of alginate through the introduction of peptide moieties (Section 9.2.4) is a common approach to improving the adhesion of encapsulated cells that are anchorage dependent for survival, e.g. osteoblasts (Evangelista et al., 2007). An extension of this approach was demonstrated in the covalent modification of alginate by YIGSR peptides prior to encapsulation of neurites, followed by the subsequent modification of the capsule via adsorption of laminin (Dhoot et al., 2004). Adhesion of the neurites was promoted. Post-modification of the encapsulated cell assembly has been attempted to improve integration with its target environment and strategies include PEGylation (Zhang et al., 2008). Surface functional molecules may also be coupled to the assembly post-encapsulation to improve immunosilence (Leung et al., 2008). An emerging initiative in the microencapsulation of cells is to encapsulate active biomolecules simultaneously with cells, with the objective of promoting a subsequent cellular response in vivo. Bioactive molecules, such as basement membrane extract BD MatrigelTM (de Guzman et al., 2008) to promote proliferation, have been incorporated with the alginate prior to formation of the Ca-alginate microcapsule. In a separate approach, extracellular matrix proteins, such as collagen and laminin, have been added during the assembly of the polycationic membrane (Cui et al., 2006). Finally, there are a number of recent studies that demonstrate that researchers are thinking ‘beyond the sphere’ in cell encapsulation by tailoring the scaffold geometry and matrix properties to suit their specific applications. There are applications where the cells themselves are ‘delivered’ as they migrate from a scaffold that is designed to co-deliver inductive molecules (Hill et al., 2006). An innovative approach is represented in the encapsulation of cells within fibres which may also contain growth factors (Wan et al., 2004). These fibres can then be self-assembled into defined and patterned structures.

9.4

Future trends

At this point the influence of the structure of alginate (M/G ratio, chain length and role of naturally derived impurities) has been well characterized. The process of assembling three-dimensional matrices through crosslinking of alginate molecules is also well understood and can be applied to control mechanical properties and stability. Modification of the assemblies is emerging as the popular route for optimizing biointegration of these matrices in their host environment. Research encompasses two approaches: the modification of alginate prior to capsule assembly and post-modification of assembled matrices. As the process of assembling alginate-based matrices for drug delivery is refined, attention will turn to tailoring the release profiles of the bioactive



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species. The nature of the interaction between the drug molecules and the scaffold matrix will define these profiles. To date, knowledge of the chemical properties of the drug molecule has been poorly explored as the delivery vehicle is designed. Introduction of functionalities that can be manipulated through change of environment or time-dependent processes offer new strategies, and already reports of pH responsive systems through the modification of alginate have appeared (Chan et al., 2008). Finally, the majority of assemblies addressed here have resulted in spherical capsules. Drug delivery is not restricted to matrices of this geometry and while the surface area is reduced in other three-dimensional geometres, there are advantages offered by drug delivery scaffolds such as self-assembled fibres (Wan et al., 2004) and semi-interpenetrating networks (Matricardi et al., 2008).

9.5

Acknowledgement

The authors are grateful to the Australian Research Council (Grant No DP0557475) for their support of this work.

9.6

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mørch, y a, holtan, s, donati, i, strand, b l, skjåk-bræk, g (2008) ‘Mechanical properties of C-5 epimerized alginates’ Biomacromolecules, 9(9), 2360–2368. mucalo, m r, yokogawa, y, suzuki, t, kawamoto, y, nagata, f, nishizawa, k (1995) ‘Further studies of calcium phosphate growth on phosphorylated cotton fibres’ Journal of Materials Science Materials Medicine, 6(11), 658–669. narang, a s, mahato, r i (2006) ‘Biological and biomaterial approaches for improved islet transplantation’ Pharmacological Reviews, 58, 194–243. nguyen, v b, wang, c x, thomas, c r, zhang, z (2009) ‘Mechanical properties of single alginate microspheres determined by microcompression and finite element modelling’ Chemical Engineering Science, 64(5), 821–829. nunamaker, e a, purcell, e k, kipke, d r (2007) ‘In vivo stability and biocompatibility of implanted calcium alginate disks’ Journal of Biomedical Materials Research, 83A, 1128–1137. omer, a, duvivier-kali, v, fernandes, j, tchipashvili, v, colton, c k, weir, g c (2005) ‘Long-term normoglycemia in rats receiving transplants with encapsulated islets’ Transplantation, 79, 52–58. orive, g, de castro, m, ponce, s, hernández, r m, gasc´on, a r, boch, m, alberch, j, pedraz, j l (2005) ‘Long-term expression of erythropoietin from myoblasts immobilized in biocompatible and neovascularised microcapsules’ Molecular Therapy, 12, 283–289. paul, w, sharma, c p (1997) ‘Infection resistant hydroxyapatite/alginate plastic composite’ Journal of Materials Science Letters, 16, 2050–2051. pavlath, a e, gossett, c, camirad, w, robertson, g h (1999) ‘Ionomeric films of alginic acid’ Journal of Food Science, 64(1), 61–63. pelletier, a, hubert, p, lapicque, f, payan, e, dellacherie, e (2000) ‘Amphillic derivatives of sodium alginate and hyaluronate: Synthesis and physicochemical properties of aqueous dilute solutions’ Carbohydrate Polymers, 43, 343–349. ponce, s, orive, g, hernández, r m, gascón, a r, canals, j m, muñoz, m t, pedraz, j l (2006) ‘In vivo evaluation of EPO-secreting cells immobilized in different alginate-PLL microcapsules’ Journal of Controlled Release, 116, 28–34. poncelet, d, lencki, r, beaulieu, c, halle, j p, neufeld, r j, fournier, a (1992) ‘Production of alginate beads by emulsification/internal gelation. I. Methodology’ Applied Microbiology and Biotechnology, 38(1), 39–45. qin, y (2008) ‘The gel swelling properties of alginate fibers and their applications in wound management’ Polymers for Advanced Technologies, 19(1), 6–14. ranade, v v, hollinger, m a (2004) Drug Delivery Systems, 2nd Edition, Boca Raton, Florida, CRC Press Inc., 63–114. ribeiro, c c, barrias, c c, barbosa, m a (2004) ‘Calcium phosphate–alginate microspheres as enzyme delivery matrices’ Biomaterials, 25, 4363–4373. remuñán-lópez, c, bodmeier, r (1997) ‘Mechanical, water uptake and permeability properties of crosslinked chitosan glutamate and alginate films’ Journal of Controlled Release, 44, 215–225. rowley, j a, madlambayan, g, mooney, d j (1999) ‘Alginate hydrogels as synthetic extracellular matrix materials’ Biomaterials, 20, 45–53. sakai, s, kawakami, k (2008) ‘Both ionically and enzymatically crosslinkable alginate–tyramine conjugate as materials for cell encapsulation’ Journal of Biomedical Materials Research, 85A, 345–351.



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vold, i m n, kristiansen, k a, christensen, b e (2006) ‘A study of the chain stiffness and extension of alginates, in vitro epimerized alginates, and periodate-oxidized alginates using size exclusion chromatography combined with light scattering and viscosity detectors’ Biomacromolecules, 7(7), 2136–2146. wan, a c a, yim, e k f, liao, i-c, le visage, c, leong, k (2004) ‘Encapsulation of biologics in self-assembled fibres as biostructured units for tissue engineering’ Journal of Biomedical Materials Research, 71A, 586–595. webber, r e, schull, k r (2004) ‘Strain dependence of the viscoelastic properties of alginate hydrogels’ Macromolecules, 37(16), 6153–6160. west, e r, xu, m, woodruff, t k, shea, l d (2007) ‘Physical properties of alginate hydrogels and their effects on in vitro follicle development’ Biomaterials, 28, 4439–4448. whistler, r l, wolfrom, m l (1962) Methods in Carbohydrate Chemistry, Vol I: Analysis and Preparation of Sugars. 2nd edition, New York and London, Academic Press Inc. wikström, j, elomaa, m, syväjärvi, h, kuokkannen, j, yliperttula, m, honkakoski, p, urtti, a (2008) ‘Alginate-based microencapsulation of retinal pigment epithelial cell line for cell therapy’ Biomaterials, 29, 869–876. wilson, j t, chaikof, e l (2008) ‘Challenges and emerging technologies in the immunoisolation of cells and tissues’ Advanced Drug Delivery Reviews, 60, 124–145. xu, j b, bartley, j p, johnson, r a (2003a) ‘Preparation and characterization of alginate hydrogel membranes crosslinked using a water-soluble carbodiimide’ Journal of Applied Polymer Science, 90, 747–753. xu, j b, bartley, j p, johnson, r a (2003b) ‘Preparation and characterization of alginate–carrageenan hydrogel films crosslinked using a water-soluble carbodiimide (WSC)’ Journal of Membrane Science, 218, 131–146. yeom, c k, lee, k h (1998) ‘Characterization of sodium alginate membrane crosslinked with gluteraldehyde in pervaporation separation’ Journal of Applied Polymer Science, 67, 209–219. zhang, y, wang, w, xie, y, yu, w, lv, g, guo, x, xiong, y, ma, x (2007) ‘Optimization of microencapsulated recombinant CHO cell growth, endostatin production and stability of microcapsule in vivo’ Journal of Biomedical Materials Research B: Applied Biomaterials, 84B, 79–88. zhang, w-j, li, b-g, zhang, c, xie, x-h, tang, t t (2008) ‘Biocompatability and membrane strength of C3H10T1/2 cell-loaded alginate-based microcapsules’ Cytotherapy, 10, 90–97. zhu, h, ji, j, lin, r, gao, c, feng, l, shen, j (2002) ‘Surface engineering of poly(dl-lactic acid) by entrapment of alginate–amino acid derivatives for promotion of chondrogenesis’ Biomaterials, 23, 3141–3148. zimmermann, u, klock, g, federlin, k, hannig, k, kowalski, m, bretzel, r g, horcher, a, entenmann, h, sieber, u, zekorn, t (1992) ‘Production of mitogen-contamination free alginates with variable ratios of mannuronic acid to guluronic acid by free flow electrophoresis’ Electrophoresis, 13(5), 269–274. zimmermann, h, zimmermann, d, reuss, r, feilen, p j, manz, b, katsen, a, weber, m, ihmig, f r, ehrhart, f, gebner, p, behringer, m, steinbach, a, wegner, l h, sukhorukov, v l, vasquez, j a, schneider, s, weber, m m, vole, f, wolf, r, zimmermann, u (2005) ‘Towards a medically approved technology for alginate-based microcapsules allowing for long-term immunoisolated transplantation’ Journal of Materials Science Materials in Medicine, 16, 491–501.


10 Functionalised nanoparticles for targeted drug delivery S. M A N J U and K. S R E E N I VA S A N, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

Abstract: Nanocarriers composed of liposomes, micelles, polymeric nanoparticles and others have shown tremendous opportunities in the field of targeted drug delivery, especially in cancer therapy. Functionalisation of nanomaterials through simultaneous assembling of chemical moieties has been a strategy of wide interest, to acquire properties such as longevity in circulation, site specificity and stimuli sensitivity. Imparting multifunctionality to nanocarriers controls their biological interaction in a desired fashion and enhances the efficacy of therapy and diagnostic protocols. Here, we attempt to review the application of various nanocarrier systems for targeted drug delivery and current strategies for the development of multifunctionality on nanocarrier systems. Key words: nanocarriers, drug delivery, micelle, liposomes, polymeric nanoparticle, gold nanoparticle, magnetic nanoparticle, hyperthermia, photodynamic therapy.

10.1

Introduction

Nanobiotechnology is a multidisciplinary field that covers a vast and diverse array of technologies coming from engineering, physics, chemistry and biology. It is the combination of these fields that has led to the birth of a new generation of materials and methods of making them. Nanomaterials, which measure 1 nm to 100 nm, allow unique interaction with biological systems at the molecular level. Among the various approaches for exploiting developments in nanotechnology for biomedical application, nanoparticulate carriers offer some unique advantages as delivery, sensing and image enhancement agents.1,2 They can engineer important advances in the detection, diagnosis and therapy of human cancers. We know that many effective bioactive agents used for pharmacotherapy exhibit side effects that limit their clinical application, so it is important to achieve selectivity in the delivery of drug molecules to target areas, in order to enhance therapeutic potential and minimise side effects. For example, cytotoxic compounds used in cancer therapy can kill cancer cells as well as normal cells. The use of pharmaceutical nanocarriers is designed to overcome most of 267 © Woodhead Publishing Limited, 2010

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these limitations.3 Even though research efforts in this area have resulted in enhanced in vitro efficacy of many drugs, both in pharmaceutical research and a clinical setting, researchers are still engaged in designing improved nanocarriers for effective and targeted delivery applications. In general, requirements in the design of nanodrug delivery systems include: (a)

design carrier systems that can incorporate different types of therapeutic agents in sufficient doses, (b) recognise disease-specific ligands that can be conjugated to drug carrier systems and achieve targeted drug therapy, (c) protect drug molecules from degradation in the body prior to their delivery at the required sites, (d) develop drug-carrier system that can release the drug at the target site at a desired or controllable rate for the duration necessary to elicit the desired pharmacological response, (e) develop drug carrier systems that are biocompatible and biodegradable so that these can be used safely in human, and (f) achieve effective intracellular drug delivery for those therapeutic agents whose receptor or site of action is intracellular. Nanocarrier systems are formulated from a variety of materials having various chemical compositions, and are engineered to carry a number of bioactive molecules in a controlled and targeted manner, making them efficient drug delivery vehicles.4 Commonly used nanocarriers in drug delivery application include liposomes,5 micelles, nanocapsules, polymeric nanoparticles, solid lipid nanoparticles, metallic nanoparticles (gold/magnetic nanoparticle/quantum dots) and others (Fig. 10.1).5–8 Some of the examples of commonly defined nanocarrier systems and their medical applications are shown in the Table 10.1. Generally, the selection of material for developing nancarriers is mainly dictated by the desired diagnostic or therapeutic effect, administration route and type of pay load. It offers considerable opportunities to improve cancer therapy by exploring the unique inherent properties of a solid tumor combined with new approaches such as local heating or reactive oxygen generation. Nanocarriers, depending on the reticular requirement, demand surface modification.10–13 Surface modified nanocarriers exhibit advantageous properties, such as prolonged circulation in the blood, demanded biodistribution, passive or active targeting to the pathological tissue site, responsiveness to local physiological stimuli (pH or temperature changes in the tissue site with pathological condition), and the ability to serve as imaging or contrast agents for various imaging systems. This review focuses mainly on different targeting methods and recent progress in the functionalization of nanocarriers for improving cancer therapy, including both imaging and site-specific targeting of drug molecules.


Functionalised nanoparticles for targeted drug delivery

269

10–3 m (mm) Spleen cut-off 200 nm

Cells Lymphocytes Erythrocytes DNA

10–6 m (μm) Lipid bilayer

Proteins

Dendrimers Micelle Emulsion Nanoparticulate Liposomes 3–5 nm 5–10 nm water in oil systems 100–150 nm Drug 100–150 nm 20–150 nm 10 m (nm) encapsulated or attached

Quanfum dots

–9

Bucky balls

Nanoparticles used in drug delivery Small molecules and atoms

10–12 m (pm)

10.1 Nanoparticle systems in drug delivery (adapted with permission from © 2007 Elsevier9).

10.2

Drug targeting

Drug targeting is a potential approach by which distribution of a drug in an organism is contrived in a manner such that its major fraction interacts exclusively with the target tissue at the cellular or sub-cellular level. Theoretically, drug targeting can improve the outcome of chemotherapy by means of one or both of the following processes: (a)

(b)

By allowing the maximum fraction of the delivered drug molecule to react exclusively with diseased cells without adverse effect to the normal cells. By allowing preferential distribution of the drug to the diseased/cancerous cells.

10.2.1 Barriers to drug targeting The main hurdles in the field of drug targeting include physiological barriers and biochemical challenges to achieving target specificity. They also include the selection of appropriate techniques for the conjugation of the targeting ligand to the nanocarriers. The challenges in drug targeting include targeting drugs to specific sites and imparting longevity in the blood to enhance bioavailability at the site of action. For intravenous administration of nanocarriers, the main barrier is that of the vascular endothelium, basement membrane14 and plasma proteins. Plasma proteins have the ability to


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Table 10.1 Commonly used nanoparticles and their medical applications Nanoparticle

Examples

Medical application

Metal nanoparticles

Quantum dots Gold nanoparticles Gold nanorods Gold nanoshells Gold nanocages Magnetic nanoparticles

Diagnostics Biosensor Molecular imaging Drug delivery

Nanotubes, nanowires

Carbon nanotubes

Biomolecular sensing Delivery of vaccines or proteins

Dendrimers

Poly(amido amine)

Drug carriers Imaging agent Gene delivery

Liposomes

PEGylated immunoliposomes

Drug delivery Gene encoding

Polymer micelles

Doxorubicin conjugated to poly(ethylene glycol)-poly (α,β-aspartic acid) PEG-PAsp(DOX)

Drug delivery of water-insoluble drugs

Ceramic nanoparticles

Silica-based nanoparticles entrapping photosensitising anticancer drug, 2-devinyl2-(1-hexyloxyethyl) pyropheophorbide

Drug delivery

Polymeric nanoparticles

PLGA (poly(D,L-lactic-coglycolic acid) Poly(lactic) acid (PLA)– polyglycolic acid (PGA)

Drug delivery Protein delivery Gene expression vector

Polysaccharide nanoparticles

Cellular nanocrystals

Targeted delivery Bioimaging

Magnetic nanoparticles

Superparamagnetic iron oxide

Magnetic resonance imaging, contrast agent

Bionanoparticleprotein-based nanosystems

Ferritin, viruses and virus-like particles Heat shock protein cages

Gene delivery Bioimaging Drug delivery Vaccine development

adversely affect the biodistribution of drug carriers introduced in the blood stream. The in vivo biodistribution and opsonisation of nanosystems in blood circulation are governed by their size and surface characteristics, such as zeta potential and hydrophilicity/hydrophobicity, so one has to be very critical in developing targeted nanocarriers and should avoid opsonisation


Functionalised nanoparticles for targeted drug delivery

271

and subsequent recognition by reticuloendothelial cells (RES). Another barrier is that of the extra cellular matrix, which has to be crossed to access the target cells in a tissue. On dealing with intracellular delivery, there are additional barriers to be crossed to allow internalisation of the systems into the specific cells.15 Also, there are a number of endocytic pathways for the cellular entry of nanosystems. Researchers have to be aware of the particular mechanisms for a better demonstration of the functionalisation of nanocarriers.The nuclear membrane creates another formidable barrier for drugs such as oligonucleotides, plasmid DNA and other low molecular weight drugs, whose site of action is located in the nucleus of a cell. During recent years, a number of cellular and molecular targets have emerged in the field of drug delivery, but poor bioavailability of the drug in the target tissue is a real problem in clinical practice due to these barriers. Thus, to develop competent targeted systems, it is necessary to successfully overcome most of the physiological barriers and deliver the drug in the target site at an optimum therapeutic level, for the required time period to elicit pharmacological action. Various techniques have been devised for the conjugation of targeting ligands onto nanocarrier systems, which include covalent and non-covalent conjugation. These approaches direct them to their target site and give them the right orientation for binding the target molecule.16 The conjugation mechanism should not adversely affect the integrity of the nanocarriers and should give a better interface to enhance the biological activity. To this end, various strategies have been developed for targeting drugs to the required tissue site in the body by the proper design of nanocarriers. The various approaches for targeting drugs in chemotherapy are shown in the Fig. 10.2. Drug targeting is mainly classified into passive targeting and active targeting. Site-specific drug targeting

Passive targeting

EPR effect

Active targeting

Tumour environment

Carbohydrate interaction

Ligand–receptor interaction

Antibody– antigen interaction

Direct local delivery

10.2 Schematic representation: various approaches for drug targeting (EPR, enhanced permeability and retention).


272


6–7 nm ECs

200–800 nm

Lumen

ECM Normal tissue, intact vasculature

Tumour tissue site, leaky vasculature

10.3 Pathological difference between normal and tumour tissue leads to better accumulation of drug in the tumour site.

10.2.2 Passive targeting Solid tumours present favourable conditions for the preferential accumulation of a variety of nanocarriers. The rapid growth of a solid tumour results in altered physiology at the tumour site, which leads to leaky and defective vasculature (gaps ~ 600 nm, Fig. 10.3) and poor lymphatic drainage. The increased vascular permeability coupled with impaired lymphatic drainage in the tumour, leads to an enhanced permeability and retention effect (EPR effect) which allows extravasations of nanocarriers and selective localisation in the inflamed tissue.17,18 Oral administration of polymeric nanocarriers could, as an example, be selectively targeted to the inflamed colonic mucosa in inflammatory bowel disease.19 Similarly, in some critical inflammatory conditions, the blood–brain barrier can be crossed to access the target sites for brain delivery.20 For enhancing the passive targeting mechanism, the physicochemical factors of the nanocarriers such as size, surface charge and surface hydrophobicity also play critical roles. Particles

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